Progress In Medical Radiation...

Progress In Medical Radiation Physics


Volume 1


Series Editor:

COLIN G. ORTON, Ph.D. Department of Radiation Oncology Wayne State University School of Medicine Detroit, Michigan

Editorial Board:

PETER R. ALMOND, Ph.D. Department of Physics M.D. Anderson Hospital Houston, Texas

JOHN S. CLIFTON, M.Sc. Department of Medical Physics University College Hospital London, England

ROY E. ELLIS, Ph.D. t Head Department of Medical Physics The General Infirmary Leeds, Yorkshire, England

J.F. FOWLER, Ph.D. Director, Gray Laboratory Mount Vernon Hospital Northwood, Middlesex, England

JAMES G. KEREIAKES, Ph.D. Eugene L. Saenger Radioisotope Laboratory Cincinnati General Hospital Cincinnati, Ohio

JACK S. KROHMER, Ph.D. Department of Radiology Wayne State University School of Medicine Detroit, Michigan

CHRISTOPHER H. MARSHALL, Ph.D. N. Y. U. Medical Center New York, New York

tDeceased

Progress In Medical Radiation Physics Volume 1

Edited by

COLIN G. ORTON Wayne State University School of Medicine Detroit. Michigan

PLENUM PRESS • NEW YORK AND LONDON

ISBN 978-1-4615-7693-8 ISBN 978-1-4615-7691-4 (eBook) DOI 10.1007/978-1-4615-7691-4

© 1982 Plenum Press, New York Softcover reprint of the hardcover 1st edition 1982 A Division of Plenum Publishing Corporation 233 Spring Street, New York, N.Y. 10013

All rights reserved

No part of this book may be reproduced, stored in a retrieval system, or transmitted, in any form or by any means, electronic, mechanical, photocopying, microfilming, recording, or otherwise, without written permission from the Publisher

Contributors

J. J. Broerse, Radiobiological Institute TNO, Rijswijk, The Netherlands

Christodoulos Constantinou, Radiation Physics Department, St. Bartholomew's Hospital, London, England. Present address: University of Wisconsin, Madison, Wisconsin

J. R. Cunningham, The Ontario Cancer Institute, 500 Sherbourne Street, Toronto, Ontario, Canada, M4X 1K9

S. John Gatley, Medical Physics Section, Department of Radiology, 3321 Sterling Hall, 475 N. Charter Street, University of Wisconsin, Madison, Wisconsin 53706

Michael Goitein, Division of Radiation Biophysics, Department of Radiation Medicine, Massachusetts General Hospital, Boston, Massachusetts 02114, and Harvard Medical School.

U. KilIat, Philips GmbH Forschungslaboratorium Hamburg, Vogt-K611nStr. 30, D 2000 Hamburg 54

B. J. Mijnheer, Antoni van Leeuwenhoek Hospital, Amsterdam, The Netherlands

Robert J. Nickles, Medical Physics Section, Department of Radiology, 3321 Sterling Hall, 475 N. Charter Street, University of Wisconsin, Madison, Wisconsin 53706

David R. White, Radiation Physics Department, St. Bartholomew's Hospital, Londbn, England

v

Preface

New developments in the application of radiation to medicine are occurring so rapidly that this is possibly the fastest growing branch of medicine today. In the past decade alone, we have seen enormous progress made in techniques used both for the diagnosis of disease, such as computerized tomography, digital radiography, ultrasonography, computerized nuclear medicine scanning, and nuclear magnetic resonance imaging, and for its treatment, such as the radiotherapeutic utilization of high-LET radiations, and the widespread application of computers to perform elegant dosimetry calculations for 3-D treatment planning and imaging.

This series will provide in-depth reviews of the many spectacular technical advances and sophisticated concepts, which are developing in medical radiation physics at such an alarming rate that it has become increasingly difficult to keep one's knowledge up-to-date. These comprehensive review articles will help to bridge the communications gap between the international research community, and the medical physicists and physicians whose responsibility it is to put these advances into clinical use. These articles should also be of value to the increasing number of physical scientists and engineers who are interested in the application of their knowledge and talents to the field of medicine.

Colin G. Orton

vii

Contents

1. Progress in Neutron Dosimetry for Biomedical Applications J. J. Broerse and B. J. Mijnheer

1. Introduction 1 2. Neutron Sources 6

2.1. Cyclotrons 7 2.2. d + T Neutron Sources 13 2.3. Reactor Neutron Beams of Special Design 15

3. Principles and Methods in Neutron Dosimetry 16 3.1. Ionization Chambers . . . . . . . . . . 17 3.2. Physical Parameters for Dosimetry with Ionization Chambers 22 3.3. Calorimeters 27 3.4. Fluence Measurements 29 3.5. Solid-State Dosimeters 30 3.6. Determination of the Gamma-Ray Absorbed Dose in a

Neutron Field . . . . . . . . . . . . . . . . . . 35 3.7. Determination of the Neutron Absorbed Dose in a Photon

Field . . . . . . . . . . . . . . . . 40 4. Assessment of Radiation Quality ....... 43

4.1. Neutron and Gamma-Ray Energy Spectra 44 4.2. Microdosimetric Parameters and Techniques 47 4.3. Lineal Energy Spectra in the Phantom 51

5. Dosimetry for Radiobiology ....... 54 5.1. Interface Dosimetry ........ 56 5.2. Absorbed Dose Distributions in Animals 61

6. Dosimetry for Radiotherapy ........ 65 6.1. Protocols for Neutron Dosimetry for External Beam Therapy 65 6.2. Absorbed Dose Distributions in the Standard Phantom 68 6.3. Absorbed Dose Distributions in the Patient . 73

7. Dosimetry for in vivo Neutron Activation Analysis 78 8. Results of Dosimetry Intercomparisons ..... 79

ix

x

9. Conclusions and Recommendations References . . . . . . . . . . . . .

2. Tissue Inhomogeneity Corrections in Photon-Beam Treatment Planning J. R. Cunningham

1. Introduction . . . . . . . . . . . . . . 2. Survey of Common Dose-Correction Methods

2.1. Methods Based on Water-Equivalent Depth 2.2. Power Law Tissue-Air Ratio Method ... 2.3. The Equivalent Tissue-Air Ratio Method . 2.4. Volume Integration of Differential Scatter-Air Ratios 2.5. Monte Carlo Calculations . . ..

3. Interface Effects-Electronic Equilibrium 4. Summary and Conclusions References . . . . . . . . . . . . . . .

3. Anthropomorphic Phantom Materials David R. White and Christodoulos Constantinou

1. Introduction 1.1. The Need for Tissue Simulation 1.2. Terminology 1.3. A Brief History

2. Radiation Properties of Real Tissues 2.1. Tissues Requiring Simulation 2.2. Radiation Characteristics

3. Simulation Procedures 3.1. Criteria for Tissue Equivalence 3.2. The Effective Atomic Number (2) Method 3.3. The Basic Data Method 3.4. The Extended Y Method 3.5. Elemental Equivalence

4. Recent Tissue Substitutes 4.1. Introduction 4.2. Base Materials and Additives 4.3. Recommended Tissue Substitutes

5. Manufacturing Procedures and Quality Control 5.1. Manufacturing Bulk Materials 5.2. Manufacturing Specialized Phantoms 5.3. Quality Control

6. Recent Phantom Studies 6.1. Radiotherapy

Contents

86 90

103 105 108 112 113 121 126 127 129 130

133 133 134 135 137 137 140 147 147 148 150 153 156 158 158 158 160 171 171 176 178 179 179

Contents

6.2. Diagnostic Radiology . . . . . . . 6.3. Nuclear Medicine and Health Physics

7. Discussion References . . . . . . . . . . . . . . . .

4. Applications of Computed Tomography in Radiotherapy Treatment Planning Michael Goitein

xi

181 183 184 189

1. Introduction . . . . . . . . . . . . . 195 2. General Studies of the Impact of CT . . 197

2.1. New England Medical Center Study 198 2.2. Massachusetts General Hospital Study 200 2.3. Royal Marsden Hospital Study 204 2.4. Other General Studies ..... 206 2.5. Discussion . . . . . . . . . . . 206

3. Site-Specific Studies of the Impact of CT 210 4. Monitoring of Tumors During Treatments and in Posttreatment

Follow-up . . . . . 211 5. Treatment Outcome 213 6. Cost Effectiveness 219 7. Patient Positioning . 221

7.1. Breathing 224 7.2. Contrast Media 225 7.3. Bolus 226 7.4. Unavoidable Discrepancies 226 7.5. From Scan to Treatment . 226 7.6. Use of CT to Select Position for Treatment 227

8. Dosimetry . . . . . . . . . . . . . . . . . . 228 8.1. x- and Gamma-Ray External Beam Therapy 228 8.2. Brachytherapy ..... 242 8.3. Charged-Particle Therapy 242

9. Tissue Characterization 248 9.1. Tomochemistry . . . . . 249 9.2. Spatial Structure 251 9.3. Radiologic-Pathologic Correlation 251

10. Specifications for CT Scanners to be Used in Treatment Planning. 252 10.1. Introduction ................... 252 10.2. Three-Dimensional Delineation of Tumors and Adjacent

Normal Structures . . . . . . . . . . . . . 253 10.3. Patient Positioning . . . . . . . . . . . . . 260 10.4. Features Necessary for the Calculation of Dose 263 10.5. Summary of Specifications ......... 263 10.6. Relationship between Diagnostician and Therapist 264 10.7. Inexpensive Scanners 266 10.8. Treatment Planning Programs . . . . . . . . . 267

xii

11. Epilogue ................... . Appendix: Review of Site-Specific Studies on the Impact of CT

AI. Prostatic Cancer . . . . . . . . . . A2. Bladder Cancer . . . . . . . . . . A3. Other Pelvic and Abdominal Diseases A4. Thorax and Chest Wall A5. Head and Neck Tumors A6. Other Sites

References . . . . . . . . . . .

s. Positron Imaging: Some Practical Considerations Robert J. Nickles and S. John Gatley

1. Introduction ..... . 1.1. Scope of the Problem 1.2. Linkage ..... 1.3. Historical Perspective 1.4. A Modest Alternative

2. Positron Imaging 2.1. Physical Background . 2.2. A Modular Positron Camera

3. Radiopharmaceutical Synthesis 3.1. General Considerations . . 3.2. Review of Recent Developments

4. Results ............ . 4.1. Practical Aspects in a Clinical Setting 4.2. Representative Images

5. Conclusions References . . . . . . . . . .

6. Modern Optical Methods for the Storage of Radiographs U. Killat

1. Introduction ..... . 2. Minification Systems . . .

2.1. The Delcomat System 2.2. The LogEtronics System 2.3. Experience with Minification Systems 2.4. Physical Limits for Minification Systems 2.5. Conclusion ........... .

3. Storage of Radiographs: The Information Theory Approach 3.1. Information Capacity of Radiographs ...... . 3.2. Information Capacity of Radiographs: Example 3.3. Information Content, Information Capacity, and Image

Quality .................... .

Contents

268 268 268 271 275 276 280 283 287

295 295 296 297 298 299 299 302 310 310 312 313 313 316 316 317

323 325 326 326 327 328 331 333 333 335

337

Contents xiii

4. Holographic Storage . . . . . . . . . . . . 340 4.1. Description of the Holographic Method 341 4.2. Concept of a Holographic Storage System 342 4.3. Image Quality in the Holographic Recording of Radiographs. 345 4.4. The Bottlenecks of Holographic Image Storage 355

5. Digital Storage of Radiographs 356 5.1. Description of the Method 356 5.2. Picture Input Devices 357 5.3. Picture Output Devices . . 359 5.4. Source Coding of Radiographs 367 5.5. Optical Disk Store . . . . . 368

6. Towards Future Hospital Communication Systems 370 References . . . . . . . . . . . . . . . . . . . 373

Index 377

1

Progress In Neutron Dosimetry for Biomedical Applications

J. J. BROERSE and B. J. MIJNHEER

1. INTRODUCTION

The use of fast neutrons for radiobiological and medical applications has increased considerably over the past 20 years. Fast neutrons are important to fundamental radiobiological studies, since they allow the exposure of relatively large multicellular systems (e.g., individual organs and tumors) to radiation with a high linear energy transfer (LET), achieving an approximately uniform dose distribution throughout the subject. With regard to risk estimates relevant to radiation protection applications, the relative biological effectiveness (RBE) of fast neutron fields for carcinogenesis is still under discussion (Bond, 1979).

The different medical applications include fast neutron radiotherapy, capture therapy with thermal and epithermal neutrons (Hatanaka and Sweet, 1975), and in vivo neutron activation analysis (IAEA, 1973). Interest in the use of fast neutrons for cancer radiotherapy was stimulated by three radiobiological findings concerned with the oxygen enhancement ratio, the sensitivity of cells at different phases of the cell cycle, and the contribution of the accumulation of sublethal damage to cell reproductive death.

J. J. BROERSE • Radiobiological Institute TNO, Rijswijk, The Netherlands. B. J. MIJNHEER • Antoni van Leeuwenhoek Hospital, Amsterdam, The Netherlands.

1

2 J. J. Broerse and B. J. Mijnheer

It has been demonstrated earlier for a number of cell systems that cells in the absence of oxygen have a much higher resistance to radiation than cells in the presence of oxygen. The quantitative factor derived to describe this effect, the oxygen enhancement ratio (OER), is in the range of 2.5 to 3.5 for conventional x rays. Studies performed in the beginning of the 1960s have shown that the OER for neutron beams of different energies is considerably smaller, namely, about 1.5 to 1.8 (Broerse and Barendsen, 1965; Fowler, 1966). This reduced oxygen enhancement ratio could offer a clinical advantage, especially for those types of tumors containing a certain fraction of anoxic cells.

Secondly, investigations of the radiosensitivity of cells in different cell-cycle stages performed with x rays have demonstrated that cells in the S phase have the lowest radiosensitivity. Studies performed by Sinclair (1968) have shown that for neutrons the variation in the response throughout the cycle is less than that found for x rays. Consequently, it was shown by Sinclair that for neutron irradiations cells in the S phase have the highest RBE. The different radiosensitivity findings in different phases of the cell cycle could have clinical significance. In a situation where a tumor contains a higher percentage of S phase cells than cells of doselimiting normal tissues, a higher RBE for tumor response and a preferential killing of tumor cells can be expected (Withers, 1973). The response to neutrons will be less affected by cyclic fluctuations in radiosensitivity, and consequently variations in response from tumor to tumor would be smaller than those observed after irradiation with x rays.

Thirdly, studies on the survival of a number of normal tissue and tumor cells have shown large variations in the intrinsic radiosensitivities of these different cell types, which is demonstrated by differences in the slopes Do of the survival curves, the extrapolation numbers N, or by the linear and quadratic components (l' and {3 of the exponential survival curves. A discussion of the mathematical fitting of survival curves can be found elsewhere (ICRU, 1979a). For neutrons, it has been demonstrated that the recovery of sublethal damage is reduced and this could have clinical advantages, especially for those combinations of tumors and dose-limiting normal tissues where the survival curve for tumor cells treated by conventional x rays shows a large shoulder. Survival curves with appreciable shoulders have been observed for specific types of experimental tumors; a comparison of neutron and x-ray survival curves showed considerable differences in RBE (Barendsen and Broerse, 1977).

Based on the radiobiological indications, fast neutrons have been introduced in a number of clinics all over the world, and these are listed in Table 1. Reviews on the present-day status of fast neutron radiotherapy can be found elsewhere (Barendsen et al., 1979; Catterall and Bewley, 1979). The first clinical results indicate that for certain types of malignancies

Progress in Neutron Dosimetry for Biomedical Applications 3

Table 1. Fast Neutron Radiotherapy Installations and Approximate Number of Patients Treated until

the Beginning of 1980

Maximum Deuteron Start of Number Reaction deuteron beam clinical of employed Location Type of machine energy (Me V) (rnA) operation patients

d+T Amsterdam Sealed tubea 0.25 18 1975 400 Glasgow Sealed tubea 0.25 30 1977 16 Hamburg Rotating targetb 0.5 8 1976 340 Heidelberg Sealed tube C 0.25 500 1978 80 Manchester Sealed tubea 0.25 30 1977 70

d+Be Chiba-shi Cyclotron 30 0.03 560 Chiba-shi Van de Graaff 2.8 1967 36 Cleveland Cyclotron 25 0.03 1977 250 Dresden Cyclotron 13.5 0.04 1972 600 Edinburgh Cyclotron 16 0.1 1978 315 Essen Cyclotron 14 0.1 1978 170 Houston Cyclotron 50 0.007 1972 720 Krakow Cyclotron 12.5 1978 35 London Cyclotron 16 0.1 1966 850 Louvain Cyclotron 50 om. 1978 170 Seattle Cyclotron 21.5 0.04 1973 360 Tokyo Cyclotron 15 0.1 1975 200 Washington Cyclotron 35 0.01 1973 300

p+Be Batavia Linac 66 0.008 1976 400

"Output: 1012 S-l.

·Output: 1 - 2 X 1012 .-1.

'Output: 3 x 1012 8-1.

(especially salivary gland tumors, cervix carcinoma, and bladder cancer), better curative results can be achieved with fast neutron beams than with conventional megavoltage x rays (Dutreix and Tubiana, 1979; Battermann and Breur, 1979). With the exception of the effects on brain tissue, damage to normal tissues does not seem to be worse than with photons (Halnan, 1979). Randomized controlled trials, in which careful assessments of normal tissue response as well as of tumor response are essential, are presently underway.

It must be mentioned that neutron radiotherapy treatments were initially performed with cyclotrons constructed for other purposes (essentially unsuitable for clinical work) or neutron generators that fall technically short of modern megavolt age x-ray machines. In consequence, unsatisfactory clinical results could reflect more the inadequacy of the delivery of neutrons to the tumor than the effects of the neutrons themselves. Recently Catterall (1976) suggested a number of factors to be considered in the


construction of a neutron machine suitable for treating patients in a way that is comparable to treatment with photons from megavoltage machines now in routine use throughout the world (see Table 2). The currently available neutron machines do not fulfill these requirements, but hospitalbased isocentric high-energy cyclotron beams are under construction.

To predict the response of irradiated biological systems and to estimate the risks of mixed neutron and gamma radiation fields, it is essential to obtain a quantitative description of the radiation fields or energy deposition processes inside an object. For purposes of radiation protection a rough characterization of the radiation field in terms of the type, energy, direction, and number of particles is sufficient in most cases. For medical and biological applications, the absorbed dose and radiation quality have to be determined. The absorbed dose is defined as the ratio of the mean energy imparted to a volume element of matter and the mass of the matter in that volume element. The concept of radiation quality needed to be introduced, since it was shown that different types of radiation have different degrees of effectiveness in the production of biological damage. Although for the final interpretation of biological data it may also be necessary to know other physical quantities like the number of ionizations produced and their correlations in space and time, only energy deposition and patterns of energy deposition will be treated in this chapter.

A review of well-established neutron dosimetry methods and calculational procedures, using the information available up to 1975, was published in ICRU Report 26 (ICRU, 1977). The increased applications of fast neutrons and the need for greater accuracy in neutron dosimetry have led to a considerable accumulation of new basic physical data for neutron dosimetry, e.g., kerma factors and W values. Also, more detailed information on absorbed dose distributions in phantoms and biological systems has become available. It is the purpose of this chapter to describe these recent developments and to evaluate the current state of the art in fast neutron dosimetry.

Table 2. Minimum Operational Requirements of Clinical Neutron Generators a

1. The beam must always be available to meet clinical requirements. 2. The output must give treatment times not exceeding 4 min. 3. The depth dose and isodose shape must be at least as good as those of 60CO gamma rays. 4. The neutron generator must be within a hospital. 5. The setup of neutron treatments must not be compromised. 6. The beam should not be fixed in one position.

"From M. Catterall. Br. J. Radiol. 49. 203-205 (1976).

Progress in Nentron Dosimetry for Biomedical Applications 5

Applications of neutron dosimetry for radiation protection will not be dealt with. In this area the most important quantity is the dose equivalent, which can be determined with specific experimental techniques. Information on this topic can be found elsewhere (ICRU, 1971; Nachtigall and Burger, 1972).

To obtain a direct comparison of biomedical results from different centers, it is important to determine the energy dissipation in the irradiated material with a sufficient degree of precision and accuracy. Experimental studies on late radiation damage in the spinal cord of rats (van der Kogel, 1979) have shown that for x and neutron irradiations a dose increase of 5% above the tolerance level can lead to considerably increased percentages of paralyzed animals (see Figure 1). There is also evidence from radiotherapy applications that the probability of tumor control is a steep function of dose: deviations of 7 to 10% in absorbed dose can be detected clinically (ICRU, 1976; Battermann et aI., 1981). Therefore, an accuracy (overall uncertainty) of ±5% in dose determination is generally considered to be

SINGLE DOSE

100 • X rays X rays

50 acute 0.065 Gy /min

-0:: ., u Ii

0 Q. • ¢ , i i .. 10 15 20 ~

25 30 35 40 a 0- 5 FRACTIONS ~ ~ 100 • • ..

. ""', .. , ! 2

50 --.-. : :

0 / • .

i i i i

10 15 20 25 30

absorbed dose ( G y )

Figure 1. Dose response curves for induction of paralysis in the rat after single and fractionated irradiations of the spinal cord with x-rays and 1S-MeV neutrons (van der Kogel, 1979).


necessary. This is a very severe demand, considering the complexities of the determination of absorbed dose distributions, e.g., due to tissue inhomogeneities. For neutron dosimetry, this goal will probably not be reached in the near future, since each of the basic physical parameters required for the determination of the neutron kerma or absorbed dose already has a systematic uncertainty of 2 % or larger. Regarding the reproducibility of the delivery of the neutron beam inside a biological object, present-day technology makes it possible to perform the monitoring with a precision of better than ±2% (Broerse et al., 1978a).

Information on neutron and photon energy spectra and dose rates is essential for neutron dosimetry because of the dependence of the instrument response on these physical parameters. The uniformity and reproducibility of the absorbed dose pattern in biological specimens, including patients, are also determined to a large extent by the inherent characteristics of the neutron sources. Consequently, specific features of different neutron sources are summarized at the beginning of this chapter. Following this, the principles and currently available methods for determining absorbed doses of fast neutrons employed in biological and medical applications will be summarized.

Since neutron fields are always accompanied by gamma rays, it will be necessary to determine the two radiation components separately. However, knowledge of the absorbed dose and kerma is of limited value without further information on radiation quality. The quality can be described by the neutron and gamma-ray energy spectra. Attempts have also been made to relate quality to the microscopic distribution of locally deposited energy such as the lineal energy (y ) spectrum or the linear energy transfer (LET) spectrum.

Energy deposition by fast neutron beams at the microscopic level and at interfaces of materials of different atomic composition will be discussed. For specific biological and clinical applications, it is important to investigate the dose distribution over the biological object for radiations with collimated and uncollimated beams. Three separate sections are devoted to dosimetry for radiobiology, radiotherapy, and in vivo activation analysis. The present adequacy of neutron dosimetry can be derived from the results of neutron dosimetry intercomparisons. The results of a number of intercomparison programs will be discussed, and recommendations for future research in neutron dosimetry made at the end of this chapter.

2. NEUTRON SOURCES

The method chosen to determine the absorbed dose in a biological specimen irradiated with neutrons will depend on the characteristics


of the neutron source; i.e., the neutron yield, the gamma-ray contribution to the total kerma rate, and the energy spectra of the emitted neutrons. When the gamma-ray contribution to the total kerma is relatively high, the gamma-ray spectrum might also be important then. If good spatial resolution is required, the size of the detector should be considered in connection with the angular distribution of the neutrons emitted by the source.

Neutron beams employed for biomedical applications can be produced by accelerators, reactors, and isotopic sources. Information on the fundamental and practical aspects of these types of sources can be found elsewhere (e.g., see Marion and Fowler, 1960, and ICRU, 1977). In addition, the characteristics of high-intensity neutron sources have been recently reviewed (Ullmaier, 1977; Barschall, 1978; Barendsen etal., 1979). In this section, emphasis is placed on new developments in high-energy cyclotrons, high-output d + T neutron generators, and reactor neutron beams of special design. In accordance with general rules (ICRU, 1977), the reaction produced by particles of type i impinging on a material of type J is described as i + J, where J is the chemical symbol of the target element. When the energy E of the particles is pertinent, the use of the notation iCE) + J is recommended, where E should be expressed in megavolts. For example, d(50) + Be neutrons means neutrons that result from 50-MeV deuterons bombarding a beryllium target.

Knowledge of the neutron energy spectra of different sources is necessary for two purposes. First, the energy distribution of the neutrons characterizes the beam quality and determines the relative biological effectiveness (RBE). The biological effects of different neutron beams can be estimated if the neutron spectrum and the neutron energy dependence of the biological effect are known. Such types of theoretical assessments can be used in addition to experimental intercomparisons employing biological dosimeters as carried out by Todd et al. (1978) and Hall and Kellerer (1979) for a number of clinical neutron radiotherapy facilities. The second reason for the neutron energy spectrum to be known is that the response of a dosimeter is generally dependent on the neutron energy. As will be discussed in Section 3.2, the relevant physical parameters determining the instrument response have to be averaged over the appropriate neutron energy spectra.

2.1. Cyclotrons

Neutron beams produced by accelerating various types of particles in a cyclotron and bombarding different target materials have been extensively employed for many purposes. The original design of these cyclotrons was intended for nuclear physics research. Consequently, these high-energy


accelerators of technical complexity and considerable dimensions are generally not located in hospitals but could be converted into a medical facility by the construction of shielding providing a fixed horizontal beam. Recently, also, a fixed vertical beam has been installed for a d(50) + Be neutron beam in Louvain-Ia-Neuve (Belgium) that facilitates patient setup. Only a few cyclotrons have been installed in hospitals, e.g., the MRC cyclotron at Hammersmith Hospital (London), which produces a fixed horizontal beam of d(16) + Be neutrons. Cyclotrons especially designed for neutron therapy with rotating treatment heads are now commercially available. Isocentric facilities with approximately the same energy as the Hammersmith beam have been installed in Edinburgh (Scotland), Essen (W. Germany), and Riyadh (Saudi Arabia).

The neutron output of a cyclotron will depend mainly on the beam current, the type and maximum energy of the accelerated particle, and the type and thickness of the target. The dose rate per unit beam current for various ions and targets as a function of particle energy is given in Figure 2. For proton- and deuteron-induced reactions, the dose rate increases roughly proportionally to the third power of the deuteron energy. For equal bombarding energies, the neutron yield from the p + Be reaction is considerably lower than that from the d + Be reaction. However, most cyclotrons can accelerate protons with almost twice the energy of deuterons, and the kerma rates obtained for the two reactions with the same machine will not show great differences. The advantage of using the p + Be reaction is that these neutrons have a better penetration than d + Be neutrons produced with the same machine if lower-energy neutrons are filtered out of the beam. The energy dissipation of the protons in the target, however, is twice that of deuterons for the same neutron output. Hospital-based cyclotrons employing the p(42) + Be and p(48) + Be reactions are now being installed in Houston (U.S.A.) and Seattle (U.S.A.), respectively. The p(66) + Be reaction is employed in the medical facility installed at the Fermi National Accelerator Laboratory in Batavia (U.S.A.). The linear accelerator at this institute, which injects protons into a higher-energy synchrotron, is simultaneously available for biomedical applications (Lawrence, 1979).

The yield of neutrons from the d + Be reaction is strongly dependent on the angle of emission; the forward peaking of the neutron beam becomes more pronounced with increasing deuteron energies. Angles at which the d + Be neutron fluences fall to half their zero-degree values are 18, 12.5, and 10° for deuteron energies of 16, 33, and 50 MeV, respectively (Meulders et aI., 1975). The angular distribution of neutrons from the p + Be reaction does not show the strong forward peaking observed with d + Be neutrons. To produce a collimated neutron beam, the shielding for the p + Be reaction should be more massive on all sides than the shielding

Progress in Neutron Dosimetry for Biomedical Applications

101

I -c( ::L

I c: E

"lI

2

C ~ " u E D

" ..c

c: 100

" Q; a.. ! 2

" is "lI

0 0

10- 1 0

i. I

101

o

o o

d (35}+Be I d (35)+li a

a o x

A

a

Ox

• p(65}+Be a p(65)+li.

• p(65)+Be x

• p(45)+Be

• p(35)+Be

• p(35)+Be • p(35)+li

A • p(26)+Be Ox o

TSO = 125cm

energy ( MeV)

9

Figure 2. Dose rates per unit beam current (cGy min -1 /LA -1) for neutrons produced by protons or deuterons incident on beryllium or lithium targets. The symbols refer to the results obtained by various investigators (Amols et al., 1977).

required for d + Be neutrons. The d + Be neutron beam will require a flattening filter if homogeneous neutron fields are required, especially at higher deuteron energies and for large field sizes.

The gamma-ray contribution to the total absorbed dose will be due to photons generated as part of the neutron-producing reaction or as a result of neutron interactions with the target, collimator, or other irradiated structures. In addition, there will be a component resulting from the interaction of the neutrons with the irradiated object. The values reported for the different reactions are generally below about 3% for open-beam


geometries. The increase in the contribution of gamma rays to the total absorbed dose when employing a collimated system will depend on the angular distribution of the neutrons. At 35 MeV, the deuteron-produced neutron beams have approximately half the gamma-ray component of the proton-produced neutrons (Amols et al., 1977). For target thicknesses in which the protons do not lose all of their energy, the gamma-ray fraction increases. When a polyethylene filter is added to a p + Be beam, a significant increase of the relative gamma-ray component of the beam will be observed. Values of 8.5 and 13% have been reported for an unfiltered and a filtered p(35) + Be beam, respectively (Quam et al., 1978).

Information on neutron spectra produced by bombarding thick beryllium targets with deuteron beams with energies between 7.5 and 26 MeV can be found elsewhere (ICRU, 1977). Neutron spectra at higher deuteron energies, up to 50 MeV, were more recently measured by Meulders et al. (1975), while Harrison et al. (1978) presented spectra for d(80) + Be neutrons. These spectra as well as those measured for lower-energy accelerators show a single maximum at about O.4Ed, where Ed is the deuteron energy. Information on the number of neutrons in the energy range below 2 MeV is scarce and partly contradictory. Time-of-flight spectra of neutrons with energies greater than 0.3 MeV reported by Lone et al. (1977) showed a relatively intense low-energy component when thick 9Be targets were bombarded with 14.8-, 18-, and 23-MeV deuterons. Cranberg (1978) suggested that this low-energy peak, which has a maximum at about 0.5 MeV and extends to about 2 MeV, may be due to (d, pn) and (d,2n) reactions in beryllium. Because RBE values have a maximum in the energy range around 0.5 MeV, the presence of a low-energy component could have a considerable influence on the biological response. It is therefore necessary to obtain more information on the differential fluence distribution of neutrons below 2 MeV.

The average neutron energy En reported by Lone et al. was about 0.33Ed for neutrons above 0.3 MeV, which is lower than the value 0.42Ed reported by Parnell (1972) for neutron spectra measured above 2 MeV. Measurements by Waterman et al. (1979a) also showed the presence of a low-energy component (see Figure 3); however, the relative contribution of these low-energy neutrons was much smaller than that reported by Lone et al. The neutron kerma rates derived from the differential fluence spectra shown in Figure 3 are in good agreement with the absorbed dose rates measured with a tissue-equivalent ionization chamber. The observation by Waterman et al. was confirmed by measurements by Graves et al. (1979), who also observed only a relatively small low-energy tail in their d(49) + Be spectrum. It should be noted, however, that these low-energy neutrons can be partly filtered out of the beam by using a layer of several centimeters of polyethylene, similar to the procedure applied to p + Be neutron sources.


u "-~ ~

> " ~

'-.

" 3!

" '>'

109

108

1010

o

d ( 28 ) + Be

d ( 16 ) + Be

i 20

neutron energy ( MeV)

i 30

11

Figure 3. Energy spectra of neutrons at 0 and 45° produced by stopping 16- and 28-MeV deuterons in beryllium (Waterman et al., 1979a).

Spectra from neutrons produced by the bombardment of thick beryllium targets with protons at different angles of emission have been reported by several groups. The results of Waterman et al. (1979a) are given as an example in Figure 4. At energies below 10 MeV the spectra exhibit a large evaporation component that increases strongly with decreasing energies. Neutron spectral distributions for p + Be neutrons measured by Lone et al. showed that this low-energy component extends to at least 0.3 MeV. Filtration of the beam by the addition of polyethylene filters results in a hardening of the beam, as demonstrated by Johnsen (1978) and Graves et al. (1979), shown in Figure 5. The spectra obtained by Johnsen differ from those measured by Graves et al. and Waterman et al. for neutron energies higher than about 20 MeV. This discrepancy may be due in part to differences in the target thickness and geometry of the measurements.


p ( 35 ) + Be

109

u ~ .. ~

> " :E "-

c:

~ 108

.!:! 1010 >.

p ( 46 ) + Be

10 30 40 50


Figure 4. Energy spectra of neutrons at 0, 15, and 45° produced by stopping 35- and 46-Me V protons in beryllium (Waterman et al., 1979a).

The use of lithium instead of beryllium as target material has been considered (e.g., see Amols et al., 1977 and Quam et aI., 1978). Measurements of 35- and 65-MeV protons and deuterons on thick beryllium and lithium targets showed that the p + Li neutrons have a larger high-energy component than p + Be neutrons. The use of beryllium instead of lithium is, however, preferred owing to its superior mechanical and physical properties. If compact cyclotrons with limited accelerating voltages are available, bombardment of deuterium gas targets with deuterons results in a reasonable neutron yield and average neutron energy for therapy applications


400 '<t

0 -X

U 300 c:

> ., :! !;; 200 "-

c:

3! ., 100 >-

0

A

B

, 0

c, , , ,

p + Be

',' .... _------------

, , 10 20


13

Figure 5. Energy spectra for p + Be neutrons at 0° without and with polyethylene filter measured by Graves et al. (curves A and B, respectively) and Johnsen (curves C and D).

(Schraube et al., 1975 and Waterman et al., 1978). The energy distribution of these neutrons contains two distinct groups: the higher-energy group arises from the reaction D(d, n) 3He, while the lower-energy one is due to the deuteron breakup reaction D(d, np) D.

Other beam and target combinations to be used for biomedical applications include targets of C, Mo, Cu, Ta, and Au bombarded with high-energy deuterons. Isochronous cyclotrons are capable of accelerating 3He ions to an energy of about 2.7 times that of the deuterons. Bombardment of a beryllium target with 3He ions produces a somewhat more penetrating neutron beam than bombardment with deuterons from the same cyclotron, but at the expense of a factor of nearly 3 in yield (Parnell et al., 1975). For the other reactions, the results of Meulders et al. (1975) indicated that the neutron yield decreases almost exponentially with the atomic number of the target. The mean neutron energies are only marginally increased with respect to that of d + Be neutrons. Specific advantages of these neutron beams for clinical applications are not to be expected.

2.2. d + T Neutron Sources

Fast neutrons produced by the d + T reaction can offer some advantages for medical purposes. The neutrons have rather high energies (about 14 MeV) and can be generated at relatively low bombarding energies (a


few hundred ke V), thus requmng relatively small and inexpensive accelerators. Because this reaction has been extensively employed, a vast amount of information on detector characteristics and radiobiological data at this energy is available.

As indicated in Table 1, d + T neutron sources are now in clinical use in some centers in Europe. The neutron output of these machines varies between 1 x 1012 and about 3 x 1012 S-1, resulting in dose rates of about 4 to 15 rad/min (cGy /min) at a distance of 1 m from the target. The existing sources consist of either a metal target impregnated with tritium that is bombarded by deuterium ions or a sealed-off tube containing a mixed beam of deuterium and tritium ions.

The source strengths and lifetimes of d + T sources with metal targets are limited by the diffusion of tritium from the active target area and sputtering of the metal during the power dissipation of the deuteron beam. Methods of improving these characteristics are the use of rotating targets, analyzing the ion beam in order to use monoatomic ions only, and increasing the number of ion sources in combination with a single target. Analysis of the deuteron beam can result in considerably increased target lifetimes, as shown for the d + T neutron generator at Hamburg (Hess and Franke, 1979). A high-intensity rotating target using a mass-analyzed D+ ion beam is under construction at the Lawrence Livermore Laboratory. The yield is expected to be 4 x 1013 s-t, with a useful target lifetime between 100 and 800 hr for a beam spot of 2-4 cm (Booth et al., 1977). Although developed for material-testing studies involving fusion reactors, the machine can be converted for clinical use. Alternative approaches for the d + T reaction are the use of tritium gas targets (DeLuca et al., 1978) or drive-in targets, where the tritium is replenished by the charged particle beam consisting of equal numbers of deuterium and tritium ions (Hillier et al., 1971; Reifenschweiler, 1972; Hendry et al., 1977). Three types of sealed-tube neutron generators have been employed for neutron radiotherapy applications: the Elliott tube (Greene et aI., 1979) in Manchester and Glasgow, the Philips tube (Broerse et al., 1977) in Amsterdam, and the Haefely tube (Schmidt and Reinhold, 1977) in Heidelberg. The last system employs a conical scandium deuteride-tritide target surrounded by a ring-shaped ion source. The mixed-beam design in the sealed-off tube has presented a number of technical problems and will require further technological improvements. For commercial reasons, the production of the Elliott and Philips tubes has been diminished or stopped; consequently, the current neutron therapy programs employing these sealed tubes in Glasgow and Amsterdam had to be terminated in 1980 and 1981, respectively.

Although for some medical applications, e.g., in vivo activation analysis, the output of existing d + T generators is adequate, a higher output,


in excess of 1013 s-t, is required to perform radiotherapy comparable to megavoltage photon radiotherapy. It should be noted that the output during a neutron burst can be a factor of 105 higher if a small d + T generator is used in the pulsed mode. This will be advantageous for certain techniques of in vivo neutron activation analysis (Oxby et al., 1980).

2.3. Reactor Neutron Beams of Special Design

Fission neutrons produced in a reactor core or in a 235U converter exposed to thermal neutrons from the core have been employed for radiobiological studies (ICRU, 1977). This type of facility can provide large fields for the simultaneous irradiation of many biological objects, which is useful for the study of effects in animal populations.

Up to now, the use of reactor beams in radiotherapy has concentrated on boron neutron capture therapy. The lOB(n, a) 7Li reaction has a very high cross section for thermal neutrons and produces secondary particles having ranges equivalent to cellular dimensions. The thermal ftuence distribution will determine the effectiveness of the exposure arrangement. Gamma rays and fast neutrons will contaminate the thermal neutron beam. By using a suitable beam portal and bismuth shielding, these components could be reduced for the MITR-II Medical Therapy Facility (Murray et al., 1978). Owing to the small relaxation length of thermal neutrons (about 2 cm), the thermal neutron ftuence will be much greater at the surface than below it. Improvements in the thermal neutron ftuence distribution can be obtained by using intermediate energies, e.g. 2-keV neutrons (Harvey and Mill, 1978). Successful therapy depends on higher accumulations of lOB in the tumor than in surrounding normal tissues. There are some indications that sufficient accumulations can be achieved in both melanoma and glioblastoma with lOB-labeled tumor-seeking organic compounds synthesized especially for this purpose (Hatanaka and Sweet, 1975). A major objective of further research should be the development of substances containing boron or other nuclides having large (n, a) cross sections, e.g., 6Li substances that are accumulated in high concentrations in different types of tumors, taking advantage of specific biochemical pathways.

A reactor neutron therapy facility is presently installed at a swimming pool research reactor in Munich (Koster et al., 1979). A 235U converter plate is placed near the reactor core at the opening of a horizontal beam line. The use of fission neutrons is planned for the treatment of superficial tumors. By applying lead filters of different thicknesses, the contribution of gamma rays to the neutron beam can be varied and this will allow the use of different mixed beams.

16 I. I. Broerse and B. I. Mijnheer

Monoenergetic neutron beams with energies between thermal (0.025 eV) and a few hundred keY, for use in radiotherapy and radiobiology and for the calibration of health physics instruments, can be produced with reactors (Schwartz, 1977; Alberts and Knauf, 1978; Harvey and Mill, 1978). Using combinations of resonant scatterers and filters, very pure beams of 2-ke V (scandium filter), 24-ke V (iron-aluminum filter), and 120-144 keY (silicon filter) neutrons can be produced. The filters preferentially transmit neutrons through certain windows in their cross sections. By combination with a second filter, unwanted neutrons transmitted through secondary windows can be filtered out.

3. PRINCIPLES AND METHODS IN NEUTRON DOSIMETRY

A number of instruments and methods are available for measuring neutron fiuence, kerma, and spectra in neutron fields (e.g., see ICRU Report 13, 1969a). A detailed discussion of the properties of the different types of neutron dosimeters used in biology and medicine is given in ICRU Report 26 (ICRU, 1977). Since that time, only a few new methods have become available for measuring fast neutron absorbed doses and these will be discussed in this section. The accuracy of the physical parameters and the uncertainties in the procedures applied to derive absorbed dose values from the readings of a number of existing dosimeters have been, however, greatly improved upon. Also, a large amount of experience has been gained from neutron dosimetry intercomparisons among several institutes using the same method or among different dosimeters in the same beam (see Section 7).

It is obvious that the choice of a particular type of dosimeter depends on a number of factors, such as the accuracy of dose determination required, the sensitivity of the system, and neutron energy dependence. Also, the size of the dosimeter and the processing of the response may be important. If the instrument is used for calibration of the beam at a reference point, then the absolute accuracy of the system will be the most important property to consider. For relative measurements, e.g., the determination of isodose lines or during in vivo dose determinations, the ease of operation and spatial resolution will be more important than absolute accuracy. If the photon component of the field has to be determined, the relative neutron sensitivity of the photon dosimeter should, in general, be as small as possible. The neutron energy dependence will be important for all dosimetry systems and this will determine the degree of detail to which the neutron energy spectrum should be known. Inaccuracies introduced by the application of conversion factors considerably different from unity, used to derive


the tissue absorbed dose from the instrument reading, can be reduced by the use of tissue-equivalent devices.

3.1. Ionization Chambers

At present, tissue-equivalent (TE) ionization chambers are generally accepted as the most practical instruments for absolute measurements as well as for the determination of relative dose distributions. Nonhydrogenous ionization chambers are often employed for the assessment of the photon components of neutron fields.

The objective of clinical neutron dosimetry is the determination of the absorbed dose in a certain reference material, for which, most commonly, muscle tissue with the composition recommended by the ICRU (ICRU, 1964) is used. The wall, central electrode, and gas filling of a TE chamber are therefore chosen to match this composition as closely as possible, although, in practice, correction factors will be necessary to account for small differences in composition. A common electrically conductive plastic used in the construction of TE ionization chambers has been a particular muscle-equivalent formulation designated as AlSO (generally supplied by the Physical Sciences Laboratory, Illinois Benedictine College, Lisle, Illinois). The elemental weight composition for AlSO plastic has been determined by Smathers et al. (1977) for a number of samples. TE gas with the composition recommended by Rossi and Failla (1956) is often employed in combination with AlSO plastic wall and central electrode material. Information on the techniques of TE ionization chamber construction and the properties and fabrication of tissue-equivalent plastic can be found elsewhere (ICRU, 1977).

The assessment of absorbed dose from a measurement with an AlSO plastic TE ionization chamber with TE gas filling can be divided into several steps. The charge Q produced within the cavity is derived from the reading R of the chamber multiplied by several correction factors IIKR, including corrections for the electrometer calibration factor, ion recombination, temperature and pressure, gas flow rate, and leakage current. Dividing the collected charge by the charge e of the electron and the mass m of gas within the cavity yields the number of ion pairs per unit mass of gas. If this quantity is multiplied by the average energy W required to produce an ion pair in the gas, then the absorbed dose in the gas cavity is obtained. The absorbed dose in the wall material adjacent to the cavity of the chamber can be derived from the energy absorbed by the gas using the gas-to-wall absorbed dose conversion factor rm,g (Bichsel and Rubach, 1978). The absorbed dose in the reference tissue can then be calculated by using the ratio of mass energy absorption coefficients in the reference tissue t and the wall material m, (lLen/ P ),/ (lLen/ P )m. If measurements are carried out in


a phantom, the quantity to be assessed is the absorbed dose at the center of tHe chamber when the chamber is replaced by phantom material. Under certain conditions (e.g., in a phantom or close to a source free in air) the effective point of measurement can be displaced and a displacement factor Kd has to be applied that is defined as the ratio of the absorbed dose for an infinitesimally small cavity to the absorbed dose measured. Consequently, the absorbed dose in tissue, D" can be derived from the reading of the TE ion chamber:

(1)

The mass of the gas in the cavity can be obtained from the absorbed dose calibration factor U e of the tissue-equivalent ionization chamber, which is defined as

(2)

where D~ is the absorbed dose in the tissue at the wall adjacent to the gas cavity of the chamber, and the subscript c refers to the photon beam used for calibration. Similar to equation (1), it can be derived that

m = ~ We (Sm,g)e( (J-Len/P)t) U e e (J-Len/ P)m e

(3)

In this equation, the gas-to-wall absorbed dose conversion factor 'm,g has been replaced by (Sm,g)" since the chambers used for biomedical applications are usually small enough to satisfy the conditions for the Bragg-Gray theorem at the photon calibration energies normally used (e.g., 60CO or 137 Cs gamma rays). It should be realized that for calibration with photons a number of correction factors similar to those used for neutron measurements should be applied. The calibration factor U e comprises specific correction factors for the finite size of the chamber and also includes the mass energy absorption coefficients in air and TE material and the exposure-absorbed dose conversion factor (Mijnheer and Williams, 1981).

The calibration of a TE ionization chamber to be used for neutron dosimetry should in principle not introduce large uncertainties, because the procedure is analogous to that used in photon dosimetry. Nevertheless, variations of up to ±8% in the calibration factors obtained with different photon beams have been reported (Broerse et al., 1978a). It is recommended, therefore, that TE ionization chambers be calibrated against an exposure standard chamber that has a calibration factor traceable to a national standards laboratory.

Neutron fields are always accompanied by photons contributing to the total absorbed dose. Consequently, equation (1) has to be generalized to

Progress In Neutron Dosimetry for Biomedical Applications 19

express the total collected charge QT in terms of the separate neutron and photon components of the absorbed dose (DN and Do):

where

QTcxAKdh = RT(TIKRhcxe(Kdh = kTDN + hTDO (4)

kT = We (Sm,g)e [(P,en/P)t/(P,en/P)m]e WN (rm,g)N (Kt/ Km)N

hT = We (Sm,g)e [(P,en/ P )t/(P,en/ P )mJc Wo (Sm,g)O [(P,en/P)t/(P,en/P)m]O

(5)

(6)

The subscript T stands for the tissue-equivalent device and Nand G refer to the neutron and photon components, respectively.

Comparisons of different TE ionization chambers employed by various groups (see Section 8) have shown that there are large systematic differences in the measurement procedures connected with, for example, calibration with photons, the gas flow rate, the collecting potential, the polarity effect,

A

c

Figure 6. Design sketches and x-ray radiographs of four different types of TE ionization chambers: A, CENF; B, Exradin; C, FWT; and D, TNO (Schraube et aI., 1980).


the correction for wall thickness, and the choice of the effective point of measurement in a phantom. In an effort to reduce the systematic uncertainties in ionization dosimetry, information has been collected on the characteristics of eight different types of tissue-equivalent ionization chambers (Broerse, 1980). Four chambers, CENF, Exradin, FWT (formerly EGG), and TNO, have been subjected to a number of performance tests (Schraube et ai., 1980). The design drawings and x-ray radiographs of the chambers given in Figure 6 show some differences between design and actual construction, mainly connected with chamber wall and central electrode position. A number of conclusions resulting from these·tests can be summarized:

1. Under defined practical conditions of ionizing potential and TE gas flow rate employing identical cables and electrometers and applying daily photon calibrations, the relative responses of the four chambers in fast neutron fields show maximum variations from the mean of less than 2 % .

2. The responses of connectors to radiation show large variations. Massive triaxial connectors show sensitivities 10 to 100 times greater than those of small coaxial connectors.

3. The wall thickness corrections for 137CS photons and 0.6-MeV neutrons are smallest for the CENF chamber, intermediate for the Exradin chamber, and largest for the FWT and TNO chambers. The wall effect for 0.6-Me V neutrons is shown in Figure 7. The differences between the thimble-type and spherical ion chambers

i (d ) In

i ( 0 )

o

0.05

0.1

0.15

o

o CENF " EXRADIN o FWT

" TNO

2 3 i 4

(il • I

wall thickness d ( mm )

5

Figure 7. Effect of wall thickness on ionization chamber reading during 0.6-MeV neutron irradiation (Schraube et aI., 1980).


1.03 o CENF chamber:0.85 R min-I

1.02

1.01

1.00 o

o EGG/GSFM chamber:11.6 R min-I

• EGG/GSFM chamber:128 R min-I

" TNO chamber: 1 .4 rad min-I

v-2

Figure 8. Relative saturation current isatl i(V), dependent on the inverse square of the chamber voltage for different ionization chambers for photon irradiation (Broerse et aI., 1979a).

can be attributed to differences in effective wall thickness, while the differences between the CENF and Exradin chambers are likely to result from the different central electrode constructions.

4. Gas flow of the four chambers without tubes connected to the outlets, does not introduce problems with pressure buildup in the flow rate region of 10 to 60 cm3 min-I. A preflush with TE gas at 60 cm3 min -1 for 5 min is sufficient to obtain a stable reading for all chambers. Replacement of TE gas with air changes the sensitivities of the four chambers irradiated with 137 Cs photons by the same factor, while with 0.6-MeV neutrons large differences are observed from chamber to chamber. For higher-energy neutrons these differences are smaller. Air-filled chambers should not be used in neutron dosimetry applications for beams with an appreciable low-energy component.

5. The operational characteristics of the four chambers tested (e.g., angular dependence and ion collection) are satisfactory, which would not have been expected on the basis of the radiographs.

Under specific experimental conditions, changes in the correction factors KR have to be anticipated. Williams (1980) and Mijnheer (1980)

22

1.02

1.01

1.00

o

o CENF chamber

o EGG/GSFM chamber

'" TNO chamber

J. J. Broerse and B. J. Mijnheer

0 ........

v- 1

Figure 9. Relative saturation current isatl i(V) dependent on the inverse of the chamber voltage for different ionization chambers for irradiation with 15-Me V neutrons (Broerse et al., 1979a).

showed that the composition and length of the gas tubes have a considerable influence on the reading of the chamber. It should be further noted that some of the correction factors, e.g., those for incomplete ion collection, can differ for the calibration situation and during measurements in the neutron beam. For photons, volume recombination (inversely proportional to the square of the voltage), and, for neutrons, columnar recombination (inversely proportional to the voltage) are predominant. Volume recombination involves separate tracks and increases with the dose rate. Columnar recombination is determined only by the ion density along each track and is independent of the dose rate. Saturation characteristics for three different TE ionization chambers are shown in Figures 8 and 9 for photons and neutrons, respectively (Broerse et al., 1979a). For photons, the dependence of the saturation curves on the dose rate is evident from the results of the EGG/GSFM chamber.

3.2. Physical Parameters for Dosimetry with Ionization Chambers

Recent experimental and theoretical studies have resulted in more accurate values for a number of the physical parameters mentioned in the previous section. The parameters recommended by the European Clinical Neutron Dosimetry Group (ECNEU, 1981) for determining the absorbed dose in a neutron beam are summarized in Table 3. The estimated uncertainties in the determination of the total absorbed dose derived from measurements with a TE ionization chamber are given in Table 4.

The energy required to create an ion pair in the chamber gas in a neutron beam, WN , depends on the types and spectra of the secondary


Table 3. Parameters Recommended by ECNEU for the Determination of the Absorbed Dose in a Neutron Beam

Parameter

kT : We WN

('m.g)N [(/Len/ P ),/ (/Len/ P )mJc

(K,/Km)N ku: GM counter

Mg/ Ar chamber hT hu: GM counter calibrated

with 60CO gamma rays Mg/ Ar chamber

Value

29.3JC I

(31.1-31.3t ICI

0.99 1.001

0.95-0.98Q

0.005-0.025 Q

0.07-0.17 Q

1.00 1.00

1.00 1-2.5 x 1O-3 ,b

(r in mm)

Reference

Goodman and Coyne, 1980 Goodman and Coyne, 1980

For 60CO, 2 MV, or 137CS (Hubbell, 1977)

Caswell et al., 1980 Mijnheer et al., 1979a

Zoetelief et aI., 1980

"The value of the parameter is dependent on the neutron spectrum. The values given indicate the approximate range of values for the neutron sources used by the ECNEU groups.

"The value for spherical chambers in a phantom irradiated with d + T neutrons. The geometrical center of the chamber is taken as the effective center of measurement.

Table 4. Uncertainties (in %) in the Determination of the Total Absorbed Dose in Tissue in a Neutron Field with the Aid of a TE Ionization Chamber

Source Symbol Uncertainty

Reading of the chamber R 0.2 Reading correction factor IlKR 0.3 Displacement correction factor Kd 0.5 Absorbed dose calibration factor ae 1.2 Ratio of average energy required to create an ion

pair for A-150/TE gas chamber WN/Wc 4.0 Ratio of gas-to-wall absorbed dose conversion

factors (rm.g)N/(Sm.g)c 2.0 Ratio of neutron kerma in tissue and that in chamber

wall (K,/Km)NQ 2.6-9.7 Photon mass energy absorption coefficients ratio (/Len/ P ),/ (/Len/ P)m 0.1

OverallQ: 5.4-10.8

"The low and high uncertainty values refer to d(16) + Be and p(66) + Be neutrons, respectively. The overall uncertainty has been calculated as the square root of the sum of the squares of the separate uncertainties that are estimates of one standard error.


charged particles. Goodman and Coyne (1980) recently presented detailed calculations of WN as a function of neutron energy, using evaluated data on W values for charged particles in TE gas. The results of their calculations for neutrons in the energy range between 0.1 and 20 Me V are shown in Figure 10. These data should be applied for the appropriate neutron spectrum multiplied by the tissue kerma factor. An uncertainty of ±4 % in the ratio WN/Wc has been estimated by Goodman and Coyne. At the time of issue of ICRU Report 26 only limited information on W values for heavy charged particles in TE gas was available. A value Wc/WN = 0.95 was recommended, with an uncertainty of ±5% for neutron energies above 1 Me V and an increased systematic uncertainty at lower energies. Obviously, the accuracy in W is improved, although the uncertainty in the W ratio for neutrons and photons is still considerably higher than that of 0.5% quoted for photon dosimetry using air-filled chambers (ICRU, 1979b).

Only limited information is available for the gas-to-wall dose conversion factor for neutrons, (rm,g)N, which is sometimes also called the effective stopping power ratio. Its calculation is complicated by the differences in the number and spectra of secondary particles and the accuracy is limited by the lack of adequate stopping power data for the gas and solid phases.

33.0

32.0

31.0

..

ii'i

100

..

..

, 'i

10 1

neu Iron energy ( MeV)

1.12

WN W

c

1.10

1. 08

1. 06

Figure 10. Histogram of energy-bin average WN for TE gas. The triangles refer to WN values at point energies of resonances or minima (Goodman and Coyne, 1980).

Progress in Neutron Dosimetry for Biomedical Applications 2S

1.10

1.05

1.00 -----.

0.95

0.90

-- Caswell,Coyne,Rondolph, 1980 ----- ICRU report 26, 1977

.2 0.85

ii' iii i • iii iii

100 10 1


Figure 11. Ratio of kerma in ICRU muscle tissue to kerma in A150 plastic based on the 1977 and 1980 kerma factors from Caswell et al. (1980).

A recent survey of Dennis (1980) suggests, however, that for most neutron energies, (rm,g)N would lie in the range of 0.95-0.99 and is most likely to be towards the higher end of this range. The uncertainty in this ratio quoted by Dennis is 1 to 2%.

Owing to differences in composition between muscle tissue and AlSO plastic, the kerma ratio (Ktf Km)N will deviate from unity. Caswell et al. (1980) recently published an updated version of their tables of kerma per unit ftuence. As can be seen from Figure 11, the new data show some changes when compared with the older data published in ICRU Report 26, especially for neutron energies in excess of 7 MeV. The main reasons for the differences are the availability of better data on the total cross section for carbon and a different treatment of the 12C(n, n') 3a reaction. The kerma values calculated by Caswell et al. extend from 0.025 e V to 30 MeV. For higher neutron energies the calculated values of Alsmiller and Barish (1977), Wells (1979), and Dimbylow (1980) can be used for kerma values up to 80 MeV. Uncertainties in the kerma ratios increase with rising energy because of a lack of cross-section information in the higher energy range. Values varying between 2.6% for d(16) + Be neutrons and 9.7% for p(66) + Be neutrons have been reported (Bewley, 1980).

Corrections for the radial displacement of the effective measuring point from the geometrical center of the chamber irradiated with fast neutrons in a phantom have recently become available (Shapiro et aI., 1976b; Zoetelief et al., 1980). Shapiro et al. inferred displacement correction factors /Cd of 0.970 for a 1.0-cm3 chamber and 0.989 for a 0.1-cm3


chamber from measurements in a phantom irradiated with d(35) + Be neutrons. Zoetelief et al. performed measurements with a set of three spherical TE chambers with different cavity radii and observed differences in the absorbed dose values when the geometrical centers of the chambers were placed at equal depths. These differences in the absorbed doses were converted into radial displacements by using the depth dose curves. When the radial displacement was plotted against the chamber radius, a linear relationship resulted, as shown in Figure 12. The radial displacements d derived for two different d + T neutron beams were (0.23 ± 0.06), and (0.30 ± 0.06),. These values are significantly smaller than the radial displacement (0.58 ± 0.06), observed for 60 Co gamma rays, which is attributed by these authors to differences in attenuation and scattering processes for neutrons and photons. Displac.;ment correction factors Kd calculated for both d + T neutron beams resulted in the same displacement correction factor of 1 - (0.25 ± 0.06)10-2" where, is expressed in millimeters. For a 1-cm3 spherical ionization chamber, the experiments of Zoetelief et al. provided a factor Kd = 0.984 ± 0.005 for d + T neutrons. For a chamber of the same dimensions, Shapiro et al. (1976b) derived a value of 0.970. This latter value has an uncertainty of 1 % (which can be assessed from their original data) and consequently the results of the two groups are probably not significantly different from each other. Recent experiments (Zoetelief et al., 1981) have demonstrated that the displacement correction factors are dependent on the photon and neutron energies. As shown in

E E 8

"" c: .~ 6 u ~ (; u ~ 4 c:

" E

" u ..2 .%

2

"" 0

0 5 10 15

cavity radius, r ( mm )

Figure 12. Radial displacement d versus cavity radius r for spherical TE ion chambers (Zoetelief et aI., 1980).


Table 5. Displacement Correction Factors Kd of Spherical Ionization Chambers for Measurements in

Phantoms with Different Types of Radiation a

Type of radiation

150-, 200-, and 300-kV x rays 137 Cs 'Y rays 60CO 'Y rays Fission neutrons (En = 1 MeV) d(2.3) + D neutrons (En = 5.3 MeV) d(0.25) + T neutrons (En = 14.2 MeV) d(O.5) + T neutrons (En = 14.8 MeV) d(50) + Be neutrons (En = 21 MeV)

1.000 ± (0.05)10-2 ,

1 - (0.22 ± 0.05)10-2,

1 - (0.37 ± 0.04)10-2,

1.000 ± (0.1)10-2,

1 - (0.25 ± 0.09)10-2,

1 - (0.25 ± 0.06)10-2,

1 - (0.25 ± 0.06)10-2,

1 - (0.21 ± 0.05)10-2,

"From J. Zoetelief et al., in Biomedical Dosimetry: Physical Aspects, Instrumentation and Calibration, International Atomic Energy Agency, Vienna (1981).

27

Table 5, for neutrons of relatively low energy, no displacement was found, whereas for neutrons in excess of 5.3 MeV, Kd shows an almost constant value. In general, it will be preferable to use relatively small ionization chambers to minimize the application of necessary corrections.

The overall uncertainty in the determination of the total absorbed dose DN + Da in a neutron field using a TE ion chamber and applying equation (4) is given in Table 4 for two neutron energy spectra. The indicated uncertainties are either estimated or taken from the references mentioned above and may therefore differ in their definition. It can be seen from Table 4 that at higher neutron energies the overall uncertainty is mainly determined by the uncertainty in the kerma ratio. It has been suggested, therefore, that the absorbed dose be expressed with AlSO plastic as the reference material (Schraube, 1980).

3.3. Calorimeters

The overall uncertainty in absorbed dose determinations with ionization chambers in neutron beams is greater than that for photon beams. It is desirable, therefore, that the results obtained with this method be compared with those derived from other dosimetry systems. Calorimetry is in principle an absolute method, since it requires no calibration in a known radiation field. The method, however, is too cumbersome to be employed for routine purposes and it should rather serve as a standard for calibration purposes. The accurate use of calorimeters depends on the precise knowledge of the thermal defect in the absorber material, i.e., the fraction of the imparted energy that does not appear as heat, which may amount to several percent.


Highly purified graphite is often employed as a "'eference material for the construction of calorimeters in photon dosimetry. The uncertainty in the determination of the absorbed dose with a calorimeter at a point in carbon is estimated to be smaller than 0.5% (ICRU, 1969b). Conversion of this photon absorbed dose in graphite to that at a point in water increases the overall uncertainty in calorimetric determination to about 2 %. Direct determination of the absorbed dose in water by calorimetry will reduce this uncertainty. Recently, Domen (1980) constructed a water calorimeter by immersing an extremely small thermistor sandwiched between two thin polyethylene films in a water bath.

In principle, the same materials used for the determination of the photon absorbed dose by calorimetric methods can be applied in neutron dosimetry. Since for biomedical applications the quantity of interest is the absorbed dose in tissue, the composition of the calorimeter material should preferably be hydrogenous. Neutron dosimetry has been performed with AlSO tissue-equivalent plastic (Bewley et al., 1974a; McDonald et al., 1981a,b) and polyethylene calorimeters (Greene and Williams, 1978). An important difference of the use of TE plastic instead of graphite as an absorbing material is that the thermal defect, which is very small in graphite, can no longer be neglected in TE plastic. For neutrons with energies up to 15 MeV, a thermal defect of (4 ± 2)% has been recommended for AlSO plastic (ICRU, 1977). Recent measurements (Goodman and McDonald, 1980) support this value to within 1%. According to ICRU Report 26 (ICRU, 1977), the overall uncertainty in the calorimetric measurement of the absorbed dose in AlSO plastic was estimated to be about 3%. Using the new values for the thermal defect, McDonald et at. (1981 b) state that the overall uncertainty has now been reduced to about 1.5%.

AlSO plastic calorimeters have recently been compared with TE ionization chambers in a number of neutron radiotherapy beams in the United States (McDonald et al., 1981a) and in Europe (McDonald et al., 1980b). If the ionization chamber readings were converted into absorbed doses in AlSO plastic using the protocols adopted by either the U.S. physicists (AAPM, 1980) or their European colleagues (ECNEU, 1981), then the differences with respect to the calorimetric value were less than about 2 % for all beams. This lends confidence to the recommended absorbed dose calculation procedures and data as outlined in both protocols. This also suggests that the uncertainties given in Table 4 should be considered as conservative estimates. It should be noted that in the overall uncertainty in the determination of the total absorbed dose in AlSO plastic the uncertainty in the kerma ratio is not included, restricting the overall uncertainty to about 5%.


3.4. F1uence Measurements

The neutron kerma in a material can be calculated at a point of interest when the differential neutron fluence spectrum t/JE = dt/J(E)/ dE at this position is known. The kerma is, according to ICRU (1977), given by

1 Joo K = - I-Ltr(E)Et/JE dE p 0

(7)

where I-Ltr is dependent on the atomic composition of the irradiated material. Kerma factors, defined as the product of the mass energy transfer coefficient I-Ltr/ P and the neutron energy E, have been calculated for a number of elements and neutron energies, as discussed in Section 3.2.

A number of methods and instruments are available for the determination of differential fluence spectra. Among these are activation and fission detectors, proton recoil spectrometers, and time-of-flight techniques [see ICRU Reports 13 (1969a) and 26 (1977)]. Because the available neutron spectrometry techniques have their limitations, it is useful to compare the results of spectrum measurements obtained by different methods. Lawson et al. (1972) compared threshold detector measurements with results obtained with a liquid scintillator for collimated d + T neutron beams. The same two techniques were compared by Bonnett and Parnell (1976) for the Hammersmith d(16) + Be neutron beam. Both groups of workers found good agreement between the two methods for energies higher than about 3 MeV. Below this energy the liquid scintillator gave higher results, probably due to the contribution of pulses from gamma rays and alpha particles in the scintillator. Schmidt and Magiera (1978) found good agreement between neutron spectra obtained from activation detectors and from measurements with a stilbene crystal for energies in excess of about 2 Me V.

Kerma values derived from the differential fluence spectra can be compared with those determined with ionization chambers. Such intercomparisons have been performed for d + T neutrons by Barschall and Goldberg (1977) using activation detectors, Thomas and Lewis (1981) making use of associated alpha-particle fluence measurements and proton recoil detectors, Nguyen et al. (1981) using scintillation detectors, and Mijnheer et al. (1981a) employing a set of activation and fission detectors. The resulting neutron kerma values have uncertainties of approximately 5%. In general, the comparison of the results of the fluence method with those of the ionization chamber method shows an agreement of better than 2 %, which is well within the overall uncertainties for both methods. For d + Be neutrons, Waterman et al. (1979a) found good agreement (within 5% for up to about 15 Me V) between the neutron kerma rates calculated from differential fluence spectra and those measured with a TE ionization

30 I. I. Broerse and B. J. Mijnheer

chamber. It should be realized, however, that both methods apply the same kerma factor for tissue and this reduces the indicated overall uncertainties in comparing the two methods.

From these comparisons of the different methods for deriving neutron kerma, it might be concluded that in the energy range of a few Me V to about 15 Me V the values of the parameters chosen for the conversion of measured ionization to absorbed dose are quite appropriate and will probably not change very much in the future. At higher neutron energies, increased knowledge of these physical parameters will be necessary for obtaining adequate accuracies in kerma and absorbed dose values. Because the main uncertainty is in the kerma factors at these energies, more data on cross sections and information on the energetics involved in the nuclear reactions are required. The most important nuclear data needed are the reaction cross sections and total cross sections for oxygen and carbon for the neutron energy range between 15 and 80 MeV. From these data the energy spectra of charged particles from these elements can be derived. Especially for neutron beams with energies higher than about 15 MeV, intercomparisons of different techniques may elucidate systematic uncertainties in the applied methods. Because kerma factors are also needed in the calorimeter and ftuence methods, these methods are hardly more accurate than that using ionization chambers. It can be concluded that the TE ionization chamber is the most practical instrument for neutron beam standardization as well as for routine neutron dosimetry for these highenergy beams.

3.S. Solid· State Dosimeters

There are reviews on the characteristics of solid-state detectors, which register changes induced by radiation in solid materials, to be used for photon dosimetry (e.g., Attix and Roesch, 1966) and neutron dosimetry (e.g., ICRU Report 26). Some of these methods have been used for in vivo dosimetry during patient treatment with fast neutron beams. In this section, recent information on the use of nuclear track detectors, silicon diodes, lyoluminescent dosimeters, thermoluminescent dosimeters, and photographic emulsions for biomedical applications will be summarized. For the latter two types of solid-state dosimeters, the neutron sensitivity is generally much smaller than the photon sensitivity and these systems can be employed to determine the photon component in a neutron beam.

Nuclear track detectors record tracks of radiation damage in insulating materials due to heavy charged particles. The charged particles are produced by neutrons in the dielectric material by either direct interaction or indirectly, e.g., by fissile material or hydrogenous material that acts as a proton radiator. After etching the tracks, their number can be determined


by the use of a microscope or by means of automatic track-counting methods. As shown by Miles and Sherwin (1977), fission fragment registration can be used as an absolute device for measuring neutron fluence because the efficiency of track registration can be calculated. As discussed earlier (ICRU, 1977), a disadvantage of solid-state track recorders is the angular dependence of their response, which limits their usefulness for absorbed dose measurements inside phantoms. The registration of tracks induced directly in different types of polymers has been studied for a number of neutron energies (Spurny and Turek, 1977; McGinley and Sohrabi, 1979). These measurements showed that for neutron energies above about 1 Me V the energy dependence of some polycarbonate foils is less than 20%. Because the photon sensitivity of the polycarbonate foils is low, the detectors give a direct measurement of the neutron component of a mixed beam and can even be used for the determination of neutron contamination of a photon beam (see Section 3.7).

Silicon diodes are sensitive to neutron radiation because silicon atoms are displaced from their normal crystal-lattice positions. The resulting change in conductivity can be used as an indication of the neutron absorbed dose. The sensitivity of silicon detectors is a rapidly varying function of neutron energy because the kerma factor for silicon differs from that in tissue. Data on the neutron energy dependence and other properties of silicon diodes can be found elsewhere (ICRU, 1977). The sensitivity of the diodes remains constant within ±20% for neutron energies between 350 KeV and about 15 MeV. For higher energies, data are available from Prichard et al. (1973), who observed a difference in sensitivity of about 16% between d(16) + Be and d(50) + Be neutrons. These authors found good agreement between depth doses measured with silicon diodes and those obtained with ionization chambers. Because changes in neutron spectra with depth are relatively small for these broad energy spectra (see Section 4), silicon diodes are useful for mapping neutron dose distributions. Frank et at. (1978) used silicon detectors for measurements in a phantom irradiated with 252Cf neutrons. Owing to their small size and low response to low-energy neutrons and photons, silicon diodes are found to be convenient and reliable integrating dosimeters for in vivo measurements in patients undergoing fast-neutron therapy. Results of these types of measurements will be discussed in Section 6. Although neutrons are much more effective in producing permanent damage to the lattice structure of silicon than are photons or electrons, the photon sensitivity is not negligible (e.g., see McCall et al., 1978).

Lyoluminescence is the effect of emitting light shown by some materials when dissolved in water after exposure to ionizing radiation. The intensity of the light output is a measure of the absorbed dose in the material. The application of lyoluminescent materials in neutron dosimetry has been


reported recently (Puite and Crebolder, 1977; Ettinger et al., 1978). The effect has been observed in a number of materials, in which saccharides and amino acids are of particular interest because of their composition, which is similar to that of soft tissue. The available data show that the neutron sensitivity for d + T neutrons is about a factor of 2 higher than that for fission neutrons. According to Ettinger et al., this change in efficiency is caused by the variable effectiveness of producing free radicals in solids. This explanation is confirmed by the relation found between the efficiency of d(16) + Be neutrons in producing Iyoluminescence in some organic solids and the effectiveness of these neutrons in creating free radicals in alanine powder. The latter process was investigated by Simmons and Bewley (1976) using the electron spin resonance technique. Lyoluminescent dosimeters may be considered as a transfer dosimetry system if the variations with neutron energy can be taken into account. However, owing to the low sensitivity of the system, relatively high doses have to be applied.

Thermoluminescent devices (TLD) have been extensively studied in neutron fields, but their application for purposes other than radiation protection is limited. Two main difficulties in the interpretation of TLD results are the low neutron sensitivity in comparison to that for photons and the dependence of this sensitivity on neutron energy. The sensitivity of the TLD is partly dependent on the efficiency of different ions in inducing thermoluminescence, which decreases with increasing LET. Multiplying the kerma values due to the various charged particles of the neutron reaction in the phosphor by the TL sensitivity of that type of ion results in a relative neutron sensitivity ku after normalization to 60CO photons. In this way, revised energy response curves have been calculated by Tanaka and Furuta (1977) for 6LiF and 7LiF (see Figure 13) and by Rinard and Simons (1979) for CaF2 and 7LiF. Some recent accurate ku measurements of TLD materials have been performed by Rossiter et al. (1977) for 14.7-MeV neutrons and by Attix et al. (1978) for neutrons produced by bombarding a thick Be target with beams of 15-, 25-, and 35-MeV deuterons. Horowitz et al. (1979) used the paired LiF thermoluminescent technique for determining the photon dose in mixed fields and observed two significant limitations: first, the variation in TL response as a function of LET from dosimeter batch to dosimeter batch due to minute variations in impurity doping concentrations and, second, the likelihood of very large TL signals in 6LiF due to thermal neutrons, compared to those due to fast neutrons. This results in the necessity of an additional dose or ftuence measurement of the thermal neutrons.

The neutron sensitivity of the TLD material strongly depends on the atomic composition of the dosimeter container. Blum et al. (1976) investigated the neutron energy dependence of CaS04:Tm in small graphite tubes with the pure phosphor and a homogeneous mixture of the phosphor and


., u c: .,

.2 .... c: ON .:: I :> E ., u c:

c:

~ ~ :> :>

Gi ~ a.

o >

7LiF T LD

i i .iI", . 'Ii iii ••• 'fi'i iii" '"

10- 3 10-2 10- 1 100

neutron energy ( MeV 1

33

Figure 13. Responses of integral thermoluminescence of 6LiF and 7LiF to neutrons as a function of energy (Tanaka and Furuta, 1977).

glucose. Schuhmacher and Krauss (1978) studied the response of CaS04:Dy encapsulated in polyethylene and lead. The introduction of the proton radiator increased the neutron sensitivity of the TL dosimeter, but the response was still smaller than that for photons and strongly dependent on the neutron energy. The ku values for the phosphor enclosed in the lead and that in graphite screen also varied with neutron energy and had a much lower value than with the proton radiator, but they were mutually different. This might be attributed to the differences in TL-LET response curves of the two phosphors or by (n, y) reactions in lead. It will be difficult to compare ku values reported in the literature if containers of different dimensions and atomic composition are used for the encapsulation of samples.

Schuhmacher and Krauss (1978) used their TL dosimeter with different shields to separate the neutron and photon components of a d(10.6) + D beam. They found good agreement (within 1 %) between total absorbed dose values measured free in air with their paired TLD system and those measured with a tissue-equivalent ionization chamber.

Busuoli et al. (1970) determined the two radiation components in a mixed n-y beam with one TL dosimeter by measuring the ratio of the two


glow peaks in LiF, which is a function of the type of incident radiation. The same one-detector TLD method has recently been applied to CaF2:Tm, which yielded the separate dose components of the d(14) + Be beam at Essen in air as well as in a phantom (Hoffmann et al., 1980). Single TL dosimeters can also be employed to measure the neutron absorbed dose in a mixed beam if use is made of the activation of the phosphor. Such a combination of activation and radiation detectors in one device has been described for CaS04:Dy by Bhatt et al. (1978). The TL contribution of beta rays from 32p, originating from the 32S(n, p) 32p reaction, was integrated after annealing the TL signal induced during the irradiation. The neutron energy dependence of such a system will follow the (n, p) reaction cross section.

Photographic emulsions can be used in neutron fields to measure the neutron absorbed dose, the photon absorbed dose, or to map neutron absorbed dose distributions. The latter application is of importance for radiotherapeutic applications of neutron beams and will therefore be discussed in Section 6. The absorbed dose resulting from (n, p) scattering in the emulsion is proportional to the number of recoil protons and their average energy. Development of the track of the recoil proton in the emulsion and counting the number and lengths of the tracks are wellestablished techniques. In principle, the technique is an absolute neutron dosimetry method if the hydrogen content of the emulsion is known. Owing to the laborious nature of the scanning process, only a limited number of applications have been reported, mainly for spectrometric purposes. Counting the number of tracks alone gives an accuracy that makes the system useful only for radiation protection purposes. The sensitivity of a photographic emulsion, expressed as an optical density, decreases with increasing LET. A low response, ku, must therefore be expected for neutrons. Measurements of the optical density can thus be used for the assessment of gamma rays in a neutron field. In practice, this can be achieved by using a film with a thin base enclosed in nonhydrogenous shields. A ku value of 1.9% for 14-MeV neutrons for a film covered with a 3-mm carbon plate has recently been reported by Hess et al. (1978). This confirms the earlier result of 2% for neutrons of this energy obtained by Bewley (1974) for a film enclosed in a 0.5-mm-thick lead screen. Such a lead filter also corrects for the energy dependence of the photon sensitivity, as discussed by Dudley (1966) and, more recently, by Scarpa et al. (1978). ku values varying between 1 and 10% have been presented by Dutreix et al. (1978) for different types of photographic emulsions in a d(50) + Be beam. It should be noted that these authors irradiated their films with their paper envelope only, without any additional screens. Their data should thus be considered as maximum values. The magnitudes of the ku values for screened photographic films seem to be comparable to those of small Geiger-Miiller (GM) counters. Owing to their convenience, GM counters, however, are usually


preferred for the determination of gamma-ray absorbed doses in neutron fields, as will be discussed in the next section.

3.6. Determination of the Gamma-Ray Absorbed Dose in a Neutron Field

It is necessary to determine the absorbed dose in tissue of neutrons and photons in a mixed field because of the differences in relative biological effectiveness between these two radiation components. An evaluation of the separate absorbed dose components can be made with a single instrument. In a tissue-equivalent proportional counter, the energy deposition events due to electrons can be separated from those resulting from protons, alpha particles, and heavy recoils with higher LET values. The microdosimetric technique of unfolding the energy deposition events will be discussed in Section 4. Using a tissue-equivalent ionization chamber, the components of the current induced by neutrons and photons can be separated by the method of stochastic current analysis introduced by Sherwin (1975). This method has proved to be successful in the measurement of the radiation components arising from a californium-252 neutron source. An uncertainty of about 4% was indicated for the ratio of the neutron and photon absorbed doses. The method is independent of photon energy but strongly dependent on neutron energy, thus requiring accurate neutron spectrometric data. The technique seems to be most useful for the determination of the neutron component in mixed fields composed mainly of gamma radiation.

Separation of pulses induced by neutrons and gamma rays in some detectors, e.g., organic scintillators, can also be achieved by using the difference in pulse rise time. This method gives good results for neutron energy spectrum determinations. Recently, information obtained by this method also became available on gamma-ray energy spectra in neutron fields (Schmidt et al., 1980). Although, in principle, scintillation devices can be used to assess absorbed doses from their responses and can even be employed to derive the neutron and gamma-ray absorbed doses separately, only a limited number of preliminary results have been reported. This is due to the complexity of the technique, which requires special procedures for calibration and the unfolding of the measured pulse height distribution.

Because the neutron and photon sensitivities of a TE chamber are almost the same, such a chamber measures, in first approximation, the total neutron plus photon absorbed dose. If a TE chamber is employed, a second detector with different sensitivities to neutrons and photons will be necessary to obtain the two components separately. Nonhydrogenous ionization chambers such as graphite-walled ones filled with CO2 gas and Mgor AI-walled chambers filled with Ar gas can be used for this purpose


because the relative neutron sensitivity will be reduced. Other dosimeters with reduced neutron sensitivities are certain types of thermoluminescent materials, photographic emulsions, and Geiger-Miiller (GM) counters.

According to ICRU Report 26 (ICRU, 1977), the quotient of the response of the dosimeter and its sensitivity to the gamma rays used for calibration, R~, is given by

(8)

where ku and hu are the ratios of the sensitivities of the dosimeter to the neutrons and photons in the mixed field, respectively, to its sensitivity to the gamma rays used for calibration. The subscript U refers to an instrument with a lower sensitivity to neutrons than to photons (a neutron-insensitive

1.0

TE/TE 0 0.9

U 0 -0

£ OJ > ~

J! ~

"j ~

:~ c ~ 0.4 c: g :> OJ c:

0.3

0.2

0.1 - Waterman et 01.( 1979 b) ---- Kuchnir et 01. ( 1975 )

o 10 20 30 40 50


Figure 14. Neutron sensitivities of C/C02, Mg/ Ar, and A1S0/TE chambers relative to the sensitivity to 60Co gamma rays as a function of neutron energy (Waterman et al., 1979b).


device). Since the publication of ICRU Report 26 (1977), a large number of new ku data have been reported. Figure 14 shows, as an example, the relative neutron sensitivities of C/C02 , Mg/ Ar, and A150/TE chambers, dependent on neutron energy, as measured by Waterman et al. (1979b) in comparison with the earlier results of Kuchnir et al. (1975). It can be seen from Figure 14 that the ku values increase with neutron energy, largely as a result of the neutron kerma ratio Kwau/ Kt• It should be noted that Mg/ Ar chambers and AI! Ar chambers have lower neutron sensitivities than do C/C02 chambers and they should therefore be preferred. The ku of a C/C02 chamber can be reduced by a factor of more than 10 by increasing the pressure inside the chamber (Zielczynski et al., 1978). The columnar recombination of ions in the gas will then be strongly increased, while the volume recombination will still be negligible.

The contribution to ionization from particles generated in the wall relative to the contribution from the gas varies with the size of the chamber and the neutron energy; ku values of nonhydrogenous detectors of different sizes and configurations are thus difficult to compare. This is illustrated in Figure 15 for a parallel-plate C/C02 ionization chamber (Makarewicz and Pszona, 1978). It should be emphasized, therefore, that any calculated or measured data are valid only for the specific geometry of the chamber used. Calculations of ku also have a rather large uncertainty, due mainly to the uncertainty in the mass stopping powers of heavy charged particles.

To minimize overall uncertainties in the absorbed doses of neutrons and photons due to the overall uncertainty in ku, it is advantageous to reduce ku, provided that the reduction is accompanied by less than a proportionate increase in the fractional overall uncertainty of ku (ICRU, 1977). Wagner and Hurst (1961) suggested the use of micro Geiger-Muller counters as photon dosimeters in neutron fields. A suitable lead/tin shield was placed around the counter in order to make its response to photons less energy dependent. Their measurements with monoenergetic neutrons showed that ku is less than 0.5% in the energy range from 0.68 to 4.2 MeV. Colvett (1974) determined experimentally that ku does not exceed 0.5% for 15-MeV neutrons. A number of recent measurements, however, resulted in higher values for d + T neutrons. Before discussing these results, the methods used to obtain the data will be considered.

Methods for deriving ku values are the lead filtration technique, the spectral difference method, the associated-particle coincidence technique, and the time-of-flight technique. ku can also be determined if the dosimeter is placed in a neutron field where the ratio of the photon absorbed dose to the neutron absorbed dose is known by means of microdosimetric methods (Ito, 1978a) or by deducing its response from measurements with a Ge(Li) photon spectrometer (Guldbakke et al., 1978). The lead filtration method, as applied by Attix et al. (1975) and Cleland and Wells (1975),

38

0 .6

0.5

0.4

0 .3

0.2

0 .1

• Greene (1974)

... Broerse (197 4)

o 0.2 em } Mokorewicz and

• 1 em Puona ( 1978 )


I

Kuehnir er 01. ( 1975 )

5 10 15

EN (MeV)

I 20

Figure 15. Relative neutron sensitivity of a parallel-plate CjC02 ionization chamber as a function of neutron energy for two spacing distances between the electrodes. The upper edge of the shaded area represents the values for a Bragg-Gray cavity, whereas the lower edge shows values for a large cavity (Makarewicz and Pszona, 1978).

has been modified in the method of data analysis used by Hough (1979). Care must be exercised in using the lead filtration method to assure that the background radiation is invariant under the filtered and unfiltered beam conditions. If this is impossible, then the background has to be determined separately for each beam condition (Waterman et al., 1977). The spectral difference method was developed by Kuchnir et al. (1975) and is based on measurements made over an angular range in neutron beams having isotropic gamma components and nonisotropic neutron components. This method has recently been applied to nonhydrogenous ionization chambers (Waterman et a/., 1979b) and GM counters (Lewis and Hunt, 1978). The associated-particle coincidence technique developed by Lewis and Young (1977) utilizes coincident pulses from a GM counter and an alpha detector measuring the alpha particle emitted simultaneously with the neutron from the 3H(d, n) 4He reaction. With this method measurements have to be


performed with the axis of the GM counter parallel to the incident neutron flux. As shown by Lewis and Young, ku values corresponding to this orientation may differ from those obtained from measurements with GM counters with their axes perpendicular to the neutron beam. The time-offlight (TOF) technique is a well-known method in neutron physics and is described elsewhere (e.g., in ICRU Report 13, 1969a). TOF methods are usually applied to pulsed neutron fields. They have been employed for the measurement of ku of GM counters in a pulsed deuteron beam of a Van de Graaff generator producing d + D and d + T neutrons (Klein et at., 1979) and in the pulsed 14-MeV deuteron beam of a cyclotron producing d + Be neutrons (Ito, 1980). The main uncertainty in this method is the contribution from non-time-correlated events due to room-scattered neutrons and photons.

ku values for GM counters of different shield designs at different neutron energies were recently summarized (Mijnheer et at., 1979a). The most commonly employed GM counters are the Philips type 18529 (also known as the Mullard type ZP 1300 or MX163) surrounded with a 1.1-mm tin/0.5-mm lead shield and the Philips type ZP 1100 surrounded with a 2-mm perforated tin shield. A comparison of ku values for GM counters in d + T neutron fields showed that ku values of 1.6 and 2.6% appear to be most realistic for the 18529 with a Pb/Sn shield and the ZP 1100 GM counter, respectively. It can be concluded that ku increases with neutron energies above 2 Me V and that, especially for the higher energies, it depends on the shield design and probably on the counter type. The neutron energy dependence is illustrated in Figure 16 (taken from Guldbakke et al., 1980), which shows some of the available data for the ZP 1100 and MX 163 counters. In addition to available data at neutron energies higher than about 2 MeV, Guldbakke et at. also determined ku values at lower energies using the TOF technique. Their data show that ku increases again with decreasing neutron energies below about 0.5 MeV. The influence of the shield design on ku has been studied by Zoetelief et at. (1978). Their measurements showed an almost linear increase in ku with increasing weights of the shield.

Although GM counters are usually surrounded by an energycompensated shield, their sensitivities still vary with photon energy, as discussed in ICRU Report 26. Because information on the photon spectrum in the neutron field is usually not available, the relative photon sensitivity hu of the GM counter is therefore taken as equal to unity. This may introduce considerable uncertainty in the determination of the photon absorbed dose in a neutron beam by means of a GM counter. GM tubes with their energy-compensating filters have a high thermal neutron sensitivity. A ku value of about 9 can be derived from the experiments of Wagner and Hurst, while a more recent determination of the thermal neutron

40

• ZP 1100 (PTS - data) o ZP 1100 ( NPL - data) I) ZP 1100 ( Mijnheer et 01. )

• MX 163/ PTFE ( PTS - data) o MX 163/ PTFE ( NPL - data) II MX 163/ PTFE ( Hough)

iii I

10- 1 i i "I

100


, • i' i I

101

EN ( MeV)

Figure 16. Relative neutron sensitivity as a function of neutron energy for different GM counters as measured by different groups (Guldbakke et al .. 1980\.

sensitivity (Lewis and Young, 1977) indicates a value of about 5. GM counters should therefore be shielded by a thermal neutron absorber that does not emit prompt gamma radiation in the neutron capture process. A commonly chosen material is 6Li in the form of the metal or 6LiF powder. A disadvantage of GM counters is their high sensitivity. Owing to their dead time of approximately 20 p.,s, accurate results can be obtained only in fields having photon absorbed dose rates less than about 0.2 rad min-1 (cGy min -1).

3.7. Determination of the Neutron Absorbed Dose in a Photon Field

High-energy photon beams obtained from electron accelerators for use in radiotherapy are usually accompanied by neutrons resulting from photon interactions in the target, shield, collimator, and flattening filter. For the same reasons as mentioned in the preceding section, separate knowledge of both components of the beam will be necessary. The magnitude of the contamination with neutrons in photon fields, however, is different from that with photons in neutron fields. In the latter situation, the ratio of the gamma-ray absorbed dose to the neutron absorbed dose

Progress in Neutron Dosimetry for Biomedical App6cations 41

measured in air amounts to several percent and increases to 10 to 20% with increasing depth in a phantom or a patient (see Section 6). As can be seen from Figure 17, the relative neutron absorbed dose in photon fields measured in air is much smaller and varies with photon energy between 10 and 45 MeV (Almond, 1979). Because the neutrons are attenuated more rapidly than the high-energy photons, the relative dose contribution of neutrons will be even lower at greater depths in a patient. The problem is therefore less important with regard to the accurate determination of the dose for treatment of a tumor, but is more important in the determination of the total body dose to the patient or the dose equivalent delivered to operators of the machine. The latter type of measurement does not require the high accuracy necessary for the determination of the absorbed dose in the beam. The problems with regard to the physical measurements

10- 1

• •

~ • " ~ " Q.

e: 0

" " 0

0 ..s: Q.

"-., 10-2 0

" " ~ " ., "

10 20 30 40

occelerotor energy ( MeV)

Figure 17. Relative neutron absorbed dose in x-ray beams of medical accelerators of various energies. The data points are measured values, the solid line shows results from calculations (Almond,1979).


and calculations of neutrons from electron medical accelerators as well as the biological interpretation of these data have been recently discussed (NBS, 1979).

Inside the photon beam, the neutrons will have a neutron energy distribution that resembles a fission spectrum with a higher energetic component. Owing to scattering, a lower energetic component will also be present. This component will contribute only a few percent to the total neutron absorbed dose inside the beam but will become larger outside the beam. The sensitivity of the detector should therefore be determined for this neutron spectrum. The detector is usually calibrated in a pure fission field eS2Cf source) and this might cause an uncertainty in the calibration factor.

The types of detectors used for these measurements can be divided into four groups: moderating spheres, silicon diodes, track detectors, and activation detectors. Moderating detectors consist of a hydrogenous moderator placed around a thermal neutron detector, e.g., an indium or gold activation detector or 6LiF tLiF thermo luminescent dosimeter. A set of spheres with different diameters can be used as a spectrometer with low-energy resolution. The characteristics of these types of detectors have been summarized elsewhere (Nachtigall and Burger, 1972); those of silicon diodes, track detectors, and activation detectors have been described here in preceding sections.

The main difficulty in the use of neutron dosimeters in photon fields is their relative photon sensitivity. All detectors will, to a certain extent, be sensitive to photons due to photoneutron production in the dosimeter itself. Photo neutron production in the carbon of the moderating sphere will enhance its response by about 25-30% for a 33-MeV photon beam (Axton and Bardell, 1979). For 16-MeV photon beams, photoneutron production in the moderator can be neglected. Silicon detectors are also sensitive to photons. Calculations and measurements by McCall et al. (1978) suggest that the apparent neutron doses measured with silicon diodes were actually due to electrons and photons, rather than to neutron interactions in the diodes. The response of fission track detectors will also be overestimated owing to photo fission induction, which has a threshold energy around 5 MeV. The photon sensitivity of polycarbonate foils caused by track registration through recoils originating from (-y, n) and (-y, p) reactions is probably very low, but further work is needed (McGinley and Sohrabi, 1979). (n, p) activation detectors have the advantage of being almost insensitive to any gamma field, since the product nuclei cannot be formed by direct photonuclear reactions in the detector material (ICRU, 1969a). Only two-step processes such as (-y, n), (-y, pn), or (-y, 2n) reactions followed by fast neutron activation are possible interfering reactions if the (n, p) reaction is applied for fast neutron measurements in a photon beam. The


use of the reactions 31P(n, p) 31Si (Price et aI., 1978) and 27 Al(n, p) 27Mg (Gur et al., 1978) has been reported. The contribution of 31Si produced by photons, relative to that produced by neutrons through the 31P(n, p) 31Si reaction, is estimated to be less than 4% in a 35-MeV photon beam (Nath et al., 1979).

It can be concluded that activation detectors employing the (n, p) reaction can be applied at all photon energies in use in medical accelerators. Moderating detectors can also give accurate results but should be corrected for photon sensitivity in photon beams of 20 MeV or higher. The use of silicon diodes to determine the neutron absorbed dose in photon beams is not recommended, owing to their response to high-energy photons. The photon sensitivity of track detectors should be further investigated before these detectors are employed in high-energy photon beams.

4. ASSESSMENT OF RADIATION QUALITY

In the preceding section, a number of methods were described that allow the separate determination of the neutron and photon absorbed doses. This can be considered as a first step in the assessment of radiation quality, which, in general, should be related to the neutron energy spectra, linear energy transfer (LET) spectra, or lineal energy (y) spectra. If the radiation quality does not show considerable variations with position in a phantom, then significant variations in relative biological effectiveness (RBE) in the irradiated body are not to be expected. Large variations in energy deposition spectra will indicate the need for radiobiological experiments to check the changes in effectiveness of the field at different positions.

Since neutrons dissipate their energy through different nuclear interactions with the various constituents of matter (the dosimeter material or the biological tissue), the energy deposition is characterized by a complex spectrum of secondary charged particles. The microscopic distributions of energy deposition can be described by the mean energy loss of the charged particles per distance traversed, the restricted linear energy transfer La (ICRU, 1977). However, a serious limitation of the linear energy transfer (LET) concept is that the LET represents only a mean value that does not account for the statistical fluctuations in the interaction of charged particles with matter. The distribution of energy deposition can be determined experimentally with proportional counters under the assumption that the ion yield is independent of energy and particle type. It is impossible, however, to derive accurate LET spectra from these measurements owing to the radial extension of the tracks, the finite range of the recoil nuclei, and energy losses due to straggling.


4.1. Neutron and Gamma-Ray Energy Spectra

The information on neutron energy spectra summarized in Section 2, was in general valid for free-in-air conditions in an open geometry without beam-defining material. Interactions of the primary neutron beam with the collimator material and inside the biological specimen or the patient will result in scattered neutrons of lower energy and photon production.

In general, experimental and theoretical data on neutron spectrum changes with position in the phantom are only available for neutron beams with energies up to 15 MeV. Measurements with an organic scintillator performed by Hannan et al. (1973) for collimated d + T neutron beams indicate that the relative contribution to the dose from neutrons with energies above 12 MeV decreases from 73 to 63% if the depth increases from 2 to 17 cm. Neutron energy spectra determined by Bonnett and Parnell (1976) for d(16) + Be neutrons and those determined by Mijnheer


et al. (1981a), Schmidt et al. (1980) and Nguyen et al. (1981) for d + T neutrons confirm the buildup of low-energy neutrons with increasing depth in the phantom along the central axis of the beam (see Figures 18 and 19). Other published measured and calculated d + T spectra (e.g., lng and Cross, 1975, and Burger and Morhart, 1978) also show the presence of a lowenergy component, although the relative contributions of neutrons in this energy range vary among the different groups of investigators. This is due to differences in experimental technique, field size, collimator design, and the material composition of the collimator and phantom. The relative contribution of the scattered component increases with field size. Neutron spectra in the penumbra of d + T neutron beams measured in a water

>- 140 D> l; I:

" I: 120 :> .. " Q.

" U I:

" :> 100

;;:

80

neu Iron energy ( MeV)

Figure 19. Energy spectra of d(16) + Be neutrons measured at different depths along the central axis in a water phantom (Bonnett and Parnell, 1976).


phantom also show an increase of the low-energy component relative to the value on the central axis of the beam.

Photons in a neutron field are produced in the neutron source by interactions of the neutrons with the collimator and other shielding material and by reactions of the neutrons in the phantom or biological object. The photon spectrum will thus generally be different for different facilities. Measurement of photon spectra in a neutron beam is complicated by the sensitivity of the photon detector to neutrons. Discrimination of the two signals from both components of the beam is possible by using a time-offlight (TOF) technique or by using the differences in pulse shape in scintillators. Cross sections for gamma-ray production by fast neutrons are available (e.g., Dickens etal., 1977). It will be difficult, however, to calculate photon spectra in extended media using these cross sections because the primary gamma rays will interact with that medium.

>.

'"

1.4

1.2

1.0

O.B

~ 0.6

depth in phantom

0.5 em ., -I: :> .. ., 0.. ., U I: ., :>

;;:

0.4

0.2

o +----,----r

0.41 0.2

O+----r---T'"

5.2 em

0.4] 0.2

o +----,----T'"

20.1 em

0.4 30.0 em

0.2

O+-----r----~----,---~~~~--r_~--.__ o 2 3 4 5 6

photon energy ( MeV)

Figure 20. Photon energy spectra measured at different depths in a TE phantom irradiated with d + T neutrons (Schmidt et al., 1980).


Elson et al. (1979) measured the gamma-ray spectra with a liquid scintillator at various distances from a 252Cf source in a tissue-equivalent medium. Schmidt et al. (1980) determined photon spectra with a stilbene crystal on the central axis of a collimated d + T beam at different depths in a phantom. As shown in Figure 20, the spectra have a main peak at 2.2 MeV due to thermal neutron capture in hydrogen and smaller peaks at higher energies due to interactions with carbon and oxygen. In addition, there is a large component of lower-energy photons resulting from Compton processes in the phantom. An average energy of about 2 Me V can be deduced for these spectra in the phantom. This is in agreement with the mean photon energy derived from a backscattering factor in lead (Schuhmacher and Krauss, 1978; Mijnheer, 1977) or from microdosimetric measurements (Menzel et ai., 1978).

Transport calculations of gamma-ray flux densities around a 252Cf source (Shapiro et at., 1976a) and in a phantom irradiated with a collimated d + T beam (Burger and Morhart, 1978) are in general agreement with the measured spectra. Small differences are to be expected because the gamma-ray spectrum from the neutron source, which is not very well known, has to be considered in these calculations.

4.2. Microdosimetric Parameters and Techniques

Recognizing that the measured spectra of energy deposition are more directly related to the biological effects of ionizing radiation than are the LET distributions has led to the introduction of microdosimetric quantities that are stochastic and concerned with energy deposition events in specific regions (Rossi, 1968). The stochastic quantity lineal energy y is the quotient of energy imparted to the matter in a volume during an energy deposition event and the mean chord length in the volume of interest (ICRU, 1977).

The radiation quality of a neutron beam can be characterized by its lineal energy spectrum; however, it would be advantageous to introduce an average value as a single parameter. Average values of the lineal energy y can be derived from the distribution of the frequency of events in y or from the distribution of absorbed dose in y. The mean of the frequency distribution of the lineal energy is

YF = LX) yf(y) dy (9)

where f(y) is the differential distribution of y. The mean of the absorbed dose distribution of the lineal energy is

YD = f" yd(y)dy (10)


where dey) is the differential distribution of the absorbed dose in y. The frequency-average lineal energy YF corresponds to the track-average linear energy transfer, while the dose-average lineal energy YD corresponds to the dose-average linear energy transfer. Previous studies (Broerse et at., 1968) have shown that the track-average LET values for neutrons are considerably lower than those of directly ionizing particles, which would have an identical relative biological effectiveness (RBE) for cell reproductive death. On the other hand, the dose-average lineal energy values YD are unsatisfactory in that the YD for neutron energies above 10 MeV is higher than that for neutron energies below lOMe V; this is incompatible with the findings of higher RBE values for the lower-energy neutrons. It should be realized that a given y event becomes less effective in producing a biological effect per unit dose when y becomes too high, since some of the energy in such a track of very dense ionization is wasted. To correct for this phenomenon, Kellerer and Rossi (1972) introduced the saturationcorrected dose-average lineal energy:

00 2

J. Yo 2/ 2 y*= -(l-e-Y YO)d(y)dy o y

(11)

where the saturation parameter Yo is 125 keV / ILm. Values for YD and y* as calculated by Caswell and Coyne (1978) for different neutron energies are shown in Figure 21.

During a meeting on the fundamental and practical aspects of fast neutron radiotherapy (Barendsen et at., 1979), the problem of a biologically and clinicaHy significant description of the radiation quality of fast neutrons

E :1.

'> " .>I. o •

o •

ii' i I

d = 1 I'm YO= 125 keV / I'm

o

•

o Yo • Y·

o 0 00 ....

i i " I

100

o

• o

o 0000

• • • •• • •••

i " i I

101


Figure 21. Comparison of YD and y* in tissue for different neutron energies (Caswell and Coyne, 1978).


was discussed. Concerning the usefulness of y *, two opposing opinions could be distinguished (Booz, 1979).

One group criticized y* because it did not take into account all the radiobiological mechanisms and gave the following arguments:

1. y * does not consider radiation mechanisms at the molecular level and therefore does not show the correct radiation quality dependence.

2. The saturation correction was said to be incorrect because it is based on the results of track-segment experiments, which reflect a much smaller scale.

3. Different biological end points show different RBE values for the same radiation, whereas y * is constant.

4. Microdosimetric spectra in general and y* in particular cannot be used in a predictive way, but only a posteriori, to explain unexpected biological results.

The speakers of the second group expressed a more practical opinion. While they appreciated a good part of the arguments on the fundamental incompleteness of y*, they nevertheless considered y* to be an empirical parameter of practical importance for the needs of clinical radiation therapy. With regard to this practical application, they said the following:

1. y* is a physical quantity of the radiation and must not be misunderstood as a measure of biological effectiveness for particular end points.

2. y* can be used to predict relative changes in RBE in the phantom but cannot predict absolute RBE values.

3. Within the range of neutron energies mainly used in neutron therapy at present, i.e., between 6 and 15 MeV, y* is roughly proportional to the initial slope of the survival curves.

4. Consequently, y* is a fair single-parameter description of radiation quality for the purpose of neutron therapy and can be used to estimate the normal tissue damage produced by different radiations.

5. y* is based on the dual radiation action hypothesis (Kellerer and Rossi, 1972) and should therefore correctly describe the dose dependence of the RBE for those biological end points that follow this theory.

The probability distributions of yare determined experimentally with proportional counters of spherical or cylindrical shape that are constructed out of tissue-equivalent plastic and filled with tissue-equivalent gas. TE proportional counters are generally operated at a pressure of 0.1 atm

so J. J. Broerse and B. J. Mijnheer

60Co y rays 14.7-MeV neutrons

, , ''''I , "'''I i' , .... 'Ii "ui • " "Ii

10- 1 10° 102

y ( keY / flm )

Figure 22. Lineal energy spectra for various radiation qualities for a l-~m sphere of tissue (Kellerer and Rossi, 1972).

(104Pa); this corresponds to a sensitive spherical volume of tissue with a diameter of 1 ~m for a counter with a diameter of 1 cm. As shown in Figure 22, the lineal energy spectra for different types of radiation extend from y values below 0.1 up to 1000 ke V I ~m for simulated regions with 1-~m diameters. There have been discussions on the relevancy of these relatively large diameters for the interpretation of radiobiological data. Some authors claim that for these interpretations y spectra for diameters as small as 10 nm should become available (e.g., see Hogeweg, 1978).

Discriminating between the low and high components of the y spectra of a mixed neutron-gamma beam can result in determining the gammaray contribution to the total absorbed dose (Bichsel, 1975; Maier et at., 1975; Leroux et at., 1978; Menzel and Schuhmacher, 1980). Special unfolding procedures can be applied in which a pure gamma-ray spectrum is fitted to the mixed field spectrum. After normalizing both spectra in the interval from 0.2 to 0.4 keY I ~m, the gamma dose fraction can be derived from the areas under both curves, as indicated in Figure 23. The gamma-ray and neutron pulses can be separated with a better resolution by using proportional counters with graphite walls and helium and carbon dioxide as counting gas, as suggested by Caswell (1960) and recently applied by August et at. (1978). A disadvantage of the proportional counter is its high sensitivity. Consequently, the device can be used only at low dose rates. Determining the absorbed dose from the event size spectra is a rather elaborate procedure, not suitable for routine measurements. Other sources of error in the microdosimetric method are the nonlinear relationship between the ion yield and the lineal energy and the matching of the spectra being measured at different gain settings to cover the large range of event sizes (Stinchcomb et at., 1980). Using the spectra derived from TE proportional counters, the ratio of the photon absorbed dose to the neutron absorbed dose, DGI DN , can be determined with an overall uncertainty of


" >-

0.4

0.3

0.2

0.1

c yc lotron-; produced .: neu trons :

;

.-....... ,

"-... \.

" .... ": ..

102

y ( keY / flm )

51

Figure 23. Procedure of separating events due to neutrons and photons in an energy deposition spectrum produced by d(10.6) + D neutrons (Menzel and Schuhmacher, 1980).

about ±10%, provided that systematic and statistical uncertainties are kept to a minimum. The proportional counter technique can thus be used as a reference method to measure the contribution of photons to the total absorbed dose.

4.3. Lineal Energy Spectra in the Phantom

The energy deposition by fast neutrons in microscopically small volumes has been determined by radiation transport calculations (Burger et al., 1978a), by means of analytical approaches (Edwards and Dennis, 1975; Caswell and Coyne, 1978), by Monte Carlo techniques (Coppola and Booz, 1975), and experimentally with proportional counters (e.g., Heintz et al., 1971; Hogeweg, 1978; Fidorra and Booz, 1978; Menzel et al., 1978). To avoid spectrum distortion due to pulse pile·up, the microdosimetric measurements had to be performed at considerably reduced dose rates in comparison to those used for radiobiological and clinical applications.

Microdosimetric measurements performed for different collimated neutron beams have shown that the lineal energy spectra do not change considerably with depth in the center of the beam (e.g., Oliver et al., 1975, for neutrons produced by 16-, 30-, and 50-MeV deuterons on beryllium; Harrison et al., 1978 for d(80) + Be neutrons) but that appreciable variations occur with lateral distance from the beam axis outside the primary field. Spectra measured with TE proportional counters by Burger et al. (1978a) and Hogeweg et al. (1978) for d + T neutrons (see Figure 24)


0.4

0.3

0.2

0.1 ._._._._.-?t

o •• , II ..

y (MeV em-I)

Figure 24. Fractional dose distribution of is-MeV neutrons at the central position in the beam at the surface (curve 1), at 7-cm depth (curve 2), and behind the shielding outside the primary beam (curve 3), as measured by Hogeweg et al. (1978).

clearly demonstrate the increasing contribution of lower-energy scattered neutrons and gamma rays behind the collimator. These observations are in agreement with neutron spectrum measurements and gamma dose measurements as described in Sections 4.1 and 6.2, respectively. In accordance with the microdosimetric findings, studies on cell survival performed by Hogeweg et at. (1978) at different depths along the beam axis in a water phantom (see Figure 25) indicate that the biological effectivity does not change with depth. Similar results were reported by Mijnheer et at. (1979b) for d + T neutrons and Menzel et at. (1978) for d + D neutrons. For positions outside the main beam, different biological results have been observed: Hogeweg et at. and Mijnheer et at. did not observe a significant change in RBE at lateral distances from the beam edge, while Menzel et at. observed a slight decrease in biological effectiveness at greater lateral distances. Measurements performed with a biological dosimeter by Bewley et at. (1978) showed that the effective dose as derived from cell survival is 30% greater than that derived from the physically measured DN and Da (see Section 6.2).

Fidorra and Booz (1978) measured energy deposition spectra with a commercial proportional counter for a collimated d(14) + Be neutron beam and derived both the frequency-average, YF, and dose-average, YD, lineal energies. They observed an increase in YF and YD for neutrons with both depth and lateral distance from the beam axis. However, YF and YD for the total energy dissipation decreased with increasing depth and lateral distance. Menzel et at. distinguished four different y intervals of different biological effectiveness correlated with energy deposition by

Progress in Neutron Dosimetry for Biomedical Applications S3

... , "

data points : . " • a t a dep th of 4 em

" .. 0 7.3 em c: .\ " • 10 em .~ " '" 13.8 em v - '" 17 ~

* " em

'" '" c:

.~ 10- 1 ":~ ,

\ ~ , ;l &l \ ,- \

\ ~ \

\ \ \ \ \ 300-kV X - rays

\ \

10-2 \ \ \

- \ \ \

\ \ \ \ \ \ \

total ( n + y ) absorbed dose ( Gy )

Figure 25. Survival of cultured cells at different depths in a phantom irradiated with IS-MeV neutrons along the central axis of the beam (Hogeweg et aI., 1978).

gamma rays, by protons with energies above and below 1 MeV, and by alpha particles and heavy recoils. In their experiments both YF and YD for the total dose decrease with lateral distance across the neutron beam; however, the saturation-corrected dose-average lineal energy y* remains nearly constant across the entire profile. These findings indicate the need for more correlated microdosimetric and radiobiological experiments. These would lead to improvements in the interpretation of energy deposition spectra and would allow one to decide on the appropriateness of the different parameters for the specification of radiation quality.

It can generally be concluded that microdosimetry provides a suitable basis for radiation quality specification for therapy purposes but that a pragmatic approach has to be adopted. It would be dangerous in clinical applications to assume that the biological effectiveness of two different neutron beams are directly comparable because their microdosimetric averages such as yp, YD, or y* are the same. Even if their event size


spectra were the same, this would not necessarily indicate that they would produce identical biological effects, since the y spectrum is largely dependent on the diameter of the sensitive site adopted. It would be even more presumptuous to assume that it would be possible, by any microdosimetric measurements or calculations, to predict the absolute biological effectiveness of any fast neutron beam employed for radiotherapy. Possible differences in the radiation quality of different fast neutron beams can be assessed by microdosimetric techniques but should be supported by comparing the responses of biological dosimeters in the different beams.

5. DOSIMETRY FOR RADIOBIOLOGY

Studies of the biological effects of fast neutron beams have been performed on a large number of biological species with different objectives. Investigations of the effectiveness of autologous bone marrow transplantations in lethally irradiated rhesus monkeys (Broerse et al., 1978b) are relevant to the treatment of victims of reactor accidents. Irradiations of Drosophila melanogaster (Goodman et al., 1972) have been performed with the aim of studying mutagenetic effects of high LET radiations. Investigations of the reproductive capacity of cells cultured in vitro and the responses of tumors and normal tissues in experimental animals have been made for the purpose of assessing the usefulness of neutrons in the treatment of cancer. Some of these irradiations must be partial body irradiations to avoid possible interfering effects due to damage to organs other than those of interest.

For all of this radiobiological research, the dose should be administered to the biological specimen in a reproducible way and the dose distribution in the specimen should be adequately described. The specification of the dose by a single set of parameters will not present a significant problem if the dose distribution is uniform. According to ICRU Report 30 (1979a), this means that the maximum to minimum dose within the object does not exceed 1.10. In a number of specific cases, e.g., the irradiation of relatively large biological objects with low-energy neutrons, it will be physically impossible to fulfill the requirements of uniform dose distribution. An appropriate choice of the irradiation geometry, however, can increase the uniformity of energy deposition in the specimen. Reviews of irradiation arrangements suitable for radiobiological experiments with fast neutrons have been published (ICRU Reports 26, 1977, and 30, 1979a).

Regarding the possible variation in the output of radiation sources with time, it should be realized that even for radioisotopes and orthovoltage x-ray machines, monitoring can serve to detect such faults as an improperly positioned source, a stuck shutter, improper filters or attenuators, and other

Progress in Neutron Dosimetry for Biomedical Applications ss

more or less subtle mechanical and electrical faults. Monitoring is indispensable when neutrons are generated at the target of an accelerator, since these complex machines and their appurtenances cannot be relied upon to operate with sufficient stability for protracted times. A discussion of monitor responses used to normalize individual irradiations and examples of monitoring arrangements can be found elsewhere (ICRU, 1977).

The energy deposited in the irradiated material by different interaction processes such as elastic and inelastic scattering and nuclear reactions depends on the nature and mass of the nuclei involved. Consequently the atomic composition of the medium, especially the hydrogen content, will determine the amount of energy transferred. It should be realized that the composition assumed for tissue, namely, that of striated muscle (ICRU, 1964), is not representative for all tissues. Information on the atomic composition of various human tissues such as bone, kidney, liver, and muscle, and of some biological specimens can be found in ICRP Publication 23 (1975) and ICRU Report 26 (1977). There are substantial differences in atomic compositions; e.g., some plant materials have a considerably lower hydrogen content than animal tissues. For the energy deposition in different tissues irradiated by fast neutron beams of different energies, Bewley (1980) calculated the relative kerma values shown in Table 6.

Table 6. Kerma Values in Different Materials Relative to That in ICRU Muscle a

Neutron spectra

% H,by Tissue p(66) + Be d+T d(16) + Be weight

25 b 14.5b 7.5 b

ICRU muscle 100.0 100.0 100.0 10.2 AlSO plastic 109.0 ± 10.6 104.2 ± 4.7 102.7 ± 2.7 10.15 Water 104.4 ± 1.8 105.6 ± 0.8 107.2 ± 0.4 11.2 Perspex (Lucite) 93.1 ± 7.1 88.4 ± 3.0 84.1 ± 1.7 8.05 ICRU bone 76.4 ± 4.5 73.2 ± 1.6 71.1 ± 1.2 6.4 Cortical bone 63.2 ± 3.2 58.3 ± 1.0 57.3 ± 1.1 4.7 Subcutaneous fatty tissue 116.4 ± 8.9 113.8 ± 5.4 114.3 ± 2.1 11.7 Brain 105.3 ± 0.4 106.1 ± 0.1 107.0 ± 0.1 11.1 Cartilage 100.3 ± 0.4 100.5 ± 0.2 100.6 ± 0.1 10.3 Connective tissue 96.1 ± 1.0 95.3 ± 0.3 94.0 ± 0.2 9.4 Kidney 100.1 ± 0.1 100.0 ± 0.1 100.0 ± 0.1 10.2 Liver 102.0 ± 0.5 102.3 ± 1.5 102.5 ± 0.2 10.5 Skin (dermis and epidermis) 100.8 ± 1.7 100.3 ± 0.6 99.8 ± 0.3 10.1

"From D. K. Bewley, in Ion Chambers for Neutron Dosimetry, J. J. Broerse, ed., EUR-6782, Harwood Academic Publishers, London (1980).

"Mean neutron energy in MeV.


The energy transfer by fast neutrons at any given point occurs through secondary particles, e.g., recoil protons and heavy nuclei originating at the given point up to a distance from the point equal to the maximum range of the secondary charged particles present. Clearly this process can result in an absorbed dose in the surface layers different from that in the deeper portions of a specimen. A lack of charged-particle equilibrium occurs at any interface between tissue and an adjacent material with a significantly different atomic composition. The resulting changes in dose distribution arising at the interface of materials of different composition will be treated in some detail.

The dose distribution over the biological specimen will depend on a number of geometrical factors such as the distance between the object and the radiation source, the relative size of the object, and absorption and scattering processes in the specimen itself as well as scattering processes from adjacent materials, including supporting structures. In the case of neutrons emitted by accelerator targets, spatial non uniformities in the kerma distributions free in air can be compensated for by slow rotation of the biological objects around or in front of the target. Conditions of constant scattering can also be achieved by performing the irradiation of mammals with separate containers for each individual animal. Specific precautions have to be taken in partial body irradiation of rodents with collimated beams.

5.1. Interface Dosimetry

Experimental conditions are usually designed to provide secondarycharged-particle equilibrium for the specimen and the dosimeter to facilitate accurate measurements of neutron kerma or absorbed dose. This means that the sensitive volume of the dosimeter should be surrounded by a layer of material of thickness equal to the maximum range of the secondary charged particles produced. For kerma determinations, the thickness of this equilibrium layer should not be much greater than that required to minimize corrections for attenuation. The situation of incomplete chargedparticle equilibrium most commonly encountered in practice is the exposure of biological specimens in free air. The number of secondary particles generated in the air, such as electrons in the case of x or gamma rays, and protons and heavier particles in the case of neutrons, is usually insufficient to provide charged-particle equilibrium in the superficial layers of the biological object. As a consequence, the absorbed dose at the surface of the specimen increases to a maximum value only at some depth within the specimen. This process, commonly referred to as buildup, can be of value in radiotherapy by producing a significant skin-sparing effect.


Even for fast neutron beams of closely comparable characteristics, different values for the entrance dose at the air-soft tissue interface have been reported. If the charged-particle buildup characteristics of different neutron beams have to be compared, it will be essential to use one set of detectors under similar experimental conditions. Zoetelief et al. (1978) employed a disk-type chamber to measure the buildup in ionization for different d + T neutron beams, with the chamber placed free in air or against a phantom. Their results are given in Figure 26. A relatively higher entrance dose is measured if the chamber is placed against the phantom. This situation is most relevant for the assessment of skin sparing for clinical applications. The increase in entrance dose is most probably due to backscattered radiation from the phantom. Also given are the results obtained with a lead filter (thickness 0.5 mm) placed in front of the chamber under similar conditions. These measurements show that the introduction of the lead filter has the most notable effect on the results obtained for the collimated beam. The decrease in the relative response of the chamber employing a lead filter can be attributed to the absorption of charged particles contaminating the beam. Similar measurements have been reported by Otte et al. (1977), while recently August et al. (1980) used tantalum foils to determine the charged-particle contamination of a d(35) + Be beam. Additional measurements by Zoetelief et al. with the disk-type chamber flushed with air resulted in lower values than those obtained for the chamber

100

90

c .. u Q; ~ 80 c .~ <; N

C 70 .~ .. >

0

~ 60

50

a 100

~ ••••• 6 free in air

.--. in front of phantom

0_0 collimated in front of phantom

)(--x in front of phantom with lead filter

(collimated ond not collimoted)

• i i , • i

200 300

build up thickness (mg cm -2)

Figure 26. Buildup for d + T neutrons under various experimental conditions measured with a disk-type TE ionization chamber (Zoetelief et af., 1978).


flushed with TE gas. It can be concluded that the measuring conditions, i.e., application of a nonhydrogenous filter, employment of a phantom, and the use of air or TE gas flushing the chamber, are of importance for the interpretation of charged-particle buildup curves.

Calculations of dose distributions across air-tissue interfaces have been performed by Bhatia and Nagarajan (1977) for 14-MeV neutrons and Chen and Chilton (1979) for neutron energies varying between 0.1 and 20 MeV. A comparison of the calculated buildup curves with the experimentally obtained results for 14-MeV neutrons shows that the curves exhibit similar trends, although the experimental data appear to be higher. Th,is may be due to the presence of lower-energy neutrons from scattering in the collimator and phantom that is not taken into account in the calculations. Bhatia and Nagarajan calculatrd, in addition, that the contribution of alpha particles generated in air to the skin dose amounts to about 5 to 10%.

Employing the same disk-type TE ionization chamber placed against a water or polyethylene phantom and with the collimator opening or the ion chamber covered with a lead layer, Mijnheer et al. (1978) compared the charged-particle equilibrium conditions for three neutron beams, d + T, d(50) + Be, and p(42) + Be with a nylon filter inserted in the collimator opening. The buildup curves, shown in Figure 27, indicate that the relative ionization at the entrance is approximately equal for comparable field sizes

~

c: .. ~ .. <>.

c: .~ 0 N

c: .~ .. .~ l! ~

100 ~O~"""t'~""""""""'"'''''' t ........ -0_ .-•• - ••• _ •••••••

. --•• - +

/'" 90 •

tf +--+ P ( 42 ) + Be; 15 em x 15 em 80

.v 0-0 d ( 0,25) + T ; 16 em x 20 em

............. d ( 50 ) + Be ; 20 em x 20 em

70

60 +

~

j

o i

500 i

1000

+

-0

1500

depth in unit density tissue (mg em -2)

Figure 27. Charged particle buildup for different neutron beams (Mijnheer et al., 1978).

Progress In Neutron Dosimetry for Biomedical AppHcations S9

for all three neutron beams. This observation is in agreement with measurements by Grant et al. (1978) for similar types of neutron sources. The depths at which the maxima are reached are in good agreement with the maximum ranges of the recoil protons (ICRU, 1977).

For the evaluation of skin damage in humans it must be realized that the critical layer of the epidermis is located at a depth between 15 and 100 #£m for most parts of the body and that it is covered by a horny layer 15 #£m thick (ICRP, 1975). Over this depth of 15 to 100 #£m, the relative doses are nearly the same for these three neutron beams.

If part of the field is covered with TE material during irradiation of the patient, one should expect a more severe skin reaction at this position than in the rest of the field, where the entrance dose will be lower. Clinical observations have shown, however, that there is no difference or only a slight difference in skin reactions between the covered and uncovered parts of the irradiated field (Battermann et al., 1981). Evidently, the increase in dose due to the addition of material causes only a minor difference in biological effect. This might be attributed to the fact that the increase in dose is due to low LET particles, fast protons and electrons having an RBE not much different from unity. The alpha particles, heavy recoils, and low-energy protons with ranges in tissues less than or comparable to 100 #£m have a much higher RBE (Broerse et al., 1968; Bewley et al., 1974b). This contribution of high-LET radiation with a relatively high RBE just below the surface will thus cause skin sparing with fast neutrons to be marginal and probably comparable to that in electron beam therapy, where the entrance dose amounts to 80-90% of the peak dose (ICRU, 1972).

Other examples of interface perturbations will be encountered during irradiations of body cavities and soft tissues adjacent or enclosed by bone (ICRU, 1977). Calculations of the variation of absorbed dose across the bone-tissue interface for 14-MeV neutrons by Bhatia and Nagaraj an (1977) show that the dose in the tissue at the interface is 12% lower than the equilibrium absorbed dose value. This value is in excellent agreement with that observed by Broerse and Barendsen (1968), who studied experimentally absorbed dose variations at the bone-tissue interface with cultured kidney cells as a biological dosimeter. It should be noted that at the equilibrium depth (-2.2 mm) the absorbed dose in bone is about 70% of that in soft tissue.

Perturbations of charged-particle equilibrium should always be considered in the irradiation of cells in monolayers on the bottom of culture dishes of different atomic composition. In orthovoltage x irradiations, excess secondary electrons will be produced in the glass of the culture flasks as compared with those produced in cells. If the total effect of all the interactions of fast neutrons with the cellular constituents has to be assessed, then a layer of tissue-simulating material should be mounted in front of the


cells. In this case, the number and energy of protons set in motion outside the cell and entering it are almost equal to the number and energy of protons produced inside the cell and emerging from it. However, in routine procedures for fast neutron irradiations of cells in monolayers, cultured in standard polystyrene flasks, the conditions of secondary-charged-particle equilibrium will not completely be satisfied (Broerse and Zoetelief, 1978). The energy deposited in the cell will depend on the geometrical arrangement, notably, the irradiation of the cells through the polystyrene bottom of the culture flasks or through the medium that has essentially the same atomic composition as water. The survival curves for the 1S-MeV neutron irradiation in the two geometries (Figure 28) show a higher level of survival for irradiations through the polystyrene bottom of the flasks. This indicates

102

~ ., u Q; Cl.

~

" u

'" .= > ~ ~

'0 c: .~

10 1 t ~

o 2 4 6

soft tissue kerma ( G y )

Figure 28. Survival curves of cells irradiated with d + T neutrons through the medium or through the polystyrene bottom of the culture flasks (Broerse and Zoetelief, 1978).

Progress in Neutron Dosimetry for Biomedical AppHcations 61

that a smaller amount of energy is absorbed in the cells in this situation. The relative absorbed doses in cell layers irradiated through water or polystyrene differ by a factor of 1.16. This ratio is somewhat lower than that expected on the basis of the kerma ratios, which amounts to 1.21. This discrepancy has to be attributed to small contributions from energy deposition processes inside the cell and from backscattered secondary charged particles. Cell irradiations with IS-MeV neutrons are generally performed with the neutron beam reaching the cells through the medium. Under these conditions, a correction of 5% has to be applied to account for the slight increase in the neutron absorbed dose.

It can generally be concluded that perturbations of charged-particle equilibrium at interfaces of materials of different compositions can lead to considerable differences in the neutron energy deposition. Accurate interface dosimetry will be necessary in these situations to explain the observed biological phenomena.

S.2. Absorbed Dose Distributions in Animals

If groups of animals are irradiated simultaneously by fast neutrons, precautions have to be taken to ensure a homogeneous dose distribution in the irradiation field. For fast neutrons emitted by accelerator targets the neutron fluence and energy are usually nonisotropic and have to be considered. A special procedure can be to place the animals at fixed directions with respect to neutron emission from the target. An example of such an exposure arrangement employed for the total body irradiation of rats is shown in Figure 29 (Broerse et al., 1978c). The animals are irradiated on a fixture that is rotated coaxially with the ion-beam axis. This procedure has the advantage of averaging fluctuations in the fluence rate around rings of constant angle due to aberrations produced by nonsymmetrical target structures. The irradiation arrangement shown has the additional advantage of multilateral irradiation, since the animal cages can rotate to maintain the vertical positions of the animals. The animal holders are located parallel to the beam axis, which implies that the distance to the cage decreases with increasing angles. However, the resulting increase in the dose rate is compensated for by decreases in the differential cross section for the neutron-producing reaction and in the average neutron energy with increasing angle. The dose distribution along the longitudinal axis of the animal as measured for neutrons from the p + T and d + D reactions at various positions are shown in Table 7. The results indicate a symmetrical dose distribution with reference to the center of the animal.

To study the effects on individual organs, e.g., spinal cord or lungs or the response of tumors in vivo, collimated neutron beams should be used. As an example, Figure 30 shows a multihole collimator constructed of


Figure 29. Experimental arrangement for multilateral irradiation of rats with 4- and D,S-MeV neutrons produced by the d + D and p + T reactions, respectively (Broerse et ai" 1978c),

Table 7. Relative Absorbed Dose Distribution Along the Longitudinal Axis of the Rat Phantom Irradiated with p + T and d + D Neutrons in the

Experimental Arrangement Shown in Figure 29

Distance along longitudinal p + T neutrons d + D neutrons axis (cm) (%) (%)

-5 93,S 91.8 -2,5 101 98,2

Center 100 100 2,5 93,S 99,2 5 85,S 87,4


Figure 30. Experimental arrangement for partial body irradiations of the spinal cord of rats with collimated d + T neutrons.

30-cm-thick steel. This arrangement allows the simultaneous irradiation of ten anesthetized rats. It should be realized that the dose rates at the ten irradiation positions can differ, e.g., owing to the geometry of the target construction. A complication can be the scattering of neutrons outside the irradiated area into other sensitive organs of the rat, e.g., the intestinal tract, producing death due tl) the gastrointestinal syndrome. As shown in Figure 31, the absorbed dose in between two circular collimator openings decreased to 20% of that in the main beam line for d + T neutrons. The introduction of tapered shielding inserts at the edges of the field produced an appreciable reduction of the dose outside the collimator openings. This procedure prevented the loss of animals due to intestinal complications.

From the two examples of irradiation arrangements described above, it can be concluded that an appropriate choice of irradiation conditions will facilitate dosimetry and will also prevent complications in the response of the irradiated object.

64

. c

" u Q; Q.

~ a "tI

"tI

" .a 0 ] a

" .~ ..5! ~

100

80

60

40

20


o total dose without insert 6. gamma dose without insert • total dose with insert

distance along vertical axis (cm)

Figure 31. Absorbed dose distributions along the vertical axis across two openings of the collimator shown in Figu~e 30.


6. DOSIMETRY FOR RADIOTHERAPY

The aim of dosimetry in radiotherapy applications is the precise statement of the absorbed dose at all points of interest in a patient during treatment. For irradiation with photon sources, this determination of the absorbed dose is usually a two-step process (see ICRU Report 23, 1973). As a first step, the absorbed dose at any point of interest in a standard water phantom is determined relative to the peak absorbed dose. In the second step, the water phantom is replaced by a patient, which causes deviations from the observed dose pattern owing to differences in the shape, size, and composition of the patient. For the determination of absorbed dose in a patient irradiated with a beam of neutrons, the procedure can be analogous to that for photons. An important difference between the two types of irradiation is that it is essential to assess the radiation quality of the beam in much more detail in neutron therapy than during photon irradiation. This implies that separate absorbed dose distributions must be determined for photons and for neutrons, while, in addition to the absorbed dose, knowledge of the neutron energy will also be required.

The results of neutron dosimetry for external beam therapy, as described in ICRU Report 26 (1977), were based mainly on experience with the Hammersmith d(16) + Be beam, the TAMVEC d(50) + Be beam, and some experimental preclinical d + T neutron beams. Since 1975, an increasing number of centers have started fast neutron radiotherapy (see Table 1) and considerable progress has been made in clinical fast neutron dosimetry. Procedures and numerical data have been evaluated and presented in European and American protocols for neutron dosimetry in external beam therapy. These protocols will be summarized and dose distributions compared for several therapy beams in the phantom as well as in the patient.

6.1. Protocols for Neutron Dosimetry for External Beam Therapy

Dosimetry protocols have been developed in Europe and in the United States to ensure compatibility of the dosimetry among neutron therapy centers. Physicists in the United States and in Europe have, in a more or less independent manner, drafted a number of recommendations concerning techniques, instruments, and basic physical parameters of neutron dosimetry. Both groups recommended the use of calibrated AlSO plastic TE ionization chambers with TE gas filling as the practical method of obtaining the absorbed dose in a TE phantom. The American centers and the European groups each adopted a specific TE ionization chamber, the first groups using the EGG 1-cm3 TE ionization chamber, and the second the Exradin O.5-cm3 TE ionization chamber (see Figure 6). Because the uncertainties in the values of the physical parameters employed to derive


absorbed dose from TE ionization chamber measurements are greater than desired (see Section 3.2), efforts are being continued to improve the data base of neutron dosimetry. The recommended values differ somewhat in the two protocols and may be subject to changes in the future.

A number of important items included in dosimetry protocols for fast neutrons, as drafted by the European Clinical Neutron Dosimetry Group (ECNEU, 1981) and Task Group No. 18 of the American Association of Physicists in Medicine (AAPM, 1980), and for photons, as drafted by the Hospital Physicists' Association (HPA) have been described earlier (Broerse et al., 1979b). One of the essential differences between the European and the American neutron dosimetry protocols is that the European group recommended that water be used as the standard phantom material for the following reasons: (a) The neutron attenuation properties of water are rather similar to those of most biological tissues; (b) its composition is constant and it is available in all centers; and (c) it is the medium on which a considerable body of information on the spatial distribution of absorbed dose in photon and electron therapy is based. The American groups, on the other hand, prefer to use muscle-equivalent liquid as the standard phantom material, with the intent of correcting for differences in elemental composition and density to derive the absorbed dose in the tissues of interest, e.g., fat, bone, or ICRU muscle tissue.

Application of the procedures and basic physical parameters recommended in the European and American neutron dosimetry protocols should result in closely comparable absorbed dose values at a reference point. The appropriateness of applying both protocols was investigated in two independent experiments. McDonald et al. (1981a,b) compared the response of a tissue-equivalent calorimeter with that of ionization chambers in a number of fast neutron therapy beams in the United States and in Europe. The ionization chamber measurements were analyzed according to the locally adopted dosimetry protocol to derive the absorbed dose in AlSO plastic. For both methods of data analysis, the resulting absorbed dose values in TE plastic agreed within 2% with the value obtained by the calorimetric method. The differences in the physical parameters recommended by both protocols partly canceled each other out. If the results would have been expressed as the absorbed dose in soft tissue, the procedures recommended in the American and European protocols would have resulted in larger differences, since the kerma factors adopted in the two protocols are different (by approximately 2% for d + T neutrons).

A second experiment was performed by Mijnheer et al. (1981c) in which either the American (Le., a TE liquid phantom and a FWT l-cm3

TE ionization chamber) or the European (Le., a water phantom and the Exradin 0.S-cm3 TE ionization chamber) procedure was followed to assess


the absorbed dose to be delivered at a prescribed position in a patient to be irradiated with clinically employed d + T and d(15) + Be neutron beams. Figure 32 shows the ratios of the total absorbed dose measured in water and in TE liquid with one type of chamber (curves la and 1b) and the total absorbed dose in water determined with the Exradin chamber relative to that in TE liquid measured with the FWT chamber (curves 2a and 2b). At a depth of 10 cm the results obtained by the two independent methods show differences of about 6.5%. The use of different phantom materials can account for approximately 0.5% of the differences, whereas the different values for basic physical parameters recommended in both protocols can account for an additional 3.5% of the difference. The remaining 2.5% difference can probably be attributed to systematic differences in the operational characteristics of the two ionization chambers. The results of this dose determination employing the two different chambers and the two protocols are unsatisfactory. They show that the European and American dosimetry physics groups should stay in close contact and maintain a continuous exchange of viewpoints.

1.10 QJ

D t (Europeon protocol)

~ ... ... 1.08 QJ

.l>

~ .l> a 1.06 J:! .2 ... a 1.04 ~ e

1.02

1.00

D t (American protocol) •••••• r·;:; ~~

~a !~. ....... .... .... .......... .

- ••••• L!> •• 2 b L!> ••••• . ...

• ~ j,::::::::::::-....... -

...... •..............

••••• ·0 D t in water phantom

,......--: • •••••••••• TEl d h D in iqui p antom • ..'0···1 b t

--:9.······· 0.98

iii' • ,

0 5 10 15 20

depth in phantom ( cm )

Figure 32. Ratio of the tissue absorbed dose measured in water and in TE liquid phantoms with one type of TE ion chamber [curve la, d + T; curve lb, d(15) + Be neutrons] and the tissue absorbed dose determined by the European procedure in a water phantom relative to that measured with the American protocol in a TE liquid phantom [curve 2a, d + T; curve 2b, d(15) + Be neutrons] (Mijnheer et al., 1981c).


To facilitate the comparison of biological and clinical results, the two groups agreed that the following information concerning the specification of absorbed dose values and irradiation conditions should be provided:

1. Beam quality. The energy of the accelerated particle, target characteristics, and neutron energy should be specified. Attempts should be made to obtain information on the radiation quality at different positions in the irradiated region.

2. Dose specification. The total and photon absorbed dose should be specified at the reference point, as defined in ICRU Report 29 (1978b), and the material for which the absorbed dose has been quoted should be mentioned. An alternative approach is to quote the neutron and photon absorbed doses.

3. Beam geometry. Information on the geometrical conditions of irradiation should be provided, e.g., source-to-surface distance, field size, phantom dimensions, material, etc.

4. Penumbra and collimation. The design and material of the collimators may affect the dose distribution at the edge of the beam and this effect should be investigated. Where relevant, the absorbed dose to shielded sites should be specified.

5. Field intensity distribution. The degree of field uniformity should be established. Scattering material may be introduced into the beam path to obtain flatter isodose curves. The effect of beam flatteners on radiation quality should be determined.

6.2. Absorbed Dose Distributions in the Standard Phantom

Relative absorbed dose distributions in the standard phantom should be related to an absolute absorbed dose measurement at the reference point, which is usually chosen at a depth of 5 cm along the central axis of the beam. In order to obtain separate dose distributions for the photons and neutrons in the standard phantom, generally the readings of two detectors have to be evaluated at any point in the phantom. It is often assumed that the physical parameters employed to calculate DN and Do from these detector readings are independent of the position of the detector in the beam. For a TE ionization chamber, such an approximation will introduce only a small error in DN and Do because WN, (rm,g)N, Kd, and (Ktf Km) N will not change appreciably for the spectrum changes encountered in therapy beams. Spectrum changes may be larger in the penumbra (see Section 4.2), but a larger uncertainty in the determination of DN and Do in this region is acceptable. If detectors with a large neutron energy dependence are employed, e.g., CfC02 chambers or some TLD materials, then the use of constant sensitivities at different positions in the phantom


will introduce large uncertainties in the determination of the absorbed dose distribution. This may explain part of the differences observed among the values quoted for the gamma-ray component in the different therapy beams (Mijnheer and Broerse, 1979).

The specification of absorbed dose distributions is usually done by means of isodose curves, which indicate the absorbed dose values relative to the peak absorbed dose. The methods of acquisition of isodose distributions for photon beams has been discussed in some detail in ICRU Reports 23 and 24 (1973 and 1976). Isodose curves are usually derived from absorbed dose data along the central axis of the beam and beam profiles measured at several depths. Because such a procedure required considerable beam and calculation times, systems for computing the dose matrix have been developed. Computer programs for the generation of dose distributions based on the scatter/air ratio (Shapiro et al., 1979) or using a modified decrement line method (Hogstrom et al., 1976) have been developed for fast neutron radiotherapy beams. The two techniques are described for photon beams in ICRU Report 24 (1976). Both methods are, however, semiempirical; the mathematical functions are based on a limited number of measurements. Transport calculations of neutron absorbed dose distributions have been reported by Burger et at. (1978b) and Bohm et al. (1978). Although these transport calculations can be helpful in understanding some of the properties of neutron beams, their accuracy for radiotherapy purposes has to be evaluated, e.g., by comparison with experimental dose distributions. Information on treatment planning procedures employed by the different neutron radiotherapy centers has recently been summarized (see Burger et a!., 1981).

A summary of relevant data concerning dose distributions of clinical fast neutron beams is given in Table 8. As can be seen, for cyclotrons the values for the depth at which the total absorbed dose (DN + Do) is reduced to half its maximum value, d(50%), increase with the average neutron energy. A comparison with photon sources shows that a deuteron energy of 30 to 35 MeV is necessary to obtain a penetration comparable to a 60CO gamma-ray source. The d(50) + Be and p(66) + Be neutron beams have depth dose characteristics comparable to a 4-MeV x-ray beam. The values for d(50%) for the d + T sources are somewhat lower than for 60CO gamma rays.

Values for the penumbral width, chosen as the distance off axis between 80 and 20% of the central axis total absorbed dose at a depth of 10 em in the phantom, are listed in the last column of Table 8. The data for the cyclotrons show that in general the penumbral width decreases with increasing deuteron energy. This may be explained by the more forward peaking of neutrons from the higher-energy deuteron reactions (see Section 2) and the decreasing scattering cross section of hydrogen at higher neutron


Table 8. Summary of Relevant Data, Including Depth Dose Characteristh:s and Penumbral Width, of Fast Neutron Therapy Sources

Reaction SSD Phantom d(50%t Field size x(20%)-x(80%)b Place (E in MeV) (em) material (em) (em x em) (em)

Dresden d(13.5) + Be 100 TE sol.c 8.0 12 x 15 2.3 Tokyo d(14.0) + Be 125 TE sol.c 8.1 10 x 10 2.3 Essen d(14.3) + Be 125 H2O 8.1 10 x 10 2.4 Edinburgh d(15) + Be 125 H2O 9.0 11x11 2.3 London d(16) + Be 117 TEsol.d 8.7 9.5 x 9.5 2.2 Seattle d(21) + Be 150 TE sol.c 9.4 10 x 10 1.8 Chiba-shi d(30) + Be 175 TEsol.d 11.7 10 x 10 1.9 Washington d(35) + Be 125 TE sol.c 11.1 10 x 10 1.6 Houston d(50) + Be 140 TEsol.c 13.2 10 x 10 2.7 Louvain d(50) + Be 157 H2O 13.6 10.5 x 10.5 1.3 Batavia p(66) + Be 153 TE sol.e 14.6 10 x 10 1.4 Glasgow d(0.25) + T 80 H2O 9.6 10 x 10 3.8 Hamburg d(0.5) + T 80 TE plastic 9.7 10.8 x 10.8 2.9

(A150) Heidelberg d(0.25) + T 100 H2O 10.6 11x11 3.6 Manchester d(0.25) + T 80 H2O 9.4 10 x 10 3.9 Amsterdam d(0.25) + T 80 H2O 9.9 9 x 11 4.6 ---------------------------------------------------------

60CO gamma rays 80 H2O 11.9 10 x 10 1.6 8-MeV x rays 100 H2O 17.3 10 x 10 0.8

a Depth at which the total n + 'Y dose is reduced to half its maximum value. "Distance off axis between 80 and 20% of the central axis total absorbed dose. <Density of 1.07 g cm-3 •

·Corrected to unit density.

energies. The large penumbra for the d + T generators reflects the larger target size needed to obtain adequate dose rates with these facilities.

The relative gamma-ray absorbed dose increases with depth in the phantom and also with field size, as can be seen from Table 9. The higher average value for the d + T generators may be due to the thinner shields defining the movable d + T beams, as compared to the better shielded, mostly fixed cyclotron beams.

Outside the useful beam, the relative gamma-ray contribution increases with increasing distance from the central axis, as can be seen from Figure 33. The resulting decrease in RBE, however, will be compensated for by the increase in RBE owing to the lower average neutron energy and the lower absorbed dose per fraction outside the beam as compared to the center. Recently Bewley and Page (1978) showed that when allowance is made for changes in the RBE of the neutron component (even with perfect shielding), the biologically effective dose outside the useful beam is much


Table 9. Gamma Dose as a Percentage of Local Neutron Plus Gamma Dose at Different Depths in a Phantom along the Central Axis of Fast Neutron Therapy

Sources a

Depth Reaction

Place (E in MeV) 2cm 10cm 20cm Neutron-insensitive device

Cyclotrons Dresden d(13.5) + Be 9.3 b 15.9b 28.0b C/C02 ion chamber Tokyo d(14.0) + Be 3.3 7.5 13.5 Mg/ Ar ion chamber Essen d(14.3) + Be 5.6 10.2 16.9 GMcounter Edinburgh d(15) + Be 4.7 7.7 12.9 GM counter London d(16) + Be 5.4 7.6 11.0 GM counter Seattle d(21) + Be 5.5 8.4 11.8 TE proportional counter Chiba-shi d(30) + Be 3.2 4.6 7.8 Teflon + C/C02 ion chamber Washington d(35) + Be 8.4 10.7 13.5 Houston d(50) + Be -5 C/C02 ion chamber Louvain d(50) + Be 6.4 7.8 8.6 GM-counter Average

value: 5.3 8.1 12.0

d + T generators Glasgow d(0.25) +T 6.7 8.9 11.7 GM counter Hamburg d(0.5) + T 7.2 9.2 12.5 GM counter Heidelberg d(0.25) + T 7.0 GM counter Manchester d(0.25) + T 14.5 15.0 19.5 C/C02 ion chamber Amsterdam d(0.25) + T 8.5 11.6 15.3 GM counter Average

value: 9.2 10.3 14.8

aFor the same field sizes, source-to-surface distance (SSD) and phantom material as listed in Table 8.

b Not included in the average due to the larger field size. cGamma dose derived by subtracting neutron dose from the total absorbed dose.

higher for d(16) + Be neutrons than for a 60Co machine. This same statement is probably valid for other fast neutron therapy beams, indicating that in neutron therapy one has to accept a greater hazard to the patient from stray radiation than with megavoltage x rays.

The buildup of absorbed dose below the surface of a phantom shows a pattern similar to that for photons alone. For the various neutron beams employed in radiotherapy the reported ratios for ionization at the surface of the phantom and the maximum ionization vary between 29 and 62% (Mijnheer and Broerse, 1979). The depth at which the maximum absorbed dose is reached varies between about 2 mm for a d(15) + Be beam and about 9 mm for a d(50) + Be beam. For p(66) + Be neutrons, the maximum

72

~ 0

" " '" 50 -e 0

1; 0

,.. +

" 0

£ D " 2 '0

'" i3 " ,.. . " '" " Q; Q.

...... ...... ... ' ,," ,," I " ,/

,I


...... ...... .",..,. ..... --

---------", .I ---------_.

I o 5

I

d ( 14.3 ) + Be - Essen d ( 50 ) + Be - louvain d (0.25) + T - Amsterdam

I 10

i 15

lateral distance from beam axis ( cm )

Figure 33. Gamma-ray contribution to the total absorbed dose as a function of the lateral distance from the beam axis for three different neutron beams for field sizes of approximately 10 em x 10 em (Mijnheer and Broerse, 1979).

absorbed dose is reached at a depth of 17 mm, reflecting the high proportion of high-energy neutrons in such a spectrum.

Separate isodose curves have to be determined for the neutrons and for the photons. Whether and at what stage the separate dose components should be combined will depend on the dose specification procedure. Because of the uncertainties in the determination of the relatively small photon dose component, the overall uncertainty in the separate neutron absorbed dose will be greater than that of the total absorbed dose. Consequently, a number of neutron radiotherapy centers recommend that the absorbed dose be primarily specified as the total dose. Alternatives are to quote the total effective dose DE = DN + Dol T, where T is a weighting factor that indicates the effectiveness of the neutron component as compared with the gamma-ray component for relevant effects on tumors and normal tissues. A number of European institutes have adopted a weighting factor of 3, however, this aspect is still under discussion.

Isodose distributions generated for the Louvain d(50) + Be cyclotron beam and the Amsterdam d + T generator for the total effective dose are

Progress in Neutron Dosimetry for Biomediatl Applications

1 ON +3 °G

d (50) + Be - Louvain

10 em x 10cm

1 ON + 3 °G

d (0.25) + T - Amsterdam

llcm x 9cm

73

Figure 34. Isodose distributions for two different neutron beams for the total effective dose DN + DG/3 (Mijnheer and Broerse, 1979).

given in Figure 34. The better penetration and sharper penumbra of the therapy installation with the higher neutron energy are evident.

Modification of standard absorbed dose distributions is often required to produce isodose lines of a specific shape, to protect areas of the patient, or to compensate for the patient contour. The use of wedge filters, beamshaping devices, flattening filters, etc. is, in principle, not different from their use in photon dosimetry, as discussed in ICRU Report 24 (1976). The method of generation of these modified isodose distributions is identical to that for the unmodified beam.

6.3. Absorbed Dose Distributions in the Patient

Dose distributions in a patient may be quite different from those in the standard water phantom owing to differences in geometry, composition, and density. The acquisition of patient data such as body contours, contours and densities of relevant organs, and target volume location is by no means easy. Modern computerized tomography (CT) scanners, however, can be very helpful in obtaining information about these characteristics.


The methods for combining patient data with the basic absorbed dose distribution are again, in principle, analogous to those applied for photon treatment. But there are some differences in practice between treatment planning of patients irradiated with photon and neutron beams. The smaller biological effect of the accompanying gamma rays per unit absorbed dose has to be taken into account. The influence of radiation scattered from the collimator and other structures may no longer be neglected, contrary to the procedure for treatment planning for some photon beams. Since, in general, a limited number of fixed field sizes is available, current sophisticated photon treatment planning techniques, which apply optimization procedures for choices of radiation beam arrangements, have to be adapted for neutron treatment. In addition, it might be necessary to take into account changes in RBE with position in the patient.

The influence of the volume on the neutron absorbed dose distribution can be ignored if a sufficiently broad region of scattering material surrounds the field. The photon contribution, however, increases with irradiated volume, owing to (n, 'Y) interactions produced by degraded neutrons.

The contour shape of a patient can be taken into account by modifying methods applied in photon dosimetry, as discussed in ICRU Report 24 (1976). These methods, e.g., the tissue/air ratio and the effective source-tosurface distance (SSD) methods, must be carefully applied, because the inverse-square law may not be valid for the neutron beam under all conditions (e.g., for low-output d + T generators).

The alteration in the dose distribution behind an inhomogeneity can be taken into account by considering differences in exponential attenuation functions. For bone and fat it is usually assumed that the distortion of dose distributions by these materials is so small that it need not be taken into account. Kerma deposited in these materials, however, is about 40% lower in bone and about 15% higher in fat, as compared to muscle tissue for most therapy neutron beams (Bewley, 1980). The change in dose distribution due to inhomogeneities adjacent to the direct radiation path requires methods of calculation other than simple exponential attenuation. A technique based on measured scatter for hydrogenous samples has been described by Weaver et al. (1979); it gave correction factors accurate to within a few percent. It should be noted that correction factors may be different for neutrons and accompanying gamma rays. This has been observed by Mijnheer et al. (1975) for measurements with and without lung-equivalent material in a phantom irradiated with a d + T neutron beam (see Figure 35).

Some examples of treatment plans for clinical fast neutron beams are given in Figures 36 and 37. The smaller penumbra of the cyclotron neutrons is advantageous during irradiations where critical organs are not far from the target area, e.g., the spinal cord during irradiation in the head and

Progress In Nentron Dosimetry for Biomedical Applications 75

3

'7 c

'E >-

Cl u

! ~ .. 0 " " .. .D (;

11 0

0.1

a

i~o.. 0 without lungs ~ • with lungs

1~o

~. o~ "~

.~~ ... , "'-;--':0, r ........... ...

..... ...... "":.,. ..... , ~ ~:tal obsorbed ~o dose rate .. "'~ ..

~ ........... _ I ............ ~ ... ~ .......

.............. .: .......... ... 0 ..................... ....

...... , ............. .... ................ -gamma ray

..... 0 dose rate

"------Iung ------01

, i i iii i • , iii' iii i

10 20 30

depth in water phantom (em)

Figure 35. Depth dose measurements of the total absorbed dose rate and gamma-ray absorbed dose rate for a collimated d + T neutron beam in a water phantom with and without lung-equivalent material (Mijnheer et ai., 1975).

neck region (see Figure 36). The treatment plan is comparable to that for a 4-MeV x-ray treatment. The poorer penetration characteristics of d + T neutrons make the use of special techniques necessary, e.g., a six-field isocentric irradiation for treatment in the pelvic area (see Figure 37).

To be certain that the patient receives the absorbed dose according to the treatment plan, the monitoring system and the patient dose should be regularly checked. Procedures for tests of the constancy of the absorbed dose monitoring system, which should consist of two independently working systems, have been described in protocols for electron accelerators (see, e.g., NACP, 1980) and can be adapted for accelerators of heavier particles. An ultimate check of the actual treatment given can only be made by using in vivo dosimetry. Entrance dose, exit dose, and intracavitary or organ dose measurements are possible, as described for photon irradiations (ICRU, 1976). In addition, it will be necessary to check the agreement between the size and position of the neutron beam and optical or mechanical beam-localizing devices. Radiographic films are very useful for this purpose

76

J DN + JDG d (50) + Be - Louvoin

JOem x JOem

-


-.. . ',.,. 4

..... A" JI.""",'..,.A ,,, ... ,,/U"A

Figure 36. Treatment plan of a lesion of the floor of the mouth by means of two equally weighted opposing fields at Louvain for d(50) + Be neutrons.

as well as for the measurement of beam homogeneity or for checking the position of shielding blocks (Dutreix et al., 1978). The position of a patient can also be checked by means of radiographic film, as has been demonstrated for the Hammersmith beam (Catterall and Bewley, 1979). Neutron radiography is, however, only useful for small target sizes and can therefore not be applied for the d + T neutron therapy beams presently employed. Only limited information is available on in vivo fast neutron dosimetry for checking how closely phantom materials simulate the average density and composition of the human body in the irradiated area. At Hammersmith hospital, in vivo dosimetry has been applied, using activation of aluminum by the 27 Al(n, p) 27Mg reaction. The rather short half-life of 27Mg (about 10 min) is a disadvantage, especially at the lower dose rates of d + T machines. Catterall and Bewley (1979) express a preference for the 11S! ( ') 115m! • . h . f' f n n, n n reactIon, slDce t e cross-sectlon curve as a unctIon a neutron energy is in reasonable agreement with the curve of tissue kerma. Counting of 24Na activity, which is also induced in aluminum, or of 56Mn produced in iron pellets has been used to check the absorbed dose in the pelvis during treatment with d + T neutrons (Mijnheer et al., 1981b). Changes in the fast neutron spectrum will affect the level of activity induced in these detectors per unit of neutron absorbed dose. Spectrum changes with depth or toward the edge of the field should therefore be taken into account. The response of CaS04:Tm TLD has been investigated for assessing DN and DG separately during neutron therapy with the d(16) + Be

Progress in Neutron Dosimetry for Biomediall Applications

.. '5~ ;-.. ~.; ~ .

; ~! ~ i .... ; \ ::!<>? 4"\, ,,)" .. ~ .:- :'~8

4 ~ '. - ? .. ~ .: -" ~ € •

'3 .. ~ .; :'3,,"I.e;;" 1 .. '5 € :-1 4'5': :

.'5': ;-1 1 1 : ; :: 80 i .. ~ ,0 , 3 .. ~ € -'34'5'; -'34'51; • '34'5"" ;' '34'567 J .. '5 6 7 l .. '5 I) .,

J .. '5 .; ~

J .. '5 oi ;0 J .. '5 6(' J .. , 61 J .. '5 67 '3 .. '5 6?' 3 .. '5 ';7 '3 .. S 67 J .. '5 ~7'

J 4' 6' »»»'»»»»)»>»»»»»» J "'5 67 • )) »)) '3 ., 61 J ») » J .. , 67 J ., 67 '3 .. , 61: , 100·· ..

t weight 0.6 ,

2 , m Ul

'" 222Z 22ZZ

" " 22

~~ " " " 22

" " 22 22

" 22 Z2

" 22 22 22

" " " " " " 22 22 22

U22 2222 222 m 222 222 222 222 m

77 nn -777 7 77777:-77771

77777 777 7777 601)6666 ;0 • Ii '5'5''5''5 . , , ,

t weight 0.8

77

1.0

Figure 37. Treatment plans for the total effective dose (DN + DG/3) for bilateral irradiation (top) and six-field irradiation (bottom) of bladder carcinoma at Amsterdam for d(O.2S) + T neutrons (field size, 11 cm x 9 cm); focus isocentric distance, 95 cm.


beam at Hammersmith (Blum et al., 1976). The uncertainty in the resulting values for DN is larger than during radioactivity counting, owing to the strong variation in TL response with neutron energy. Smith et al. (1976) inserted silicon diodes into the esophagus to measure the corrections to tumor dose resulting from intervening lung tissue during therapy with the TAMVEC d(50) + Be neutron beam. Silicon diodes are well suited for in vivo fast neutron dosimetry, but their accuracy is comparable only to the activation method if special precautions are taken.

7. DOSIMETRY FOR IN VIVO NEUTRON ACTIVATION ANALYSIS

In vivo neutron activation analysis is employed for the determination of elemental composition in patients and laboratory animals. The subjects are exposed to a low absorbed dose of neutrons (usually below 1 rad = 1 cGy) and the induced radioactivity is measured either subsequently or immediately by analysis of prompt gamma radiation. A number of elements, e.g., calcium, phosphorus, sodium, chlorine, oxygen, and nitrogen, can be determined employing total and partial body irradiations (IAEA, 1973 and 1979).

Both thermal neutron reactions, e.g., 48Ca(n, 'Y) 49Ca, and fast neutron reactions, e.g., 40Ca(n, a) 37 Ar (Lewellen et al., 1979) are employed. Regardless of the type of reaction involved, fast neutrons must be used to obtain an approximately uniform ftuence over the irradiated subject. A bilateral irradiation using neutrons produced by a cyclotron, d + T generator, or isotopic neutron source is usually performed.

Calibration of these systems is usually done by measuring the activity induced in anthropomorphic phantoms filled with aqueous solutions containing known amounts of the elements under consideration (see e.g., Williams et al., 1978). The absolute accuracy will thus depend on the reproducibility of repeated measurements and not on the accuracy of the determination of the neutron absorbed dose. Monitoring the irradiation and assessment of the uniformity of the neutron ftuence over the human subject determine to a large extent the absolute accuracy of these measurements. In addition, the effects of differences in shape, size, and composition between the phantom and the patient should be considered (Williams et al., 1978; Spinks, 1979).

Monitoring the absorbed dose delivered to a patient is in principle identical to that applied during radiotherapy procedures. The monitor, however, must be capable of accurately measuring much lower doses. A number of devices are employed: large ionization chambers and LiI scintillators (Lewellen et al., 1979), bottles containing NaN03 solution (Spinks,


bu tyrote forearm --.j....L.....!..!...--~~w. .uppart

water bath -+---.:~_

-----

__ 17cm -

wax reflector

•• _--:-:--tt---i-i-- 2 52C f

-- .... .., ........ I

delivery tube

Figure 38. Irradiation geometry for neutron activation analysis of calcium in the arm, incorporating a water bath to act as a combined premoderator and reflector. The dotted lines indicate the position of a second wax reflector (Smith and Tothill, 1979).

1979), copper foils (Williams et at., 1977), 238U fission counters, and BF3 counters (Elliott et aI., 1978). If these detectors are placed near the patient, the body may affect the reading of the monitor and this should be considered (see, e.g., Williams et at., 1977).

The uniformity of the neutron fluence distribution within the body is affected by (a) the neutron energy, (b) the construction of the irradiation facility, and (c) the shape, size, and composition of the various regions of the body (Elliott et at., 1978). Methods for determining neutron fluence have been discussed elsewhere (see Section 3.4). Special patient enclosures incorporating a water bath acting as a combined premoderator and reflector are often employed to provide optimal fluence uniformity and to allow for different body sizes (e.g., see Figure 38).

8. RESULTS OF DOSIMETRY INTERCOMPARISONS

During the past seven years considerable efforts have been made to compare the results obtained by various groups performing fast neutron


dosimetry. The most important motive for the performance of these neutron dosimetry intercomparisons is to allow for an adequate evaluation and comparison of biological and clinical results obtained by different groups. Additional aims of neutron dosimetry intercomparisons are to obtain information on the adequacy of neutron dosimetry and on the advantages, disadvantages, corrections, and systematic errors involved in various methods. Comparisons of absorbed dose measurements with different types of detectors, e.g., ionization chambers, calorimeters, and fluence-measuring devices, are already discussed in Section 3.

The different neutron dosimetry intercomparison programs can be divided into two classes: intercomparisons of dosimetry systems performed at specific locations and intercomparisons of the neutron beams actually used for biomedical applications. In the first type of intercomparisons, such as the International Neutron Dosimetry Intercomparison, INDI (ICRU, 1978a) and the European Neutron Dosimetry Intercomparison Project, ENDIP (Broerse et ai., 1978a), all participants brought their systems to central locations (Brookhaven National Laboratory for INDI, Institut fur Strahlenschutz GSF and Radiobiological Institute TNO for ENDIP). In the second type of intercomparisons, including those carried out under the auspices of Task Group No. 18 of the AAPM, the former American Neutron Dosimetry Physics Group, NDPG (Almond and Smathers, 1977), and the Japan-U.S. Cooperative Cancer Research Program, CCRP (Ito, 1978b), the dosimetry systems were taken to each institution and, where possible, reciprocal visits among institutes were made.

The NDPG and CCRP measurements were made primarily on cyclotron-produced neutrons, with fairly high dose rates, larger than 20 rad min -1 (cGy min-I) for tissue kerma in air and absorbed doses at depth. The results of the intercomparisons of neutron dosimetry at the institutes cooperating within the NDPG in the United States are given in Table 10. It can be seen that all of these groups agree within 2 % in their total dose measurements. It should be emphasized, however, that the American groups involved in neutron radiotherapy all use a common set of commercial TE ionization chambers.

The INDI and ENDIP studies were generally performed with monoenergetic neutron beams produced by the p + T, d + D, and d + T reactions at relatively low kerma rates, which varied between 5 and 80 rad h -1 (cGy h -1). The quantities intercompared were the soft tissue kerma free in air and the absorbed dose at three depths in a water phantom for neutrons as well as photons. The results of INDI and ENDIP were basically the same. Although it is recognized that the calculation of a mean value for the participants' results has limited relevance, this procedure was used in order to allow a quantitative comparison of the ENDIP results. In Table 11 the results of the groups participating in the ENDIP sessions at

Tab

le 1

0.

Sum

mar

y o

f Dos

imet

ry I

nter

com

pari

sons

bet

wee

n U

.S.

Inst

itut

ions

"

Par

tici

pant

s v.

w.

NR

L

FN

L

GL

AN

TA

M

DA

H-T

AM

V

NR

L

v.w

. G

LA

NT

A

NR

L

GL

AN

TA

M

DA

H-T

AM

V

FN

L

FN

L

MD

AH

-TA

MV

L

ocat

ion

v.w

. F

NL

G

LA

NT

A

FN

L

GL

AN

TA

N

RL

M

DA

H-T

AM

V

Bea

m

d(21

.S)

+ B

e p

(66

)+B

e d(

2S)

+ B

e p(

66)

+ B

e d(

2S)

+ B

e d(

3S)

+ B

e d(

SO)

+ B

e D

etec

tors

T

E io

n ch

ambe

rs

TE

I.C

. T

E I

.C.

TE

I.C

. T

E I

.C.

Tis

sue

kerm

a in

air

99

%

>99

.5%

99

.5%

>

99.5

%

>9

9.9

%

Dos

e at

dep

th in

pha

ntom

99

% a

t 2

cm

98%

at

10 e

m

60

eo

cal

ibra

tion

99

%

>99

.5%

99

%

>99

.5%

>

99

%

"Fro

m P

. W

oott

on a

nd J

. E

enm

aa;

priv

ate

com

mun

icat

ion

(197

8).

S/K

v.

w.

v.w

.

d(21

.5)

+ B

e

TE

I.C

. T

E c

alor

imet

er

>9

8%

at

Scm

99.5

%

l 1 s· Z I =

=

'"' r S - = .. 1:1:1 f :&. [ ~

"CI = I i Q

a I-

"


Table 11. Number of Evaluated ENDIP Results with Relative Differences Jx from the Mean a

Neutron Site of energy

intereomparison (MeV) Condition Ax ,,;;;5% 5% < Ax,,;;; 10% Ax> 10%

GSF 15.1 Free air KN: 6/12 3/12 3/12 K,o': 8/12 2/12 2/12

GSF 5.25 Free air KN: 7/11 3/11 1/11 K,o': 8/11 3/11 0/11

TNO 15 Free air KN: 11/12 1/12 0/12 K,o': 10/13 3/13 0/13

TNO 15 5 em depth DN: 6/12 5/12 1/12 D,o': 10/13 2/13 1/13

TNO 15 10em depth DN: 5/1z" 6/12 1/12 D,o': 9/13 3/13 1/13

TNO 15 20 em depth DN: 3/12 8/12 1/12 D,o': 9/13 3/13 1/13

TNO 5.5 Free air KN: 8/8 0/8 0/8 K,o': 8/9 1/9 0/9

TNO 5.5 5 em depth DN: 4/8 4/8 0/8 D,o': 6/9 3/9 0/9

TNO 5.5 10em depth DN: 2/8 5/8 1/8 D,o': 7/9 2/9 0/9

TNO 5.5 20 em depth DN: 4/8 2/8 2/8 D,o': 7/9 2/9 0/9

"From J. J. Broerse et al. (eds.), A European Neutron Dosimetry Intercomparison Project (ENDIP) Results and Evaluation, EUR-6004, Commission of the European Communities, Luxembourg (1978a).

GSF and TNO are grouped for relative differences from the mean below 5%, from 5 to 10%, and in excess of 10%. An extensive analysis of the ENDIP results has been made (Broerse et al., 1978a); only a few main conclusions will be summarized:

1. In general, the variations in total kerma and total absorbed dose are smaller than those observed for the neutron kerma and the neutron absorbed dose.

2. The results showed relatively small variations (e.g., standard deviations of ±3 to 3.5% for ENDIP-TNO) for the free-in-air conditions, but larger variations (±4 to 6% for ENDIP-TNO) are observed for measurements in the phantom for the total absorbed dose.

3. The neutron kerma or absorbed dose values show standard deviations on the order of ±7 to 8% for measurements in a phantom. These variations seem to be in accordance with the relatively large

Progress In Neutron Dosimetry for Blomedleal AppUeations 83

systematic uncertainties quoted by the participants. Only for a few specific situations were maximum differences up to 20% observed.

4. The values reported for the gamma-ray kerma and absorbed dose showed large variations of up to 100% from the mean value. These variations are not acceptable for the measurements in the phantom, where relatively large photon contributions are measured (up to 25% of the total absorbed dose).

5. In the ENDIP and INDI studies, the participants employed hydrogenous ionization chambers of different design for the determination of the total absorbed dose, and they applied different basic physical parameters characterizing the detector response for identical experimental conditions.

To exclude the influence of the introduction of differing values for the basic parameters, the relative responses of the participants' dosimeters were also compared. In Table 12, the variance analysis for total kerma and total dose values and for the reduced instrument responses is given for the ENDIP results obtained at TNO by participants using TE chambers flushed with TE gas. It can be seen that the standard deviations for instrument response are of the same magnitude as those calculated for absorbed dose and kerma. This implies that, in addition to the inconsistencies in basic physiCal parameters, there are also large systematic differences in measurement procedures (see Section 3.1).

Table 12. Variance Analysis of ENDIP Results at TNO Rijswijk G

Measurement condition - b D, s s (%) R'/Nb s s (%)

5 MeV free in airc 56.1 1.7 3.0 55.0 1.2 2.1

5cm 60.5 2.3 3.8 59.6 2.1 3.6 10cm 29.2 1.1 3.7 28.9 1.1 3.9 20 em 6.9 0.27 3.9 6.8 0.25 3.6

15 MeV free in airc 56.8 2.1 3.7 56.5 1.6 2.9

5cm 64.8 3.5 5.4 64.5 3.2 5.0 10cm 37.3 1.9 5.2 37.2 1.8 4.8 20 em 12.7 0.79 6.2 12.6 0.71 5.6

"From J. J. Broerse et at. (eds.), A European Neutron Dosimetry Intercomparison Project (ENDIP) Results and Evaluation, I EUR-6004, Commission of the European 'Communities, 'Luxembourg (1978a).

~. Values in rad per 10' monitor units. ·Under these conditions the values refer to the mean total kerma.


The different intercomparisons have shown that it is equally important to adopt uniform basic parameters so as to standardize the experimental techniques employed by different groups for the determination of absorbed dose. A better understanding of the systematic uncertainties in detection techniques is required before future neutron dosimetry intercomparisons at central locations will be profitable. Although each institute should be free to develop instruments of special design, it has been recommended that all neutron radiotherapy centers have the same type of ion chamber to check their other dosimeters. It would be profitable if international bodies were involved in testing the characteristics of the common dosimeter systems. At the initiative of the Commission of the European Communities, a committee involved in the collection and evaluation of neutron dosimetry data (CENDOS) has been formed. In connection with this program a study of possible operational errors has been performed for a number of ENDIP participants whose results showed large discrepancies. The groups from CENF, Fontenay-aux-Roses, GSF, Neuherberg, and TNO, Rijswijk, participated in a measuring session in 1977 at GSF, Neuherberg, employing the ENDIP setup. The three participating groups each used different ionization chambers: CENF and TNO used homemade ones and GSF employed commercially available chambers. The results of this CENDOS small-scale neutron dosimetry intercomparison (Broerse et al., 1979a) led to the following conclusions:

1. As can be seen from Table 13 for free-in-air exposure conditions, the 1977 CENDOS results for the responses of the ion chambers show a much better agreement than those of 1975 ENDIP at GSF. However, none of the three groups are aware of fundamental changes in the experimental techniques applied.

2. The CENDOS ion-chamber measurements in the water phantom show larger discrepancies than for free-in-air conditions. These differences could be attributed to the choice of the effective measuring point (see Section 3.2) being the geometrical center for one group, at approximately three-quarters the radius of the gas cavity to the front of the chamber for another group, whereas the third group applied a correction factor of 0.975. When a consistent displacement correction is applied, the results of the three groups are in close agreement, as shown in Figure 39.

3. Comparing the measuring techniques employed by the CENDOS participants and the calculational procedures on site appeared to be very useful. In this way initial differences could be resolved.

The results of the sQlall-scale CENDOS neutron dosimetry intercomparison once again emphasized the need for uniform procedures and


Table 13. Relative Responses of Ionization Chambers and GM Counters Obtained during a 1977

CENDOS Measuring Session, Including 1975 Free-in-Air Measurements a

Type of Experimental condition detector CENF GSFM

15 MeV free in airb TE chamber 1.084 1.000

GM 0.037 0.048 free in air TE chamber 1.002 1.000

GM 0.045 0.048 5-cm depth TE chamber 0.993 1.000

GM 0.106 0.111 10-cm depth TE chamber 0.996 1.000

GM 0.136 0.145 252Cf free in air TE chamber 1.040 1.000

GM 0.316 0.311

85

TNO

0.923 0.037 0.993 0.034 0.951 0.091 0.948 0.120 1.036 0.303

OJ. J. Broerse et al .• A Small Scale Neutron Dosimetry Intercomparison. EUR-6567. CENDOS. Rijswijk (1979a).

"Results of 1975 study.

techniques for measuring chamber response and determining the appropriate corrections. In addition to the introduction of a common type of dosimeter, a consistent set of basic physical parameters should be used and an agreement should be reached on the procedure for converting instrument readings into absorbed dose values. All of these steps were considered to be of great importance for improving the consistency of neutron dosimetry for biological and medical applications.

The various intercomparison programs have stimulated much activity in the collection of new dosimetry data and drafting protocols for neutron dosimetry applications. It is gratifying that these efforts have already achieved their aims to a certain extent. Intercomparisons of neutron dosimetry performed among different groups, namely, Amsterdam-Edinburgh-Essen, Essen-Hamburg-Heidelberg, and CENDOS-EdinburghLondon-New York, have generally resulted in agreements for the measured values of total absorbed dose to within a few percent.

Repeated intercomparisons at time intervals of several years will provide the participants with the possibility of checking the accuracy and precision of their neutron dosimetry and the long-term appropriateness of their monitoring and exposure arrangements. If more accurate values for the basic physical parameters become available, they should be incorporated in the neutron dosimetry protocols after careful evaluation. As discussed in Section 6.1., special attention should be given to ensure that

86

+14

+12

+10

+8 ... -+0 ... ... ...

c +4 II

~

" ~ +2 c D

"

... 00

... II II ...

II II II 0

E 0 E

II 0-0-II

D ~

-2 II

0 II II II II

u c l!! -4 ~

... o !il ~ 0

" -6

-8

-10 o ••• , •• , , , , i i

~~ o;EEEo;eEE ~ ~ ~ II U U U II U U u

~_~~1I)2~~an2~ .. ." 0"'''' 5.5 MeV 15.1 MeV

...

...

II

t ... II

II 0 II ... .. 0 ---N

N

NO

00

J. J. Broene and B. J. Mijnheer

II II

II ... f

... ...

0

o 0

0

...

... ... II

II II

000

... CENF

II II GSF

o TNO

II

, i , , • , , , , , • i • i , , i • ~ ~

·DEEE·DEEE UUUUUUUU

~""~~~lt')~~

5.3 MeV 14.8 MeV GSF TNO TNO

~

DEE ~ II U U U e Lt) 0 ",- -.,,--'" 15 MeV

GSF CENOOS

1977 Figure 39. Results of three groups, namely, CENF, Fontenay-aux-Roses, GSF, Neuherberg, and TNO, Rijswijk, in three consecutive neutron dosimetry intercomparisons. The results are expressed with reference to the mean values observed for each separate experimental condition. During ENDIP 1975, GSF performed two independent measurement series, indicated by the different symbols.

INOl1973 ENOIP 1975

protocols for neutron dosimetry followed by different groups show as much consistency as possible.

9. CONCLUSIONS AND RECOMMENDATIONS

The increased use of fast neutrons for biomedical applications and the need for greater accuracy in neutron dosimetry have led to a considerable accumulation of new basic physical parameters. Although the principles and methods in neutron dosimetry as discussed in ICRU Report 26 (1977) are still valid, this review chapter can be considered as updating and


extending the information on basic neutron dosimetry data and experimental procedures.

Tissue-equivalent (TE) chambers are generally accepted as the most practical instruments for absolute measurements as well as for the determination of relative dose distributions in a neutron field. The largest uncertainties in the determination of total absorbed dose with TE ionization chambers are connected with W ratios, gas-to-wall absorbed dose conversion factors, and neutron kerma ratios between dosimeter materials and tissue (see Table 4). At neutron energies in excess of 15 MeV, an increased knowledge of these physical parameters will be necessary for obtaining satisfactory accuracy in kerma and absorbed dose values. The most important nuclear data required are the reaction cross sections and total cross sections for oxygen and carbon for the neutron energy range between 15 and 80 Me V and the correlated energy spectra of charged particles from these elements. Also, more information is necessary for correction for the finite size of the chamber during measurements in air (wall absorption) and in a phantom (displacement correction). A careful error analysis will be essential to quantitatively assess the overall uncertainty in the total absorbed dose determination with TE ion chambers.

The increased knowledge of the thermal defect in AlSO plastic gives an overall uncertainty of about 1.5% in the calorimetric determination of absorbed dose in this material. Direct determination by calorimetry of absorbed dose in a material resembling ICRU muscle tissue more closely than AlSO plastic should be investigated.

Future neutron radiotherapy installations will to an increasing extent consist of accelerators producing a considerable number of neutrons with energies higher than 20 MeV. Especially for these higher neutron energies, neutron dosimetry intercomparisons should be performed with TE ionization chambers, calorimeters constructed of different materials, and differential fluence-measuring devices.

The preferred method for the determination of the gamma-ray absorbed dose in a neutron field depends on a number of factors, e.g., the absorbed dose rate, the neutron energy, and the required measuring time. GM counters are suitable for routine use if photon absorbed dose rates can be reduced to less than about 0.2 rad min -1 (cGy min -1). Knowledge of the relative neutron sensitivity ku for GM counters is now increasing, but information is required at energies higher than 15 MeV. Nonhydrogenous ionization chambers are also suitable for routine use at higher dose rates. However, the much higher ku values compared to the GM counter and the strong variation in ku with the neutron energy makes a much more detailed knowledge of neutron spectra necessary. A number of data have become available for ku values of nonhydrogenous ionization chambers.


Mg/ Ar chambers as well as AI/ Ar chambers have a lower neutron sensitivity than C/C02 chambers and should therefore be preferred. Nonhydrogenous chambers with even lower ku values should be developed. The TE proportional counter technique can be used as a reference method for measuring the contribution of photons to the total absorbed dose. The device can only be used at low dose rates and is not suitable for routine measurements, but an overall uncertainty of about ±10% can be obtained. The different techniques applied to determine the photon absorbed dose in a neutron field should be compared in neutron beams of different energies and with a varying contribution of photons.

Activation detectors employing the (n, p) reaction are well suited for the determination of the neutron absorbed dose in a photon field. An estimate of the energy distribution of the neutrons, however, is necessary. Moderating detectors can also give accurate results but should be corrected for photon sensitivity in photon beams of 20 MY or higher. The use of silicon diodes and fission track registration directly in the photon beam should not be recommended because of their response to high-energy photons. The registration of tracks induced directly inside polymers, e.g., in polycarbonate foils, may have a low photon sensitivity, but this needs further investigation.

Knowledge of the neutron spectrum at a reference point is necessary for the determination of relevant physical parameters determining the instrument response. Changes in the neutron spectrum at different positions in a phantom or patient inside the primary neutron field are small and the resulting changes in these physical quantities can be ignored for measurements with TE ionization chambers. If the variation in dosimeter response with neutron energy is large, e.g., for nonhydrogenous ionization chambers or some types of solid-state devices, then this variation in the neutron spectrum should be taken into account during relative absorbed dose determinations. More information on the photon spectrum in the neutron field is required to determine the photon sensitivity of some detectors, e.g., GM counters, in a neutron field.

Microdosimetry provides a suitable basis for radiation quality specification for therapy purposes, but a pragmatic approach has to be adopted. It is impossible to derive the absolute biological effectiveness of any fast neutron beam employed in radiotherapy from microdosimetric measurements or calculations alone. Possible differences in the radiation quality of different fast neutron beams can be assessed by microdosimetric techniques, but should be supported by comparing the responses of biological dosimeters in the different beams.

Perturbations of charged-particle equilibrium at interfaces of materials of different composition can lead to appreciable differences in energy deposition by neutrons. For the interpretation of biological

Progress in Neutron Dosimetry for Biomedical AppHcations 89

results it should be realized that the short-range high-LET secondaries reach equilibrium at much smaller depths than the high-energy lowLET particles. It has been shown that the measuring conditions (Le., the application of a nonhydrogenous filter, employment of a phantom, and the use of air or TE gas to flush a TE ionization chamber) are of importance for the interpretation of charged-particle buildup curves at airtissue interfaces.

An appropriate choice of experimental conditions for animal irradiations will facilitate the dosimetry to be performed and may also prevent complications in the response of the irradiated object. Dosimetry and irradiation techniques should therefore be considered as an essential part of the experimental design before radiobiological experiments with fast neutron beams are started.

Protocols for neutron dosimetry for external beam therapy have been established in Europe and the United States. Adoption of these protocols will reduce systematic differences in measurement procedures and ensure compatibility of the dosimetry among therapy centers engaged in coordinated trials of fast neutron beam therapy. Experiments with TE ionization chambers applying the measurement procedures recommended in the two protocols show differences in total absorbed dose of up to 7%. These results clearly indicate that the European and American dosimetry groups should remain in close contact and maintain a continuous exchange of viewpoints. The specification of absorbed dose and the description of the irradiation conditions should be provided according to the recommendations given in ICRU Report 29 (1978b).

The methods for the determination of isodose distributions in a patient are in principle analogous to those applied for photon treatment, but they must be carefully applied. The gamma-ray contribution to the total absorbed dose must be taken into account, while the inverse-square law may not be valid for the neutron beam under all conditions. The influence of bone, fat, lung, or other inhomogeneities on the dose distribution should be established for a particular neutron beam.

Silicon diodes and activation detectors, e.g., those employing the 27 Al(n, p) 27Mg, 27 AI(n, a) 24Na, 56Fe(n, p) 56Mn, or 1151n(n, n') 115mIn reac-tion, are well suited for in vivo fast neutron dosimetry during radiotherapy. TLD in vivo dosimeters should be further developed to obtain accuracies comparable to those of silicon diodes and activation detectors, but they may be applied for assessing neutron and photon absorbed doses simultaneously.

The accuracy of dosimetry for in vivo neutron activation analysis depends to a large extent on the monitoring of the irradiation and the uniformity of the neutron fluence distribution within the body. In general, monitoring can be carried out with a sufficient degree of precision. Because


none of the irradiation facilities currently used are identical, relative neutron ftuence measurements must be carried out for each setup.

The analyses of the results of dosimetry intercomparisons have shown that it is equally important to adopt uniform basic parameters so as to standardize the experimental techniques employed by different groups for the determination of absorbed dose. Adoption of a common ionization chamber as well as common protocols will reduce the systematic uncertainties. The protocols for neutron dosimetry in biomedical applications should be as consistent as possible. If more accurate values for the various physical data be<- )me available in the future, they should be incorporated into the protocols after careful evaluation. The accuracy of fast neutron dosimetry can be further improved through the assistance of standards laboratories, which have recently undertak..:n the establishment of standard neutron fields for absorbed dose calibrations in biomedical applications.

ACKNOWLEDGMENTS

The authors want to express their gratitude to Drs. G. W. Barendsen, J. J. Coyne, B. Hogeweg, and J. Zoetelief for stimulating discussions, Dr. A. C. Ford for assistance in the correct use of the English language, Mr. J. Ph. de Kler for his cooperation in the preparation of the figures, and Miss M. C. von Stein for her expert assistance in the typing and retyping of this chapter.

REFERENCES

AAPM, Protocol for Neutron Beam Dosimetry, Task Group No. 18, Fast Neutron Beam Dosimetry Physics Radiation Therapy Committee, AAPM Report No.7. American Association of Physicists in Medicine (1980).

Alberts, W. G., and K. Knauf, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 265-274, Commission of the European Communities, Luxembourg (1978).

Almond, P. R.., in Proceedings of a Conference on Neutrons from Electron Medical Accelerators (H. T. Heaton and R. Jacobs, eds.), NBS Special Publication 554, pp. 129-138, National Bureau of Standards, Washington, D.C. (1979).

Almond, P. R., and J. B. Smathers, Physics intercomparisons for neutron radiation therapy, Int. J. Radiat. Oneol. Bioi. Phys. 3, 169-176 (1977).

Alsmiller, R. G., and J. Barish, Neutron kerma factors for H, C, N, 0 and tissue in the energy range of 20-70 MeV, Health. Phys. 33, 98-100 (1977).

Amols, H. I., J. F. Dicello, M. Awschatom, L. Coulson, S. W. Johnsen, and R. B. Theus, Physical characterization of neutron beams produced by protons and deuterons of various energies bombarding beryllium and lithium targets of several thickness, Med. Phys. 4, 486-493 (1977).


Attix, F. H., and W. C. Roesch (eds.), Radiation Dosimetry, Vol. II, Academic Press, New York (1966).

Attix, F. H., R. B. Theus, and C. C. Rogers, in Proceedings of the Second Symposium on Neutron Dosimetry in Biology and Medicine, Vol. I (G. Burger and H. G. Ebert, eds.), EUR-5273, pp. 329-340, Commission of the European Communities, Luxembourg (1975).

Attix, F. H., S. G. Gorbics, and R. B. Theus, Neutron-Gamma Sensitivity Ratio for CaF2 : Mn and LiF (TLD-700) TLDs and Air Filled Graphite Ion Chamber, Report No. COO-1105-260 (1978).

August, L. S., R. B. Theus, and P. Shapiro, in Proceedings of the Sixth Symposium on Microdosimetry, Vol. I (J. Boozand H. G. Ebert,eds.), EUR-6064, pp.441-449,Harwood Academic Publishers, London (1978).

August, L. S., P. Shapiro, and L. A. Beach, The light-ion flux mixed with collimated fast-neutron beams, Phys. Med. Bioi. 25, 573-576 (1980).

Axton, E. J., and A. G. Bardell, in Proceedings of a Conference on Neutrons from Electron Medical Accelerators (H. T. Heaton and R. Jacobs, eds.), NBS Special Publication 554, pp. 109-119, National Bureau of Standards, Washington, D.C. (1979).

Barendsen, G. W., and J. J. Broerse, Differences in radiosensitivity of cells from various types of experimental tumours in relation to the RBE of 15 MeV neutrons, Int. 1. Radiat. Oncol. Bioi. Phys. 3, 211-214 (1977).

Barendsen, G. W., J. J. Broerse, and K. Breur (eds.), High LET Radiations in Clinical Radiotherapy, Pergamon Press, Oxford (1979).

Barschall, H. H., Intense sources of fast neutrons, Annu. Rev. Nucl. Part. Sci. 28, 207-237 (1978).

Barschall, H. H., and E. Goldberg, Response of tissue-equivalent ionization chamber to 15 MeV neutrons, Med. Phys. 4,141-144 (1977).

Battermann, J. J., and K. Breur, in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 17-22, Pergamon Press, Oxford (1979).

Battermann, J. J., G. A. M. Hart, and K. Breur, Dose-effect relations of tumour control and complication rate after fast neutron therapy for pelvic tumours, Br. 1. Radiol., in press.

Bewley, D. K., Private communication (1974). Bewley, D. K., in Ion Chambers for Neutron Dosimetry (J. J. Broerse, ed.), EUR-6782,

pp. 209-215, Harwood Academic Publishers, London (1980). Bewley, D. K., and B. C. Page, On the nature and significance of the radiation outside the beam

in neutron therapy, Br. 1. Radiol. 51, 375-380 (1978). Bewley, D. K., E. C. McCullough, B. C. Page, and S. Sakata, Neutron dosimetry with a

calorimeter, Phys. Med. Bioi. 19, 831-842 (1974a). Bewley, D. K., N. J. McNally, and B. C. Page, Effects of the secondary charged particle

spectrum on cellular response to fast neutrons, Radiat. Res. 58, 111-121 (1974b). Bewley, D. K., B. Cullen, and B. C. Page, in Proceedings of the Third Symposium on Neutron

Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 227-234, Commission of the European Communities, Luxembourg (1978).

Bhatia, D. P., and P. S. Nagarajan, Computed dose distributions across air-tissue, polythenetissue and graphite-tissue interfaces for 14 MeV neutrons, Radiat. Res. 69, 197-209 (1977).

Bhatt, R. C., A. R. Lakshmanan, B. Chandra, and A. S. Pradham, Fast neutron dosimetry using sulphur activation in CaS04: Dy TL dosimeters, Nucl. Instrum. Methods 152, 527-529 (1978).

Bichsel, H., in Proceedings of the Second Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5273, pp. 191-226, Commission of the European Communities, Luxembourg (1975).


Bichsel, H., and A. Rubach, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 549-563, Commission of the European Communities, Luxembourg (1978).

Blum, E., J. D. Heather, A. D. R. Beal, and D. K. Bewley, Calcium sulphate phosphor for clinical neutron dosimetry, Health Phys. 30, 257-261 (1976).

Bohm, J. K., H. M. Eissa, G. Hehn, G. Pfister, and P. Stiller, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 145-163, Commission of the European Communities, Luxembourg (1978).

Bond, V. P., The risk from fast neutron exposure, Health Phys. 37, 391-395 (1979). Bonnett, D. E., and C. J. Parnell, in Basic Physical Data for Neutron Dosimetry (J. J. Broerse,

ed.), EUR-5629, pp. 43-47, Commission of the European Communities, Luxembourg (1976).

Booth, R., J. C. Davis, C. L. Hanson, J. L. Held, C. M. Logan, J. E. Osher, R. A. Nickerson, B. A. Pohl, and B. J. Schumacher, Rotating target neutron generators, Nucl. Instrum. Methods 145,25-39 (1977).

Booz, J., in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 147-150, Pergamon Press, Oxford (1979).

Broerse, J. J. (ed.), Ion Chambers for Neutron Dosimetry, EUR-6782, Harwood Academic Publishers, London (1980).

Broerse, J. J., and G. W. Barendsen, A fast neutron source for radiotherapy, Nature 206, 208-209 (1965).

Broerse, J. 1. and G. W. Barendsen, in Proceedings of the First Symposium on Microdosimetry (H. G. Ebert, ed.), EUR-3747, pp. 353-383, Commission of the European Communities, Brussels (1968).

Broerse, J. J., and J. Zoetelief, Dosimetric aspects of fast neutron irradiations of cells cultured in monolayer, Int. J. Radiat. BioI. 33, 383-385 (1978).

Broerse, J. J., G. W. Barendsen, and G. R. van Kersen, Survival of cultured human cells after irradiation with fast neutrons of different energies in hypoxic and oxygenated conditions, Int. 1. Radiat. Bioi. 13, 559-572 (1968).

Broerse, J. J., D. Greene, R. C. Lawson, and B. J. Mijnheer, Operational characteristics of two types of sealed-tube fast-neutron radiotherapy installations, Int. 1. Radiat. Oncol. BioI. Phys. 3, 361-365 (1977).

Broerse, J. J., G. Burger, and M. Coppola (eds.), A European Neutron Dosimetry Intercomparison Project (ENDIP) Results and Evaluation, EUR-6004, Commission of the European Communities, Luxembourg (1978a).

Broerse, J. J., D. W. van Bekkum, C. F. Hollander, and J. A. G. Davids, Mortality of monkeys after exposure to fission neutrons and the effect of autologous bone marrow transplantation, Int. 1. Radiat. Bioi. 34, 253-264 (1978b).

Broerse, J. J., S. Knaan, D. W. van Bekkum, C. F. Hollander, A. L. Nooteboom, and M. J. van Zwieten, in Late Biological Effects of Ionizing Radiation, Vol. II, pp. 13-27, International Atomic Energy Agency, Vienna (1978c).

Broerse, J. J., J. Zoetelief, G. Burger, H. Schraube, and A. Ricourt, A Small Scale Neutron Dosimetry Intercomparison, EUR-6567, CENDOS, Rijswijk (1979a).

Broerse, J. J., B. J. Mijnheer, J. Eenmaa, and P. Wootton, in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 117-123, Pergamon Press, Oxford (1979b).

Burger, G., and A. Morhart, in A European Neutron Dosimetry Intercomparison Project (ENDIP) Results and Evaluation (J. J. Broerse, G. Burger, and M. Coppola, eds.), EUR-6004, pp. 165-174, Commission of the European Communities, Luxembourg (1978).


Burger, G., E. Maier, and A. Morhart, in Proceedings of the Sixth Symposium on Microdosimetry, Vol. I (J. Bool, and H. G. Ebert, eds.), EUR-6064, pp. 451-468, Harwood Academic Publishers, London (1978a).

Burger, G., A. Morhart, and H. Schraube, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 125-143, Commission of the European Communities, Luxembourg (1978b).

Burger, G., A. Breit, and J. J. Broerse (eds.), Proceedings of the Workshop on Treatment Planning for External Beam Therapy with Neutrons, Urban and Schwarlenberg, Munich (1981).

Busuoli, G., A. Cavallini, A. Fasso, and O. Rimondi, Mixed radiation dosimetry with LiF (TLD-100), Phys. Med. Bioi. 15, 673-681 (1970).

Caswell, R S., Neutron-insensitive proportional counter for gamma-ray dosimetry, Rev. Sci. Instrum. 31, 869-871 (1960).

Caswell, R S., and J. J. Coyne, in Proceedings of the Sixth Symposium on Microdosimetry, Vol. II (J. Bool and H. G. Ebert, eds.), EUR-6064, pp. 1159-1171, Harwood Academic Publishers, London (1978).

Caswell, R. S., J. J. Coyne, and M. L. Randolph, Kerma factors for neutron energies below 30 MeV, Radiat. Res. 83, 217-254 (1980).

Catterall, M., Radiology now, fast neutrons-clinical requirements, Br. 1. Radiol. 49, 203-205 (1976).

Catterall, M., and D. K. Bewley, Fast Neutrons in the Treatment of Cancer, Academic Press, London (1979).

Chen, S. Y., and A. B. Chilton, Depth-dose relationships near the skin resulting from parallel beams of fast neutrons, Radiat. Res. 77, 21-33 (1979).

Cleland, M. R, and N. Wells, in Proceedings of the Second Symposium on Neutron Dosimetry in Biology and Medicine, Vol. I (G. Burger and H. G. Ebert, eds.), EUR-5273, pp. 397-421, Commission of the European Communities, Luxembourg (1975).

Colvett, R D., in Annual Report on Research Project, USAEC Report COO-3243-3, pp. 152-154, U.S. Department of Commerce, Springfield, Va. (1974).

Coppola, M., and J. Bool, Neutron scattering and energy deposition spectra, Radiat. Environ. Biophys. 12, 157-168 (1975).

Cranberg, L., Neutrons from deuteron bombardment of Be and D, Phys. Med. Bioi. 23, 335-336 (1978).

DeLuca, P. M., R P. Torti, G. M. Chenevert, N. A. Detorie, J. R Tesmer, and C. A. Kelsey, Performance of a gas target neutron source for radiotherapy, Phys. Med. Bioi. 23, 876-887 (1978).

Dennis, J. A., in Ion Chambers for Neutron Dosimetry (J. J. Broerse, ed.), EUR-6782, pp. 223-237, Harwood Academic Publishers, London (1980).

Dickens, J. K., G. L. Morgan, G. T. Chapman, T. A. Love, E. Newman, and F. G. Perey, Cross sections for gamma-ray production by fast neutrons for 22 elements between Z = 3 and Z = 82, Nucl. Sci. Eng. 62, 515-531 (1977).

Domen, S. R, Absorbed dose water calorimeter, Med. Phys. 7, 157-159 (1980). Dudley, R. A., in Radiation Dosimetry, Vol. II (F. H. Attix and W. C. Roesch, eds.),

pp. 325-387, Academic Press, New York (1966). Dutreix, A., A. Bridier, and A. Wambersie, in Proceedings of the Third Symposium on Neutron

Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 645-657, Commission of the European Communities, Luxembourg (1978).

Dutreix, J., and M. Tubiana, in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 243-249, Pergamon Press, Oxford (1979).


Dimbylow, P. J., Neutron cross sections and kerma values for carbon, nitrogen and oxygen from 20 to 50 MeV, Phys. Med. Bio[.15, 637-649 (1980).

ECNEU, European Protocol for Neutron Dosimetry for Externa[ Beam Therapy, European Clinical Neutron Dosimetry Group, Br. 1. Radio[., in press.

Edwards, A A, and J. A Dennis, The calculation of charged particle fluence and LET spectra for the irradiation of biologically significant materials by neutrons, Phys. Med. Bio[. 20, 395-409 (1975).

Elson, H. R, T. A. Stupar, A. Shapiro, and J. G. Kereiakes, Measured neutron and gamma spectra from californium-252 in a tissue-equivalent medium, Med. Phys. 6, 426-431 (1979).

Elliott, A, I. Holloway, K. Boddy, J. K. Haywood, and D. Williams, Neutron uniformity studies related to clinical total body in vivo neutron activation analysis, Phys. Med. Bio[. 23, 269-281 (1978).

Ettinger, K. V., A Takavar, and J. R Mallard, in Proceedings of the Third Symposium on Neutron Dosimetry in Bio[ogy and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 689-702, Commission of the European Communities, Luxembourg (1978).

Fidorra, J., and J. Booz, in Proceedings of the Sixth Symposium on Microdosimetry, Vol. I (J. Booz and H. G. Ebert, eds.), EUR-6064, pp. 483-496, Harwood Academic Publishers, London (1978).

Fowler, J. F., in Cu"ent Topics in Radiation Research, Vol. 2 (M. Ebert and A. Howard, eds.), pp. 303-364, North-Holland Publishing, Amsterdam (1966).

Frank, H., J. Seda, and A. Skubal, in Proceedings of the Third Symposium on Neutron Dosimetry in Bio[ogy and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 769-779, Commission of the European Communities, Luxembourg (1978).

Goodman, L. J., and J. J. Coyne, W. and neutron kerma for methane-based tissue equivalent gas, Radiat. Res. 82, 13-26 (1980).

Goodman, L. J., and J. C. McDonald, Measurement of the thermal defect of A-150 plastic, Radiat. Res. 83, 491-492 (1980).

Goodman, L. J., S. A. Marino, and R E. Mills, in Annua[ Report on Research Project, USAEC Report COO-3243-1, pp. 39-51, U.S. Department of Commerce, Springfield, Va. (1972).

Grant, W. H., P. R Almond, V. A Otte, and J. B. Smathers, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (c. Burger and H. G. Ebert, eds.), EUR-5848, pp. 303-314, Commission of the European Communities, Luxembourg (1978).

Graves, R G., J. B. Smathers, P. R. Almond, W. H. Grant, and V. A. Otte, Neutron energy spectra of d(49) + Be and p(41) + Be neutron radiotherapy sources, Med. Phys. 6, 123-128 (1979).

Greene, D., and P. C. Williams, in Proceedings of the Third Symposium on Neutron Dosimetry in Medicine and Bio[ogy (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 635-644, Commission of the European Communities, Luxembourg (1978).

Greene, D., C. H. Gill, and R E. Jones, in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 136-137, Pergamon Press, Oxford (1979).

Guldbakke, S., R. Jahr, M. Cosack, H. Klein, and H. Scholermann, in Proceedings of the Third Symposium on Neutron Dosimetry in Bio[ogy and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 821-834, Commission of the European Communities, Luxembourg (1978).

Guldbakke, S., R Jahr, H. Lesiecki, and H. Scholermann, Neutron response of Geiger-Miiller photon dosimeters for neutron energies between 100 keVand 19 MeV, Hea[th Phys., 39, 963-969 (1980).


Gur, D., J. C. Rosen, A. G. Bukovitz, and A. W. Gill, Fast and slow neutrons in an 18 MV photon beam from a Philips SL/75-20 linear accelerator, Med. Phys. 5, 221-222 (1978).

Hall, E. J., and A. M. Kellerer, in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 171-174, Pergamon Press, Oxford (1979).

Halnan, K. E., in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 251-255, Pergamon Press, Oxford (1979).

Hannan, W. J., D. Porter, R. C. Lawson, and R. Railton, Spectrum measurements in 15 MeV neutron therapy beams, Phys. Med. Bioi. 18, 808-820 (1973).

Harrison, G. H., C. R. Cox, E. B. Kubiczek, and J. E. Robinson, Dosimetric properties of d(80) + Be and d(80) + (Be + Ta) neutrons, Radiat. Res. 74, 1-9 (1978).

Harvey, J. R., and A. J. Mill, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 275-290, Commission of the European Communities, Luxembourg (1978).

Hatanaka, H., and W. H. Sweet, in Biomedical Dosimetry, pp. 147-148, International Atomic Energy Agency, Vienna (1975).

Heintz, P. H., M. A. Robkin, P. Wootton, and H. Bichsel, In-phantom microdosimetry with 14.6 MeV neutrons, Health Phys. 21, 598-602 (1971).

Hendry, G. 0., J. L. Hilton, and J. L. Tom, Neutron source development at The Cyclotron Corporation, Int. I. Radiat. Oncol. Bioi. Phys. 3, 367-372 (1977).

Hess, A., and H. D. Franke, in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 137-138, Pergamon Press, Oxford (1979).

Hess, A., H. K. Kraus, and H. D. Franke, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 181-188, Commission of the European Communities, Luxembourg (1978).

Hillier, M., P. D. Lomer, D. S. Stark, and J. D. L. H. Wood, A 14 MeV neutron tube for radiotherapy, Br. I. Radiol. 44, 716-719 (1971).

Hoffmann, W., G. Moller, J. Rassow, and P. Meissner, Neutron und Pion Dosimetrie mit Hochtemperaturpeaks von Thermolumineszens-Materialen, Medizinische Physik 1,157-165 (1980).

Hogeweg, B., Microdosimetric measurements and some applications in radiobiology and radiation protection, Ph.D. thesis, Radiobiological Institute TNO, Rijswijk (1978).

Hogeweg, B., J. Zoetelief, and J. J. Broerse, in Proceedings of the Sixth Symposium on Microdosimetry, Vol. I (J. Booz and H. G. Ebert, eds.), EUR-6064, pp. 507-516, Harwood Academic Publishers, London (1978).

Hogstrom, K. A., A. R. Smith, P. R. Almond, V. A. Otte, and J. B. Smathers, Computer dosimetry for flattened and wedged fast-neutron beams, Med. Phys. 3, 26-30 (1976).

Horowitz, Y. S., S. Freeman, and D. Dubi, Limitations of the paired LiF TLD-600, 700 technique for the estimation of gamma ray dose in mixed n-gamma radiation fields: the effect of thermal neutrons, Nucl. Instrum. Methods 169, 317-320 (1979).

Hough, J. H., A modified lead attenuation method to determine fast neutron sensitivity ku of a photon dosimeter, Phys. Med. Bioi. 24, 734-747 (1979).

Hubbell, J. H., Photon mass attenuation and mass energy absorption coefficients for H, C, N, 0, Ar and seven mixtures from 0.1 keV to 20 MeV, Radiat. Res. 70, 58-81 (1977).

IAEA, Neutron Activation Analysis, International Atomic Energy Agency, Vienna (1973). IAEA, Nuclear Activation Techniques in the Life Sciences, International Atomic Energy

Agency, Vienna (1979). ICRP, Publication 23, Report of the Task Group on Reference Man, International Commission

on Radiological Protection, Pergamon Press, Oxford (1975).


ICRU, Report lOb, Physical Aspects of Irradiation, International Commission on Radiation Units and Measurements, Washington, D.C. (1964).

ICRU, Report 13, Neutron Fluence, Neutron Spectra and Kerma, International Commission on Radiation Units and Measurements, Washington, D.C. (1969a).

ICRU, Report 14, Radiation Dosimetry: X Rays and Gamma Rays with Maximum Photon Energies between 0.6 and 50 Me V, International Commission on Radiation Units and Measurements, Washington, D.C. (1969b).

ICRU, Report 20, Radiation Protection Instrumentation and its Application, International Commission on Radiation Units and Measurements, Washington, D.C. (1971).

ICR U, Report 21, Radiation Dosimetry: Electrons with Initial Energies between 1 and 50 Me V, International Commission on Radiation Units and Measurements, Washington, D.C. (1972).

ICRU, Report 23, Measurements of Absorbed Dose in a Phantom Irradiated by a Single Beam of X or Gamma Rays, International Commission on Radiation Units and Measurements, Washington, D.C. (1973).

ICRU, Report 24, Determination of Absorbed Dose in a Patient Irradiated by Beams of X or Gamma Rays in Radiotherapy Procedures, International Commission on Radiation Units and Measurements, Washington, D.C. (1976).

ICRU, Report 26, Neutron Dosimetry for Biology and Medicine, International Commission on Radiation Units and Measurements, Washington, D.C. (1977).

ICRU, Report 27, An International Neutron Dosimetry Intercomparison, International Commission on Radiation Units and Measurements, Washington, D.C. (1978a).

ICRU, Report 29, Dose Specification for Reporting External Beam Therapy with Photons and Electrons, International Commission on Radiation Units and Measurements, Washington, D.C. (1978b).

ICRU, Report 30, Quantitative Concepts and Dosimetry in Radiobiology, International Commission on Radiation Units and Measurements, Washington, D.C. (1979a).

ICRU, Report 31, Average Energy Required to Produce an Ion Pair, International Commission on Radiation Units and Measurements, Washington, D.C. (1979b).

lng, H., and W. G. Cross, Spectra and dosimetry related to neutron irradiations of the human body, Phys. Med. Bioi. 20, 906-917 (1975).

Ito, A., in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 605--618, Commission of the European Communities, Luxembourg (1978a).

Ito, A., in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 113-124, Commission of the European Communities, Luxembourg (1978b).

Ito, A., private communication (1980). Johnsen, S. W., Polyethylene filtration of 30 and 40 MeV p-Be neutron beams, Phys. Med. Bioi.

23,499-502 (1978). Kellerer, A. M., and H. H. Rossi, The theory of dual radiation action, Curro Top. Radiat. Res. Q.

8,85-158 (1972). Klein, H., S. Guldbakke, R. Jahr, and H. Lesiecki, The fast neutron sensitivity of a Geiger

Miiller counter photon dose meter by the time-of-flight technique, Phys. Med. Bioi. 24, 748-755 (1979).

Koster, L., A. Breit, G. Burger, H. Kneschaurek, and H. Schraube, in High LET Radiations in Clinical Radiotherapy (G. W. Barelldsen, J. J. Broerse, and K. Breur, eds.), pp. 140-141, Pergamon Press, Oxford (1979).

Kuchnir, F. T., C. J. Vyborny, and S. L. Skaggs, in Biomedical Dosimetry, pp. 107-115, International Atomic Energy Agency, Vienna (1975).


Lawrence, G. A., in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 37-42, Pergamon Press, Oxford (1979).

Lawson, R. C., D. Porter, and W. J. Hannan, Threshold detector measurements with sealed tube DT neutron generators, Phys. Med. Bioi. 17, 771-782 (1972).

Leroux, J. B., Y. Herbaut, J. Rouillon, and J. Giroux, in Proceedings of the Sixth Symposium on Microdosimetry, Vol. 1 (J. Booz and H. G. Ebert, eds.), EUR-6064, pp. 535-548, Harwood Academic Publishers, London (1978).

Lewellen,!' K., W. B. Nelp, R. Murano, H. E. Palmer, and G. M. Hinn, A system for measuring total body calcium in man using the 40Ca(n, a) 37 Ar reaction, Phys. Med. Bioi. 24, 124-135 (1979).

Lewis, V. E., and J. B. Hunt, Fast neutron sensitivities of Geiger-Muller counter gamma dosemeters, Phys. Med. Bioi. 23, 888-892 (1978).

Lewis, V. E., and D. J. Young, Measurements of the fast neutron sensitivities of Geiger-Muller counter gamma dosemeters, Phys. Med. Bioi. 22,476-480 (1977).

Lone, M. A., C. B. Bigham, J. S. Fraser, H. R. Schneider, T. K. Alexander, A. J. Ferguson, and A. B. McDonald, Thick target neutron yields and spectral distributions from 7Li(d/p, n) and 9Be(d/p, n) reactions, Nucl. Instrum. Methods 143, 331-344 (1977).

Maier, E., P. Knesewitsch, and G. Burger, in Proceedings of the Second Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5273, pp. 263-286, Commission of the European Communities, Luxembourg (1975).

Makarewicz, M., and S. Pszona, Theoretical characteristics of a graphite ionization chamber filled with carbon dioxide, Nucl. Instrum. Methods 153,423-428 (1978).

Marion, J. B., and J. F. Fowler (eds.), Fast Neutron Physics, Interscience Publishers, New York (1960).

McCall, R. c., T. M. Jenkins, and G. D. Oliver, Jr., Photon and electron response of silicon-diode neutron detectors, Med. Phys. 5, 37-41 (1978).

McDonald, J. C., I -Chang Ma, J. Liang, J. Eenmaa, M. Awschalom, J. B. Smathers, R. Graves, L. S. August, and P. Shapiro, Calorimetric and ionimetric dosimetry intercomparisons,!. U.S. neutron radiotherapy center, Med. Phys., 8,39-43 (1981a).

McDonald, J. C., I-Chang Ma, B. J. Mijnheer, and J. Zoetelief, Calorimetric and ionimetric dosimetry intercomparisons, II. d + T neutron source at the Antoni van Leeuwenhoek Hospital, Med. Phys., 8, 44-48 (1981b).

McGinley, P. H., and M. Sohrabi, in Proceedings of a Conference on Neutrons from Electron Medical Accelerators (H. T. Heaton and R. Jacobs, eds.), NBS Special Publication 554, pp. 99-107, National Bureau of Standards, Washington, D.C. (1979).

Menzel, H. G., H. Schuhmacher, and H. Blattmann, in Proceedings of the Sixth Symposium on Microdosimetry, Vol. I (J. Booz and H. G. Ebert, eds.), pp. 563-578, Harwood Academic Publishers, London (1978).

Menzel, H. G., and H. Schuhmacher, in Ion Chambers for Neutron Dosimetry (J. J. Broerse, ed.), EUR-6782, pp. 337-344, Harwood Academic Publishers, London (1980).

Meulders, J. P., P. Leleux, P. C. Macq, and C. Pirart, Fast neutron yields and spectra from targets of varying atomic number bombarded with deuterons from 16 to 50 MeV, Phys. Med. Bioi. 20, 235-243 (1975).

Mijnheer, B. J., Determination of the average energy of gamma rays in a collimated 14 MeV neutron field, Radiochim. Acta 24,227-229 (1977).

Mijnheer, B. J., in Ion Chambers for Neutron Dosimetry (J. J. Broerse, eds.), EUR-6782, pp. 125-126, Harwood Academic Publishers, London (1980).

Mijnheer, B. J., and J. J. Broerse, in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 109-115, Pergamon Press, Oxford (1979).


Mijnheer, B. J., and J. R Williams, Determination of absorbed dose and kerma in a neutron field from measurements with a tissue-equivalent ionization chamber, Phys. Med. Bioi., 26,57-69 (1981).

Mijnheer, B. J., J. E. Broers-Challiss, and J. J. Broerse, in Proceedings of the Second Symposium on Neutron Dosimetry in Biology and Medicine, Vol. I (G. Burger and H. G. Ebert, eds.), EUR-5273, pp. 423-443, Commission of the European Communities, Luxembourg (1975).

Mijnheer, B. J., J. Zoetelief, and J. J. Broerse, Build-up and depth-dose characteristics of different fast neutron beams relevant for radiotherapy, Br. I. Radiol. 51, 122-126 (1978).

Mijnheer, B. J., P. A. Visser, V. E. Lewis, S. Guldbakke, H. Lesiecki, J. Zoetelief, and J. J. Broerse, in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 162-163, Pergamon Press, Oxford (1979a).

Mijnheer, B. J., H. Haringa, H. Gorter, and B. F. Deys, in High LET Radiations in Clinical Radiotherapy (G. W. Barendsen, J. J. Broerse, and K. Breur, eds.), pp. 161-162, Pergamon Press, Oxford (1979b).

Mijnheer, B. J., H. Haringa, H. J. Nolthenius, and W. L. Zijp, Neutron spectra and neutron kerma derived from activation and fission detector measurements in a d + T neutron therapy beam, Phys. Med. Bioi., 26, 641-655 (1981a).

Mijnheer, B. J., Tj. Wiederdink, and J. J. Battermann, In vivo dosimetry in the pelvis during fast neutron therapy, Br. I. Radiol., 54, 56-60 (1981b).

Mijnheer, B. J., J. R Williams, and J. J. Broerse, in Biomedical Dosimetry: Physical Aspects, Instrumentation and Calibration, International Atomic Energy Agency, Vienna, in press (1981c).

Miles, J. C., and A. G. Sherwin, Fission detectors as secondary standards for the measurement of fast neutron fiuence, Nucl. Instrum. Methods 146, 503-508 (1977).

Murray, B. W., O. L. Deutsch, and G. L. Brownell, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 379-388, Commission of the European Communities, Luxembourg (1978).

Nachtigall, D., and G. Burger, in Topics in Radiation Dosimetry (F. H. Attix, ed.), Supplement 1, pp. 385-459, Academic Press, New York (1972).

NACP, Nordic Association of Clinical Physics, Procedures in external radiation therapy dosimetry with electron and photon beams with maximum energies between 1 and 50 MeV, Acta Radiol. Oncol. Rad. Phys. Med. 19,1-25 (1980).

Nath, R, K. W. Price, and G. R. Holeman, in Proceedings of a Conference on Neutrons from Electron MedicalAccelerators (H. T. Heaton and R Jacobs, eds.), NBS Special Publication 554, pp. 87-97, National Bureau of Standards, Washington, D.C. (1979).

NBS, Special Publication 554, Proceedings of a Conference on Neutrons from Electron Medical Accelerators, National Bureau of Standards, Washington, D.C. (1979).

Nguyen, V. D., M. Chemtob, B. Hogeweg, and J. J. Broerse, Neutron kerma values derived from neutron energy spectra, to be published (1981).

Oliver, G. D., Jr., W. H. Grant III, and J. B. Smathers, Radiation quality of fields produced by 16,30 and 50 MeV deuterons on beryllium, Radiat. Res. 61, 366-373 (1975).

Otte, V. A., J. Horton, E. Goldberg, D. Tripler, and J. B. Smathers, Skin sparing with d-T neutrons, Br. I. Radial. 50, 449 (1977).

Oxby, C. B., B. Oldroyd, and I. D. McCarthy, Experiences with a small pulsed neutron generator, Phys. Med. Bioi. 25, 137-141 (1980).

Parnell, C. J., A fast neutron spectrometer and its use in determining the energy of some cyclotron-produced fast neutron beams, Br. I. Radiol. 45, 452-460 (1972).

Parnell, C. J., B. C. Page, and E. J. Jones, Fast neutrons produced by bombarding a beryllium target with 40 MeV helium-3-ions, Phys. Med. Bioi. 20, 125-127 (1975).


Price, K. W., G. R. Holeman, and R. Nath, A technique for determining fast and thermal neutron flux densities in intense high-energy (8-30 MeV) photon fields, Health Phys. 35, 341-351 (1978).

Prichard, H. M., A. R. Smith, and P. R. Almond, Silicon diodes as dosimeters in fast neutron therapy, AAPM Q. Bull. 7, 92 (1973).

Puite, K. J., and D. L. J. M. Crebolder, Lyoluminescence dosimetry in photon and fast neutron beams, Phys. Med. Bioi. 22,1136-1145 (1977).

Quam, W. M., S. W. Johnsen, G. O. Hendry, J. L. Tom, P. H. Heintz, and R. B. Theus, Dosimetry measurements of 26,35 and 45 MeV p-Be and p-Li neutron beams, Phys. Med. Bioi. 23, 47-54 (1978).

Reifenschweiler, 0., in Proceedings of the First Symposium on Neutron Dosimetry in Biology and Medicine, Vol. II (G. Burger, H. Schraube, and H. G. Ebert, eds.), EUR-4896, pp. 821-837, Commission of the European Communities, Luxembourg (1972).

Rinard, P. M., and G. G. Simons, Calculated neutron sensitivities of CaF2 and 7LiF thermoluminescent dosimeters, Nucl. Instrum. Methods 158, 545-549 (1979).

Rossi, H. H., in Radiation Dosimetry, Vol. 1 (F. H. Attix and W. C. Roesch, eds.), pp. 43-92, Academic Press, New York (1968).

Rossi, H. H., and G. Failla, Tissue-equivalent ionization chambers, Nucleonics 14, 32 (1956). Rossiter, M. J., V. E. Lewis, and J. W. Wood, Response of thermoluminescent dosimeters to

fast (14.7 MeV) and thermal neutrons, Phys. Med. BioI. 22, 731-736 (1977). Scarpa, G., A. Aceto, M. Moscati, and C. Furetta, in Proceedings of the Third Symposium on

Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 737-746, Commission of the European Communities, Luxembourg (1978).

Schmidt, K. A., and G. Reinhold, The Haefely-GFK fast-neutron generator, Int. 1. Radiat. Oncol. Bioi. Phys. 3, 373-376 (1977).

Schmidt, R., and E. Magiera, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 443-454, Commission of the European Communities, Luxembourg (1978).

Schmidt, R., E. Magiera, and W. Scobel, Neutron and gamma spectroscopy for clinical dosimetry, Med. Phys. 7, 507-514 (1980).

Schraube, H., in Ion Chambers for Neutron Dosimetry (J. J. Broerse, ed.), EUR-6782, pp. 345-350, Harwood Academic Publishers, London (1980).

Schraube, H., A. Morhart, and F. Griinauer, in Proceedings of the Second Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5273, pp. 979-1003, Commission of the European Communities, Luxembourg (1975).

Schraube, H., J. Zoetelief, J. J. Broerse, and G. Burger, in Ion Chambers for Neutron Dosimetry (J. J. Broerse, ed.), EUR-6782, pp. 131-150, Harwood Academic Publishers, London (1980).

Schuhmacher, R., and O. Krauss, in Proceedings of the Third Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 713-725, Commission of the European Communities, Luxembourg (1978).

Schwartz, R. B., in Neutron Standards and Applications (c. D. Bowman, A. D. Carlson, H. O. Liskien, and L. Stewart, eds.), NBS Publication 493, pp. 250-254, National Bureau of Standards, Washington, D.C. (1977).

Shapiro, A., B. I. Lin, J. P. Windham, and J. G. Kereiakes, Transport calculations of gamma-ray flux density and dose rate about implantable californium-252 sources, Phys. Med. Bioi. 21, 509-523 (1976a).

Shapiro, P., F. H. Attix, L. S. August, R. B. Theus, and C. C. Rogers, Displacement correction factor for fast-neutron dosimetry in a tissue-equivalent phantom, Med. Phys. 3, 87-90 (1976b).


Shapiro, P., L. S. August, and R. B. Theus, Computer generation of dose distributions for a fast neutron therapy beam, Med. Phys. 6, 12-20 (1979).

Sherwin, A. G., in Proceedings of the Second Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5273, pp. 341-358, Commission of the European Communities, Luxembourg (1975).

Simmons, J. A., and D. K. Bewley, The relative effectiveness of fast neutrons in creating stable free radicals, Radiat. Res. 65, 197-201 (1976).

Sinclair, W. K., in Biophysical Aspects of Radiation Quality, Second Panel Report, pp. 39-54, International Atomic Energy Agency, Vienna (1968).

Smathers, J. B., V. A. Otte, A. R. Smith, P. R. Almond, F. H. Attix, J. J. Spokas, W. H. Quam, and L. J. Goodman, Composition of A-150 tissue-equivalent plastic, Med. Phys. 4, 74-77 (1977).

Smith, A. R., J. H. Jardine, G. L. Raulson, P. R. Almond, and D. D. Boyd, In vivo measurements of lung corrections'for fast-neutron therapy, Med. Phys. 3, 391-396 (1976).

Smith, M. A., and P. Tothill, Development of apparatus to measure calcium changes in the forearm and spine by neutron activation analysis using californium-252, Phys. Med. Bioi. 24,319-329 (1979).

Spinks, T. J., Effect of size and composition of the body on absolute measurement of calcium in vivo, Phys. Med. Bioi. 24, 976-987 (1979).

Spurny, F., and K. Turek, Fast neutron dosimetry by means of different solid state nudear track detectors, Nucl. Instrum. Methods 147, 143-147 (1977).

Stinchcomb, T. G., F. T. Kuchnir, and L. S. Skaggs, Comparison of the microdosimetric event-size method and the twin-chamber method of separating dose into neutron and gamma components, Phys. Med. Bioi. 25, 51-64 (1980).

Tanaka, S., and Y. Furuta, Revised energy responses of 6LiF and 7LiF thermoluminescence dosimeters to neutrons, Nucl. Instrum. Methods 140, 395-396 (1977).

Thomas, D. J., and V. E. Lewis, Standardization of neutron fields produced by the 3H(d, n) 4He reaction, Nucl. Instrum. Methods, 179, 397-404 (1981).

Todd, P., J. P. Geraci, P. S. Furcinitti, R. M. Rossi, F. Mikage, R. B. Theus, and C. B. Schroy, Comparison of the effects of various cyclotron produced fast neutrons on the reproductive capacity of cultured human kidney (T-l) cells, Int. 1. Radiat. Oneal. BioI. Phys. 4, 1015-1022 (1978).

Ullmaier, H. (ed.), High energy and high intensity neutron sources, Nucl. Instrum. Methods 145, pp. 1-128 (1977).

Van der Kogel, A. J., Late Effects of Radiation on the Spinal Cord, Ph.D. thesis, Radiobiological Institute TNO, Rijswijk (1979).

Wagner, E. B., and G. S. Hurst, A. Geiger-Muller gamma-ray dosimeter with low neutron sensitivity, Health Phys. 5, 20-25 (1961).

Waterman, F. M., F. T. Kuchnir, L. S. Skaggs, G. O. Hendry, and J. L. Tom, Dosimetric properties of neutron beams from the D-D reaction in the energy range from 6.8 to 11.1 MeV, Phys. Med. Bioi. 23, 397-404 (1973).

Waterman, F. M., F. T. Kuchnir, and L. S. Skaggs, Comparison of two independent methods for determining the n/ gamma sensitivity of a dosemeter, Phys. Med. Bioi. 22, 880-888 (1977).

Waterman, F. M., F. T. Kuchnir, L. S. Skaggs, R. T. Kouzes, and W. H. Moore, Neutron spectra from 35 and 46 MeV protons, 16 and 28 MeV deuterons, and 44 MeV protons on thick beryllium, Med. Phys. 6, 432-435 (1979a).

Waterman, F. M., F. T. Kuchnir, L. S. Skaggs, R. T. Kouzes. and W. H. Moore, Energy dependence of the neutron sensitivity of C-C02, Mg-Ar and TE-TE ionization chambers, Phys. Med. Bioi. 24, 721-733 (1979b).

Weaver. K. A., J. Eenmaa, and P. Wootton, Correction factors for neutron dose changes caused by inhomogeneities, Med. Phys. 6, 21-25 (1979).


Wells, A. H., A consistent set of kerma values for H, C, Nand 0 for neutrons of energies from 10 to 30 MeV, Radiat. Res. 80, 1-9 (1979).

Williams, E. D., K. Boddy, and J. K. Haywood, Measurement of neutron fluence in total body in vivo neutrons activation analysis, Phys. Med. Bioi. 22, 1003-1006 (1977).

Williams, E. D., K. Boddy, I. Harvey, and J. K. Haywood, Calibration and evaluation of a system for total body in vivo activation analysis using 14 MeV neutrons, Phys. Med. Bioi. 23,405-415 (1978).

Williams, J. R., Problems in neutron dosimetry associated with incorrect composition of tissue-equivalent gas, Phys. Med. BioI. 25, 501-508 (1980).

Withers, H. R., Biological basis for high-LET radiotherapy, Radiology 108, 131-137 (1973). Wootton, P., and J. Eenmaa, private communication (1978). ZieJczynski, M., M. Makarewicz, S. Marjanska, and A. Wysocka, in Proceedings of the Third

Symposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 597-604, Commission of the European Communities, Luxembourg (1978).

Zoetelief, J., J. J. Broerse, and B. J. Mijnheer, in Proceedings of the Third S)'mposium on Neutron Dosimetry in Biology and Medicine (G. Burger and H. G. Ebert, eds.), EUR-5848, pp. 565-578, Commission of the European Communities, Luxembourg (1978).

Zoetelief, J., A. C. Engels, J. J. Broerse, and B. J. Mijnheer, Effect of finite size of ion chambers used for neutron dosimetry, Phys. Med. Bioi., 25, 1121-1131 (1980).

Zoetelief, J., A. C. Engels, and J. J. Broerse, in Biomedical Dosimetry: Physical Aspects, Instrumentation and Calibration, International Atomic Energy Agency, Vienna, in press (1981).

2

Tissue Inhomogeneity Corrections in Photon-Beam Treatment Planning

J. R. CUNNINGHAM

1. INTRODUCTION

The International Commission on Radiological Units and Measurements (ICRU) in Report 24(1) has suggested that the degree of accuracy required for the delivery of tumor dose in radiotherapy is ±5%. This number is largely based on an interpretation of the steepness of the response-versusdose curve for a rather limited number of tissues. Although there is in fact no unequivocal way of knowing what accuracy is required, this is probably reasonable. There are a number of steps in the delivery of a prescribed dose to a patient. These include (a) the calibration of a dosimeter, (b) its use in dose determination in a water phantom, (c) the calculation of dose in a "water-equivalent" patient and its correction to make allowance for tissue inhomogeneities, and, finally, (d) the actual delivery of the treatment. This chapter deals with the third of these steps.

In Report 24,(1) it is shown, based on an analysis by Loevinger and LoftuS(2) that the uncertainty in delivering a desired dose to a point in a water phantom can be made about as small as ±2.5%, but only under rather optimum conditions. Assuming that the uncertainties are random in character, this would imply an accuracy of about ±4.3% for the combination

J. R. CUNNINGHAM • The Ontario Cancer Institute, 500 Sherbourne Street, Toronto, Ontario, Canada M4X lK9.

103

104 J. R. Cunningham

of the next two steps. If the calculations for a water-equivalent patient have an accuracy of ±2%, this would in turn imply an accuracy of about ±3.8% for the procedures of correcting doses for tissue inhomogeneities. It will be shown that this is possible, but requires both detailed and accurate anatomical information as well as a rather sophisticated calculation procedure.

In an attempt to assess the importance of anatomical information and calculation methods, Sontag et al.(3,4) carried out the following experiment. A chest section of an Alderson Rando phantom was loaded with thermoluminescent dosimeters (TLD) and the assembled phantom was irradiated with a beam from a cobalt unit as indicated in Figure 1. The doses registered by the TLD dosimeters were compared with doses calculated for the same locations by a number of different methods. In the first comparison the doses were calculated assuming water equivalence, that is, ignoring the internal structures. The average absolute difference between calculated and measured values was just under 10% (9.9%); the maximum difference (in the posterior region of the lung) was about 32%.

Next, the internal structures were delineated using an atlas of anatomical cross sections and a simple method was used to make allowances for these structures (the effective attenuation coefficient method). There was

--, ')

I ,-------:------------\

I

/ , , ,

/ t I I. I \

. I \

/'

I Estimated I \ I .,

\/ \ \ \

\

I '-I \

I • \ I I I I I ,

,~! • ," /\ ; )~\\ I I '

I \ I \

• I , .... , \

" I I ~ I \ '" J , " I (~/ ...... ,..!

I

/

0

0 0

0 0

0 0

0 0

\x 0 0 , , , ~~-

, \ I

\

Figure 1. Outline of an Alderson Rando phantom, section 17. The solid circles, crosses, and squares show the location of TLD dosimeters during irradiation by a beam (15 x 15 em 2) from a cobalt treatment unit. The dashed lines are the outlines of lung and bone structures as inferred from an atlas of anatomical cross sections. The solid lines are the outlines of these structures derived by transmission radiography of the section.

Photon-Beam Treatment Planning 105

greatly improved accuracy for those points that were correctly judged to be within the lung, but there was in fact a decrease in the overall accuracy. The average absolute error was now just over 12%, with a maximum of about 42 %. For another comparison the outlines of the lungs, etc. were determined by scanning across the (disassembled) section with a narrow cobalt beam and were thus "correctly" determined. The simple dose-correction method used above now gave an average (absolute) accuracy of 3.4%, with a maximum dose difference of 6.9%. Although this is very good, it is not quite within the 3.8% target for this part of the procedure. As a last comparison, a much more sophisticated dose-correction algorithm was used along with the "correct" anatomical information. The average difference was now only 1.6% and the maximum difference for any point was about 4%.

Next, only the 13 points actually within the lung structures were considered. The simple dose-correction method now gave an average absolute difference of about 5 %, while the more detailed calculation gave an average absolute difference of 2.3%.

The conclusion from this experiment was that in correcting dose calculations for tissue inhomogeneities the most important step is to obtain correct anatomical information and next is the use of an accurate dose calculation method. The importance is in that order, but to meet the overall target of ±5% in dose accuracy, both are required. This chapter deals with methods of allowing for tissue inhomogeneities in dose calculations.

2. SURVEY OF COMMON DOSE· CORRECTION METHODS

Over the years a number of methods have been devised for correcting tissue doses for the presence of inhomogeneities. Some of these are given in Table 1; they are listed in order of increasing complexity. They have also been grouped according to whether or not they take into account the position (with respect to a point of calculation) of a structure and its lateral extent.

All of the methods, with the exception of the last two, are most easily applied if the dose is first calculated under the assumption that there is complete water equivalence and then a correction factor C is determined:

C = Dose in heterogeneous phantom (1) Dose at same place in homogeneous water phantom

All of the methods in the first two groups were originally developed for hand (non-computer) calculations and are discussed at some length in ICRU Report 24. (1) They will be reviewed and compared here.


Table 1. Algorithms for Making Corrections to Absorbed Doses in Heterogeneous Phantoms

path Algorithm length

1. Linear attenuation coefficient Yes

Ratio of TARs Yes Effective SSD Yes Isodose shift Yes

2. Power law TAR (Batho) Yes 3. Equivalent TAR Yes 4. Volume integration of

differential scatter-air ratios (Beaudoin) Yes

5. Monte Carlo Yes

0)

P =\

Method can take into account the

field size

No Yes Yes Yes Yes Yes

Yes Yes

b) I I I I

position of structure

No No No No Yes Yes

Yes Yes

shape of electronic structure equilibrium

No No No No No No No No No No Yes No

Yes No Yes Yes

i Polystyrene

Figure 2. (a) Diagram showing a phantom of unit relative density containing a region of relative density 0.25. Point P is at a depth of 9 cm below the surface and at a depth of 1.4 cm below the low-density region. The water-equivalent or radiological depth to point P is 5.4 cm. (b) Experimental arrangement for comparing simple methods for making corrections to absorbed ck>se calculations to allow for the presence of a large, low-density object in a water phantom. The phantom is made up of a 2.8-cm layer of waterlike material (polystyrene) followed by a 4.8-cm layer of cork (density 0.25 g/cm3 ) and followed again by layers of waterlike material. Relative doses are measured at points along the central ray both with the cork in place and with it replaced by a layer of waterlike material of the same thickness.


As a means of describing and comparing the methods, they will be applied to the situation shown in Figure 2(a). A beam of radiation is incident on a phantom containing a low-density region. The beam is from a cobalt unit and the field size is 10 cm x 10 cm at a source-to-surface distance (SSD) of 80 cm. Point P is at a depth of 9 cm below the surface along the central ray of the beam; it is at a depth of 1.4 cm below the lower boundary of the low-density inhomogeneity. The inhomogeneity is 4.8 cm thick along the central ray and has a density relative to that of water of 0.25. Thus point Pis at a water-equivalent depth of 2.8 x 1.0 + 4.8 x 0.25 + 1.4 x 1.0 = 5.4 cm. This will be referred to as the radiological or water-equivalent depth.

The configuration of Figure 2(a) is simulated by the experimental arrangement shown in Figure 2(b). A phantom is made of a layer of polystyrene (p = 1.04 g/cm3) of thickness 2.8 cm followed by layers of cork (p = 0.25 g/cm3), making a thickness of 4.8 cm, followed by a further thickness of over 10 cm of the polystyrene. Measurements were made with a Farmer-type ionization chamber at a series of points along the central ray both in the inhomogeneous phantom and with the cork replaced by polystyrene to produce a (reference) homogeneous phantom. The ratios of the

1.18

'- 1.14

~ ~ I c:: 1.10 I ~ pOlystyreny

t 1.06 ~

1.02

0.98

o 2 4

polystyrene

~point P

6 8 10 12 14 16 Depth (em)

Figure 3. Graph showing correction factors for points in the phantom of Figure 2(b) when irradiated by a 10 x 10-cm2 beam from a cobalt unit. The solid line was obtained from measurements with a Farmer-type ionization chamber. The dashed lines were obtained from calculations using an effective linear attenuation coefficient. The upper curve is from equation (2a) and the lower from equation (2b). The crosses are from a ratio of tissue-air ratios, equation (3). The circles are from the effective SSD method, equation (4a). The triangles and the dotted line were obtained using the isodose-shift method, equation (6), and the squares from the power law (Batho) method, equation (7).


former to the latter, for each depth, constitute the experimental correction factors [equation (1)] for this phantom. These data are shown as the solid line of Figure 3. They will be compared to correction factors calculated by each of the methods of Table 1. Strictly speaking, one should also allow for the fact that polystyrene has a density slightly above (about 4%) that of water; however, since its electron density (number of electrons per unit volume) is very close to that of water, corrections for depth in polystyrene are ignored in this experiment.

2.1. Methods Based on Water-Equivalent Depth

2.1.1. EFFECTIVE ATTENUATION COEFFICIENT METHOD

For point P of Figure 2 the (water-) equivalent depth is 5.4 cm (see above), while the actual depth is 9 cm. Because of the low density of the cork, there is the equivalent of 9.0 - 5.4 = 3.6 cm of water missing. This would give rise to a higher dose at point P because of the lower attenuation. In this simple method it is assumed that the dose is increased by a certain fraction for each centimeter of water-equivalent material missing. This quantity plays the role of an attenuation coefficient for the radiation beam. Suggested values for a number of beam energies are given in the third column of Table 2. Using this method, the correction factor for the dose at a point such as P would be

C = ell-'(d-d') (2a)

For point P of Figure 2(a), p,' is 0.05 (from Table 2 for cobalt radiation), d is 9.0 cm, and d' is 5.4 cm. Thus C = e O.05x3.6 = 1.197. Correction factors calculated by this method for points along the beam axis for the phantom of Figure 2(a) are plotted as the upper dashed curve of Figure 3. This procedure, as can be seen, does not take into account the field size or depth.

Table 2. Parameters Useful for Correcting Dose Data for Tissue Inhomogeneities·

SSD Effective attenuation Fractional Radiation (em) coefficient J.L isodose shift

HVL-2mmCu SO 0.105 0.7 60CO 80 0.05 0.67 4 MeVp 100 0.04 0.65 25 MeVp 100 0.02 0.5

"From ref. 5, Table X-6.


Frequently the calcula,tion is carried out as a linear rather than exponential correction:

C = 1.0 + JL'(d - d') (2b)

This is the first term of the expansion of the exponential. It gives a slightly smaller correction than does equation (1) and is shown as the lower dashed curve of Figure 3. For our point P the factor would be C = 1.0 + 0.05 (3.6) = 1.18.

2.1.2. RATIO OF TISSUE-AIR RATIOS

A slightly more sophisticated correction factor may be obtained by using a ratio of two tissue-air ratios:

C = T(d', Wd )

T(d, Wd ) (3)

where, as before, d is the depth and d' is again the water-equivalent depth. Wd represents the dimensions of the cross section of the beam at depth d. The size of the beam of Figure 2(a) is 10 x 10 cm2 . The source-to-surface distance is 80 cm and the beam size at point P is 10 cm x 89 cm/80 = 11.125 x 11.125 cm2 • The tissue-air ratio for this beam for a depth of 9cmisO.759andO.898fordepth5.4cm,givingC = 0.898/0.759 = 1.183. Correction factors calculated by this method for the phantom of Figure 2(a) are shown as crosses in Figure 3. This method of calculation takes into account to some extent both the field size and depth through the use of the tissue-air ratio. It does not take into account the lateral dimensions of the inhomogeneity nor its position with respect to the point of calculation.

2.1.3. EFFECTIVE SSD METHOD

The effective SSD method is entirely equivalent to the ratio of tissue-air ratios method but makes use of percentage depth doses rather than tissueair ratios. It can be most easily described by referring to Figure 4. This diagram shows an isodose chart superimposed on the phantom of Figure 2(a). A dose-correction factor is to be calculated for point P.1f the phantom were homogeneous, the percentage depth dose could be read off directly as 59.7. Let this be represented symbolically by P(d, Wo, F), where d is the depth to point P, Wo represents the field size, and F is the source-to-surface distance (SSD). To take into account the inhomogeneity, the isodose chart is "slid" down so that the depth to point P is equal to the water-equivalent depth d'. This is indicated by the dashed lines in Figure 4. The physical justification for this move is that the phantom material is considered to be compressed (or expanded) so that it is homogeneous and water equivalent.


Figure 4. Diagram depicting the application of the effective SSD method to the calculation of dose at points such as P and Q in the inhomogeneous phantom of Figure 2(a).

Symbolically, the new percentage depth dose would be P(d', Wb, F'), where d' is the water equivalent depth, Wb is the field size, which should be adjusted slightly for the divergence of the beam in going from F to F', and P' is the new (effective) source-lO-surface distance. For our point P, the water-equivalent depth d' is 5.4 cm. Normally, in using this method one does not alter the percentage depth dose to allow for the divergence of the beam, and the dashed isodose lines of Figure 4 are obtained directly by moving the solid isodose lines downward a distance d - d' = 3.6 cm. For point P, the new percentage depth dose can be read off as 76.5. This value is, symbolically, P(d', Wb, F). It should be adjusted, however, so that it takes into account the new source-to-surface distance F' rather than F. This involves inverse-square corrections and would be

( F+d')2(F'+d )2 P(d', W o, F') = P(d', W o, F) F + dm F' + d7

The percentage depth dose is expressed as a percentage of the dose at the reference point (at depth dm ), and this point has been moved from distance F + dm to F' + dm • The correction factor thus becomes

C = P(d', Wo,F') (F + dm )2 P(d, Wo, F) F' + dm

where F' + d' = F + d.

= P(d', W o, F) (F + d')2 P(d, Wo,F) F + d

(4a)


For point P the correction factor is evaluated to be (76.5/59.7)· (85.4/89)2 = 1.18.

Correction factors for points along the axis of the phantom of Figure 2 are shown as open circles in Figure 3. They are very close to the values obtained (crosses) using the ratio of tissue-air ratios method.

By using the relationship between tissue-air ratios and percentage depth doses, which in this notation is

T(d, Wd) (F + d m)2 P(d, Wo, F) = 100 T(dm , Wdm ) F + d

it can be shown that the correction factor of equation (4a) reduces to

c = T(d', Wd)( T(dm , Wdm )) T(d, Wd ) T(dm , W dm )

(5)

(4b)

which differs from equation (3) only by the term in parentheses. This term is a ratio of backscatter factors and corrects for the (neglected) divergence of the beam discussed above and takes into account the differences between the crosses and circles of Figure 3.

Although the two methods, the effective SSD method and the ratio of tissue-air ratios method, can be shown to be formally the same, at least on the central ray, they may be different in practice. This is because one frequently manipulates an isodose chart, which can be "slid" along ray lines, as indicated in Figure 4 for point Q. This makes a difference for points such as Q that are off the central ray and the ratio of percent depth doses used in the correction factor [equation (4a)] in this way takes into account the lateral position of the point of calculation.

In the computer environment, storage of beam data on ray line and depth coordinate systems makes this method of correction for inhomogeneities (including surface curvature) very convenient.

2.1.4. THE ISODOSE-SHIFT METHOD

The isodose-shift method is essentially a simplification of the effective SSD method and was introduced largely for the manual manipulation of isodose lines. In its application the isodose lines are moved only a fraction n of the distance d - d' and the corrected-dose values can be read off directly, eliminating the need for the inverse-square term of equation (4a). The value of n depends on the beam energy. Selected values for n are given in column 4 of Table 2. The correction factor using this method can be expressed as

where d" = d - n(d - d').

c = P(d", Wo, F) P(d, Wo,F)

(6)


Calculated correction factors using this method for points along the central ray of the phantom of Figure 2 are plotted as triangles and a dotted line in Figure 3.

2.2. Power Law Tissue-Air Ratio Method

A correction factor involving tissue-air ratios raised to a power that depends on density was proposed by Batho(6) and Young and Gaylord(7) for points lying below an inhomogeneity. It was generalized by Sontag and Cunningham(S) to include points within an inhomogeneity. The general form of the expression for the correction factor is

T(dl, Wd )Pl-P2

C = T(d2 , wd P2 (7)

where PI is the density of the material in which the point of calculation lies and d l is the depth at which it is within this material. P2 is the next or overlying material and d2 is the depth below its upper surface. This can be made clearer by an example. For point P in the phantom of Figure 2, PI = 1.0, d l = 1.4, P2 = 0.25, d2 = 6.2, T(1.4, 11.125) = 1.0245, T(6.2, 11.125) = 0.868, and therefore C = (1.0245)°·75/(0.868)°·75 = 1.132. For a point inside the inhomogeneity, say, at depth 6 cm, d l = 3.2, PI = 0.25, d2 = 6 and P2 = 1.0, T(3.2, 10.75) = 0.972, T(6, 10.75) = 0.873, and C = (0.972)-0.75/(0.873)° = 1.022. Calculated correction factors using this method for points along the axis of the beam shown in Figure 2 are plotted as open squares in Figure 3. It can be seen that this method gives results that are markedly different from all of the previously discussed methods but they are in considerably better agreement with experimental data. This will be discussed further in a later section.

This method of correction is categorized separately in Table 1 because it takes into account one more factor, namely, the position of the inhomogeneity with respect to the point of calculation. In this sense the configuration of scattered photons is considered. This shows up in ways that may be somewhat surprising at first sight. For example, as seen from Figure 3, for points just inside the low-density structure the correction factor drops below 1.0-as low as 0.97 in our example. All the previously mentioned correction methods predict correction factors greater than 1.0 in low-density regions. The Batho result is reasonable from a physical point of view because at such points there will be an increase in dose due to primary radiation, but this may be more than compensated for by the decreased amount of scattered radiation. The experiments confirm this behavior. The method, as it stands, does predict a discontinuity in the dose that must not be real. This can be seen, for example, by considering points (say at a depth of 2.5 cm in the Figure 2 phantom) that are not overlain by an inhomogeneity. A correction


factor of 1.0 is predicted, yet there must be a reduced amount of scatter coming from the low-density region immediately beyond. This is observed in the experimental determination, as shown in Figure 3. There is a similar discontinuity at the bottom of the low-density region, where there will be photons scattered back from the high-density region below it, and this is not taken into account either.

A word of caution is required in the application of equation (7) so near to an interface that d1 is less than the buildup distance for that energy. In this case a value for the tissue-air ratio determined by extrapolating data from greater depths must be used.

2_3. The Equivalent Tissue-Air Ratio Method

Several of the methods just described use ratiQs of tissue-air ratios. One of them uses such ratios directly. It was shown that the effective SSD method reduces to a ratio of tissue-air ratios and another uses tissue-air ratios raised to a power. It can be seen, by reference to Figure 3, that for the phantom and beam size chosen, the power law correction factors were, by a considerable margin, the closest to the experimentally determined values. Nevertheless, one can choose situations where this agreement is not so good. An example of this is shown in Figure 5, which is a repeat of Figure 3 but for a larger field size. Correction factors obtained by ratios of tissue-air ratios are plotted as crosses, as squares by the power law method, and the experimentally

1.18

1.14 I I I

... POlystyren~ .!:? 1.10 I

~ t: ~ 1.06 ~ :::: a 1.02

0.98

b C

o 2 4 6 8 10 12 14 16 Depth !em}

Figure 5. Graph showing correction factors for points along the axis of the phantom of Figure 2(b) when irradiated by a 20 x 20-cm2 beam from a cobalt unit. The crosses are obtained for the ratio of tissue-air ratios method, the squares from the power law tissue-air ratio method, and the solid line is from experiment. The solid circles are from the equivalent tissue-air ratio method.


determined correction factors are again shown as a solid curve. The power law correction factors agree well with experiment for points below the inhomogeneity (this is the region for which Batho originally derived it), but disagree for points within it. In fact, within the low-density region, the simple ratio of tissue-air ratios gives results that are almost as good. For much larger field sizes, such as are used for half and total body irradiation, Van Dyk et al.(9) have shown that errors as great as 20% may result from the use of the "generalized Batho method" within lung tissue. It would thus appear that for small field sizes the power law tissue-air ratio method is very good, while for very large field sizes the simpler ratio of tissue-air ratios method is better. Neither method takes into account the three-dimensional shape of structures.

Sontag and Cunningham(3) have suggested yet another correction factor using ratios of tissue-air ratios. It is a variation of the simple ratio of tissue-air ratios but in its numerator not only is the depth altered to take into account inhomogeneities but so is the field size. They have called it the "equivalent tissue-air ratio method." The alteration of the field size follows from the idea, first proposed by O'Connor,(10) that a beam irradiating a homogeneous but non-water-equivalent phantom is equivalent to a beam irradiating a water phantom with all linear dimensions such as depth and field size scaled in proportion to the density (or electron density) of the non-water-equivalent material.

The physics of the equivalent tissue-air ratio method will be discussed in some detail here, not only because its application is relatively new and has been shown to be accurate and practical but also because its examination gives a useful insight into the interaction of a broad beam of radiation with an inhomogeneous medium.

Consider a beam of radiation incident on a phantom such as shown in the left of Figure 6. If we consider primary and once-scattered photons only, a tissue-air ratio could be calculated from the following expression, letting the radiation be monoenergetic and the fluence <1>:

d -I-"b ,

<I>(e-l-'d~Eab+ f e-l-'an~~~E~bdV) T(d, Wdh = p v dw b p (8a)

<I>(e-l-'d~~Eab+ IdS)

The meanings of some of the parameters used in this expression are illustrated in Figure 6. The first term in the numerator is the absorbed dose from primary photons that are attenuated along a path of length d. IL is the linear attenuation coefficient and ILl p is the mass attenuation coefficient for these photons. Eab is the average energy absorbed as a result of each photon interaction. The second term is the integration of the dose from oncescattered photons over the irradiated volume. The distance a is the depth to

Photon-Beam Treatment Planning

I hv

! I I I

lIT ! ~bJ !

-.., €=2 Q t- r I I

0)

20

E = I --. I I

b)

I hv

~

2d

2b

e

Q'

115

I I I I I I I I I I I I I I I I I I I I 1-2r I I I

Figure 6. Parameters used in the calculation of absorbed dose from primary and once-scattered photons. The diagram on the left shows a beam irradiating a phantom whose electron density relative to that of water is E = 2. The diagram on the right shows conditions that would produce equal doses at points Q and Q'.

the site of scattering, n is the number of electrons per unit volume in the phantom, dul dw, is the (Klein-Nishina) cross section for scattering a photon through an angle f) expressed per electron and per unit solid angle. IL' is the linear attenuation coefficient for the scattered radiation and b is the path length from the scattering site to point Q. p,' I p is the mass attenuation coefficient for the scattered radiation and E~b is the average energy absorbed from each interaction of the scattered photons.

The denominator in this expression is the absorbed dose from primary and once-scattered photons in a small (reference) mass of phantom material of radius dm in air. The integral Iuds is exactly like the integral in the numerator but the integration is over the small (reference) volume v. The denominator, if divided by (p,I p )Eab, would be numerically equal to the factor Aeq discussed by Johns and Cunningham.(S)

To illustrate the scaling procedure, let equation (Sa) be applied to the beam on the left of Figure 6. A circular beam of radius r is irradiating a material of electron density e relative to that of water. For simplicity we have assumed that the beam does not diverge, although this is not necessary. Next let it be assumed that the attenuation coefficients of the material can be expressed in terms of the attenuation coefficients of water as IL = ILwe etc.


Similarly, n = nwE, while 1-'1 p, 1-" 1 p, Eab, E~b' and Aeq would not be altered. The exponents would then be I-'wEd, I-'wEa, and I-' ~Eb. Since the lines a and b can be oriented in all directions, this procedure suggests the scaling of all linear dimensions by E, in which case I-'wEd = I-'wdw, etc. Also, we may substitute for the volume element dV = dVw1E3. Rearranging the factors slightly allows us to write

) ( ) 1 ( -I'- d 1 T(d, r 1 = T dw, Ww 1 = Aeq e w w + (1-'Ip)Eab

(8b)

Equation (8b) was obtained from equation (8a) by substitution of variables. The same result would have been obtained directly by scaling all linear dimensions in equation (8a) by the factor E. This derivation has been applied to once-scattered radiation only, but since the procedure can be interpreted as a linear coordinate transformation and all angular relationships are preserved, it would also apply to mUltiply scattered photons. This procedure does require, however, that there be some constant factor E that relates all attenuation coefficients in one material to those of the other. It can therefore only be applied accurately in the energy range where Compton processes are dominant.

The equivalent tissue-air ratio method for correcting for inhomogeneities operates on the assumption that at any point in an inhomogeneous phantom it is possible to find an "equivalent" homogeneous phantom, with relative electron density e, and that a correction factor can be obtained accordingly using an "equivalent" tissue-air ratio

c = T(d',;) T(d, r)

(9)

In this factor, d is the depth of the point of calculation and r is the radius of an equivalent circular field as applied to the patient. (For equivalent circular fields see Johns and Cunningham (5) or Supplement 11 of the British Journal of Radiology(1l).) d' and; are the scaled versions of these two parameters.

The problem in applying the method is that of choosing the proper scaling procedure for the two parameters. Sontag and Cunningham(4) have suggested that primary and scatter components be considered separately. That is,

T(d',;) = T(d', 0) + S(d',;) (10)

The proper depth to use for the primary component is clearly the radiological, or water-equivalent, depth, and d' as used above has this meaning. It is


not so clear, however, that this same depth should be used in the scatter component, since the effect of scattering structures must depend on their geometrical arrangement. Sontag and Cunningham(4) have chosen to account for these geometrical factors entirely through the scaling of the field size. The method of determining the scaling factors will be discussed later, but first it will be useful to discuss some of the implications of a few of the experimental points shown on Figure 5, referring to the phantom of Figure 2(b).

Consider first a point at a depth of 2.5 cm. It is within the waterequivalent material but there is a large low-density structure below it but near to it. This would not alter the primary component but would decrease some of the radiation scattered back to the point. The experimentally determined dose-correction factor is 0.973. It will be useful to determine what degree of scaling to the field size is required to account for this. Let the radius of a circular beam that is equivalent to this (20 x 20-cm2) field be r = 11 cm. The denominator of the correction factor is T(2.5, 11) =

(0.875 + 0.150) = 1.025. The scatter term that would give the observed correction factor would be 0.122, which implies an equivalent field size, of ; = 7 cm, that is, T(d', ;) = T(2.5, 7). This is perhaps surprisingly small, but not unreasonable. Consider next a point well within the low-density structure, at a depth of 7 cm. The radiological depth to this point is d' = 2.8 + (7.0 - 2.8)0.25 = 3.85 cm, with a zero-area tissue-air ratio T(3.85, 0) = 0.800. The observed correction factor from Figure 5 is 1.057 and so C = 1.057 = [0.800 + S(3.85, ;)]/0.889, giving S(3.85,;) = 0.140, implying an equivalent field radius ; = 5.43. This field size, as would be expected, is considerably smaller than the one obtained for the previous point. It is still considerably larger, however, than 11.0 x 0.25 = 2.75 cm, which would be the limiting value for; in a region where the relative density is 0.25. Consider two more points, first at a depth of 9 cm, where the radiological depth is 5.4 cm and the measured correction factor is 1.110. The inferred effective beam radius is now 6.54 cm, which shows it is increasing again. Finally, at a depth of 15 cm it is 10.3, almost back up to the actual field size.

For the phantom shown in Figure 2(a) the radiological d~pths would be the same as those for Figure 2(b), but the alterations in the effective field size would be less, allowing for the fact that the lateral dimensions of the inhomogeneity do not extend to the edges of the field.

USE OF CT

The equivalent tissue-air ratio method lends itself particularly well to input using computed tomography (CT) information. In this context Sontag(12) and Sontag and Cunningham(4) define the (scaled) beam radius as

118 .J. R. Cunningham

follows:

r = re

with

(11)

where the eijk are the (relative) electron densities of pixels in a series of CT images of the irradiated volume. The Wijk form a set of weighting factors, which are intended to express the relative importance of each of the eijk

elements in affecting the dose due to scattered radiation at the point of calculation.

There is no unique or correct set of weighting factors for use in equation (11) because they would be slightly different for each point of calculation and each particular configuration of tissues. They would have a form, however, that is similar to the "isoeffect" curves shown in Figure 7. These lines show the results of an experiment carried out by Andrew et al. (13) in which an ionization chamber was positioned as shown in a water tank and readings were taken with and without polystyrene foam rings of various radii placed at various depths. The curves indicate the change of dose when a unit volume (1 cm3 ) of water is replaced by air. The numbers, when divided by 1000, express the percentage change. The curves show that replacing water by air at some locations causes an increase in the dose, while at other locations it causes a decrease.

A direct evaluation of equation (11) would imply an integration over the entire irradiated volume for each point of dose calculation. As this was not considered to be practical, a compromise procedure was adopted that resulted in an integration only over an area. This procedure can be described by referring to Figure 8. On the left of this diagram is depicted information as contained in six CT slices. For the example, dose calculations are to be made in the plane of slice 3, which is shown shaded. As a first step all of the density information contained in the six slices is coalesced into a single "effective" slice, as shown on the right side of the diagram. This is done by assigning a weighting factor Wk to each slice, with the result that the "coalesced" slice consists of the (thus) weighted average of all pixels that have the same i and j indices:

(12)

Photon-Beam Treatment Planning

0

2

4

6 ~

~

" '- 8 ~ ~ 1:::110

12

14

16

18

12 10 8 6 4

Figure 7. Isoeffect curves for scattered radiation in a water tank. An ionization chamber is positioned as shown and readings taken with and without rings of polystyrene foam displacing water. The curves show the effect of replacing 1 cm3 of water with air and become percentages when divided by 1000. The radiation is fro~ 60CO.

--+--z

Figure 8. Schematic diagram illustrating the steps used in the equivalent tissue-air ratio method to approximate integration over a volume by an integration over a plane. The density information contained in six CT slices shown on the left is first coalesced into an effective single slice, as shown on the right. This slice is a weighted average of all six slices; the ones close to the plane of dose calculation (shaded) count most strongly. As a final step, a weighted average electron density is formed for each point of dose calculation.


The form chosen for the weighting factors Wk is shown in Figure 9. The "coalesced" plane is considered to be at an effective distance Zeff from the plane of calculation. This step involves a number of approximations and assumptions. Its experimental implications are discussed at some length by SontagY2)

The next step, which is to determine the weighted average density and thus the equivalent beam radius and consequently the equivalent tissue-air ratio, is performed for each of the points at which dose calculations are made:

L L eijWij(Zeff)

; = ri = r_i--,-j ---L L Wij(Zeff) i j

(13)

The procedure that was adopted by Sontag and Cunningham(4) for the evaluation of Wij(Zeff) is described briefly in their paper and a practical

30

25

15

10

5

Weighting factors for CT slices in Z direction

o I 2 3 4 5 6 7 8 9 10 II 12 13 14 15 Distance from plane of calculation(cm)

Figure 9. Density weighting factors for CT slices. They are obtained by calculating differences of scatter-air ratios. Zeft is the "effective distance" for the coalesced slice (see Figure 8) averaged over the irradiated volume.


method for determining values for the weighting factors Wij(Zeff) is suggested. It was also shown by Sontag(12) that the exact form of these weighting factors is not very critical. They must at least approximate the form shown in Figure 7.

This method is an improvement on all the previously mentioned methods. It does, at each point, take into account, to some extent, the density and size of all structures in an irradiated medium as well as their position with respect to the point of calculation. The shape of the external contour on the exit side of the beam is treated like an inhomogeneity. This method was applied to the experimental phantom and the results for a 20 x 20-cm2 cobalt beam are shown in Figure 5 by the solid circles.

Numerous other experiments, some of which are described by Sontag and Cunningham (4) and others by Sontag, (12) indicate that a general average accuracy of about ±2.5% can be expected. It appears to be applicable to high-energy beams (25 MeVp , for example) as well as to cobalt radiation. The tests of accuracy are based on measurements both in anatomical phantoms and geometrical phantoms specially designed to "fool" the method. It, like any method short of Monte Carlo methods, does not take into account losses of electronic equilibrium near interfaces between structures. This probably constitutes the major remaining shortcoming of this technique.

2.4. Volume Integration of Differential Scatter-Air Ratios

As long ago as 1968 Beaudoin(14) showed that it was possible to use scatter-air ratios, cast into a differential form, to calculate the absorbed dose at a point in an inhomogeneous medium. Cunningham(lS) showed how scatter-air ratios could be derived from tissue-air ratios and how the scatter-air ratio could in tum be differentiated so as to produce a quantity describing the total amount of scattered radiation reaching a point from a volume element. The process can be described by referring to Figure 10. A beam of radiation is incident on a water phantom. The beam is circular in cross section and of radius R. The dose at point P at depth d due to scattered radiation is given by

(14)

DAis the dose in air (to a small mass of phantomlike material) at point P, but in the absence of the rest of the phantom. The scatter-air ratio S(d, R) is an empirical quantity that expresses the total amount of scattered radiation reaching point P (at depth d) from all of the irradiated volume. This radiation includes once- and multiply scattered photons.

For calculation of dose in an inhomogeneous phantom it would be useful to have a measure of the scatter that arises in a volume element such as


R

Figure 10. Diagram showing the param~ters involved in deriving differential scatter-air ratio data. Point P is at a depth d on the axis of a circular beam of radius R. Volume element dV is a distance z in front of P and is at a radius r away from the axis. Its volume is given by r ac/> ar az and the differential scatter-air ratio for this value of z and r is a measure of the radiation that interacts first in this volume and eventually reaches point P.

dV in Figure 10, and reaches point P. This can be obtained by the following procedure, which is analogous to differentiating S(d, R) with respect to depth, radius, and an angle 4>.

The first, and easiest, step is to use the symmetry of the circular beam and to consider only a sector of it. If the angular width of the sector is 1l4>, then the scattered radiation from such a sector is

IlS = ~S(d R) 1l4> 217 '

(15)

Such a sector is indicated in Figure 10. Next consider a circular beam of radius'1 and its scatter-air ratio S(d, '1), and next a smaller beam of radius'2 and its scatter-air ratio S(d, '2). The difference should be a measure of the amount of scattered radiation coming from the annular space between the beams:

IlS S(d, '1) - S(d, '2)

Il, '1 - '2 (16)


These two steps can be combined to give an expression for the amount of scatter coming (to point P) from a columnar volume element extending from the surface to great depths but having a cross section, at depth d, given by r tlr tl4>:

tl2S(d, r) S(d, r) - S(d, r - tlr)

r tl4> tlr 21Tr tlr (17)

The next step in the process is not as direct but can be described by referring to Figure 11. First, let the depth (to point P) be d 1. The scatter-air ratio will be S(dl, r). Next imagine that the depth is decreased to d2 • The scattered radiation reaching point P will be changed in two ways: Some of it will be missing, but at the same time scatter from all of the remaining material will be increased. The increase should be the inverse of the attenuation of the primary that occurred in the thickness d 1 - d2 • This attenuation can be expressed by a ratio of zero-area tissue-air ratios T(d2, 0)/T(d1, 0). The scattered radiation that comes from the layer of

Figure 11. Diagram showing the parameters used in differentiating scatter-air ratios with respect to depth as part of the process of forming differential scatter-air ratios.

bod

~r -----11 I

I


material of thickness d l - d 2 to a point at depth d is therefore given by

6.S 1 ( T(d l , 0) ) 6.d = d 1 - d2 S(dt, r) - T(d2 , 0) S(d2 , r)

= T(dt, 0) (S(dt, r) _ S(d2 , r)) d 1 - d2 T(dt, 0) T(d2,0)

(18)

This process may be extended to apply to a more generalized layer, such as 6.z of Figure 11. This gives the amount of scattered radiation reaching point P (at depth d) from a scattering layer a distance z in front of it:

6.S(d, z, r) = T(d, 0)( S(z, r) _ S(z - 6.z, r))

6.z 6.z T(z,O) T(z - 6.z, 0) (19)

The use of zero-area tissue-air ratios for the attenuating factor here cannot be strictly correct because it assumes that all radiation coming from above the (scattering) layer is attenuated like primary radiation. However, to the goodness of this approximation, we may combine all three steps to produce a statement of the relative dose at a point due to scattered radiation that originated in a unit volume such as dV of Figure 10:

6.S = 6.3S(d, z, r) = T(d, 0) (S(z, r) - S(z, r - 6.r)

6. V r 6.4> 6.r 6.z 27Tr 6.r 6.z T(z, 0)

_ S(z - az, r) - S(z - az, r - ar») T(z - 6.z, 0)

(20)

It is useful to note that the differentiation with respect to depth produces a quantity, inside the parentheses, that is independent of depth and therefore forms a quantity that could, like scatter-air ratios themselves, be tabulated for a given radiation quality.

The practical accumulation of numerical values for 6.S/6. V is in fact not easy, because of the inevitable imprecision of the original measurements of the tissue-air ratios. We are dealing here with second-order differences and smoothing procedures must be employed. Whatever data processing takes place, however, must be subject to the boundary condition

f d rR 6.S S(d,R)=27T -ooJo 6.Vrdzdr (21)

That is, the integration of the differential scatter-air ratio over the entire irradiated volume must equal the scatter-air ratio for that beam. The integration is over radius and depth but not angle because of the symmetry conditions of the circular beam.

A set of values for 6.S/6. V for cobalt radiation is shown in graphical form in Figure 12. In producing these data considerable smoothing was


.2 .2

3 2 o 234

Radius r (em)

Figure 12. Graphical presentation of differential scatter-air ratio data. The lines connect volumes from which the amounts of scattered radiation (per unit volume) reaching point Pare equal. This diagram is somewhat similar to Figure 7. The numbers represent the depthindependent part of equation (20).

carried out. The lines in Figure 12 connect regions that produce equal doses from scattered radiation at point P. The shape of these "isoscatter" lines is somewhat similar to the "isoeffect" lines of Figure 7 but their meaning is quite different. The isoeffect data of Figure 7 indicates the change in dose (at a point) that is brought about by replacing a unit volume of water by air, while the isoscatter data of Figure 12 indicates the dose due to radiation that is scattered from a volume element in a homogeneous phantom and ultimately reaches a point.

The integration in equation (21) is analogous to the second term in equation (8b), which is an analytical expression, using once-scattered photons only for the same quantity, the scatter-air ratio.

To make calculations for a nonhomogeneous medium, in analogy to the procedure for describing once-scattered radiation, it is necessary to know the nature of the material at the site of the scattering volume so that the scatter term given by equation (20) can be assigned a weight. A reasonable value for such a weighting factor would be £ the relative electron density at the scattering site. Using differential scatter-air ratios, the dose at a point in an inhomogeneous medium could then be calculated by

(22)

DAis the dose in air at the location of the point of calculation and the expression inside the parentheses is the equivalent of a tissue-air ratio for


the inhomogeneous medium. The first term is the zero-area tissue-air ratio and the second is the integration of the scatter, here replaced by a sum. The depth d' is the radiological depth to the point of calculation. a I is the radiological path length to the scattering volume and b' is the radiological distance from the scattering volume to the point of calculation. The first exponential represents the difference in the attenuation along the path to the scattering volume due to non-water-equivalent material. The second exponential represents the difference in attenuation of the scattered radiation. /.A- I is the linear attenuation coefficient for the scattered photons.

The evaluation of equation (22) implies a summation over all of the irradiated volume for each point of calculation. It also implies the implementation of a ray-tracing procedure to examine the material along the paths a and b for each volume element used in the summation. This is not yet a practical procedure for routine dose calculations in treatment planning. Beaudoin(14) explored some of its possibilities and showed good agreement with experiment for the conditions he tested. Sontag(12) showed that the method did not give good results for large beams in non-water-equivalent but homogeneous phantoms, presumably because multiply scattered radiation is handled as if it were once-scattered radiation. Larson and Prasad(16) have explored with considerable success the possibility of carrying out the complex summation procedures using macromodular computer systems.

2.S. Monte Carlo Calculations

In principle, Monte Carlo calculations should be capable of givihg the best possible answers to the problem of making dose calculations in homogeneous beams. The essence of the method is that a random sample of photon paths is generated and used to predict the behavior of the large number of photons that constitutes a radiation beam. The problem is that a very large number of such photon "histories" are required.

The probability that a photon will travel a particular path is determined by an appropriate expression involving a random number. For example, suppose a photon has a given energy. The probability that it will go a distance x before interacting and then interact between x and x + ax is given by

p, the probability, must be between 0 and 1, if p is chosen as a random number between these limits, a value for this path length x can be calculated:

1 x = --lnp

/.A-

Assuming that the interaction is by the Compton effect, one must next select


on a random basis the energy of the scattered photon, and thus its scattering angle fJ, and then a random azimuthal scattering angle <p. The process is continued for further interactions until the photon is absorbed by the (randomly chosen) photoelectric effect or else exits from the phantom. An extensive discussion and bibliography concerning Monte Carlo methods for photon interactions is given by Raeside. 17 In order to describe the spatial pattern of radiation dosage it is necessary to keep track of the locations of all of the reactions considered and to continue generating such photon histories until, at a point of calculation, the number of photons contributing to the dose is no longer limited by statistical uncertainty. Webb/8 for example, showed that to limit the uncertainty in the dose distribution for a 20 x 20-cm2 beam of cobalt radiation to ±2% it was necessary to follow the histories of some 1.8 x 107 photons.

Such an effort is entirely possible with current computer technology but is not practical for routine treatment planning. Monte Carlo calculations are practical, however, for studying the behavior of photon interactions in complicated situations. They form an excellent substitute for, or an extension of, experiments that are difficult to carry out.

The absorbed dose connected with photon interactions is actually a result of electrons that are set into motion by the photon interactions. It is possible to extend Monte Carlo calculations to also follow the electrons, but this is a much more difficult task because each electron suffers thousands of collisions and some approximation must be made. Berger,19 for example, had the electron deposit its energy along averaged but randomly generated electron tracks. Webb(18) uses a similar approach but assumes that each electron deposits its energy along a straight-line track.

Monte Carlo calculations are without doubt the most powerful method available for studying the complex behavior of the interaction of a broad beam of radiation with an absorbing medium.

3. INTERFACE EFFECTS-ELECTRONIC EQUILIBRIUM

Only one of the methods mentioned above even potentially takes into account the effect of inhomogeneities on the pattern of electron motion. This is the Monte Carlo method.

Epp et al.(20) and Nilsson and Schnell(21) have shown that the presence of air cavities, such as that in the larynx, can cause a disruption of electronic equilibrium such that the dose to the far inside surface of the larynx may be 10% or more below the equilibrium dose. The factors governing this behavior are not well known. Presumably both the depth and lateral dimensions of the cavity are important. Leung et al. (22) have shown that for very-high-energy radiation, where the electron tracks may be several


centimeters in length, lung tissue lying in the buildup region may extend this region by several centimeters.

No general rules may be given at this time, but some general statements can be made that help to put the problem into perspective. Consider the arrangement depicted in Figure 13. A layer of tissue is followed by a layer of bone, which is again followed by a layer of tissue. Imagine the radiation to be from a cobalt unit. The kerma (energy transferred by the photon interactions) is represented by the dashed line. There is a discontinuity in the kerma at each interface. The magnitude of the discontinuity is given by the ratio of the average mass energy absorption coefficients, which for tissue to bone for this energy is 4 %.

The absorbed dose to tissue, bone, and tissue is represented by the solid lines. There is also a discontinuity in the dose at each interface. The magnitude of this discontinuity is given by the ratio of the average stopping powers for the electron spectrum for bone compared to that of tissue. For these two materials, for this energy of radiation, this ratio is about 7%.

Frequently one is not, in fact, interested in the actual dose to bone itself but rather in that to tissue-such as blood-forming tissues-within small structures within bone. The dose to such tissues within bone would be related to the kerma in bone, but the absorbed dose would be proportional to the stopping power in the tissue. The magnitude of the dose would depend on the size of the cavities containing them. This was discussed long ago in some detail by Spiers(23) for low-energy radiation and is also discussed by Johns and Cunningham(S). The shaded region in Figure 13 is intended to indicate the range of variation the absorbed dose to tissues within bone might take. At the present time, no method other than Monte Carlo has

depth

Figure 13. Diagram showing the behavior of kerma and absorbed dose near tissue-bone interfaces. The dashed lines represent kerma, and the solid lines absorbed dose. The shaded region represents the possible range of absorbed dose to tissue in small cavities within the bone.


been developed that can take into account dose variations resulting from loss of electronic equilibrium at interfaces.

4. SUMMARY AND CONCLUSIONS

Sontag et al., (3) in an experiment in which dose calculations and measurements were made in an inhomogeneous anatomical phantom, showed that the most important step was to obtain correct anatomical information. They also showed, however, that if an accuracy of better than ±5% was to be achieved, it was also necessary to use an accurate method of dose calculation.

In this chapter, various calculational schemes have been reviewed. They are listed in Table 1 and have been discussed in order of increasing complexity. The first group of methods (attenuation coefficient, ratio of tissue-air ratios, effective SSD, and isodose-shift methods) can all be shown to be variants of the ratio of tissue-air ratios method. They are all based on the radiological depth to the point of calculation and do not take into account the relative position of inhomogeneities and the point of calculation. Use of any of them can frequently lead to errors of 10 to 15%.

Next in order of sophistication is the power law tissue-air ratio method developed by Batho. This method does take into account the distance between an inhomogeneity that is traversed by the primary beam and the point of calculation and is very accurate for small field sizes. It becomes increasingly inaccurate for large field sizes, particularly for calculations at points within an inhomogeneity. Its use, however, will only occasionally lead to errors exceeding 5 to 10%.

Next in order of complexity is the equivalent tissue-air ratio method, which is also a ratio of tissue-air ratios. It does take into account the full three-dimensional structure of the irradiated body for each point of calculation. To do this rigorously would require an integration over the whole (three-dimensional) volume irradiated. This method lends itself, however, to the use of an (approximate) procedure for reducing the volume integration to an area integration, and this step makes the method practical at the present time. Its use will rarely give rise to errors exceeding 5%, but its application does require three-dimensional anatomical information. This is available only by means of a device such as a CT scanner. The CT scanner does, however, give better resolution and more structural detail than are actually required for dose calculation. Although a simpler device may well be developed, it must be realized that such a step would be at the expense of the detail needed for beam localization.

Methods of the future include the volume integration of differential scatter-air ratios and Monte Carlo calculations. The former have in fact


been implemented by Larsen and Prasad(16) using a computer with specially constructed microprocessors. This method also involves assumptions about the behavior of multiply scattered radiation and some of these assumptions are not always obvious. It is not known at the present time whether it is more accurate than the equivalent tissue-air ratio.

None of the methods just mentioned take into account the loss of electronic equilibrium near interfaces between dissimilar structures. Only the Monte Carlo method can do this. It should be, without doubt, the most accurate method discussed, but its application requires the generation of enormous numbers of photon histories and its use for treatment planning is not practical with present-day technology. Its usefulness is principally to provide reference data to which results of other methods can be compared.

REFERENCES

1. ICRU, Report 24, Determination of Absorbed Dose in a Patient Irradiated by Beams of X or Gamma Rays in Radiotherapy Procedures, International Commission on Radiation Units and Measurements, Washington, D.C. (1976).

2. R. Loevinger and T. P. Loftus, Uncertainty in the delivery of absorbed dose, in Ionizing Radiation Metrology, International Course, Varenna, Italy, 1974 (E. Casnati, ed.), pp. 459-473, Editrice Compositore, Bologna (1977).

3. M. R. Sontag, J. J. Battista, M. J. Bronskill, and J. R. Cunningham, Implications of computed tomography for inhomogeneity corrections in photon beam dose calculations, Radiology 124, 143-149 (1977).

4. M. R. Sontag and J. R. Cunningham, The equivalent tissue-air ratio method for making absorbed dose calculations in a heterogeneous medium, Radiology 129,787-794 (1978).

5. H. E. Johns and J. R. Cunningham, The Physics of Radiology, 3rd Ed., Charles C. Thomas, Springfield, III. (1969); ibid., 4th Ed. (in press).

6. H. F. Batho, Lung corrections in cobalt 60 beam therapy, J. Can. Assoc. Radiol. 15, 79-83 (1964).

7. M. E. J. Young and J. D. Gaylord, Experimental tests of corrections for tissue inhomogeneities in radiotherapy, Br. J. Radiol. 43, 349-355 (1970).

8. M. R. Sontag and J. R. Cunningham, Corrections to absorbed dose calculations for tissue inhomogeneities, Med. Phys. 4,431-436 (1977).

9. J. Van Dyk, J. J. Battista, and W. D. Rider, Half-body radiotherapy: The use of computed tomography to determine the dose to lung, Int. J. Radiat. Oncol. BioI. Phys. 6, 463-470 (1980).

10. J. E. O'Connor, The variation of scattered x-rays with density in an irradiated body, Phys. Med. BioI. 1, 352-369 (1957).

11. Central Axis Depth Dose Data for use in Radiotherapy, Br. J. Radiol. Suppl. 11 (1972). 12. M. R. Sontag, Ph.D. thesis, University of Toronto (1979). 13. J. W. Andrew, J. Van Dyk, and H. E. Johns, Use of scattered radiation measurements in

radiotherapy dose calculations based on computed tomographic (CT) images, in Proc. Soc. Photo-Opt. Instrum. Eng. 173, 342-347 (1979).

14. L. Beaudoin, M.Sc. thesis, University of Toronto, 1968. 15. J. R. Cunningham, Scatter-air ratios, Phys. Med. Bioi. 17,42-51 (1972).


16. K. B. Larson and S. C. Prasad, Absorbed dose computations for inhomogeneous media in radiation treatment planning using differential scatter-air ratios in Proceedings of the Second Annual Symposium on Computer Applications in Medical Care, Washington, D.C., U.S.A., 5-9 November 1978, pp. 93-99, IEEE, New York (1978).

17. D. E. Raeside, Monte Carlo Principles and Applications, Phys. Med. Bioi. 21,181-197 (1976).

18. S. Webb, The absorbed dose in the vicinity of an interface between two media irradiated by a 60CO source, Br. I. Radiol., 52, 962-967 (1979).

19. M. J. Berger, Absorbed doses near an interface between two media, in NBS Report 10550, pp. 38-48, National Bureau of Standards, Washington, D.C. (1971).

20. E. R. Epp, M. N. Lougheed, and J. W. McKay, Ionization buildup in upper respiratory air passages during teletherapy with cobalt 60 radiation, Br. I. Radiol. 31, 361-367 (1958).

21. B. Nilsson and P. O. Schnell, Buildup studies at air cavities with thin thermoluminescent dosemeters. Acta Radiol. Ther. Phys. Bioi. 15, 427-432 (1976).

22. P. M. K. Leung, B. Seaman, and P. Robinson, Low-density inhomogeneity corrections for 22-MV x-ray therapy, Radiology 94,449-451 (1970).

23. F. W. Spiers, II. Dosage in irradiated soft tissue and bone, Br. I. Radiol. 24, 365 (1951).

3

Anthropomorphic Phantom Materials

DAVID R. WHITE and CHRISTOOOULOS CONSTANTINOU

1. INTRODUCTION

1.1. The Need for Tissue Simulation

Since the introduction of x rays into the medical environment at the end of the last century and the emergence of a multitude of associated diagnostic and therapeutic techniques, experimentalists have endeavored to replicate the effects within and around irradiated body tissues using so-called "tissueequivalent" materials. In general, these materials have to provide the same attenuation properties as the real tissues being simulated, for the radiation modalities being investigated.

As measuring equipment and clinical techniques have become more sophisticated, so the types of tissues needing simulation and the degree of agreement required between replicating material and real tissue have dramatically increased. Radiation physicists are no longer satisfied to use simple blocks of solid materials or tanks of liquids, but demand composite arrays of materials or "body phantoms" having ever-increasing degrees of realism.

It is apparent that these materials are needed in three specific aspects of medical radiation physics.

DAVID R. WHITE and CHRISTODOULOS CONSTANTINOU • Radiation Physics Department, St. Bartholomew's Hospital, London, England. Christodoulos Constantinou's present address: University of Wisconsin, Madison, Wisconsin.

133

134 David R. White and Christodoolos Constantinoo

Firstly, in radiotherapy, radiation dose distributions frequently have to be derived when new or specialized treatments are being developed. As the radiations to be used could be either photons, charged particles, or neutrons, then solids, liquids, gels, and powders simulating the principal human tissues for these radiations should be available. Reliable bolus and buildup materials are also necessary once the patient treatments begin. It can be argued that depth dose data in water are not, by modern standards, adequate for many radical treatment regimes when the routine use of computed tomography (CT) and computerized treatment planning permits the evaluation of heterogeneous distributions. In support of this, anthropomorphic body phantoms are commonplace in many radiotherapy departments.

Secondly, in diagnostic radiology, patient dosimetry and radiological protection studies necessitate the use of either partial body sections or full anthropomorphic phantoms so that doses may be derived at specific body sites. Recent progress in the design of test objects (see Section 6.2) for assessing the radiographic performance of such x-ray procedures as mammography, dental radiography, and computed tomography, has utilized the new materials now being produced. In fact, the introduction of CT machines has been a stimulus for the formulation of more varied materials having improved simulation properties. The extraordinary sensitivity of these machines can, if inferior materials are used, result in the test objects being checked by the machines, rather than the converse.

Lastly, in nuclear medicine and associated radio nuclide studies, simulation materials are being used in dosimetric studies, the production of counting standards, and for quality assurance. For example, studies of the dosimetry of internally absorbed radionuclides may require the use of materials simulating a range of body tissues and organs. As counting standards, the use of anthropomorphic chest phantoms are proving invaluable in such studies as the assessment of 239pU in lung, while performance checks on emission computed tomography (ECT) machines are currently performed with body phantoms containing radionuclide sources at selected sites.

It is unlikely that the requirement for these materials will diminish in the foreseeable future. On the contrary, it is likely that modern medical radiation equipment and techniques will become even more discerning and require the development of more elaborate anthropomorphic materials and phantoms. As radiation physicists, it is our duty to ensure that we have developed the mathematical formulation procedures and manufacturing skills that will enable us to satisfy this demand.

1.2. Terminology

The terms equivalent, quasiequivalent, and substitute will have definite, precise meanings in this study. A material that has mass interaction data

Anthropomorphic Phantom Materials 135

over a given energy interval and a mass density within 1 % of similar data for a tissue is said to be a tissue equivalent material. If this agreement is achieved by formulating the material so that its elemental composition and proportions match those of a tissue, the material is said to be elementally equivalent. Difficulties in the manufacture of solid elementally equivalent materials result from the high carbon and low oxygen content of polymers and resins compared to tissues. Any such material that matches all of the elements in a tissue except the C and 0 group, which is matched as a single C + 0 entity, is elementally quasiequivalent. The general term substitute is used to cover both of these categories and any material claimed to approximate a tissue in a given radiation response.

The contraction coefficient is used in a general sense to describe both partial and total mass attenuation (f.L/ p) and energy absorption coefficients (f.Len/ p). Similarly, the term power is generally employed to describe partial and total electron and proton mass stopping powers (S/ p) and electron mass angular scattering powers (02 / pi). All of these quantities are discussed in Section 2.2.

1.3. A Brief History

The historical development of tissue substitutes is outlined in the "family tree" depicted in Figure 1; the important materials introduced during the past 70 years are given, together with the original references. (1-40)

All tissue substitutes appear to stem from water, suggested by Kienbock(1) as a muscle substitute in 1906. Seven years later, the first solid (Bakelite) was used and during the next 50 years basic materials such as wax, pressdwood, glass, aluminum, and many polymers were employed in a diversity of experimental arrangements. Although mixtures of powders had been used in 1926 in the form of bolus alba,*(6) it was not until 1937 that the first formulated solid substitute was produced, namely Siemens' wax. (9) The intensive use of both water and wax in dosimetric studies from the mid-1920s to the late 1930s had indicated the low photon attenuation properties of pure wax when compared to water and muscle. Siemens' wax, composed of paraffin wax and a corrective high atomic number (Z) filler, magnesium oxide, was the first of a long series of wax-based substitutes, which included the muscle substitutes Mix D,(15) Harris wax (41) and M3,(22) and the bone substitute Cameron wax.(32) Just as Siemens' wax initiated the wax-based products, so was, in 1956, the polyethylene-based Markite(21) the forerunner of a series of polymer- and rubber-based materials that included the electrically conducting Shonka plastics, (23) Temex, (26,27) and Rando. (28.29) Foamed solid materials for use as lung substitutes were introduced in

*Bolus alba is a mixture of two parts flour and one part china clay by weight.

136 David R. White and Christodoulos Constantinou

l~'r----------------------------------------------------. WATER

II,

Liquids &. Gels Powders 1910- -

I Talc'3'

Bakelite'2i

I 19201- -Wax 14I _______ ---, n

19301- IvolrYl51 I ~~~~~6' -

foams l Rice'" _ 1940rl'ressdwood'8' Siemens' Triolein"ol

~ _____ --, Wax'91

I Rice"" Plaster of Paris Zellstoff"21 Mixtures

19501- G,lassll3,'41 Mix 0"51 I -Plexiglas"6-191 I I I 1

Polystyrene Markl'te '2" 1211 Lincolnshire M3'221 Rossi Liq, and Gel Bolus'20' 19601- AI ~'24'251 A150'23, I I I I -

, Temex & Rando Spiers Frigerio

I 1'26-291 I Cameron Liq!301 Gel'3"

Wax l321 I I 19701- Frigerio'33'& Weber -Witt Liqs'341 Powder '35,

EVA-28 '36, Egoxiesl36, l N1'361 I 6110'40) Ethoxyethanol'371 Barts Liqs, Gels & Powders'36-39,

1980~------------~-----~~------~

Figure 1. The development of tissue substitutes since 1906.

1944(12) and subsequent products included Rando epoxy resins,(28) Temex latex,(27) adiprene,(42) and polyurethane (LN1).(36)

Only three pure liquid systems, that is, liquid devoid of additives, appear to have been used as substitute materials, with water, triolein, (10) and ethoxyethanol(37) simulating, respectively, muscle, fat, and adipose tissue. In 1956, Rossi and Failla(21) introduced a four-component liquid mixture in an attempt to elementally replicate an approximate formula for soft tissue, namely, (CSH lO0 1SN)n' Following this pioneering work, other originators have formulated similar muscle substitutes, with Frigerio and Sampson(33) and Goodman(43) producing useful formulations. The difficult task of simulating liquid bone has been attempted, notably, by Spiers and Chesters(30) and Witt and Cameron. (34) A logical development from formulated liquid substitutes is the production of gels simulating muscle. Rossi and Failla(21) used gelatin and Frigerio(31) used agar in their systems, while more recently Laponite has been used in the thixotropic gel BL8. (39)

As for both solids and liquids, powdered systems used predominantly as bolus in radiotherapy treatments also fall naturally into two categories. Pure particulate substitutes, devoid of additives, make up the first group, while the specially formulated mixtures of powders make up the second group. Talc(3) and rice(7) are examples of muscle substitutes of the first category, while bolus alba, (6) the rice-sodium bicarbonate mixtures of Spiers/ll)


Lincolnshire bolus,(20) and Weber powder(35) are examples of formulated muscle substitutes.

For further details on the constituents, elemental composition, photon and electron interaction properties of many of these materials, the reader's attention is drawn to two review articles. A report issued by the Hospital Physicists' Association(44) considers in detail the characteristics of 33 materials, while a review by White(45) discusses 64 adipose, cortical bone, fat, lung, and muscle substitutes.

Since 1970, research and development work on tissue simulation has been confined to a few laboratories in North America and Europe. Work at the Argonne National Laboratory,(46) the Illinois Benedictine College,(47) the Lawrence Livermore Laboratory,(48) and Oak Ridge National Laboratory(49) has indicated some interesting liquid and polymer-based substitutes and body phantoms. An extensive development program at St. Bartholomew's Hospital, London, England, has resulted in the production of over 160 new substitute materials. (36,38) Eighteen different tissues and body organs and tissue components have been simulated, including adipose tissue, muscle, skin, cortical bone, brain, lung, breast, and fat. Some 114 solid and powdered substitutes based upon acrylics, epoxy resin systems, nylon-6, polyethylene, polyurethane, and TPX (methylpentene polymers) have been formulated, together with nearly 50 liquid and gel substitutes.

In this contribution, emphasis will be given to the progress that has been made since 1970 in the production of anthropomorphic phantom materials for use with photons, electrons, protons, and neutrons. After the radiation properties of the principal tissues requiring simulation (Section 2) and the simulation procedures (Section 3) have been discussed, 50 of the best and most useful tissue substitutes will be presented (Section 4). Calculated, comparative interaction data will be given for the selected substitutes, together with a discussion of the available measured interaction data. Detailed information on the manufacture of bulk materials and phantoms will be given (Section 5). In order to illustrate the diverse applications of these systems, an account will be presented (Section 6) of recent experimental studies that have used phantoms manufactured from these new materials.

2. RADIATION PROPERTIES OF REAL TISSUES

2.1. Tissues Requiring Simulation

The wide range of tissue substitutes available at any time is usually due to a combination of the needs of the experimentalist and the techniques and expertise of those involved with simulation. In fact, the first 60 years of medical radiation measurements forced the evolution of significant numbers


of substitutes for muscle, bone, lung, and fat, with one or two substitutes being prepared for such tissues as blood or breast. These were, of course, considered acceptable, given the equipment and methods available at that time, and it is only in more recent years that the general improvement in technology has necessitated the simulation of more tissues.

Another factor that has strongly influenced the types of tissues being simulated has been the sparse published data on the elemental compositions of human tissues and organs. For tissue substitutes to be formulated, reliable elemental compositions and mass densities must be available, and prior to 1975, only a few such sources of information existed. For example, the reports of Woodard(SO) in 1962, describing the elemental composition and mass density of cortical bone, and an International Commission on Radiation Units and Measurements (ICRU) report(Sl) published by the National Bureau of Standards (NBS) in 1964, giving the elemental composition of striated muscle and compact bone, appear to be the only sources of useful data. Even with these two reports, a conflict in the agreement between the Woodard and ICRU formulas made the task of simulating bone very difficult and resulted in many workers believing, quite erroneously, that the ICRU version was the correct one. An earlier publication of the International Commission on Radiological Protection (ICRP)(S2) listing the compositions of 36 organs and tissues also proved to be of little use in tissue simulation, as the important elements C, H, N, and 0 were excluded from the elemental tabulations. Fortunately, in 1975 the ICRP(53) published an excellent revision of the reference man concept, which included a wealth of information on organ and tissue masses, mass densities, and water, fat, and protein contents of some 112 important biological systems. The elemental data were comprehensive and included a tabulation of the concentration of 51 elements in 81 organs, tissues, and tissue components. An analysis of this data by White and Fitzgerald(54) has shown that the compositions of 69 of these organs and tissues are of a form that can be readily used in tissue simulation studies.

Variability between tissues of the same type from different subjects due to age, sex, etc., is a problem that is always present in simulation exercises. If a tissue or organ is known to vary significantly, then, assuming reliable formulation procedures to be available, a selection of substitutes spanning the known range of compositions might be required.

Current and future requirements for tissue substitutes have been taken into consideration in the tissues discussed in this contribution. The 15 tissues and organs selected for simulation will now be described.

2.1.1. PRINCIPAL SOFT TISSUES

Four types of soft tissues have been considered, namely, adipose tissue, blood, muscle, and skin. The elemental compositions and mass densities for


adipose tissue, blood, and skin have been derived from the latest ICRP reference man publication,(53) while the ICRU muscle formula(51) was retained because of its widespread acceptance. Adipose tissue is included in preference to the tissue component "fat." The term "fat" is frequently used loosely in the literature and sometimes quoted when adipose tissue is really intended. Adipose tissue as used here is strictly as defined by the ICRP, being that body tissue composed of 80% fat, 15% water, and 5% protein (by weight).

2.1.2. PRINCIPAL SKELETAL TISSUES

The three skeletal materials considered were cortical bone, inner bone, and red marrow. The elemental composition and mass density used for cortical bone is as defined by Woodard,(50) and those for red marrow as suggested by ICRP.(53) Yellow marrow has not been included because of its similarity to adipose tissue.

Inner bone,(36.38) a mixture of 22.4% cortical bone and 77.6% red marrow (by weight), represents an approximate composition of the inner core of bone and marrow found in certain bony structures (spongiosa). It should be noted that the composition used in this study is different from previous inner bone formulas(36) because of the use of red marrow in place of soft tissue, which was employed originally. Further collaborative work(55) on the specification of all skeletal materials is in progress, which, it is hoped, will yield a clearer picture of the compositions of skeletal tissues from different body sites.

2.1.3. BODY ORGANS

The compositions and mass densities of the five body organs evaluated, namely, brain, kidney, liver, lung, and thyroid, were all taken from the ICRP reference man document. (53) They represent some of the most important organs present in the body and were thought to usefully extend the range of available substitutes with particular application to experimental studies in nuclear medicine and diagnostic radiology.

2.1.4. AVERAGE TISSUES

Whenever tissues or organs are known to vary in composition, the use of substitutes having average compositions often provide a satisfactory solution to the problem of what composition should be simulated. An example is breast tissue, which is known to vary between the extremes of adipose and soft tissues. A formula based upon 50% fat and 50% water (by


weight) was introduced(36) as a basis for the production of test objects for assessing a radiographic performance in mammography.

In this average tissue category, average breast tissue is included, together with total soft tissue and total skeleton, the latter two being derived from ICRP sources. (53)

2.1.5. ELEMENTAL COMPOSITIONS

All of the calculated radiation interaction data presented in the next section are based upon the elemental compositions of the 15 organs and tissues discussed previously and that are listed in Table 1. For each of the tissues, the compositions (in percentage by weight) are given for the 11 principal elements, H, C, N, 0, Na, Mg, P, S, CI, K, and Ca. Where appropriate, data for additional elements are also quoted. Details of mass densities (kg m -3) and references to the data sources are included in the text.

For the purposes of simulation, particularly in the formulation of elementally equivalent materials, the fact that for certain tissues the sums of the constituent percentage weights do not equal 100% necessitates small corrections in the composition data. This has been achieved by correcting the carbon and oxygen contents to achieve the requisite 100%. Such minor adjustments, which take into account other published tissue compositions, (56) are justified, since carbon and oxygen are the most common elements and small changes in their absolute values will produce minimal changes in mass densities and radiation properties.

2.2. Radiation Characteristics

The interaction processes occurring within an irradiated tissue will be dependent upon the types of radiation being employed, the radiation energies, and the composition of the tissue. In order to characterize the magnitude of all these processes in both tissues and the derived substitutes, basic interaction data have been calculated for photons, electrons, protons, and neutrons. Detailed accounts of the computations have been given elsewhere, (36,38,54) so only outlines of the procedures will be given in subsequent sections.

For convenience, the radiation characteristics of tissues will be considered for each type of radiation in turn, but it must be borne in mind that once a given primary beam passes into a tissue, the resulting interactions inevitably produce secondary radiations of the other types. For example, an x-ray beam produces secondary electrons, while a neutron beam yields fluxes of scattered protons and gamma rays.

>- ~ T

able

1.

Ele

men

tal C

ompo

siti

ons

of t

he P

rinc

ipal

Org

ans

an

d T

issu

es

.. =

"C:I =

Ele

men

tal c

ompo

siti

on (

perc

enta

ge b

y w

eigh

t)

a =

.;

Mas

s ~

Oth

er

dens

ity

"1:1

Tis

sue

H

C

N

0 N

a M

g P

S C

I K

C

a el

emen

ts

(kg

m -3

) R

efer

ence

s :r

II

Pri

ncip

al s

oft

tiss

ues

~ a A

dipo

se t

issu

e 12

.00

64.0

0 0.

80

22.9

0 0.

05

0.02

0.

07

0.12

0.

03

92

0

53

a: B

lood

10

.00

9.82

2.

91

76.3

9 0.

18

0.03

0.

18

0.27

0.

16

0.01

F

e(0.

05)

1060

53

!

Mus

cle

10.2

0 12

.30

3.50

72

.89

0.08

0.

02

0.20

0.

50

0.08

0.

30

0.01

1

00

0-1

04

0

51

:I.

Ski

n 10

.00

22.7

0 4.

62

61.5

0 0.

18

0.01

0.

03

0.16

0.

27

0.08

0.

02

1100

53

;.

Pri

ncip

al s

kele

tal

tiss

ues

Cor

tica

l b

on

e 3.

39

15.5

0 3.

97

44.1

0 0.

06

0.21

10

.20

0.31

22

.20

1850

50

In

ner

bo

ne

8.66

40

.31

2.58

40

.35

0.08

0.

06

2.30

0.

46

0.23

4.

97

1120

3

6,3

8

Red

mar

row

10

.18

47.4

8 2.

18

39.6

7 0.

01

0.03

0.

15

0.10

0.

17

Fe(

O.O

l)

1030

53

Bod

y or

gans

B

rain

10

.70

15.3

3 1.

29

71.4

0 0.

18

0.34

0.

17

0.23

0.

30

1030

53

K

idne

y 10

.30

12.9

0 2.

74

73.2

4 0.

20

0.01

0.

16

0.24

0.

19

0.01

F

e(O

.Ol)

10

50

53

Liv

er

10.0

0 14

.40

2.83

71

.85

0.10

0.

02

0.03

0.

29

0.20

0.

25

0.01

F

e(0.

02)

1070

53

L

ung

9.90

12

.30

2.80

74

.00

0.18

0.

01

0.08

0.

22

0.26

0.

19

0.01

F

e(0.

04)

26

0-1

05

0

53

Thy

roid

10

.00

13.6

0 2.

20

73.5

0 0.

22

0.01

0.

08

0.17

0.

12

0.04

F

e(O

.Ol)

; 10

60

53

1(0.

06)

Ave

rage

tis

sues

A

vera

ge b

reas

t 11

.70

38.0

4 50

.26

96

0

36

Tot

al s

oft

tiss

ue

10.5

0 23

.33

2.50

62

.87

0.11

0.

01

0.13

0.

20

0.13

0.

20

0.02

10

30

53

Tot

al s

kele

ton

7.00

22

.70

3.90

48

.60

0.30

0.

10

7.00

0.

20

0.10

0.

20

9.90

14

00

53

~

~ ~


2.2.1. PHOTONS

Photons in the energy interval 1 a ke V-I 00 MeV interact by an energydependent combination of photoelectric absorption, coherent and incoherent scattering, and pair production. For photon attenuation in muscle, at lake V the photoelectric effect is the most important process, with coherent and incoherent interactions making up only 4 and 3%, respectively, of the total effects. At 30 keY the importance of photoelectric (39%) and incoherent (49%) processes has become reversed and the coherent effect is at its peak value (12%). For an energy of 1 MeV incoherent scattering in muscle predominates (>99%), while at 100 MeV pair production and incoherent interactions comprise 84 to 16%, respectively, of the total attenuation effect. With the exception of the incoherent process, which is independent of atomic number (Z), the other effects depend upon Z to varying degrees. All processes depend upon electron density (no) (electron/kg), which in turn is influenced by the high electron density of the hydrogen content (3-12% by weight) of tissues. Consequently, the high-Z constituents and hydrogen content of a tissue, together with its mass density, are the important factors that govern the transmission of a beam of photons through a given thickness of the tissue.

For the 15 tissues listed in Table 1, mass attenuation coefficients (IL/ p) and energy absorption coefficients (lLen/ p) have been calculated at 33 energy points in the interval 10 keV-I00 MeV using the elemental cross sections of Hubbell(S7) and Storm and Israel(S8) and the mixture rule. At each energy, the value of IL/ P is calculated from

t!:. = LWi(t!:.) PiP i

(1)

where Wi is the proportion by weight of the ith element having an attenuation coefficient given by (IL/ P )i. A similar expression was used for lLen/ p.

Figure 2 illustrates the range of attenuation data for adipose tissue, muscle, and cortical bone and represents the spread of coefficients that the complete 15 selected tissues have for this energy interval. As expected, at low energies, where the photoelectric effect predominates, adipose tissue, with its low-Z content, has the lowest attenuation properties; cortical bone, with its high-Z content, gives the highest attenuation. These differences appear small in the figure owing to the logarithmic nature of the graph, but at 10 keY the attenuation coefficient for cortical bone is 9.6 times that for adipose tissue. At intermediate energies (-1 Me V) owing to the higher hydrogen content of adipose tissue as compared to that of cortical bone, the roles are reversed. At high energies, where pair production becomes important, the roles reverse again, with cortical bone giving slightly greater attenuation than adipose tissue. All of the waterlike tissues in the group of


1 - Adipose Tissue

2 - Muscle

3 - Cortical Bone

1cr3~~~~~~~~uw~~~~~~~~~~

1~ 1~ 1if 1d 1~ ENERGY (MeV)

143

Figure 2. Photon mass attenuation coefficients for adipose tissue, muscle, and cortical bone.

15 (blood, brain, kidney, lung, and thyroid) yield results very close to those for muscle, which is also included in Figure 2. Skin, red marrow, liver, average breast, and total soft tissue have data that lie on curves between those of muscle and adipose tissue. Similarly, inner bone and total skeleton have mass attenuation coefficients between those of muscle and cortical bone.

2.2.2. ELECTRONS AND PROTONS

Charged particles lose energy while traversing tissues through collision and radiation processes. For electrons from 10 to 500 keY, electronic collisions are the major cause of energy loss, while for protons it is the primary process for the energy interval 1 0 ke V-I 000 MeV. Above 500 ke V electrons start to lose energy by radiation (bremsstrahlung) processes. For muscle, even at 1 Me V the electron radiation loss accounts for only 1 % of the interactions; this rises to 16% at 20 MeV, while at 100 MeV collision (48%) and radiation processes (52%) are approximately equally important. Collision stopping powers have an inverse dependence on Z, resulting in tissues with significant concentrations of low-Z elements (e.g., hydrogen) having the larger stopping powers. Conversely, electron radiation stopping powers are weakly Z dependent.

Angular scattering differs between electrons and protons, with electrons being scattered through large angles, whereas protons are confined to small angles.

Electron collision and radiation stopping powers, (S/ P )001 and (S/ P )rad, respectively, have been derived for the previously mentioned tissues for the


same 33 energy points in the intervallO keV-lOO MeV as used for photons. Collision data were calculated from first principals,(59.60) while the radiation data were derived from elemental radiation stopping powers and the use of the mixture rule.(59.61) Total electron stopping powers (Sf P)e were obtained from equation (2):

(Sf P)e = (Sf P )col + (Sf p )rad (2) -2

Electron mass angular scattering powers (0 f pI) have been calculated at the same energies, using calculated elemental data and the mixture rule as suggested by ICRU:(62)

02 = L w.(02)

pI i' pI i (3)

where Wi is the proportion by weight of the ith element having mass angular -2

scattering powers (0 f p/)i' For the analysis of the proton interaction characteristics of the simu

lated tissues, elemental stopping powers were derived for 38 energy points in the intervallO keV-lOOO MeV from published sources.(38.63-65) By using the mixture rule, the total proton stopping powers (Sf p)p of a given tissue was calculated from

(S) (S) - = LWi-P pip p.i

(4)

where Wi is the proportion by weight of the ith element having a proton stopping power given by (Sf p )P.i'

~102 E :a; :::E (f)

ffi 10'

~ (!) z &10° ~ (f)

(f)

~

1 - Adipose Tissue

3 - Cortical Bone

:::E10'L-~LU~~J-LLUil~-L~~~-L~LU~-L~~~ 10-2 10-' 10° 10' 103

ENERGY (MeV)

Figure 3. Proton and electron mass stopping powers for adipose tissue, muscle, and cortical bone.


Figure 3 illustrates the spread of electron and proton stopping powers versus energy for adipose tissue, muscle, and cortical bone. Differences are seen to be very small for both sets of data. For electrons, below 500 ke V, the higher hydrogen content of adipose tissue produces the larger stopping powers; cortical bone with its low hydrogen content produces the smallest powers. Once radiation effects become significant, the curves for adipose tissue and cortical bone cross over. The data for muscle, together with those for the other 12 tissues in the group, fall between these two extremes but very close to the data for adipose tissue and so, for reasons of clarity, have not been plotted.

For protons, as the collision process is the major contributor to the interactions, the stopping powers for adipose tissue form the upper curve for the entire energy interval. Again, as for electrons, the data for muscle and the other tissues fall between the extremes of these powers but close to the values for adipose tissue.

2.2.3. NEUTRONS

The uncharged particles, neutrons (100 eV-30 MeV), lose their energy traversing tissue through a complex series of interactions that includes elastic and inelastic scattering, captive processes, and fragmentation of the nucleus, or spallation. Certain elements have resonance regions within this energy interval (e.g., 0.3-9 MeV for oxygen and 2-9 MeV for carbon). For energies up to 14 MeV, elastic scattering with hydrogen is the major contributor to the loss of energy, while the capture processes l H (n, y) 2H and 14N(n, p) 14C are significant at low and thermal energies. Once again, as with photons, electrons, and protons, the interactions are dependent upon the hydrogen content of the tissue. For neutrons the effects are very strongly dependent upon the hydrogen concentration, with the nitrogen content being important at certain low energies. In soft tissue containing -10% by weight hydrogen, a 1 % change in this concentration can cause a -10% change in the magnitude of the resulting interactions.

To characterize the magnitude of neutron interactions, kerma factors have been used. The kerma K is defined as dEtr / dm, where dEtr is the sum of the initial kinetic energies of all the charged particles liberated by a neutron in a mass dm. The kerma factor is the ratio K/4> (E), where 4> (E) is the particle ftuence. It has been shown by the ICRU(66) to be equal to the product of the mass energy-transfer coefficient ILtr/ P and the neutron energy Em or

K ILtr Kerma factor = -- = - E 4> (E) p n (5)


Elemental kerma factors at 98 energy points in the interval! keV-30 MeV have been extracted from ICRU Report 26,(66) and corresponding data for each tissue calculated by the application of the mixture rule:

(6)

where Wj is the proportion by weight of the ith element having a kerma factor (K/</> (E))j.

The kerma factors for adipose tissue, muscle, and cortical bone are shown graphically in Figure 4, where the data are seen to increase sharply as the hydrogen concentrations increase. Kerma factors for red marrow may be considered as an illustration of the insensitivity to carbon and oxygen content, which is vital to the formulation of substitutes based upon polymers or resins with low oxygen concentrations. Muscle and red marrow both have essentially the same hydrogen content (10.2% by weight) and carbon plus oxygen concentration, -86% by weight. The carbon/oxygen ratios are different for the two tissues, being 12.3/72.9 (= 0.17) for muscle and 47.5/39.7 (= 1.20) for red marrow. Except for the resonance regions, from 360 eV to 1.85 MeV the kerma factors for red marrow are within 1 % of those for muscle, indicating the similarity in the interactions for carbon and oxygen over this energy interval and their possible interchange in substitute materials.

1 - Adipose Tissue

2 - Muscle

3 - Cortical Bone

Figure 4. Neutron kerma factors for adipose tissue, muscle, and cortical bone.


3. SIMULATION PROCEDURES

3.1. Criteria for Tissue Equivalence

For a material to be acceptable as a tissue substitute, the absorption and scattering of radiation within a given thickness or mass of the material must be the same as that obtained for a comparable thickness or mass of tissue. Strictly, the agreement should extend to both the magnitude of the interactions and, where appropriate, the angular distribution of any secondary radiation. Each type of radiation employed will require, over the energy interval of interest, the evaluation of different interaction processes. Materials for use with photons, electrons, protons, and neutrons must take into consideration four categories of interaction data, namely,

1. Mass attenuation and energy absorption coefficients (ILl p, lLenl p). 2. Electron mass stopping and angular scattering powers

[(SI p) .. 021 pI]. 3. Proton mass stopping powers [(Sip )p]. 4. Neutron cross sections or kerma factors [0', KI1J(E)].

Wherever possible, partial rather than total interaction data for the tissue and the proposed substitute should always be matched, because balancing discrepancies in competing partial processes could nullify, resulting in apparently acceptable total data.

If equal masses of the tissue and substitute are required to have the same volume, then the mass densities must be identical. Linear interaction data may be used directly, but it has been found that because of the variations in mass density of many materials, it is more practical to base formulation procedures on mass interaction data and to consider mass densities separately.

The agreement necessary between these data for the tissue and the substitute is dependent upon the subsequent application of the substitute. Certain modern techniques such as computed tomography can readily distinguish differences of less than 1 % in photon linear attenuation coefficients, while many dosimetric procedures for both photons and particulate radiations can indicate differences of 1-2% in interaction characteristics. Even bearing in mind the variability of tissues, it seems unreasonable by modern standards to have to accept substitutes with simulation discrepancies in their interaction data exceeding 5 %. If it is possible to produce substitutes with discrepancies within 1 %, then this should be achieved, not only because such materials are more appropriate, but the high-precision procedures necessary for their formulation will


enable families of substitutes to be reliably produced to cover tissues known to have a range of compositions.

So wide are the uses of tissue substitutes now that for a material to be of general application an energy interval of at least 10 ke V -100 Me V should be considered for most interactions. Extensions to lower energies (-1 ke V or less) are necessary for neutrons, and extensions to higher energies are necessary for protons (-1000 Me V). In many specialized applications the matching of all these quantities is not obligatory and many useful substitutes based upon selected or weighted groups of interaction data and restricted energy intervals may be formulated. For example, the consideration of photon attenuation coefficients from 60-90 ke V can result in acceptable substitutes for use in radiographic phantoms in computed tomography, while the same data from 10-40 keY can yield useful materials for mammographic phantoms. Problems occur, often due to a misunderstanding of the term "tissue equivalent," when experimentalists use substitutes for radiation modalities and energy intervals that were not considered by their originators.

An obvious solution to the problem of formulating tissue substitutes for general use is to make substitutes that are elementally equivalent to and of the same mass density as the tissue. Then, as long as there are no "phase" differences or changes in the interaction processes due to chemical bonding, such a product could be used with complete confidence for all types of radiation and energy ranges. In practice, the production of liquid substitutes by this method is relatively straightforward; gels are somewhat more difficult to produce but can be achieved with persistence. Solid, elementally equivalent materials are exceedingly difficult to formulate, and have only been successfully achieved in a few limited cases.

The principal methods of simulating tissue substitutes, using either the indirect approach of matching interaction data or directly by means of elemental equivalence, will now be discussed in detail.

3.2. The Effective Atomic Number (Z) Method

One of the first analytical methods of formulating tissue substitutes was based on the use of so-called effective atomic numbers Z and has been used extensively since 193i67 ) for selecting materials for use with photons. The concept is that for a given Z-dependent partial interaction process, an irradiated compound may be said to be acting as if it were an element with an atomic number Z derived from an equation of the form

(7)

where CXi is the fractional electron contribution of element Zi in the


compound and the Z exponent x is a constant dependent upon the interaction process under investigation.

Other definitions of Z have been proposed in the literature, such as weighting according to relative atomic proportions or relative masses of the constituents, (68-70) but the relationship quoted in equation (7) is the one that appears to be most favored in tissue simulation studies.

In 1946 Spiers(3) showed that the partial mass attenuation coefficient is directly proportional to the product of the electron density no (electrons/kg) and the effective atomic number raised to the power x. For, say, the photoelectric mass attenuation coefficient /-LT/ p,

/-LT/ P ex: noZx (8)

Conventionally, the simulation procedure was to formulate a compound comprising a mixture of constituents such that its Z was as close as possible to that of the tissue. This was often performed only for the photoelectric process with Z exponents x being given a value in the range 2.94_3.4.03,35) Unfortunately, no was frequently only used if incoherent processes were being considered. Although the Z dependence of other partial processes such as coherent scattering and pair production have been discussed in the literature(35,70) they were only rarely used prior to 1974.(36)

A simple example should help to indicate the problems and highlight some possible solutions. Assume that the suitability of acrylics (e.g., Plexiglas) (C5H80 2)n, polycarbonate (e.g., Lexan) (C16H1403)m and polystyrene (C8H8)n as potential muscle substitutes in the energy interval 10-40 keY is being considered and the agreement between photoelectric mass attenuation coefficients for muscle and these materials is to be established. Using Zexponents (x) of 3.0 and 3.5, the ratios ZmadZmus have been listed in Table 2, together with a selection of the ratios (/-LT/ P )matl (/-LT/ P )mus at specific photon energies. It can be seen that the Z data indicate that the photon attenuation discrepancies for acrylics, polycarbonate, and polystyrene are approximately 13, 15, ahd 25%, respectively, when, in fact, the discrepancies are really 42,50, and 64% at 20 keY. Included in the table are the ratios Y mati Y mus> where Y = noZx. Agreement between these latter ratios and the attenuation coefficient ratios is seen to be much closer than with the Z ratios, with the agreement improving as the Z exponent increases from 3.0 to 3.5.

These results serve to illustrate the point that the use of single Z values to characterize extended energy ranges will not, in general, be satisfactory. The exclusion of the electron density no from the considerations and the value given to the Z exponent are the primary factors causing the apparent disparities in the data. The way in which the concept of effective atomic number can be extended and used more fruitfully will be discussed in Section 3.4.


Table 2. Some Z, Y, and II-T/P Data for Acrylics, Polycarbonate, and Polystyrene

Materials

Parameters Conditions Acrylics Polycarbonate Polystyrene

Zmat/Zmus x = 3.0 0.87 0.84 0.76 x = 3.5 0.86 0.85 0.75

Yma.IYmus x = 3.0 0.64 0.56 0.43 (Y = noZ') x = 3.5 0.58 0.50 0.36

(/LTI P )ma.1 (/LTI P )mus 10 keY 0.60 0.51 0.37 15 keY 0.59 0.51 0.37 20 keY 0.58 0.50 0.36 30keV 0.57 0.49 0.35 40 keY 0.56 0.48 0.34

3.3. The Basic Data Method

In an attempt to develop a more general solution to the problem of tissue simulation for different types of radiation the basic data method was introduced in 1974.(36.71) Designed originally for the complete range of photon and electron interactions, the method has three main aims:

1. The procedures have to precisely establish the relative weights of the components of a substitute having one or more selected interaction data within predefined limits.

2. To eliminate the arbitrary selection of corrective additives, the procedure has to be able to select from a library of compounds those most suitable for addition to a specified base material.

3. The procedure must enable one to formulate two-component substitutes comprising a base material and a corrective compound.

Using these guidelines, the following analytical procedure may be developed for photons and electrons only.

A tissue substitute S composed of two compounds A and B precisely simulates a tissue X at one energy and for one partial interaction process. If the corresponding coefficients (or powers) are given by Cs, CA , CB , Cx and the mixture rule is obeyed, Cx = Cs and

(9)

where WA and WB are the proportions by weight of A and B. As WA + WB = 1, then

(10)


For the proportions defined by equation (10), the coefficient (or power) at the appropriate energy for the substitute is identical with that for the tissue. The agreement at other energies and other partial processes depends upon the choice of materials. A method for screening the corrective compounds may be based upon the slopes of the plots of the coefficients (or powers) versus energy. Many of the important partial photon and electron interaction data when expressed in log-log plots against energy form linear graphs over certain energy intervals. For compounds, the slopes of these linear graphs depend upon the slopes of similar graphs for the elements making up the compound, because for many interaction processes the slopes for the elements are Z dependent. With the exception of the slopes for incoherent processes and electron angular scattering, whose slopes show a minimal change, all other partial photon and electron effects show significant variations in their slopes with atomic number. This same reasoning can be applied to a substitute made up of two compounds; the final slope for a given process will depend upon the compounds used and their proportions (by weight) in the substitute. Typical log-log plots of partial coefficients (or powers) versus energy for a tissue X and a substitute S composed of compounds A and B are illustrated in Figure 5. The data are plotted over an energy interval described by a minimum energy Eo and a midpoint energy Em. The slope of the substitute, ms, is given by ms = In(e~/ ex }/In(Em / Eo}, where e~, and ex are as shown in Figure 5. Using a similar expression for the slope of the tissue, mx, the permissible variation in the slope of the substitute may be calculated from

(11)

where f is the fitting ratio, given by f = e~/e'x, with e~ and e'x being the coefficients (or powers) for the substitute and tissue at energy Eo.

Figure 5. Log-log plots of interaction data versus energy for a tissue X and a substitute S comprising a base material A and a corrective compound B.

A - Base material B - Corrective compound S - Substitute X -Tissue

In (energy)


If the substitute is assumed to be made up of a known base material A and an unknown compound B, it may be shown with reference to Figure 5 that the slopes ms, mA, and mB are related by

(12)

where K = Em/Eo and {3A and {3B, the fractional contributions to the coefficients (or powers) at Em, are given by {3A = WACA/Cs and {3B = WBCB/Cs.

From equation (12) it follows that if the substitute gives perfect simulation at Em (i.e., Cs = Cx ), then

mB = (In Kms - (WACA/Cx)KmA) (In K)-l 1 - (WACA/Cx )

(13)

If the permissible spread in ms is set by means of the fitting ratio, then the permitted variation in mB may be calculated. Using equations (10)-(13), it is now feasible to select and formulate two-component substitutes having coefficients (or powers) for a given interaction process within defined limits. The procedure requires six definite mathematical steps, which lends itself to computerization. Once the tissue and base material have been identified, the procedure is as follows:

1. Select the fitting ratio /. 2. Calculate the limits of the slope of the substitute, ms [equation (11)]. 3. Calculate permissible slopes (mB) of added material for a range of

fractional weights [equation (13)]. 4. Find a compound that is complementary to the base material (i.e., if

the base material has coefficients, or powers, less than those for the tissue, the corrective compound B must have corresponding larger values).

5. Calculate relative fractional weights of base and corrective materials, which together give perfect simulation at a midpoint energy [equation (10)].

6. Check that the proposed added material has an acceptable slope for the fractional weights required. If this test is positive, the material is suitable for this interaction.

The sequence must be performed on all the partial interaction processes thought to be important for the simulation exercise and a common formulation of base plus corrective materials found that satisfies the fitting requirements of each process.

The energy intervals that have been used for this method, for photon attenuation and absorption, have been 10-150 keV for photoelectric and coherent effects, with exact fitting at 40 keV, and 5-20 MeV for pair


production effects, with exact fitting at 10 MeV. As the form of incoherent data is not strongly Z dependent, fitting for this effect is at a single energy only (1 Me V). For electrons, collision stopping powers in the energy interval 10-500 keV have been used, with exact simulation at 60 keV, while for radiation stopping powers the interval is 1-100 MeV, with exact fitting 10 MeV. For the same reason as given for incoherent effects, angular scattering powers may be evaluated at a single energy (1 MeV).

To date, the method has not been used in the formulation of substitutes for protons or neutrons, although, at least for protons, the method appears feasible.

3.4. The Extended Y Method

The realization that there were severe limitations with the application of effective atomic numbers Z, discussed in Section 3.2, led to a thorough investigation of the Z dependence of photon and electron interactions, (36,72)

with special reference to the implications in tissue simulation studies. The derived Z exponents for both photon and electron interactions were shown to be strongly influenced by elemental grouping as well as energy dependent. For example, a Z exponent suitable for photoelectric attenuation processes at, say, 10 keV in muscle would not necessarily be acceptable for cortical bone for the same interactions and energy. Electron interactions were also included in the evaluation, but yielded poor results for collision and radiation processes. The investigation showed that, at least for photons, groups of two or three Y values (Y = noZX), each characterizing the magnitude of a given process at a specific energy, could usefully describe that effect if manipulated correctly.

A simulation procedure, called the extended Y method(36,73) evolved from the study and paralleled the aims and reasoning developed for the basic data method. The analytical procedures, for photons only, use Y values and associated data in the following fashion:

A substitute S simulating a tissue X for a given partial interaction process and energy is composed of a base material A and a corrective compound B. If the corresponding Yvalues are Ys, Yx , YA , and YB and if the mixture rule is obeyed, from equations (8) and (9) we have

(14)

If the Z exponent x has been correctly chosen to characterize the partial process at the required energy, equation (14) establishes the relative weights of A and B (WA, WB) that give perfect simulation at that energy.

In order to screen possible additives B for suitability, the Y concept must be applied to coefficient slopes (m). As the attenuation coefficients at two different photon energies will be characterized by two Y values, each


with different Z exponents, then, with reference to Figure 5,

C~ = ko(no)sZ~o and Cs = km(no)sZ'Sm

where ko and km are constants, (no)s is the electron density for the substitute S; XO and Xm are the Z exponents at energies Eo and Em, and C~ and Cs refer to photon interactions only.

As the slopes may be expressed in the form Kms = C~/ Cs and, from equation (7), Z~ = aAZA + aBZ;', it follows that

Kms = I/IAKmA + I/IBKmB (15)

where I/IA = aAZ""f'/(aAZ""f' + aBZ;r), I/IA + I/IB = 1, and aA and aB are the fractional electron contents of materials A and B in substitute S. Equation (15) gives the relatior.ship between the slopes of A, B, and S in terms of the energy ratio K (= Em/Eo) and I/IA and I/IB, the fractional components of Z'Sm.

If the ratios (R) of the Z powers (ZX) at two energies are considered, Rs = Z~o/Z'Sm (and similar expressions for RA and RB). Then, as K = (ko/ km)Rs, substitution in equation (15) gives

(16)

Equations (15) and (16) show that the ratios R may be used instead of slopes m to screen the additives. To make the procedure complete, two further equations must be derived, one establishing the permissible variation in the attenuation coefficients for the tissue and the substitute at the minimum energy Eo and the other specifying the permissible variation in R B •

As the slopes ms and mx are given by Kms = C~/ Cs and Kmx = C'x/Cx, and at energy Em, Cs = Cx, then

Kms/Kmx = C's/C'x = Rs/Rx

But (the fitting ratio) f = C~/C'x, so

Rs/Rx = f (17)

From equation (16) it follows that the permissible variation of RB may be derived from

RB = Rs - (aAZ_~/Z!:)RA 1- (aAZ-;('/Z?,,)

(18)

As for the basic data method, its implementation requires six definite mathematical actions once the tissue and base material have been identified:

1. Select the fitting ratio f. 2. Calculate the limits of the Z power ratio for the substitute, Rs

[equation (17)].


3. Calculate the acceptable Z power ratios for the corrective material (RB ) for a range of electron fractions [equation (18)].

4. Find a material that is complementary to the base material (Le., if the base material has Y values smaller than the tissue, the added corrective compound must have correspondingly larger values).

5. Calculate the relative weights and electron fractions of materials A and B that together give perfect simulation of the tissue at energy Em [equation (14)].

6. Check that RB for the proposed corrective material is acceptable for the relative weights and electron fractions required. If the test is positive, the material is suitable for this interaction.

The sequence must be repeated for all the relevant partial interaction processes, and a common formulation of the base material A and corrective compound B found that satisfies the fitting requirements of each process.

Details of suitable Z exponents for use in this method for the Zdependent effects (photoelectric, coherent, and pair production) are to be found in the literature.(n) Incoherent processes are evaluated by considering electron densities when it may be shown that the values for materials S, A, and B [(noh, (nO)A, and (nO)B] are given by equation (19):

1 aA aB --=--+--(no)s (nO)A (nO)B

(19)

Up to now this method has only been used successfully for photon interactions, and its application to particulate radiations requires further investigation.

An interesting formulating procedure, somewhat similar to the extended Y method, for both photon and electron interactions has been introduced independently by Geske. (74-77) Partial interaction processes are characterized by .Keff, given by

- Zf X eff = L wi A ,

I I

where Wi is the fractional mass of the ith atom with atomic number Zi and relative atomic mass Ai, within the substitute material. An improvement in

(76) --the procedure uses the quantity Yeff =' pXeff, where p is the mass density.(77) In both cases the .Keff or Yeff values for the substitute and tissue being simulated are considered for respectively coherent scattering, photoelectric absorption, incoherent scattering, and pair production, with suitable Z exponents being employed. Single formulations that satisfy the requirements for the four partial photon processes have been derived and are said also to show good simulation properties for the partial electron interaction processes.


3.5. Elemental Equivalence

Compared to the popular effective atomic number method for formulating tissue substitutes, the elemental equivalence method is a relative newcomer, having been introduced by Rossi and FaiIla(21) only 25 years ago. The basic concept is very attractive, as a successful simulation by this technique produces a material that can be used, without reservations, with all types of radiations. The method is particularly important with regard to tissue substitutes for protons and neutrons, which are not, as yet, served by any of the simulation procedures described in this section.

When Rossi and Failla introduced their liquid system, they attempted to reproduce an approximate formula for soft tissue, (CsH40018N)no using a mixture of water, glycerol, urea, and sucrose, which had the formula CsH37.s018No.97. The method of arriving at this formulation was not explained in their original publication. Much later, in 1969, Frigerio and Sampson(33) used a simplified version of Gibbs' method of canonical components(78) and produced some excellent liquid systems that had the same elemental composition as the ICRU version of muscle.(S1) Even more recently, an extensive analysis of elemental equivalence has been undertaken by Constantinou, (38) and as a result over 35 tissue substitutes have been formulated. The principal findings of this research will form the basis of this section.

The method devised by Constantinou was based upon three criteria similar to those used for the basic data and extended Y methods. Firstly, the number of components in a tissue substitute should be kept to a minimum. Secondly, the most appropriate components for addition to a specified base material should be chosen from a library of compounds. Lastly, the procedure should precisely establish the proportions (by weight) of the components of a tissue substitute having the important elemental proportions generally within 0.5% (0.1 % for hydrogen) of those for the tissue.

The procedure may be resolved into a sequence of six steps:

1. From the given elemental composition of the tissue to be simulated, derive its empirical formula and/ or the number of moles needed per kilogram(79) for each constituent element.

2. Decide upon the compounds that will be used to satisfy the requirements for trace elements (Le., elements whose proportions in the tissue are less than 0.5% by weight).

3. Subtract the quantity of C, H, N, and 0 already introduced with the compounds selected in 2 from the number of moles per kilogram derived for each element in 1. Derive the empirical formula for the remaining C, H, N, and 0 and any other elements not satisfied in 2.

4. Rearrange the formula derived in 3 as the sum of two elemental groupings (parts A and B). Consider water, gelling agent, or a


selected polymer or resin system as an obligatory component (part A) for a liquid, gel, or solid substitute, respectively. Attempt to find a compound with elemental contents equal to those of part B.

5. If no single compound having the same composition as part B is found, select compounds with elemental contents bracketing those of part B. Solve a set of simultaneous equations satisfying the requirements for C, H, N, 0, etc. given in 3. Derive the proportions by weight of the compounds employed.

6. Calculate the mass density (p) of the formulated substitute using equation (20), assuming that the volumes of pure substances are additive (this is an approximation that sometimes leads to small differences between calculated and measured density values):

P = (L Wi)-l

i Pi (20)

where Wi is the proportion by weight of the ith constituent compound with mass density Pi.

As it is not always possible to find a single rigorous solution in each simulation exercise, this procedure has not been computerized. The procedure does lend itself to interactive computing, saving time by doing the repetitive calculations (e.g., solving sets of simultaneous equations) but leaving it to the originator to choose the best compound on the basis of compatibility, solubility, availability, and likely cost factors.

The problem of matching the density to within 1 % of the real tissue proves, in practice, not to be a difficult one. In the case of liquids it is nearly always possible to find alternative combinations of liquid components to formulate substitutes of lower, equal, or higher densities than that of the real tissue. In the case of cold-cure resin-based substitutes, the use of phenolic microspheres with a mass density of 200 kg m -3 solves the problem. Relatively small quantities of these hollow, nitrogen-filled microspheres added to the resins reduces the bulk mass densities to below unity.

A minor modification of this method involves the use of polymers and resins in the formulation of solid tissue substitutes. In these instances the partial replacement of oxygen by carbon is inevitable owing to the fact that these materials are very rich in carbon and poor in oxygen content. Consequently, in this quasiequivalent method, the elemental composition of the tissue and, in particular, the important hydrogen and nitrogen contents are matched, together with as much of the oxygen concentration as the polymer or resin permits. The outstanding oxygen content is now added to the remaining carbon requirement and the sum satisfied as if it were a single carbon entity.

158 David R. White and Christodonlos Constantinon

The success of all these simulation procedures in producing viable tissue substitutes will be evident in the next section, where a selection of the best materials formulated by these methods will be presented.

4. RECENT TISSUE SUBSTITUTES

4.1. Introduction

In this section the composition and radiation properties of 50 of the best and most recent tissue substitutes will be discussed. Most of the substitutes presented are the results of rigorous applications of the basic data, extended Y and elemental equivalence methods undertaken at St. Bartholomew's Hospital, London, England, since 1970. A total of 43 solid, liquid, gel, and powder systems from this work is included, of which 11 are elementally equivalent, and ten quasiequivalent materials. In addition, seven substitutes from other sources are included that either fulfill the strict simulation requirements discussed earlier or are considered to be potentially useful systems. A number of these were originated prior to 1970, but a combination of excellent simulation properties and, frequently, the continuing improvements being made by the originators provides forceful arguments in favor of their inclusion. The additional group includes three solid materials, namely, the conducting muscle and bone substitutes A150 and BllO(23.40) and the muscle material designed to match low-energy linear attenuation data, Griffith urethane. (48) The remaining four products are liquid and gels, comprising the muscle systems Frigerio liquid(33) and gel, (31) the bone substitute Witt liquid,(34) and, despite its longevity, water.

4.2. Base Materials and Additives

For a material to be acceptable either as a base material or as a corrective additive in a tissue substitute it should possess a number of important qualities. It must have a well-defined composition, but must not be carcinogenic (or suspected to be so), corrosive, toxic, explosive, volatile, deliquescent, nor unpleasant to use. Useful advice on the toxicity and carcinogenic properties of many compounds are to be found in the literature. (80--82) The material should be inert and not adversely affected by radiation. Compounds with melting points below 75°C should be excluded as additives in cold-cure polymers and resins. Solids of lower melting points could, if soluble in water, be used for the formulation of liquid substitutes; but liquids with boiling points below 65°C should be excluded.

In the search for suitable solid base materials a number of modern polymers and resins have been reported.(47,48.83.84) Polymers such as poly-ethylene, (C2H4)no TPX, (C6Hdno nylon-6, (C6HllNO)n, etc., although


resulting in excellent end products, are difficult to manipulate and necessitate specialized and elaborate manufacturing equipment. However, cold-cure systems such as acrylics, epoxy resins, polyesters, and certain polyurethanes are easier to handle with relatively modest equipment.

Four specially designed epoxy resin systems have been used extensively in the formation of many of the new solid substitutes presented here. They have the advantage of being rich in hydrogen (7.90-11.30% by weight) and nitrogen (1.60-5.62% by weight), essential in the production of substitutes for particulate radiations, as well as being tolerant to the addition of large quantities of powdered fillers. The constituents and elemental compositions of the four epoxy systems, designated CB1, CB2, CB3, and CB4, have been described in detail by White et ai.(83)

In order to modify and correct the mass densities of epoxy resin-based substitutes, the addition of low-density (-200 kg m -3) phenolic microspheres has proved most successful when substitutes with densities down to -600 kg m -3 are easily manufactured. (83) For the lower densities required for lung substitutes, the additiolJ of small quantities (-1 % by weight) of the foaming agent DCll07 produces substitutes with densities down to 300 kg m-3 .(85)

A wide range of particulate fillers has been added to these epoxy resin systems. Useful additives have included polyethylene (for increasing hydrogen concentrations and changing viscosities), the polymers acrylics, PTFE, or Teflon, PVC, and the high atomic number compounds, aluminum oxide, calcium carbonate, and calcium orthophosphate.

The manufacturing procedures for epoxy resin systems will be discussed in detail in Section 5.1.1.

Regarding liquid and gel substitutes, water is the preferred base material for many reasons. It is an important constituent of many tissues; it is one of the best solvents, has a high hydrogen content (11.19% by weight), and is readily available and cheap. Suitable additives, both miscible liquids and water-soluble solids, are to be found in abundance in the literature and reliable formulas are usually readily available. Solubility data, conversely, are sparse and often have to be derived experimentally for potentially important additives. Ethanol, glycerol, glucose, and ethylene glycol are particularly useful as corrective additives.

Organs and tissues are basically composed of fat, protein, water, and trace elements. By mixing fat and protein substitutes with water in the desired proportions, substitutes for many biological systems could be made. A suitable protein substitute in the form of gelatin, derived from bone by the hydrolysis of collagen, is already available. Unfortunately, all attempts to formulate fat substitutes that are miscible in water have so far failed.

The use of gelatin, however, has facilitated the formulation of many gel substitutes. These products usually find only short-term applications

160 David R. White and Christodonlos Constantinon

because of water loss and deterioration due to bacteriological action. Sealing gels in polyethylene bags and using trace quantities of bacteriostatic agents (e.g., sodium azide) preserves them for longer periods. Other gelling agents such as agar (a sulfated polysaccharide) and guar (another polysaccharide gelled with borax) have some uncertainties in their elemental composition and caution should be exercised in their use.

Two comprehensive, computer-based libraries of materials including organic and inorganic compounds, polymers, resins, and liquids have been compiled specially for use with computerized formulation procedures. A "compound library" specifically for the basic data and extended Y methods contains approximately 1040 materials,(36) while a more modest library for use with the elemental equivalence method contains some 280 products. (38)

4.3. Recommended Tissue Substitutes

Before a description is given of the recommended formulations, reference must be made to the identification codes and designations adopted. As the bulk of the substitutes stem from two separate treatises on tissue simulation, the codes employed by the authors in their original communications are used.

All the substitutes originating from the application of the basic data and extended Y methods(36) use a three- or four-character code of two letters designating the tissue being simulated followed by an identifying number (e.g., AP6, MS15, etc.; AP = adipose, MS = muscle).

The substitutes developed from the equivalence and quasiequivalence methods(38) use a two-part code; up to four letters designate the tissue being simulated and these are followed by sets of characters, describing the form of the end product, and an identifying number (e.g., AP/SF1, MS/SR4, IB/L1, BRN/G1, HB/P2, etc.; SF = solid flexible, SR = solid rigid, L = liquid, G = gel, P = powder).

The remaining seven substitutes are either quoted as originally communicated (e.g., AlSO) or, when the materials were not named, are referred to by the name of the primary author, followed by the type of material (e.g., Griffith urethane, Witt liquid, etc.).

Table 3 lists 50 recommended tissue substitutes and gives a general description of the constituents, their mass densities (kg m -3), and the original reference for each material. Mass densities for the powdered systems are quoted assuming a homogeneous mixture and do not take into account the air volumes present.

Suppliers of the components of the substitutes, together with details of constituent percentage weights and elemental compositions, may be found by consulting the references cited.


Table 3. Recommended Tissue Substitutes

p Tissue substitute Description (kg m -3)

Adipose tissue APjSF1 Flexible solid based on epoxy CB3 with.fillers 920

of glucose, polyethylene, and phenolic micro-spheres. A four-component formula is avail-able for trace elements.

AP3 Flexible solid using epoxy CB3. Fillers are 900 Teflon, polyethylene, and phenolic micro-spheres.

AP6 Rigid solid using low-exotherm epoxy CB4. 920 Fillers as for AP3.

Ethoxyethanol APjL2

Ethylene glycol monoethyl ether (C4H IO0 2). 930 Water-based substitute containing urea, 920 propanol, and phosphoric acid. A fourcomponent formula is available for trace ele-ments.

Blood BL/L2 Water-based substitute containing urea,

ethylene glycol, and acetic acid. Trace elements are available (five components).

Muscle A150 Polymer-based (electrically conducting) sub

stitute comprising polyethylene, nylon, carbon, and calcium fluoride.

Epoxy CB2 A three-component epoxy system curing to a semiflexible solid.

Griffith urethane Polyurethane-based material having calcium carbonate as filler.

MS/SR4 Rigid solid using epoxy CB4 and fillers urea, polyethylene, and phenolic microspheres. A five-component formula for trace elements is available.

MS15 Flexible solid using epoxy CB3 and fillers aluminum oxide, polyethylene, and phenolic microspheres.

MS20 Rigid end product made up of epoxy CB2 and fillers magnesium oxide, polyethylene, and phenolic microspheres.

MS22 Polymer-based material composed of TPX and magnesium oxide.

Frigerio liquid Water-based substitute containing urea, ethylene glycol, and glycerol. A sixcomponent formula for trace elements is available.

1060

1120

1100

1080

1060

1000

1000

1020

1070

161

References

38

83

83

37,44 38

38

23,86

83,84

48

38

83

83

71

33

______________________________ (continuedoverleaf)

162

Tissue substitute

MS/Ll

Water BL8

Frigerio gel

MS/Gl

MS/G2

SKI

SK2

B110

HB/SR4

SB3

SB4

Witt liquid

HB/P2

SB7

IB/SRI

David R. White and Christodoulos Constantinou

Table 3. (cont.)

p

Description (kg m-3 ) References

Muscle (cont.) Water-based substitute containing urea, ethylene glycol, urea, and acetic acid. A sixcomponent formula for trace elements is available. H 20. Thixotropic, clear gel comprising water, glycerol, ethanol, and Laponite XLG. A gel comprising water, ammonium formate, ethylene glycol, agar, and sodium carbonate. A water-gelatin gel containing ethanol and, if required, a six-component formulation for trace elements. As MS/Gl, but urea and propanol replace ethanol.

Skin

1070

1000 1000

1120

1060

1050

Rigid solid based on epoxy CBl. Fillers are 1100 aluminum fluoride and phenolic micro-spheres. A polymer-based substitute composed of 1090-1190 TPX (or polyethylene) and magnesium car-bonate.

Cortical bone A polymer-based, electrically conducting, material made up of nylon, polyethylene, carbon, and calcium fluoride. Rigid end product comprising epoxy CB2, urea, calcium oxide, calcium hydrogen orthophosphate, magnesium sulfate, and sodium sulfate. Rigid end product comprising epoxy CB2 and calcium carbonate. Flexible solid based on CB3. A single filler, calcium carbonate, is used. Saturated solution of dipotassium hydrogen orthophosphate in water. A seven-component powdered system, including calcium orthophosphate, glucose, and urea. A two-component powdered system made up of polyethylene and calcium sulfate.

Inner bone Epoxy resin-based (CB2) solid having fillers of calcium orthophosphate, polyethylene, and sodium nitrate.

1780

1670

1840

1750

1720

1960

1720

1150

38

1 39

31

38

38

83

71

40

38

83

83

34

38

37

38


Tissue substitute

IB7

IB/Ll

RM/SR4

RM/L3

RM/G1

BRN/SR2

BRN/L6

KD/Ll

LV/Ll

LN/SR4

LN1

LN9

Table 3. (cont.)

p Description (kg m ~3)

Inner bone (cant.) Rigid solid based on epoxy CB4. Fillers are 1120 calcium carbonate, polyethylene, and phenolic microspheres. Water-based substitute compnsmg 1140 dipotassium hydrogen orthophosphate, sodium nitrate, phosphoric acid, urea, and ethylene glycol.

Red marrow Rigid solid using epoxy CB4 and fillers of 1030 ammonium nitrate, polyethylene, and phenolic microspheres. A five-component formula for trace elements is available. Water-based substitute containing urea and 1040 glycerol. A five-component formula for trace elements is available. A water-gelatin gel containing glucose. Trace 1070 elements may be added using a fourcomponent formulation.

Brain Epoxy resin-based (CB2) solid using fillers 1040 of acrylics and polyethylene. Formula (fivecomponent) for trace elements is available. Water-based substitute containing urea, 1040 ethanol, and glycerol. A four-component for-mula for trace elements is available.

Kidney Water-based substitute containing sodium 1050 chloride, dipotassium hydrogen orthophos-phate, urea, and ethylene glycol.

Liver Water-based substitute containing potassium hydrogen sulfate, sodium chloride, urea, ethanol, and glycerol.

Lung Foamed rigid epoxy (CB4) system. Fillers include urea, polyethylene and the foaming agent DC1107. A five-component formula is available for trace elements. Flexible polyurethane foam having aluminum oxide as filler. Foamed flexible epoxy (CB3) system. Fillers include aluminum oxide and the foaming agent DC1107.

1060

300

300

300

163

References

87

38

38

38

38

38

38

38

38

38

36

85

------------------------------(continuedoverleaf)


Tissue substitute

TH/L2

BR12

AV.BR/L2

BR13

BR14

TST/L3

TSK/SF3

TSK/Ll

Table 3. (cont.)

Description

Thyroid Water-based substitute containing urea, ethylene glycol, and acetic acid. Threecomponent formula for trace elements is available.

Average breast Rigid solid using low-exotherm epoxy CB4. Fillers are calcium carbonate, polyethylene, and phenolic microspheres. Water-based substitute containing ethanol and pentanediol. Water-based substitute containing ethoxyethanol. Water-based substitute for fixing breast tissue saqlples. Additives are ethoxyethanol and formaldehyde solution.

Total soft tissue Water-based substitute containing urea, ethanol, and ethylene glycol. A fivecomponent formula is available for trace elements.

Total skeleton Flexible solid based on epoxy CB3. Fillers include calcium hydrogen orthophosphate, calcium orthophosphate, and acrylics. A three-component formula is available for trace elements. Water-based substitute containing diammonium hydrogen orthophosphate, dipotassium hydrogen orthophosphate, and glucose.

p (kg m -3) References

1080 38

970 83

960 38

970 88

950 89

1040 38

1360 38

1360 38

The tissue substitutes fall into four categories, namely, principal soft tissues, principal skeletal tissues, body organs, and average tissues.

4.3.1. PRINCIPAL SOFT TISSUES

The first 22 substitutes listed in Table 3 (AP/SFl through SK2) represent adipose tissue, blood, muscle, and skin. Solid substitutes are quoted for all the tissues except blood; liquid systems are available for all the tissues with the exception of skin; gels are only available for muscle. Of the 22 substitutes in this category, six are elementally equivalent systems


(BL/L2, Frigerio liquid and gel, MS/Ll, MS/G 1, and MS/G2) and two are quasiequivalent (AP /SFI and AP /L2).

4.3.2. PRINCIPAL SKELETAL TISSUES

Thirteen substitutes simulating the principal skeletal tissues, cortical bone, inner bone, and red marrow, are given in Table 3 (B 11 0 through RM/Gl). Solids, a liquid, and some powdered systems simulating cortical bone are presented, together with solid and liquid substitutes for inner bone and red marrow. A gelled material for red marrow is also included. Five elementally quasiequivalent materials are in this category (HB/P2, IB/SRl, RM/SR4, RM/L3, and RM/Gl).

4.3.3. BODY ORGANS

The eight formulations given in Table 3 are the recommended substitutes for brain, kidney, liver, lung, and thyroid (BRN/SR2 through TH/L2). Solid substitutes are quoted for brain and lung, while liquid systems are given for brain, kidney, liver, and thyroid. The group includes three elementally equivalent liquids (BRN/L6, LV /Ll, and TH/L2) and three quasiequivalent materials (BRN/SR2, KD/Ll, and LN/SR4).

4.3.4. AVERAGE TISSUES

Finally, seven substitutes for average breast, total soft tissue, and total skeleton are listed in Table 3 (BR12 through to TSK/Ll). Liquid substitutes, including a breast tissue fixative (BRI4), are quoted for each tissue; solids are only available for average breast and total skeleton. In this group, two products are elementally equivalent (A V.BR/L2 and TST /L3).

4.3.5. CALCULATED RADIATION CHARACTERISTICS

To assess the suitability of the 50 tissue substitutes presented in the previous section, interaction data for photons, electrons, protons, and neutrons have been computed for each substitute, using the calculation procedures detailed in Sections 2.2.1-2.2.3 and the elemental compositions in Section 7. For each material the calculated interaction data have been compared with similar data for the tissue being simulated. Six sets of ratios have been derived:

Photons ("'/ P )subS/ ("'/ P )tis and ("'en/ P )subs/ ("'en/ P )tis at 33 energy points in the interval 10 keV-I00 MeV.


Electrons (S/P)e,subs/(S/P)e,tis and (02/pl)subs/(02/pl)tis at 33 energy points in the interval 10 keV-100 MeV.

Protons (S/ p )p,subs/(S/ P )p,tis at 38 energy points in the interval 10 keV-100 MeV.

Neutrons [K/ <P (E)]subs/[K/ <P (E)]tis at 98 energy points in the interval 1 keV-30 MeV.

Putting into perspective such a large quantity of data (over 10,000 coefficients, powers, etc.) is very difficult, so a method of classifying the results has been devised, The tissue substitutes are classified as A, B, or C, depending upon the magnitude of the discrepancy in a given process when compared to the corresponding real tissue:

Class A .,; 5% errors

5% errors < Class B < 20% errors

Class C ~ 20% errors

In addition, discrepancies within 1 % are classified as A *. A complete listing of this classification for photon, electron, proton, and

neutron interaction ratios is given in Table 4. The classification is applied to each radiation type and interaction ratio. For photons, the ratios for Il-/ pare quoted in two energy ranges, namely, 10-99 keY and 100 keV-100 MeV, so separating the photoelectric region, frequently poorly simulated by quasiequivalent substitutes, from the incoherent and pair production region. If the ratios for Il-en/ P data are of a different classification; these classes are indicated in parentheses. Similarly for electrons, (S/ P)e ratios are classified, with 0 2/ pi data, if different, given in parentheses. Single classifications are used for protons [(Sf p )p], but neutrons [K/ <P (E)] are listed in two energy intervals, namely, 1-99 keY and 100 keV-30 MeV.

Inspections of Tables 3 and 4 show that many of the substitutes are tissue equivalent, being Class A * for one or more radiation types and having mass densities within 1 % of the real tissue.

Elementally equivalent substitutes give, predictably, the best results for all types of radiation. This form of categorization is by definition very stringent, and it must be remembered that some of the Class B materials (e.g., epoxy CB2, Griffith urethane, MS/SR4, 'BRN/SR2, LN/SR4) have better classification~ over restricted energy intervals. It is strongly recommended that for primary and secondary radiations being employed experimentally, only Class A materials are used if a serious attempt is being made to achieve a high degree of realism. If for some reason a material in Class B is thought to be appropriate for a given experiment, its interaction data should be calculated over the required energy range, using the elemental compositions listed in the Appendix, to ascertain whether it is acceptable over the

Tabl

e 4.

C

alcu

late

d R

adia

tion

Cla

ssif

icat

ions

a >

~

(S/ P

)e.s

ubs/

(S

/ P )

•• sub

s/

.. Q

"15

(IL/

P )s

ubs/

(IL/

P ),;

s (S

/ P )

•. tis

(S/ P

) •. ti

s [K

/ t/J (

E)]

sub

s/[K

/ t/J (

E)]

tis

Q a Q ..

Tis

sue

bein

g 10

0 k

eV-

lOk

eV-

lOk

eV-

10

0k

eV-

"15 ::r

sim

ulat

ed

Sub

stit

ute

Pha

se

10-9

9 ke

Y

100

MeV

10

0 M

eV

1000

MeV

1

-99

keY

30

MeV

;:;.

"1:

1 ::r

Pho

tons

b JIO

E

lect

rons

C

Pro

tons

N

eutr

ons

~ A

dipo

se t

issu

e A

P/S

Fl

Soli

d B

A

A

*(A

) A

A

* A

a

AP

3 So

lid

A*(

A)

A*

A*

A

B

B

:: A

P6

So

lid

A

A

A

B

C

C

; E

thox

yeth

anol

L

iqui

d B

A

A

*(A

) A

B

B

:I

.

AP

/L2

L

iqui

d C

B

A

(B)

A

A*

A

!-

Blo

od

BL

/L2

L

iqui

d A

A

A

*(A

) A

* A

* A

*

Mus

cle

AlS

O

Sol

id

C

B

B

B

A

B

Ep

ox

yC

B2

S

olid

C

B

B

B

A

B

G

riff

ith

uret

hane

So

lid

B

B

B

B

B

B

MS

/SR

4

Sol

id

C

B

B

B

A

B

MS1

S So

lid

A

B(A

) A

(B)

A

A

B

MS

20

Solid

A

A

A

B

B

B

M

S22

S

olid

A

B

(A)

A(B

) B

B

B

F

rige

rio

liqu

id

Liq

uid

A*

A*

A*

A*

A*

A*

MS

/Ll

Liq

uid

A*

A*

A*

A*

A*

A*

Wat

er

Liq

uid

A

A(A

*)

A*(

A)

A

B

B

BL

8 G

el

A*(

A)

A*

A*

A

C

B

Fri

geri

o ge

l G

el

B

A*

A*

A*

A

A

MS

/Gl

Gel

A

A

* A

* A

* A

* A

* M

S/G

2

Gel

A

A

* A

* A

* A

* A

* ~

~

(con

tinu

ed o

verl

eaf)

-.

I

~

~

Tabl

e 4.

(c

ont.)

(SIp

) •. s

ubs/

(S

I p ) •

. sub

sl

(IL

l P )s

ubs/

(IL

l P )ti

s (S

I P) •

. tis

(SI P

) •. ti

s [K

I .,,

(E)]

subs

/[K

I .,,(

E)]

tis

Tis

sue

bein

g 10

0 k

eV-

10

keV

-lO

keV

-1

00

keV

-si

mul

ated

S

ubst

itut

e P

hase

10

-99

keY

lO

OM

eV

100

MeV

lO

OO

MeV

1

-99

keV

3

0 M

eV

0 .. <I!

Pho

tons

b E

lect

rons

C

Pro

tons

N

eutr

ons

is:

Ski

n S

K1

Sol

id

A

A

A

B

C

C

!II' S

K2

Sol

id

B

A

A(B

) A

A

B

~

Cor

tica

l bo

ne

Bll

O

Sol

id

A*

A

A

B

B

B

i H

B/S

R4

S

olid

B

A

A

B

C

..

C

=

~

SB

3 S

olid

A

A

A

A

C

B

('

)

SB

4 S

olid

A

A

A

B

B

B

=- :I.

Wit

t li

quid

L

iqui

d A

(B)

A

A

A

C

C

~ H

B/P

2

Sol

id

~

Q

(pow

der)

A

A

* A

* A

* A

* A

t

SB

7 S

olid

'" (')

(p

owde

r)

A

A

A(A

*)

B

C

C

Q

Inne

r bo

ne

IB/S

R1

S

olid

B

B

B

B

A

B

I =

IB

7 S

olid

A

B

(A)

A(B

) A

B

C

::t

o =

IB/L

l L

iqui

d B

A

(B)

A

A

A*

A

Q =

Red

mar

row

R

M/S

R4

S

olid

C

B

(A)

A(B

) R

M/L

3

Liq

uid

C

B

A(B

) R

M/G

l G

el

C

B

A(B

)

Bra

in

BR

N/S

R2

S

olid

C

B

B

B

RN

/L6

L

iqui

d A

* A

* A

*

Kid

ney

KD

/Ll

Liq

uid

A

A*

A*

Liv

er

LV

/Ll

Liq

uid

A*

A*

A*

Lun

g L

N/S

R4

S

olid

C

B

B

L

NI

Sol

id

A

A

A

LN

9 S

olid

A

B

(A)

A(B

)

Thy

roid

T

H/L

2

Liq

uid

A(B

) A

*(A

) A

*

Ave

rage

bre

ast

BR

I2

Sol

id

A

A

A

AV

.BR

/L2

L

iqui

d A

* A

* A

* B

RI3

L

iqui

d A

* A

* A

* B

RI4

L

iqui

d A

* A

* A

*

Tot

al s

oft

tiss

ue

TS

T/L

3

Liq

uid

A*

A*

A*

Tot

al s

kele

ton

TS

K/S

F3

S

olid

A

A

A

T

SK

/Ll

Liq

uid

B

A*(

A)

A*(

A)

• Cla

ss A

* ,..

1 %

err

ors;

Cla

ss A

,.. 5

% e

rror

s; 5

% e

rror

s <

Cla

ss B

< 2

0% e

rror

s; C

lass

C;.

20%

err

ors.

·C

lass

ifica

tion

for

(/L

en/P

).u.

,./(/

Len

/P).

;. is

giv

en in

par

enth

eses

if d

iffe

rent

fro

m q

uote

d cl

asse

s. 'C

lass

ific

atio

n fo

r (f

J'/p

l).u

bs/

(82/p

l)'d

is g

iven

in p

aren

thes

es if

dif

fere

nt f

rom

quo

ted

clas

ses.

A

A*

B

A*

B

A*

B

A

A*

A*

A*

A*

A*

A

B

A

B

C

B

A

A*

A*

B

C

A*

A*

A

A

A

A

A*

A*

A

A

A

B

A

B

B

B

A*

A*

A

B

C

B

A*

C

A*

A

A

A*

B

B

> i .. C

l ..., Cl 51 Cl .; ~ ."

::r

III ~ 51 == ~ :I.

III ;;- ~

=- \C


energy range of interest. Materials should not be considered for use with those types of radiations where a C classification occurs.

4.3.6. MEASURED RADIATION CHARACTERISTICS

As the majority of recommended tissue substitutes presented here are relatively new, reliable measured data are sparse, with comparative techniques still under development. The relevant data available include investigations with low- and medium-energy photons, ISO-MeV protons, and fast neutrons.

A rapid method of measuring narrow-beam linear attenuation coefficients (iJ-) for liquids and solids have been reported, (90) employing a collimated beam of fluorescent x rays. Essentially monoenergetic radiation is selected by means of a high-resolution, lithium-drifted germanium detector and multichannel analyzer. The precision and accuracy of the method is claimed to be within 1-2%. Using this technique, a group of 38 systems, including base materials and tissue substitutes, have been evaluated at 9.88 and 17.44 ke V, and 6 human tissues similarly assessed at seven selected energies in the range 9.88-59.32 keV.(91) Early prototype versions of some epoxy resin-based materials listed in Table 3 were included in the evaluation, and they showed that problems with the variable chlorine concentration in these products could cause detectable effects at these low energies. Variations of 0.1 % chlorine by weight can be easily detected at 9.88 keY, which indicates that the method can be used as a screening procedure for bulk samples of resin. The maximum discrepancy between measured and calculated iJ-/ p values for all the new substitutes tested was 5%, occurring with the lung substitute LNl. The electrically conducting muscle substitute AlSO gave corresponding discrepancies of only 1 % at 9.88 and 17.44 keY. As expected, owing to inhomogeneities and minute air volumes in the tissue samples tested, discrepancies between measurement and calculation were greater than with the substitutes. Of the 23 determinations of attenuation coefficients, 16 were within 5% of the calculated data; the remainder had discrepancies of 5-25%, the larger values occurring in the 10-keV region.

A series of experiments comparing some of the recommended substitutes with real tissues using particulate radiations has been recently reported.(38) In one set of measurements, muscle liquid MS/L1 was compared with human muscle, beef steak, and pork. In another series, brain liquid BRN/L6 and solid BRN/SR2 were compared with human brain tissue. Both investigations utilized special thin-walled cells to contain the tissue samples. These were square in cross section (10 cm x 10 cm), of varying thicknesses, and fabricated from either the muscle (MS/SR4) or the brain (BRN/SR2) substitutes.


Using a lS0-MeV proton beam of S-cm-diameter and ionization detectors, relative depth doses were measured in MS/Ll. A S-cm-thick cell of human muscle was then immersed in MS/Ll in the plateau region of the depth dose curve, and the measurements in the liquid behind the cell repeated. The human tissue was replaced first by beef steak and then by MS/SR4, and the measurements repeated. The results indicated that the maximum difference in the position of the Bragg peak between any two sets of data was less than 1 mm. Similar agreement was found in the comparison of BRN/L6, BRN/SR2, and human brain tissue. (It should be noted that MS/SR4 and BRN/SR2 are Class A materials for protons from 0.5 to 1000 MeV).

These experiments were duplicated with a 9.5 cm x 9.5 cm beam of fast neutrons (maximum energy, 17 MeV; mean energy, 7.5 MeV). Central axis depth doses in MS/Ll and then with S-cm samples of beef steak, together with human muscle and pork immersed in turn in the liquid, gave data in very close agreement. The difference between any two measurements at the same point with different materials was generally of the order of ±O.S% and was in no case greater than 0.8%, which did hot prove to be statistically significant. Comparing BRN /L6 and unfixed and formalin -fixed brain tissue gave neutron depth doses behind any real or substitute brain sample not differing by more than 1 %.

The results from measurements that have been completed to date on this new generation of tissue substitutes, although encouraging, are far from complete and more experimental evaluations for other radiations and energies are needed.

5. MANUFACTURING PROCEDURES AND QUALITY CONTROL

The establishment of reliable manufacturing procedures and associated quality control tests is an essential part of tissue simulation. Elaborate and precise analytical methods of formulating substitutes are of little practical use if complementary manufacturing expertise has not been carefully evolved.

Recently developed manufacturing procedures will be described in this section. Emphasis will be given to the manipulation of cold-cure resins and polymers, as techniques for these products are well advanced and within the scope of most medical physics laboratories.

5.1. Manufacturing Bulk Materials

As an introduction to the technology of fabricating anthropomorphic phantoms, the manufacture of bulk substitute materials will be discussed


in two categories: "solids" and then "liquids, gels, and powders." It should be noted that the manufacture of natural and synthetic rubbers, favored by some commercial organizations, will not be considered.

5.1.1. SOLIDS

The major problems associated with the manufacture of most solid substitutes are the result of the need for homogeneous and, except for lung substitutes, air-free end products. The problems first came to light with the wax-based formulations. Here the conflicting requirements of the high-viscosity liquid phase necessary to minimize settlement or flotation of the particulate fillers and the low viscosity needed for efficient removal of air inevitably led to poor materials. Mixing the molten wax and fillers continuously and pouring them into molds just as the solidification process begins and the viscosity starts to rapidly increase diminishes these problems but still generally produces inhomogeneous and air-pocketed samples that are unacceptable for many applications.

These defects can be significantly reduced by the use of cold-cure resins and polymers (acrylics, epoxy resins, polyesters, and certain polyurethanes). Viscosities are easier to control by the careful choice of the type and grade of product used or the addition of finely ground powders, and the pre-cure handling times usually allow for removal of most of the trapped air.

Figure 6 shows a vacuum-mixing apparatus that, although adopted originally for epoxy resin-based substitutes,(83) has been used successfully for all of the cold-cure products mentioned earlier. It must be emphasized that for reasons of safety the apparatus should be shielded by a metal mesh or polycarbonate implosion guard and housed in a fume cupboard. As epoxy resins are employed in 19 of the 23 recommended solid substitutes, full details of the use of this apparatus will now be given for these base materials.

Accurate quantities of liquid resin, hardener, and fillers, measured most conveniently on a digital, top-loading balance, are dispensed into a Pyrex reaction vessel of a volume at least twice that of the mixture. To facilitate corrections of the masses during dispensing, the heavy viscous resin should be added first, followed by the light, low-viscosity hardener or diluent. The powders are then added, in order of decreasing mass densities. The liquids and powders are given a short manual mix in order to "wet" all the powdered fillers. The ground-glass lid is attached to the reaction vessel and a twin-bladed rotor passed through the central stirrer gland and connected to the variable-speed electric stirrer. The components are then thoroughly mixed for 5 min, ensuring that the blades of the rotor are deep in the fluid. The blades are raised to just below the top surface


Figure 6. Vacuum-mixing apparatus for cold-cure resins and polymers: A, reaction vessel; B, twin-bladed rotor; C, stirrer gland; D, air intake control; E, vacuum control; and F, variable-speed stirrer (0-1400 rpm).

of the fluid and, while still stirring the mixture, the system is evacuated to a pressure of not less than 10-2 Torr. The reduction in pressure causes the fluid to expand and rise in the vessel, but the trapped gases soon escape as the rotating blades break the resulting foam. After 3-5 min, the foam collapses and contracts. At this point the rotor is returned to its previous position deep in the fluid and the mixing and evacuation continued for a further 20 min, when the stirrer may be stopped and the reaction vessel brought back to atmospheric pressure. The mixture is then generally homogeneous and air-free and can be poured carefully into molds made of Teflon, silicon rubber, or waxed metal or plastic. To avoid damaging or distorting the sample, it should preferably be left for 48 hr before being demolded. If, however, it is imperative that the hardening time be reduced, after 15-20 hr, when the initial curing process is over, heating the sample to 80aC in an oven for 3 hr should complete the hardening.

As the curing process is moderately exothermic, the heat evolved must be removed as quickly and as efficiently as possible; otherwise excessive shrinkage (>0.1 %) and distortion occur. The use of metal molds placed in contact with cold plates and allowing cool air via a fan to blow over the mold is advantageous when masses greater than 1 kg are being cast. In any


event, large castings (e.g., transverse body sections) should use the lowexotherm resin CB4 and be preferably less than 5 kg in mass.

The particle sizes of the fillers employed have an important influence over the viscosities of the pre-cure mixtures. Mean particle sizes should generally be no greater than 50 JLm, but viscosity problems can occur when high loadings of very finely ground powders are used. For example, if a 5-JLm grade of calcium carbonate is used to make the cortical bone substitute SB3, only about 50% by weight can be added, instead of the required -67%, before the mixture becomes a dough. Establishing the ranges of particle sizes by sieving is recommended.

The preparation of polymer-based substitutes such as AlSO or BllO necessitates equipment beyond the scope of most medical physics departments. An interesting report by Spokas(47) has recently described the developments that have taken place at the Illinois Benedictine College, where the Shonka plastics(23) have been manufactured over the past 20 years. Screw-type mixers with either twin parallel or concentric screws were noted quite early on to be superior to the roller-type mixers for producing homogeneous products. Both of these are able to continuously circulate the components of the substitute, so the simple expedient of increasing the mixing time progressively improves the uniformity. The investigations led to the reciprocating screw-type machine, which is still currently used, where the ingredients in powder form are fed steadily into the mixer at one end and the mixed plastic is extruded through a nozzle at the other end. Using these techniques, superb electrically conducting substitute materials are formed, primarily for use as the walls of ionization chambers.

The development of lung substitutes is still in progress at a few centers, with polyurethane foams (LN1) and foamed epoxy resin systems (LN9) being two such results of these investigations. With LN1, (36) the constituents minus the activator are thoroughly mixed together. The activator is then added and quickly stirred into the mix. As the foaming action normally starts. in less than 30 sec, depending upon the temperatures of the ingredients, the mixture has to be poured rapidly into its closed metal mold and the mold firmly sealed so that the mixture can foam and overpack to the required mass density. With the foamed epoxy system LN9,(85) all the components, except the foaming agent DCll07, are dispensed into a beaker. The components are thoroughly mixed with a mechanical mixer, which, as only small volumes (0.1-0.2 kg) are involved, usually takes 1-2 min. The foaming agent is then added and the fluid mixed mechanically for a further 10 sec. The fluid is then rapidly poured into a mold as the foaming action normally commences immediately. Both the polyurethane and foamed epoxy lung substitutes should be prepared in a well-ventilated fume cupboard.


5.1.2. LIQUIDS, GELS, AND POWDERS

Mixing several liquid and solid compounds to make liquid and gel substitutes does not present any serious problems once the formulation procedure is complete and it has been established that the components are compatible. Difficulties arise owing to the sparse solubility data available on organic liquids and many compounds, and frequently apparently perfect, elementally equivalent substitutes have to be rejected because the addition of one of the ingredients cause the precipitation of another. Consequently, a certain amount of trial and error is inevitable.

When liquid substitutes are being manufactured, firstly the required quantity of distilled water should be measured into a beaker. The inorganic compounds necessary to introduce the trace elements are then transferred and stirred into solution one by one. It is important to wait until each is completely dissolved before adding the next, otherwise intermediate precipitates may be formed that are often difficult to bring back into solution. Urea, used in most liquid substitutes to satisfy the nitrogen requirement, is added after the trace elements and is followed by any other organic liquids, each again being stirred separately into solution.

In the case of gelatin-based gel substitutes (e.g., MS/G1, MS/G2, RM/G1, etc.) the water component with the trace elements already dissolved in it should be heated up to -80aC and then the required quantity of gelatin added and dissolved. Once a clear, uniform solution is obtained, it should be left to cool and approach room temperature, when the remaining organic component and a bacteriostatic agent (if necessary) are added. The whole mixture can then be transferred into the reaction vessel of the vacuum-mixing apparatus described in Section 5.1.1. Stirring for a short time under vacuum conditions ensures the release of any air trapped in the mixture. The substitute can now be poured into polyethylene bags (heat sealed to inhibit water loss) and left to gel before use.

A slightly different procedure has to be adopted for the Laponite-based gel substitute (BL8). The water component is dispersed into a beaker. The glycerol, ethanol, and Laponite powder are dispensed, in that order, into a reaction vessel and thoroughly stirred (mechanically) for 60 sec in order to "wet" the powder. The water component is then added to the reaction vessel, given a 20-sec fast mechanical stir, and then evacuated for about 10 sec or until the rapid increase in volume has subsided. As the contents of the reaction vessel are now rapidly thickening, the mixture must be quickly poured into suitable containers, such as soft metal gel tubes, and left to gel. After 20-40 min, depending upon the temperature of the mixture, homogeneous, air-free, clear gel samples should be produced. Gel times can be usefully extended by cooling (not freezing) the water below room temperature before use.


Powdered systems are relatively easy to prepare by dispensing all the components together into a suitable container and then ensuring adequate mixing by mechanical aids. Prior sieving of the powders is desirable. All powdered systems suffer from variable packing and the separation of components of differing mass densities during use, so periodic mechanical mixing is recommended.

5.2. Manufacturing Specialized Phantoms

Whenever complete phantoms, rather than bulk materials, are required, the manufacturing procedures of necessity become more involved. Fortunately, cold-cure resins permit the production of quite ambitious phantoms by a combination of multiple casting and machining techniques.

As the items being fabricated and their associated molds become more geometrically complex, trapping air as the molds are being manually filled at atmospheric pressure becomes a distinct possibility. The apparatus shown in Figure 7 was designed to overcome this problem. Essentially, the mixed and evacuated resin mixture is poured manually into the dispenser positioned vertically above the mold. On replacing the Perspex bell jar (and implosion shield), the mold can be evacuated and only filled by releasing

Figure 7. Vacuum-pouring apparatus for cold-cure resins and polymers: A, vacuum control; B, air intake control; C, handwheel for rotating table; D, rotating table (23 cm in diameter); E, mold; F, motorized clamp; G, vacuum seals; H, resin dispenser; I, vacuum gauge; J, Perspex bell jar (45 cm in diameter, with rubber seals); K, steel base; and L, implosion shield.


Figure 8. Partially completed mammographic test strip.

the clamped filling tube when the level of vacuum is such that the residual air present is minimal.

Such an apparatus is ideal for the production of radiographic test pieces similar to the mammographic test strip shown partially completed in Figure 8. The base of the test strip would first be made by partly filling a metal mold, which accurately forms the base and sides of the strip, with an average breast substitute (BR12). Bulk materials prepared in the adipose and muscle substitutes, say, AP6 and MS20, are machined into the square sectioned filaments and wafers shown in Figure 8. These structures, together with the silicon filaments that represent calcifications, are fixed in position on the machined top surface of the strip, with small quantities of BR12 used as the adhesive. When the resin has hardened and the items are firmly fixed, the strip can be replaced in the mold and "topped up" with a new batch of BR12. Pouring over such an irregular arrangement would necessitate the use of the vacuum-pouring apparatus to avoid air being trapped between the items as the fluid flows around them. The complete test strip contained with a tissue specimen in a replicated breast block is shown in Figure 9 (center, back).

Figure 9. Mammographic phantoms.


This basic technique has been used in the manufacture of a variety of radiographic phantoms for use in mammography, computed tomography, and image intensification. In fact, multiple casting can be used to produce anthropomorphic body sections up to 7 cm thick, examples of which are shown in Figure 11, Section 6.2. From bulk samples of suitable substitute materials, the required organs (e.g., lungs) and bones are cut and machined to shape and orientated correctly on a flat plate in a mold that defines the external body contours. A muscle substitute is poured into the mold up to the top level of the organs and allowed to harden. Once hardened, a strip of an adipose substitute can be similarly added, if required, and then both sides of the block are machined flat to provide the completed anthropomorphic section.

5.3. Quality Control

To maintain the quality of manufacture, especially when many batches of material are being made over a prolonged period of time, certain quality control tests must be performed. The two categories of investigation currently employed, density determinations and homogeneity tests, are outlined in this section.

5.3.1. DENSITY DETERMINATIONS

Mass density determinations, the simplest yet one of the most effective quality control techniques, can be performed by direct measurement or by the use of density bottles. Casting or machining rigid solids into cylinders or cubes and directly determining their mass and volume provides mass density data with an estimated error of ±O.5%. The technique should also be used for lung substitutes, with the largest convenient sample being considered.

The mass densities of flexible or rubber-like materials are more difficult to assess accurately and direct measurements should be confirmed by the use of density bottles. Trapped air can produce erroneous results and evacuating the density bottle and its contents in a vacuum-dessicator is recommended. Density bottles, the ideal tool for mass density determinations of liquids and gels, may be used for powdered systems if care is exercised in the choice of liquid employed and the solubility properties. Bulk densities of powders have to be estimated by determining the mass of powder packing a container of known volume.

5.3.2. HOMOGENEITY TESTS

Inhomogeneities, sometimes indicated as discrepancies in mass density determinations if gross, are for many applications very serious and have


to be identified and corrected. This is especially true for materials to be used for radiographic phantoms.

Xeroradiography, using low-energy x rays and fine-focus tubes, forms the basis of a powerful technique for investigating both inhomogeneities due to aggregations of high atomic number particulate fillers or pockets of trapped air. The smallest detectable sizes are -100 /Lm.

Poor dispersion of low atomic number fillers is rather more difficult to investigate. Computed tomography can be of use in these cases, where the high sensitivity for low-contrast objects can indicate unacceptable macroscopic areas. Because the CT picture element (pixel) sizes are -1 mm, resolutions are relatively low. (The CT technique, together with the lowenergy photon attenuation measurements described in Section 4.3.6, are also useful for checking the simulation properties of a formulated substitute.)

Optical transmission microscopy may be used if the substitute can be produced as a thin, translucent film, while optically opaque materials may be investigated by reflection microscopy. Careful sample preparation is obligatory for the two methods. Resolutions are excellent in both of these techniques, but can be improved still further by scanning electron microscopy, a homogeneity test of value if a very high degree of uniformity is necessary.

6. RECENT PHANTOM STUDIES

With the introduction of many new substitute materials during the past few years, a number of interesting phantoms have been fabricated and used in a variety of projects spanning both routine and research investigations. In this section emphasis will be given to the diversity of these phantoms rather than to detailed accounts of individual models, so that a broad picture emerges of the scope and potential of the new tissue substitutes.

6.1. Radiotherapy

Historically, tissue substitutes formed into composite anthropomorphic phantoms have been used for many decades to investigate the dosimetry of specific radiotherapy treatment regimes. From Westman's(92) early attempt at making a pelvic phantom composed of wax and bolus alba for gynecological measurements, the technology has developed to the extent that complete body phantoms containing real skeletons are now readily available.

Basically two types of body phantoms are being used: (a) the sliced, anthropomorphic body phantom, which attempts to realistically represent


the body tissues and their geometrical arrangement so that dose distributions may be measured at specific transverse planes, and (b) the regular geometry phantom, which relies on regular arrangements of sheets or blocks of tissue substitutes, sometimes immersed in liquids, to evaluate the effects of inhomogeneities of varying dimensions and at varying positions.

Recent investigations(38) using cobalt-60 gamma rays and 10-MeV electrons have set out to measure the changes that occur in central-axis depth dose curves when different thicknesses of lung and bone are introduced. Using ionization detectors in a water phantom, the bone substitutes SB3 and HB/SR4, both in sheet form, were found to have the same effect on the depth dose curves when placed in the beams near the entry port. These data, together with the changes in depth doses when sheets of the lung substitute LN/SR4 were interposed in the beam, were presented in a tabulation of "isodose shift factors."

Fast neutrons are now being used routinely in a number of radiotherapy centers, so the compilation of comprehensive sources of depth dose data, with and without the presence of heterogeneities, is essential for reliable treatment planning. In a program of measurements paralleling the investigations on cobalt-60 photons and 10-MeV electrons, the effects of bone and lung on the depth dose curves for fast neutrons (mean energy = 7.5 MeV) have been evaluated. The results showed that the presence of up to 3 cm of the cortical bone substitutes HB/SR4 and SB3 did not alter the neutron depth doses in the muscle liquid MS/Ll behind the bone. This is said to be due to the higher mass density of bone combined with its lower hydrogen content compared to muscle resulting in similar interactions (per centimeter) to muscle. The effect of lung was investigated using LN /SR4 (Class A for 1 keV-2 MeV neutrons) in the muscle liquid MS/Ll and, again, a tabulation of "isodose shift factors" were obtained. (93) Interestingly, the neutron depth dose measurements in water alone were closer to those in MS/Ll than the data in "unit density muscle" quoted in British Journal of Radiology Supplement 11.(94)

Pretherapeutic dosimetry on a lS0-MeV proton beam was recently reported, (38) when a selection of the new elementally equivalent and quasiequivalent materials were employed to establish some basic physical parameters. Using water, the muscle substitute MS/Ll, and the brain substitute BRN/L6, such data as the positions of the Bragg peak, peak height, peak/plateau ratios, and beam widths were determined using ionization and solid-state detectors. The effects of cortical bone and lung were investigated by introducing various thicknesses of each type of solid tissue substitute (HB/SR4 and LN/SR4) into the liquids. (These two products are class A materials for 0.5-100 MeV and 4-1000 MeV protons, respectively.) Since the accurate positioning of the Bragg peak is essential in pituitary proton therapy, a typical treatment was simulated, using a real


skull immersed in the liquid BRN/L6. The pituitary fossa was positioned at a depth that was indicated from previous measurements and showed that if the necessary bone corrections were accurately established and applied, the spatial error in the position of the Bragg peak could be made as small as 0.5 mm.

6.2. Diagnostic Radiology

Three areas of experimental diagnostic radiology have recently attracted the attention of the phantom designer, namely, mammography, dental radiology, and computed tomography.

The widespread public interest in breast cancer has led to considerable efforts to develop reliable techniques for early diagnosis both in the United States and in Europe. An outcome of this has been the production of specialized x-ray machines and image receptors for visualizing the soft tissue structure and calcifications in the breast. In order to assess the performance of these systems, mammography phantoms have been designed in many centers for use in dosimetry, quality assurance, and surveying different mammographic procedures. Phantoms have been designed by Stanton et ai.(95) in the United States and Tonge and Davis(96) and White and Tucker(97) in England. Figure 9 shows a group of epoxy resin-based mammography phantoms, including one (center, front) used during 1976-1977 in a national survey of mammography procedures in the United Kingdom.(98) Also included in the photograph (center, back) is a phantom designed jointly by the physics departments of St. Bartholomew's Hospital, London, and Memorial Sloan-Kettering Hospital, New York. This phantom contains a fixed specimen of breast tissue, together with replicated structures, and is being used as a research tool in the optimization of the mammographic technique.

An interesting dental phantom was introduced by Baracs(99) in 1977 for assessing radiographic performance and patient dose. Figure 10 shows a photograph of the complete phantom. A set of real teeth are embedded in an epoxy resin-based skeletal substitute. The teeth, which are drilled and packed with different dental fillings, contain artificial soft tissue volumes of differing dimensions to simulate the range of carious volumes found in human teeth. Dental films taken in the normal way are assessed for quality and the smallest carious volume detectable. A realistic geometry is thought to be important so that the complex absorption and scatter distributions produced during dental radiography are replicated.

The assessment of the radiographic performance and dosimetric characteristics of CT scanners requires the use of carefully constructed phantoms made from homogeneous materials having well-defined radiation properties. Useful guidelines on the physical parameters that should be


Figure 10. Dental phantom.

considered have been published by the American Association of Physicists in Medicine (AAPM)(100) and both anthropomorphic and regular geometry phantoms have been developed. (39) Figure 11 shows an array of epoxy resin-based phantoms for measuring such parameters as alignment, lowcontrast sensitivity, resolution, modulation transfer function (MTF), and dose profiles. All of the test pieces are embedded in conical plugs, which can be positioned at different locations in a variety of anthropomorphic skull, thorax, abdomen, and pelvic body sections and a set of regular geometry phantoms. Air volumes created around each plug upon insertion into the sections are minimized by the use of the gel BL8. A selection of these phantoms have been used to assess the performances of 15 differing CT machines in the United States(lOl) and a group of 29 EMI CT machines in the United Kingdomyo2)

Figure 11. A range of phantoms for computed tomography.


6.3. Nuclear Medicine and Health Physics

Despite the elaborate array of equipment and techniques available for the detection of absorbed radionuclides, the phantoms used in both routine and research studies do not appear to have developed to the level of sophistication found in diagnostic radiology and radiotherapy. All too frequently, substitutes with poor simulation properties are employed in arrangements that bear little resemblance to reality. To illustrate how the new materials and manufacturing techniques may be used to improve this situation, a few investigations in nuclear medicine and closely related projects in health physics will be outlined.

The dosimetry of internally absorbed radio nuclides can be investigated by the use of liquid-filled shell phantoms, formed into the shape of the organ being considered. By positioning radiation detectors such as thermoluminescent devices inside the phantom and adding known activities of radionuclides to the liquid organ substitute, doses can be measured at selected sites. The technique was reported in 1974(36) for a liver phantom containing technetium-99m, when isodose distributions within the phantom were plotted. A comprehensive evaluation of organ doses from ingested radionuclides is apparently under development at Oak Ridge National Laboratory.(49) Adult and pediatric phantoms based on the Fisher-Snyder mathematical models of the body have been fabricated.

A quantitative investigation of the performance of an emission tomography scanner has been undertaken with a range of specially designed solid phantoms.(103) Spatial resolutions were measured using 1-mmdiameter line sources positioned within head and body phantoms composed of the new epoxy resin-based substitutes. By having removable inserts, simulated lesions could be introduced into the phantoms, yielding data on minimum detectable lesion sizes.

Lung burdens of plutonium-239 are usually assessed experimentally with x-ray detectors of large areas (normally> 200 cm2) placed over the anterior surface of the chest. As the photons emitted from this radio nuclide are in the energy interval 13.6-20.2 keY, the flux of photons emerging from the chest are strongly dependent on the types and geometries (internal and external) of the tissues found in the thorax. The calibration of the detecting system for plutonium assessments in humans is frequently performed with anthropomorphic chest phantoms containing uniformly radioactive "lungs." Serious errors in the estimates of pulmonary contamination can result if tissue substitutes giving poor simulation at low photon energies

d · . . I d' . d h (104 105) T . I are use 10 1Oappropnate y ImenSlOne p antoms.' wo natlOna laboratories, one in Europe and the other in the United States, have been involved during the past few years in the fabrication of these specialized phantoms. At the Lawrence Livermore Laboratory (U.S.A.)

184 David R. White and Chrlstodoulos Constantlnou

Figure 12. A thorax phantom for assessing lung burdens of plutonium-239. Reproduced by permission of the National Radiological Protection Board, Didcot, England.

polyurethane-based substitutes(48) are being designed for this work, while at the National Radiological Protection Board (England) some of the epoxy resin-based substitutes and the polyurethane foam LN1 are being used(106) (Figure 12). Both models contain real skeletons and have removable outer shells in order to match as closely as possible the geometry and composition of the phantom to those of the human subject being investigated.

7. DISCUSSION

The study of the formulation and manufacture of anthropomorphic phantom materials has interested investigators for over 70 years. Until recently, these studies, although presenting the experimentalist with a multitude of solid and liquid systems and a few gels and powders, considered only a limited range of tissues. The new generation of substitutes now being produced simulate a much wider range of tissues. Furthermore, the use of modern materials in these formulations, with their improved mechanical and physical properties, is permitting the development of more relevant anthropomorphic phantoms and experimental techniques. Imaging procedures such as computed tomography are indirectly playing a leading role in promoting this work; highlighting simulation and homogeneity deficiencies with some formulated substitutes and, by the nature of their displays, helping experimentalists to appreciate the differing attenuation properties of the many components of the human body.

To facilitate the formulation of tissue substitutes to be used over extended energy intervals, it is imperative that reliable elemental composi-


tions and mass densities be available (see Table 5). Prior to 1975, suitable composition data were surprisingly sparse, but during that year ICRP introduced their reference man report, (53) which significantly improved the situation. Composition data, appropriate to tissue simulation studies, for 69 tissues and organs were presented, enabling many "new," previously unused biological systems used in simulation exercises to be considered. Analyses of the photon and electron interaction properties for these tissues have followed, and it is hoped that similar evaluations for other radiations will be attempted in the future. Despite the broad approach and usefulness of the reference man document, reliable elemental data and mass densities on diseased tissues and calcifications are still badly needed.

A gradual improvement in the analytical methods of simulation has been evident during the past few years. The procedures are now more rigorous, have greater precision, and lend themselves to computerization. The arbitrary choice of corrective fillers has been eliminated by the screening techniques provided by the basic data and Y methods. The application of effective atomic numbers (.2), originally used as single numbers characterizing interaction processes over extended energy intervals, has also been modified. By using groups of two or three Y values, each characterizing specific energies and based upon .2 values and electron densities (no), improved simulations are possible. More development work is required on these parameters, especially in their application to particulate radiations.

The emergence of cold-cure r~sins, which replaced wax as the most popular base material for tissue substitutes, is having a profound effect on both the quality of formulated substitutes and the design of dosimetric and radiographic phantoms. Basic formulas are now better defined, which in turn has led to more precise formulations. The problems of adequate filler dispersion and the exclusion of air pockets can be minimized by using vacuum-mixing and vacuum-pouring and by the careful choice of the types and grades of base materials and corrective fillers employed. As these newer and better materials and manufacturing techniques have been developing, so, correspondingly, have anthropomorphic phantoms been improving, with phantoms for checking radiographic performance in such procedures as computed tomography and mammography being the first to show this improvement. There is still room, of course, for advances in both materials and phantom design. The production of more solid substitutes for neutron applications, substitutes devoid of particulate fillers but using corrective liquids miscible in the base material, and the development of gel substitutes should be investigated. Similarly, the production of better lung substitutes would be advantageous. Regarding phantoms, the improvement of quality assurance phantoms and anthropomorphic counting standards used in nuclear medicine should be given very serious consideration.

I-'

QC

~

Tab

le 5

. E

lem

enta

l C

ompo

sitio

ns o

f the

Rec

omm

ende

d T

issu

e Su

bsti

tute

s

Ele

men

tal

com

posi

tion

(%

by

wei

ght)

T

issu

e su

bsti

tute

H

C

N

0

Na

Mg

P S

Cl

K

Ca

Oth

er e

lem

ents

Adi

pose

tis

sue

AP

/SF

1 11

.96

75.5

0 0.

80

11.1

1 0.

05

0.02

0.

07

0.45

0.

03

0.02

A

P3

11.0

8 73

.55

1.22

9.

66

0.69

F

(3.8

0)

AP

6 8.

36

69.1

4 2.

36

16.9

4 0.

14

F(3

.07)

E

thox

yeth

anol

11

.18

53.3

1 35

.51

AP

/L2

12

.12

29.2

9 0.

80

57.4

0 0.

05

0.00

2 0.

18

0.12

0.

03

0.00

2

Blo

od

BL

/L2

10

.01

9.82

2.

91

76.3

7 0.

18

0.00

2 0.

20

0.27

0.

14

0.00

4 1:1

.. M

uscl

e ~.

Q,

A15

0 10

.20

76.8

0 3.

60

5.90

1.

80

F(1

.70)

!II'

Ep

ox

yC

B2

9.

54

71.2

5 3.

03

16.0

3 0.

15

~

Gri

ffit

h ur

etha

ne

8.82

66

.40

3.55

19

.50

1.72

i

MS

/SR

4 9.

50

70.2

8 3.

48

15.5

5 0.

08

0.02

0.

18

0.50

0.

12

0.30

0.

01

.. =

MS1

5 9.

75

63.1

6 0.

94

16.0

2 0.

53

Al(

9.60

) Q

,

MS

20

8.12

58

.35

1.78

18

.64

13.0

3 0.

09

(":l =-

MS

22

10.8

1 64

.45

9.82

14

.92

a· ... F

rige

rio

liqu

id

10.2

0 12

.30

3.50

72

.89

0.07

0.

02

0.20

0.

32

0.08

0.

39

0.01

=

Q

,

MS

/Ll

10.2

0 12

.30

3.50

72

.90

0.07

0.

02

0.20

0.

32

0.09

0.

39

0.01

=

g.

Wat

er

11.1

9 88

.81

... B

L8

11.1

0 12

.71

74.1

6 0.

10

0.69

0.

002

Li(

0.02

); A

l(0.

004)

; ("

:l =

Si(

1.22

); F

e(O

.OO

l)

=

;. F

rige

rio

gel

10.0

0 12

.00

4.00

73

.30

0.40

0.

20

0.10

=

M

S/G

1 10

.20

12.5

1 3.

50

73.0

0 0.

07

O.oz

0.

20

0.09

0.

39

0.Q

1 ~ =

MS

/G2

10

.35

12.3

1 3.

50

73.0

4 0.

07

0.02

0.

20

0.09

0.

39

0.Q

1 =

=

Ski

n ;..

S

KI

7.36

60

.03

4.53

10

.94

0.12

F

(11.

55);

Al(

5.47

) ~

SK2

9.66

62

.26

18.6

4 9.

44

.. =

'C:I

Cor

tica

l b

on

e =

a

B11

0 3.

55

36.7

2 3.

97

4.53

26

.29

F(2

4.93

) =

.. H

B/S

R4

4.

45

29.0

9 3.

88

31.9

3 0.

06

0.21

10

.00

0.32

0.

06

19.9

9 'C

:I =-SB

3 3.

10

31.2

6 0.

99

37.5

7 0.

05

27.0

3 ;;.

'o::

l S

B4

3.72

32

.46

0.53

36

.12

0.30

26

.87

=- Ii> W

itt

liqu

id

4.70

56

.80

10.9

0 27

.90

~ H

B/P

2

3.41

19

.22

3.97

40

.71

0.06

0.

21

10.2

0 0.

32

21.8

0 a

SB7

4.66

27

.74

31.7

8 15

.92

19.9

0 ==

Inne

r b

on

e i :I

. IB

/SR

1 8.

73

63.1

9 2.

36

17.8

3 0.

06

2.62

0.

12

5.09

Ii>

;;

"

IB7

6.86

59

.01

2.08

24

.12

0.12

7.

81

IB/L

l 8.

65

17.2

7 2.

58

60.8

3 0.

06

2.49

4.

99

Red

mar

row

R

M/S

R4

10

.08

73.5

7 2.

16

13.7

7 0.

01

0.00

3 0.

03

0.14

0.

11

0.15

R

M/L

3

10.1

7 12

.77

2.22

74

.24

0.08

0.

03

0.15

0.

17

0.17

R

M/G

1

10.2

0 9.

38

2.36

78

.18

0.08

0.

03

0.15

0.

17

0.17

Bra

in

BR

N/S

R2

10

.69

72.3

3 1.

28

14.5

9 0.

18

0.01

0.

36

0.06

0.

30

0.01

B

RN

/L6

10

.68

15.1

4 1.

29

71.6

7 0.

18

0.34

0.

17

0.23

0.

30

Kid

ney

KD

/Ll

10.4

0 11

.35

2.74

74

.50

0.18

0.

19

0.28

0.

25

Liv

er

LV

/Ll

10.1

8 14

.40

2.83

71

.80

0.11

0.

24

0.18

0.

29

Lun

g L

N/S

R4

9.

70

70.2

6 2.

80

16.3

0 0.

17

0.Q

1 0.

12

0.22

0.

11

0.19

0.

01

Si(

0.50

) L

N1

6.00

51

.44

4.29

30

.72

Al(

7.55

) ~

ac

LN

9 10

.10

64.0

9 1.

00

15.2

4 0.

54

Al(

8.17

); S

i(0.

84)

......

(con

tinu

ed o

verl

eaf)

Tis

sue

subs

titu

te

H

C

N

0

Thy

roid

T

H/L

2

10.0

1 13

.58

2.20

73

.52

Ave

rage

bre

ast

BR

12

8.68

69

.95

2.37

17

.91

AV

.BR

(L2

11

.79

37.8

6 50

.41

BR

13

11.1

9 38

.55

50.2

6 B

R1

4

11.1

1 38

.87

50.0

2

Tot

al s

oft

tiss

ue

TS

T(L

3 10

.46

23.3

3 2.

59

62.5

4

Tot

al s

kele

ton

TS

K(S

F3

7.16

45

.50

3.08

26

.12

TS

K/L

l 7.

45

4.64

2.

94

66.9

3

Tabl

e 5.

(c

ont.)

Ele

men

tal

com

posi

tion

(%

by

wei

ght)

Na

Mg

P S

Cl

K

0.22

0.

08

0.14

0.

19

0.14

0.11

0.

Q1

0.13

0.

20

0.13

0.

20

0.31

0.

12

7.02

0.

16

0.47

0.

15

0.32

7.

00

0.13

10

.15

Ca

Oth

er e

lem

ents

1(0.

06)

0.95

0.02

10.0

3

'"""

QC

Q

C

0 II> :l.

~

!ID ~ ~ II>

::I

~

("'l :r

:I • ... Q

~ =

e. =

... ("'l =

::I i ::I ::.

::I =

=


It is imperative that all the progress made during the past few years be sustained and enhanced so that the requirements of the experimentalist can be fully met in the future.

REFERENCES

1. R. Kienbock, On the quantimetric method, Arch. Roentgen Ray 11, 17-20 (1906). 2. F. T. Christen, Messung und Dosierung der Roentgenstrahlen, Lucas, Grafe and SiIlen,

Hamburg (1913) [Abstract in Arch. Roentgen Ray 18, 280 (1913)]. 3. O. Jiingling, Untersuchungen zur chirurgischen Roentgentiefentherapie, Strahlen

therapie 10, 576-584 (1920). 4. L. Baumeister, Roentgen-ray measurements, Acta Radiol. 2,418-429 (1923). 5. A. Fernau, The absorption of (3 and 'Y rays of radium in bone and ivory, Strahlentherapie

19, 149-152 (1925). 6. W. Stenstrom, Physics and radiology, Acta Radiol. 7, 547-558 (1926). 7. R. T. Pettit and R. S. Landauer, A rice phantom for depth dose measurements, Radiology

21,484-487 (1933). 8. G. Failla, The measurement of tissue dose in terms of the same unit for all ionizing

radiations, Radiology 29,202-215 (1937). 9. P. Ott, Zur Rontgenstrahlenbehandlung oberflachlich gelagerter Tumoren, Strahlen

therapie 59, 189-223 (1937). 10. H Triibestein, The absorption and scattering of monochromatic Roentgen rays in water,

triolein, blood, skeletal muscle and subcutaneous tissue of man, Strahlentherapie 60, 330-354 (1937).

11. F. W. Spiers, Materials for depth dose measurements, Br. 1. Radiol. 16, 90-97 (1943). 12. L. Edling, Experiences with rotation therapy, Acta Radiol. 15,427-443 (1944). 13. F. W. Spiers, Effective atomic number and energy absorption in tissues, Br. 1. Radiol.

19,52-63 (1946). 14. F. W. Spiers, The influence of energy absorption and electron range on dosage in

irradiated bone, Br. 1. Radiol. 22, 521-533 (1949). 15. D. E. A. Jones and H. C. Raine, Letter, Br. 1. Radiol. 22, 549-550 (1949). 16. A. F. Holloway and H. Janzen, Preliminary depth does studies with a 70 MeV Syn

chrotron, Br. 1. Radiol. 25, 393-394 (1952). 17. J. F. Fowler, Absorbed dose near bone: a conductivity method of measurement, Br. 1.

Radiol. 30, 361-366 (1957). 18. J. L. Haybittle, Dosage distributions from "hairpin" of radioactive tantalum wire, Br.

1. Radiol. 30,49-52 (1957). 19. B. Markus, Energy determination of fast electrons from depth dose curves, Strahlen

therapie 116,280-286 (1961). 20. D. D. Lindsay and B. E. Stern, A new tissue-like material for use as bolus, Radiology

60, 355-362 (1953). 21. H. H. Rossi and G. Failla, Tissue equivalent ionization chambers, Nucleonics 14, 32-37

(1956). 22. B. Markus, Uber den Begriff der Gewebeaquivalenz und einige 'wasserahnliche' Phan

tomsubstanzen fur Quanten von 10 keV bis 100 MeV sowie schnelle Elektronen, Strahlentherapie 101, 111-131.(1956).

23. F. R. Shonka, J. E. Rose, and G. Failla, Conducting plastic equivalent to tissue, air and polystyrene, Prog. Nucl. Energy Ser. 12 1, 184-187 (1958).


24. P. N. Goodwin, Calorimetric measurements of bone/tissue absorption ratios, Radiology 75, 112-115 (1960).

25. B. Markus, Ionisationsdosimetrie und Dosisverteilungen schneller Elecktronen in Knochengewebe, Strahlentherapie 113, 379-393 (1960).

26. A. J. Stacey, A. R. Bevan, and C. W. Dickens, A new phantom material employing depolymerised natural rubber, Br. 1. Radiol. 34, 510-515 (1961).

27. A. J. Stacey, The absorption and scattering of radiation in heterogeneous media, Master's thesis, University of London (1972).

28. S. W. Alderson, L. H. LanzI, M. Rollins, and J. Spira, An instrumented phantom system for analog computation of treatment plans, Am. 1. Roentgenol. 87, 185-195 (1962).

29. S. W. Alderson, Radiotherapy phantom, U.K. Patent Specification No. 1,025,962 (1966).

30. F. W. Spiers and M. S. Chesters, Theoretical and experimental dosimetry based on bone structure, Some Aspects of Internal Irradiation, pp. 423-434, Pergamon Press, London (1962).

31. N. A. Frigerio, Neutron penetration during neutron capture therapy, Phys. Med. BioI. 6,541-549 (1962).

32. J. R. Cameron, Factors affecting the measurement of bone mineral content by the direct photon absorption technique, USAEC Progress Report COO-1422-4, U.S. Atomic Energy Commission, Madison, Wisconsin (1965).

33. N. A. Frigerio and M. J. Sampson, Tissue equivalent phantoms for standard man and muscle, Report ANL-7635, Argonne National Laboratory, Argonne, Illinois (1969).

34. R. M. Witt and J. R. Cameron, Bone standards, USAEC Progress Report COO-1422-42, U.S. Atomic Energy Commission, Madison, Wisconsin (1969).

35. J. Weber and D. J. van den Berge, The effective atomic number and the calculation of the composition of phantom materials, Br. 1. Radial. 42, 378-383 (1969).

36. D. R. White, The formulation of substitute materials with predetermined characteristics of radiation absorption and scattering, Ph.D. thesis, University of London (1974).

37. D. R. White, Tissue substitute materials, presented at the 4th International Conference on Medical Physics, Ottawa, July 1976.

38. C. Constantinou, Tissue substitutes for particulate radiations and their use in radiation dosimetry and radiotherapy, Ph.D. thesis, University of London (1978).

39. D. R. White, R. D. Speller, and P. M. Taylor, Evaluating performance characteristics in computerised tomography, Br.l. Radiol., 54, 221-231 (1979).

40. J. J. Spokas and D. R. White, A conducting plastic simulating cortical bone, Phys. Med. Bioi., to be published.

41. J. H. Harris, W. J. Tuddenham, L. Stanton, F. Glauser, and E. P. Pendergrass, The development of a chest phantom for use in radiologic dosimetry, Radiology 67, 805-814 (1956).

42. R. T. Rogers, A phantom material to represent lungs, Br. 1. Radial. 43,441-444 (1970). 43. L. J. Goodman, A modified tissue equivalent liquid, Health Phys. 16, 763 (1969). 44. Phantom Materials for Photons and Electrons, The Hospital Physicists' Association,

Scientific Report Series-20, London, England (1977). _ 45. D. R. White, Tissue substitutes in experimental radiation physics, Med. Phys. 5, 467-478

(1978). 46. N. A. Frigerio, L. L. Lavoy, and R. S. Stowe, Tissue-equivalent materials for human

phantoms and dosimetry, presented at the 24th Annual Meeting of the Radiation Research Society, San Francisco, California, July 1976.

47. J. J. Spokas, Preparation of conducting solid mixtures, U.S. Department of Energy Technical Report COO-323-40, Lisle, Illinois (1978).


48. R. V. Griffith, A. 10. Anderson, and P. N. Dean, Further realistic torso phantom development, UCRL-50007-76-1, Livermore, California (1976).

49. s. M. Garry, P. S. Stansbury, and J. W. Poston, Measurement of absorbed fractions for photon sources distributed uniformly in various organs of a heterogeneous phantom, ORNL-4979, 33-39, Oak Ridge National Laboratory, Oak Ridge, Tennessee (1974).

50. H. Q. Woodard, The elementary composition of human cortical bone, Health Phys. 8, 513-517 (1962).

51. ICRU, Report lOb, Physical Aspects of Irradiation, Handbook 85, Washington, D.C. (1964).

52. ICRP, Publication 2, Permissible Dose for Internal Radiation, Pergamon Press, London (1959).

53. ICRP, Report 23, Reference Man: Anatomical Physiological and Metabolic Characteristics, Pergamon, Oxford (1975).

54. D. R. White and M. Fitzgerald, Calculated attenuation and energy absorptiolJ coefficients for ICRP Reference Man (1975) organs and tissues, Health Phys. 33, 73-81 (1977).

55. H. Q. Woodard, private communication (1979). 56. Y. S. Kim, Human tissues: chemical composition and photon dosimetry data, Radiat.

Res. 57, 38-45 (1974). 57. J. H. Hubbell, Photon cross sections, attenuation coefficients and energy absorption

coefficients from 10 keY to 100 GeV, NSRDS-NBS 29, National Bureau of Standards, Washington, D.C., (1969).

58. E. Storm and H. I. Israel, Photon cross sections from 1 keY to 100 MeV for elements Z = 1 to Z = 100, Nucl. Data Tables A7, 565-681 (1970).

59. M. J. Berger and S. S. Seltzer, Tables of energy losses and ranges of electrons and positrons, National Research Council Report 1133, National Academy of Sciences, Washington, D.C. (1964).

60. Y. S. Kim, Density effect in dEl dx of fast charged particles traversing various biological materials, Radiat. Res. 56,21-27 (1973).

61. L. Pages, E. Bertel, H. Joffre, and L. Sklavenitis, Pertes d'energie, parcours et rendement de freinage pour les electrons de 10 keY 11 100 MeV dans les elements simples et quelques composes chimiques, Report CEA-R-3942 (Centre d'Etudes Nucleaires de Saclay, Gif-sur-Yvette, France) (1970).

62. ICRU, Report 21, Radiation Dosimetry: Electrons with Initial Energies between 1 and 50 MeV, Washington, D.C. (1972).

63. U. Oldenburg and J. Booz, Mass stopping power and path length of neutron produced recoils in tissue and tissue equivalent materials. 1. Neutron Energy,,;; 6 MeV, EUR 4786e, Commission of the European Communities (1972).

64. H. W. Barkas and M. J. Berger, Tables of energy losses and ranges of heavy charged particles, NASA SP-3013, National Aeronautics and Space Administration, Washington, D.C. (1964).

65. M. J. Berger and S. M. Seltzer, Additional stopping power and range tables for protons, mesons and electrons, NASA SP-3036, National Aeronautics and Space Administration, Washington, D.C. (1966).

66. ICRU, Report 26, Neutron Dosimetry for Biology and Medicine, Washington, D.C. (1977).

67. W. V. Mayneord, The significance of the roentgen, Acta Int. Union Cancer 2, 271-282 (1937).

68. T. Henriksen and J. BaarH, The effective atomic number, Radiat. Res. 6, 415-423 (1957).

69. G. J. Hine, Secondary electron emission and effective atomic numbers, Nucleonics 10, 9-15 (1952).


70. R. C. Murty, Effective atomic numbers of heterogeneous materials, Nature 207, 398-399 (1965).

71. D. R. White, The formulation of tissue substitute materials using basic interaction data, Phys. Med. Bioi. 22, 889-899 (1977).

72. D. R. White, An analysis of the Z-dependence of photon and electron interactions, Phys. Med. BioI. 22, 219-228 (1977).

73. D. R. White, Effective atomic numbers in the formulation of tissue substitute materials for photons, Radiat. Res. 76, 23-31 (1978).

74. G. Geske, Das Konzept der effecktiven Dichte von Phantommaterialien fur die Elektronendosimetrie und eine einfache Methode ihrer Messung, Radiobiol. Radiother. 16, 671-676 (1975).

75. G. Geske and J. Geske, Entwicklung und Aufbau eines Korperanalogen Oberkorperphantoms fur elektronendosimetrische Zwecke, Radiobiol. Radiother. 18, 187-199 (1977).

76. G. Geske, Ein einfaches Verfahren zur Bestimmung der Zusammensetzung von gewebeaquiviilenten Phantommaterialien, Strahlentherapie 153,95-100 (1977).

77. G. Geske, Ein verbessertes Verfahren zur Nachbildung von Wasser und biologischen Geweben beziiglich Iinearer Wechselwirkungseffekte mit Photonen und schnellen Elektronen, Strahlentherapie 155,407-415 (1979).

78. E. Hutchinson, Chemistry: The Elements and Their Properties, Saunders, Pa. (1964). 79. A. Holderness and J. Lambert, A New Certificate Chemistry; 5th Ed., Heinemann

Educational Books, London (1977). 80. E. Browning, Toxicity and Metabolism of Industrial Solvents, Elsevier, New York (1965). 81. Precautions for Laboratory Workers Who Handle Carcinogenic Aromatic Amines, The

Chester Beatty Research Institute, London (1966). 82. The Carcinogenic Substances Regulations, Her Majesty's Stationery Office, London

(1967). 83. D. R. White, R. J. Martin, and R. Darlison, Epoxy resin based tissue substitutes, Br.

J. Radiol. 50, 814-821 (1977). 84. D. R. White, R. J. Martin, R. M. Witt, and M. E. Masterson, Letter, Med. Phys. 5,

453-454 (1978). 85. D. R. White and C. Constantinou, Lung substitute materials, to be published (1979). 86. J. B. Smathers, V. A. Otte, A. R. Smith, P. R. Almond, F. H. Attix, J. J. Spokas, W.

M. Quam, and L. J. Goodman, Composition of A150 tissue-equivalent plastic, Med. Phys. 4, 74-77 (1977).

87. Unpublished derivation from IBl (see ref. 83) comprising (by weight) epoxy CB4 (69.32), phenolic microspheres (3.42), polyethylene (7.76), and calcium carbonate (19.50).

88. Unpublished derivation from ethoxyethanol (see ref. 37) comprising (by weight) water (27.27) and ethoxyethanol (72.73).

89. Unpublished derivation from ethoxyethanol (see ref. 37) comprising (by weight) water (24.28), ethoxyethanol (71.72), and 40% formaldehyde solution (4.00).

90. L. H. J. Peaple and D. R. White, The measurement of attenuation coefficients at low photon energies using fluorescent X-radiation. United Kingdom Atomic Energy Authority Report, AERE-R 9051, Harwell, England.

91. D. R. White, L. H. J. Peaple, and T. J. Crosby, Measured attenuation coefficients at low photon energies (9.88-59.32 keY) for 44 materials and tissues, Radiat. Res., 84, 239-252 (1980).

92. A. Westman, A simplified dosimetric method in gynaecological deep roentgentherapy, Acta Radiol. 3, 68-75 (1924).


93. C. Constantinou, Fast neutron depth dose data and lung correction factors for use in radiotherapy, to be published (1979).

94. Central Axis Depth Dose Data for Use in Radiotherapy, Supplement 11, British Institute of Radiology, London (1972).

95. L. Stanton, T. Villafana, J. L. Day, and D. A. Lightfoot, A breast phantom method for evaluating mammography technique, Invest. Radiol. 13, 291-297 (1978).

96. K. A. Tonge and R. Davis, A phantom designed to compare the quality of various mammographic images, Br. J. Radiol. 51, 731-733 (1978).

97. D. R. White and A. K. Tucker, A test object for assessing image quality in mammography, Br. J. Radiol., 53, 331-335 (1980).

98. M. Fitzgerald, D. R. White, E. White, and J. Young, Mammography practice and dosimetry in Britain, Br. J. Radiol., 54, 212-220 (1981).

99. S. E. Baracs, Some physical investigations in dental radiology, Master's thesis, University of London (1977).

100. Phantom for Performance Evaluation and Quality Assurance of cr Scanners, AAPM Report No.1, American Association of Physicists in Medicine, Chicago, I\linois (1977).

101. R. D. Speller, D. R. White, L. Rothenburg, K. Pentlow, T. Morgan, C. Showalter, and T. B. Shope, An evaluation of CT systems from ten manufacturers, to be published (1981).

102. R. D. Speller, D. R. White, A. J. Coleman, R. Veerappan, and P. Waller, A Survey of 29 EMI CT Scanners in Britain, Br. J. Radiol., 54, 232-240 (1981).

103. A. T. Elliott, G. A. Sandison, N. E. Hanson, and D. R. White, Quantitative investigation of the performance of emission tomography equipment, presented at the 6th International Conference on Information Processing in Medical Imaging, INSERM, Paris (1979).

104. D. Newton, F. A. Fry, B. T. Taylor, M. C. Eagle, and R. C. Sharma, Interlaboratory comparison of techniques for measuring lung burdens of low-energy photon emitters, Health Phys. 35, 751-771 (1978).

105. D. Newton and D. R. White, Attenuation of 13-20 keV photons in tissue substitutes and their validity for calibration purposes in the assessment of plutonium in lung, Health Phys. 35, 699-703 (1978).

106. F. A. Fry, B. M. R. Green, A. Knight, and D. R. White, A realistic chest phantom for the assessment of low-energy photon emitters in human lungs, Proceedings of Fourth Conference of International Radiation Protection Association, Paris, 2,475-478 (1977).

4

AppUcatlons of Computed Tomography In Radiotherapy Treatment Planning

MICHAEL GOITEIN

1. INTRODUCTION

In the early 1970s, computed tomography was a laboratory curiosity and a speculative industrial venture. Today, with a Nobel prize awarded to Cormack and Hounsfield, computed tomography (CT) is an accepted, almost venerable, technique. In parallel with the rapid clinical assimilation of computed tomographic scanners, there has been serious public concern about the economic implications of their widespread proliferation. This, in turn, has fostered an almost unprecedented effort to assess the clinical role and value of CT. The application of CT to the patient with malignant disease has been extensively investigated. CT has proven to be of major value in this application. In this paper, the role of body CT in the radiotherapeutic management of such patients will be reviewed with regard both to the clinical impact of CT and its influence on technical factors of treatment.

The management of a patient with a suspected malignancy is a multifaceted and complex problem. What are its main components and, in particular, in which areas is CT likely to have useful application? Among the first tasks is that of evaluating the patient's disease. In this phase, one

MICHAEL GOITEIN • Division of Radiation Biophysics, Department of Radiation Medicine, Massachusetts General Hospital, Boston, Massachusetts 02114, and Harvard Medical School.

195

196 Michael Goitein

is concerned with the detection of a tumor, the determination of its malignancy, or histology, and the assessment of the extent of its spread, that is, its clinical stage. In all three areas, CT has something to offer. The detection of soft tissue masses has proven to be one of the major contributions of CT scanning, which is all the more impressive in that previously available diagnostic techniques have generally imaged soft tissue masses poorly, if at all. The value of CT in establishing tumor pathology is less clear, but some interesting work is being attempted in this area. CT certainly has a significant role to play in assessing the extent of local disease, including the evaluation of spread of gross disease to contiguous organs, the assessment of regional nodal involvement in some instances, and, occasionally, the detection of distant metastatic spread.

These aspects of the patient's early work-up are often thought of as being purely diagnostic in nature. However, they blur imperceptably into the next two categories, which are clearly therapeutic. These are the assessment of the anatomic distribution of known and suspected disease and the choice of therapeutic technique. The latter, of course, depends upon knowledge of the tumor pathology and clinical stage. In helping to establish these and, particularly, in providing anatomic information, which, for example, would permit an assessment of the resectability of disease, CT can play an important role in forging a treatment strategy. This may often point to a combined approach in which surgical and radiotherapeutic approaches to local disease may complement one another and chemotherapy may be invoked for distant spread.

If radiation therapy is selected as, at least, a component of the treatment, the treatment approach must be formulated. This may include the choice of brachytherapy or external beam therapy and the selection of the radiation quality for external therapy, including the choice of chargedparticle beams for all or part of the treatment. For a given kind of external beam, the number, angulation, and cross-sectional shape of the external beams must be determined with a view to fully encompassing the target volume and minimizing the dose to adjacent sensitive tissues. In addressing these questions, conventional techniques have often been hampered by a lack of confidence in the anatomic relationships of the relevant tissues and in the three-dimensional distribution of disease. The use of CT has improved our ability to make these assessments and, as a consequence, in addition to its obvious role in defining a target volume, has opened the door to the design and implementation of relatively complex treatment plans. These, in particular, may well involve the use of noncoplanar beams, which, in turn, implies a need for fully three-dimensional treatment planning capabilities.

In designing a treatment plan as outlined above, the choice of the relative weightings of the various beams, their angulation, and the need

Applications of CT in Radiotherapy Treatment Planning 197

for auxiliary dose-modifying devices such as wedges or compensators all require the calculation of the dose distribution within the patient for each treatment configuration considered. The definition of the external patient contour by CT scan and the measurement of the distribution of absorption coefficients of the tissues make such dose calculations more accurate than is possible with conventional techniques.

One treatment tactic is the choice of patient position during therapy. A judicious choice can improve the margin between the target volume and sensitive adjacent structures, leading to a reduced probability of complication from the treatment. CT can help in making this choice.

CT has a particular role to play in charged-particle therapy, both with the more prevalent electron beams and with heavy charged particles such as protons, heavy ions, and pi mesons. CT is probably essential for a major fraction of patients treated with heavy charged particles.

CT can be invaluable in monitoring the response of a tumor during the course of therapy and in the follow-up of patients after the completion of their therapy. In the first case, there may be significant changes of plan as a result of monitoring the tumor's response to treatment. Radiation fields may have to be reduced, shifted, or enlarged to follow a tumor's regression or a redisposition of tissues as a result of the initial radiation; tumors previously thought to be unresectable may respond to therapy to the point that surgery can then be attempted, and the total dose might be adjusted based on the tumor response. In the follow-up of tumors, CT may enable the earlier detection of tumor recurrence, with consequent improved salvage or palliation.

In this review, the application of CT across the broad range of problems outlined above will be discussed. We will begin with a review of the clinical studies that have been undertaken to investigate the impact of CT on the process of developing a treatment plan (in the broad sense) for cancer patients. There are also many studies on the application of CT to specific disease sites and treatment problems and these are summarized in the Appendix.

2. GENERAL STUDIES OF THE IMPACT OF CT

The studies discussed here are designed to examine the influence of CT on the process of treating patients with malignant disease. Primarily, they attempt to define what influence the CT scan had on the conduct of treatment. The studies fall into two categories, retrospective and prospective. In the former, the records of patients are retrospectively reviewed and an effort is made to determine how the patient would have been treated without the information from a CT scan, and this is compared to the actual

198 Michael Goitein

treatment using all information, including that of the CT scan. The prospective studies attempt to avoid the bias that can be introduced in a retrospective review by first working up a patient without the benefit of CT and only when preliminary pretreatment decisions have been established and documented is a CT scan performed and its incremental influence on the treatment assessed.

2.1. New England Medical Center Study

The first comprehensive analysis of the influence of CT scanning in planning radiation therapy was a retrospective study by Munzenrider and colleagues.(l) Seventy-six cases were available for review. The patients in the study covered a broad range of disease sites. They were grouped according to the following categories: head and neck (8 patients), chest wall and breast (3 patients), thorax (21 patients), abdomen (27 patients), and pelvis (17 patients). Tumors were clearly demonstrated in 48 of the 76 patients (63%), suggested in 25/76 (33%), and not seen in 3 patients (4 %). The largest proportion of merely suggestive scans was in the abdomen and pelvis, where they constituted approximately one-half of the cases. These workers observed that in 32 patients the CT scan demonstrated areas of tumor involvement not suggested from other data. In 19 of these, direct extension of tumor into previously unsuspected areas was seen, unsuspected involvement of regional nodes were found in 6, while other areas of tumor involvement were seen in 7. The single most critical question that could be assessed was whether or not the tumor was adequately encompassed in the treatment field. Coverage of the tumor as determined from all available studies, except CT scan, relative to tumor coverage with CT data was clearly inadequate in 15 patients (20%), marginal in 20 patients (27%), and adequate in 40 patients (53%). These conclusions, broken down by an anatomic region, are presented in Table 1.

Table 1. Tumor Coverage without CT Data Compared to Coverage with CT Data a

Region Inadequate Marginal

Head and neck 2 Chest wall and breast 1 Thorax 5 6 Abdomen 9 6 Pelvis 1 5

Totals 15 20

"From Munzenrider et al. in ref. 1.

Adequate

6 3

10 10 11

40


The CT scans were responsible for alterations in the volume of tissue irradiated with high-level doses. For the 75 patients the treatment volume would have been unchanged in 41 (55%), increased in 16 (21%), and decreased in 18 (24%).

The value of CT was considered for each anatomic region. In the area of the head and neck, there was relatively little effect on the total volume treated or tumor coverage and it was concluded that the role of CT scanning might be to define the volume of greatest tumor involvement for purposes of boost or interstitial treatment techniques. In the chest wall and breast, the CT scan was thought to be helpful in determining the extent of disease, both for the primary tumor and for involved internal mammary nodes, and in assisting the treatment planning process. In the thorax, the CT scan was very helpful in detecting unsuspected areas of tumor involvement, even though plain films and tomograms were available. Tumor coverage was often inadequate without CT and the CT data was felt to have a significant impact on the treatment process. In the abdomen too, the CT was thought to have a significant impact on the treatment process, with a large proportion of the cases involving a change in treatment plan on the basis of the CT scan. In the pelvis, the CT scan was judged to be of lesser value because the treatments were primarily large-field techniques that involved the irradiation of the total pelvis. This question is further discussed below.

An assessment was made of the overall importance or "essentiality" of the CT data for treatment planning. Overall, treatments were judged to be essential in 41 patients (55%), helpful for a further 23 patients (31 %), and unnecessary for 11 patients (15%). Broken down by site, these data are shown in Table 2.

CT scans were obtained for purposes of treatment "follow-up" in 22 patients. (These cases represent an assessment of the tumor response to the initial "definitive" course of radiation therapy rather than follow-up scans at long times after treatment with the purpose of determining the

Table 2. Essentiality of CT Data for Treatment Planninga

Region Essential Helpful

Head and neck 2 4 Chest wall and breast 2 2 Thorax 14 4 Abdomen 16 7 Pelvis 7 6

Totals 41 23

"From Munzenrider et al. in ref. 1.

Unnecessary

2

3 2 4

11

200 Michael Goitein

possibility of tumor recurrence.) In 11 of these, tumor progression or activity was clearly shown by the CT scan, while no progression or activity was seen in 6 patients. No conclusion could be reached for five patients. As a result of these data, four patients received additional treatment that would not have been given without CT data (two received additional radiation therapy, one began chemotherapy, and one had chemotherapy changed because of demonstrated tumor activity). In four patients additional treatment that might have been given was withheld (in three, suspected recurrent mediastinal Hodgkin's disease was not confirmed, in the fourth chemotherapy was discontinued when massive recurrence of rhabdomyosarcoma was demonstrated in a recently irradiated area). Thus CT was responsible for changes in patient management after the primary therapy had been given to 8 out of 22 patients (36%).

2.2. Massachusetts General Hospital Study

The results of a prospective study of 77 patients have recently been reported.(Z) In this study the patients were selected for evaluation on the basis of a radiation therapist's request. They were first evaluated using conventional methods, including history and physical examination, review of laboratory studies, review of histopathologic material, and analysis of imaging studies, which could include regular radiographic studies, polytomographic series, xerograms, lymphangiograms, arteriograms, air contrast studies, radioisotope studies, and ultrasound. At this stage the patient's problem was defined, the treatment goals outlined, and the radiation therapy planned. The patient then underwent a pretreatment simulation procedure, the resulting portal films were marked with the estimated tumor volume, and the proposed field outlined. Where appropriate, computergenerated treatment plans were prepared using a transverse contour of the patient and a superimposed tumor volume. Pre-CT questionnaires were filled out by a physician and a physicist. These included a statement of the planned therapeutic strategy, including a statement of the intended dose and fractionation scheme, and a full specification of the treatment plan parameters. A CT scan was then performed, with sections taken at appropriate planes throughout the volume being considered for treatment as well as in regions just superior and inferior to that volume. After consultation between the radiologist and the radiation therapist, tumor volumes were outlined in all relevant sections of the study. The CT information was used to reassess the pattern of local extent of disease and the relations between the tumor-bearing tissues and uninvolved normal tissues. The patient's problem and treatment goals were then reconsidered and the treatment plan revised when necessary. Post-CT questionnaires, defining the changes made, were filled out by the physician and the physicist. Changes made subsequent to the CT scan formed the basis of the analysis.


Patients had a broad range of primary diseases and were divided into the following categories: thorax (16 patients), abdomen (14 patients), pelvis (41 patients), and extremities (6 patients).

Forty of the 77 patients (52%) experienced some change in their therapy as a result of the CT scan. The reasons for the changes are summarized in Table 3. In four cases (5%), the change was one of treatment modality. In two patients planned radiation was replaced by surgery. In the other two cases, metastatic disease was identified by the CT scan and chemotherapy instituted. In the remaining 36 cases the changes were of radiation therapeutic technique and consisted of (a) changes in the radiation quality (use of electrons, use of protons, changes in radiation energy) or in the range or angulation of treatment beams, (b) changes in the shape or size of one or more radiation fields [most of the changes (63%) were in this category], or ( c) changes in the total dose delivered. In a few instances more than one of these modifications were involved.

In 32 patients (42%) the pre-CT treatment plans did not adequately encompass the target volume as assessed with the aid of the CT scan. These cases were categorized as either constituting a "miss" or a "marginal miss." In the case of the former, some portion of the target volume lay outside the edge of one or more treatment fields, the edge being the geometric projection of the field-collimating device. A marginal miss was defined as the situation in which the target volume lay inside the field, but within 0.5 cm of its margin. There were 27 misses recorded in the study and 10 marginal misses. Thus, an out-and-out miss constituted 73 % of the instances of inadequate tumor coverage.

The patients evaluated in the study often had a two-phase treatment in which an initial course of radiation was delivered to a comprehensive field that included both the primary tumor and regional lymph node areas,

Table 3. Changes in Treatment as a Result of CTa

Number of Number of patients with change in patients with CT-induced Percentage quality, field

changes/ of patients number, or shape number with any treatment angulation or total

Region in group change modality of fields size dose

Thorax 7/16 44 1 3 4 0 Abdomen 12/14 86 1 4 7 2 Pelvis 18/41 44 2 2 16 1 Extremities 3/6 50 0 1 2 0

Total 40/77 52 4 (5%) 10 (13%) 29 (38%) 3 (4%)

"From Goitein et al. in ref. 2.

202 Michael Goiteln

Table 4. Analysis of Inadequate Tumor Coverage a

Region

Thorax Abdomen Pelvis Extremities

3 2 3 2

Main Field

Marginal Miss miss

1 0 2 1

Total 10 (13%) 4(5%)

"From Goitein et al. in ref. 2.

Boost Field

Marginal Miss miss

1 1 6 0

10 4 0 1

17 (22%) 6(8%)

followed by a "boost" treatment to the tumor or involved area. It is of interest to know whether inadequate tumor coverage was primarily associated with boost treatments or with the comprehensive fields. This analysis was made and the results are presented in Table 4. As one might suspect, more misses were recorded for the boost fields, where field margins are intended to be relatively tight around the target volume. However, somewhat over one-third of the misses were recorded for the comprehensive field. It is worth noting that in the abdomen and pelvis the misses were most often recorded for the boost fields (20 out of 27 cases).

It should be clear from the above that not all the misses that would have occurred had CT scanning not been available would have resulted in a total lack of dose to the volume of missed tumor. In many cases, for example, the entire tumor would have received a significant dose from the comprehensive field, and the missed section of tumor would only have lost the boost portion of the dose. A quantitative analysis was made of the dose deficit that would have been experienced in those cases of misses or marginal misses by a portion of the tumor and the results are expressed in the scatter plot of Figure 1. In this plot the percentage of the tumor volume outside the field is shown as a function of the dose deficit that it would have received. Only seven cases appear at the 100% dose deficit end-in these seven instances a portion of the tumor (ranging from 2 to 40% of its volume) would have received essentially no dose. The majority of the data cluster in the region of a 10-30% dose deficit to 5-40% of the tumor volume. Implications of these data for treatment outcome are discussed below.

In four patients the changes in treatment were made owing to normal tissue considerations alone. Of course, the changes made in the other patients also often resulted in changes in the total volume of tissue taken to high dose. The distribution of treatment volume changes was analyzed


60 1::1 ~ P T THORAX

~ 50 P PELVIS

~ A ABDOMEN E EXTREMITIES

I... ( I MARGINAL MISS 40 A

::;) E <:)

P P

~ ::;)

30 A A P

~

~ A Pp

P

Q;; 20 A P P

~ A P

~ to (PI EA (TI T (PI

~ ('\PIP T (E) T AP (PI , \

P 0

0 to 20 30 40 50 60 70 80 90 100

DOSE DEFICIT (%)

Figure 1. Scatter plot indicating the distribution of dose deficit versus the volume experiencing that deficit for patients whose tumor coverage was inadequate in either comprehensive or boost fields or both. From Goitein et al. in ref. 2.

and is presented in Figure 2. 85% of the changes were increases in the treated volume. The average change in volume, if the two exceptional patients with large volume increases are excluded from consideration, was an increase of 18%.

An overall subjective assessment of the value of CT scans in the work-up of the patient was made, with the CT scan significance categorized as being major, minor, or none. Overall, the CT scan was thought to have

-50

NUMBER OF PATIENTS

10

o +50 +tOO 300 850

" INCREASE IN VOLUME IRRADIATED

Figure 2. Histogram of the number of patients as a function of increase in irradiated volume due to CT. From Goitein et al. in ref. 2.

204

Region

Thorax Abdomen Pelvis Extremities

Total

Table 5. Significance of Changes in Therapy Resulting from CT Scanning a

major

4 9

13 2

28 (36%)

Number of patients for whom the CT insignificance was

minor

2 2 7 1

12 (16%)

a From Goitein et at. in ref. 2.

Michael Goitein

none

10 3

21 3

37 (48%)

major significance for 36% of the patients, minor significance for a further 16%, and essentially no significance for 48% of the cases. These data are broken down by anatomic region in Table 5. CT scans were of approximately equal value in all the anatomic regions studied.

2.3. Royal Marsden Hospital Study

Hobday et al. have reported on their prospective study of 123 patients. (3) Conventional localization and CT scanning were performed independently on the patients. Conventional localization made use of the radiation therapy simulator, on which orthogonal radiographs of the treatment area were taken. Body contours were made using plaster of Paris strips or flexible lead wire. The treatment volume was localized on the radiographs using information from clinical examinations, surgical findings, and conventional imaging techniques. The CT scans were performed taking care to simulate as much as possible the treatment situation. A flat insert was used on the CT scanner, and bolus bags, normally required for a good-quality scan, were left out in order to avoid distortion of the body contour. Scans were performed both under suspended respiration to obtain good diagnostic images and with selected sections repeated under conditions of quiet respiration in order to simulate the treatment condition. Attention was also given to details such as the degree of bladder filling (see below).

The assessment of the tumor coverage and the changes in treatment are summarized in Table 6. Inadequate tumor coverage was defined as a situation in which the "growth extended up to or beyond the margins of the treatment volume." Overall, an incidence of 29 out of 123 patients (24 %) were judged to have had inadequate coverage of the tumor when the plan had been made without the benefit of a CT scan. Substantial

Reg

ion

Pelv

is

Bla

dder

and

pro

stat

e R

ectu

m

Cer

vic

and

ute

rus

Abd

omen

P

ancr

eas

Ret

rope

rito

neum

T

hora

x H

ead

and

Nec

k

aFro

m H

obda

y e/

al.

in r

ef. 3

.

Tabl

e 6.

Su

mm

ary

of R

esul

ts o

f the

Roy

al M

arsd

en H

ospi

tal S

tudy

a

Inad

equa

te

Num

ber

of

tum

or

case

s co

vera

ge

48

8 7

3 10

1

10

7 9

4 30

6

9

Cha

nge

due

to

norm

al ti

ssue

3/5

2

/4

3

Tre

atm

ent

volu

me

chan

ged

9 3 1 9 2 8

Man

agem

ent

chan

ged

1 0 4

Tre

atm

ent

chan

ged

18/6

5 (2

8%)

13

/19

(68

%)

8/3

0 (

27%

) 3

/9

(33%

)

~ I 2. ~ if

~ f i a. I N =

CIt

206 Michael Goitein

changes were noted in all anatomic regions studied. Normal tissue changes, based on kidney and spinal cord localization, were also made in a number of cases. Extra irradiation was given to five patients with pelvic disease as a result of the CT scan. The greatest frequency of change was seen in the abdomen. Overall, treatment changes were made for 47 of the 123 patients (38%).

A comparison was made of the body contour obtained with CT relative to that obtained with conventional measurement techniques using plaster of Paris strips or lead wire, and a discrepancy in outline of more than 1 cm was seen in 22 of 123 cases (18%). Six of these cases (5%) would have led to changes of 5% or more in the midpoint tumor dose and were therefore judged to be "significant."

2.4. Other General Studies

Ragan and Perei4) studied the efficacy of CT in a treatment planning analysis of 45 patients. This was a prospective study and the analysis was primarily in terms of the dose distributions achieved before and after CT scanning information. Parameters were developed that attempted to characterize nonuniformity of dose and a "local efficiency factor." The nonuniformity factor is defined as "the average absolute deviation from the average dose within the target area," and the local efficiency factor as "the ratio of the integral surface dose from the target to that of the patient contour." Using these measures of treatment adequacy Ragan and Perez found that 31 of the 45 treatments could have been improved. In many cases they considered this improvement to be only marginal, but in appr6ximately 20% of the patients the radiotherapist felt diat there was significant improvement in the dose distribution because of the enhanced definition of tumor extent. They considered that the greatest .gain was achieved in lesions of the brain, lung, and retroperitoneal tissues. Fewer advantages were seen for lesions of the oral cavity, oropharynx and hypopharynx, breast, and pelvis.

Schlager et al. (5) have studied the application of CT to 98 patients. They found that the treatment plan was changed in 38% of the cases. They presented a number of case reports and have concluded that "the CT scan has been of unquestionable value in radiation therapy treatment planning."

2.S. Discussion

The first thing to be said about these studies is that the frequency with which treatment plans had to be changed on the basis of CT information is quite astounding. Table 7 summarizes the conclusions of these studies. Overall, from a quarter to a half of all plans were significantly affected by

Aut

hors

Mun

zenr

ider

et a

l. a

Goi

tein

et a

l. b

Hob

day

et a

l. C

Reg

an a

nd P

erez

d

Sch

lage

r et

al.

e

Tab

le 7

. Su

mm

ary

of S

tudi

es o

f the

Inf

luen

ce o

f cr f

or P

atie

nts

Und

ergo

ing

Rad

iati

on T

hera

py

Num

ber

of p

atie

nts

75

77

123 45

98

Tum

or c

over

age

Inad

equa

te,

20%

M

argi

nal,

27%

Mis

s, 3

1%

Mar

gina

l, 1

0%

Inad

equa

te,

26%

Ove

rall

ass

essm

ent

Ess

enti

al,

55%

H

elpf

ul,

31 %

Maj

or s

igni

fica

nce,

36%

M

inor

sig

nifi

canc

e, 1

6%

Cha

nge

in p

lan,

38

%

Sig

nifi

cant

im

prov

emen

t, 2

0%

Cha

nge

in p

lan,

28%

aln

ref

. 1.

<I

n re

f. 3

. 'I

n r

ef.

5.

bIn

ref.

2.

dIn

ref.

4.

i f a. ~ 5" j i .. Ii ! I- IIC

I ~ S

208 Michael Goitein

CT scan information. Of course, one wishes to know precisely what constitutes a "significant" change of plan and the skeptic might suggest that the changes involved were of a relatively modest nature. However, a major proportion of the changes were due to inadequate tumor coverage in the pre-CT treatment plan and it is hard to think that such an error in the treatment plan is other than important.

Perhaps the strongest criticism that can be made of all the above studies is that the CT scan findings are accepted as "correct." Of course, where a CT reading was negative or highly equivocal, it probably would not have been used as a basis of a treatment plan change. However, when a positive interpretation of a tumor mass was made, it was generally accepted in the analysis as being correct. There is, as emphasized below, a lack of pathological correlation with CT findings that would allow one to assess the frequency of false positive CT readings. Certainly, investigators are aware that edematous or fibrotic changes may exist in and adjacent to a tumor and may be misread as tumor. There have been anecdotal cases in which CT findings suggested tumors larger than were found upon surgical exploration. Thus, there may be some overreading of the CT scan, which will tend to enhance the value of the CT scan as reported in these studies. On the other hand, it seems quite unlikely that the correction due to this effect is a major one.

A second area of criticism of these studies relates to the selection of the patients. In none of the studies were patients randomly selected. This, of course, is not a criticism of the studies as such, but presents a problem of extrapolating their findings to a general population of radiation therapy patients. The studies were generally limited to the evaluation of patients being assessed for curative radiation therapy and did not speak to the value of CT to the palliative patient. This is a very important category of patient and the potential benefit of CT in such cases may be substantial. Local control of disease can be of major value to such patients. In the Massachusetts General Hospital (MGH) study there were patients whose eventual treatments based on the CT scan would have had to be considered essentially palliative but for whom it was considered that the CT scan offered significantly improved palliation. Moreover, in studies of specific anatomic sites (see Appendix to this chapter), palliative patients have been considered and it has been demonstrated that CT scans have had an important role in evaluating the possibility of attempting curative treatment, with some patients being "converted" from palliative to curative intent by virtue of their CT scan and others experiencing the reverse. The fact that the studies gave results so very consistent amongst themselves lends support to the proposition that patient selection bias could not have been very strong in the individual studies. Nevertheless, it must be admitted that the studies probably selected patients with more difficult management prob-


lems. It will be an important goal of future research to refine the definition of the best population for the application of CT scanning.

Changes in the overall management of patients, including changes of treatment modality (such as from radiation therapy to combined radiation therapy and surgery), constitute a very important impact of CT. This category is certainly underestimated in the above studies. This is because such management decisions are often made before the point at which a patient would become eligible for the studies, but, nevertheless, would sometimes have been based upon a prior CT scan finding. This function of CT scanning may be one of the most important contributions it has to offer.

In considering these studies on the application of CT to radiation therapy, it should be recognized that they have generally been performed with early generation equipment, either with scanners taking several minutes per slice or, at best, with I8-sec scans. Moreover, the spatial resolution and display capabilities of these scanners have generally been inferior to what is presently available. It is unclear how great a benefit will accrue from the use of faster scanners with improved spatial resolution, but preliminary experience suggests that this is likely to be significant.

The conclusions from these studies are very much dependent upon the treatment policies in effect in the individual institutions. For example, Munzenrider et al. (1) and Ragan and Perez(4) concluded that CT scanning was of significantly lesser interest in the assessment of pelvic disease. On the other hand, the Massachusetts General Hospital study(2) and that of Hobday et at. (3) found a major role for CT in the pelvis. This is because the first two studies were limited to the evaluation of the treatment of the whole pelvis through generous portals. In the MGH study the major proportion of the changes due to CT were seen in the boost portion of the treatment. It is therefore important that the detailed conclusions of these and similar studies be carefully assessed with respect to the institutional treatment policies. The availability of alternative imaging techniques is also an important consideration. For example, the relative value of CT and ultrasound in planning radiation therapy remains an important question that has yet to be critically analyzed. It is entirely possible that the conclusions of the above studies may not have been based upon the best available ultrasound techniques as part of the "conventional" capabilities of the institution.

It is of great interest to determine what anatomic sites or disease situations would profit from the application of CT scanning. As already indicated, this depends in part upon the treatment policies of the institution. In the Appendix a number of studies are presented that relate to specific anatomic sites and, as a reading of the studies indicates, the application of CT is a complex matter, with its contributions varying widely from situation

210 Michael Goitein

to situation. What seems to be clear is that no strong patterns are emerging that would suggest CT to have an overriding value or lack of value in any specific site. There is perhaps a suggestion that abdominal and pelvic diseases are strong candidates for CT assistance, but the data indicating that this advantage is substantially greater than at other sites are not compelling. The present data favor the proposition that CT scanning is of great value over a broad range of anatomic sites and treatment problems.

Criticisms of CT scanning that are frequently voiced are (a) that only gross disease is, at best, imaged, (b) that the radiation therapist will be falsely encouraged to design the radiation field too tightly around what may not be an accurate assessment of the disease process, and (c) that there will be a consequent increased incidence of marginal failure due to CT. These suggestions are not without merit, but should be thought of as cautionary rather than prognostic. Of course, only gross disease is seen in CT scanning. This is true of almost every imaging technique and is well known and understood by practicing therapists. One of the major considerations in developing a treatment plan is the pattern of likely spread of disease, and treatment fields are always designed to encompass not only the known gross disease but also the areas thought to be at risk for microscopic involvement. It is interesting that the data from the above studies suggest that CT scanning results in no net reduction in treatment volume and may result in a net increase. Specifically, Munzenrider et al. found, on the average, no net change and Goitein et al. observed a net increase in the treated volume as a result of CT scanning. It is all in all unlikely that the inclusion of an additional diagnostic tool, especially one as accurate as CT, will degrade the information available and therefore worsen the plan of treatment. It is certainly possible that in some instances a false negative reading of CT could lead to a treatment failure, but the weight of the evidence strongly suggests that the net influence will be positive.

3. SITE-SPECIFIC STUDIES OF THE IMPACT OF CT

A variety of studies have begun to chart the application of CT in specific disease sites. The design and thrust of these studies have varied widely, reflecting the broad range of applications of CT to specific circumstances. Obviously, there is a wide overlap between primarily diagnostic studies and those with therapeutic applications. There is a vast literature on the former,(6,7) which would neither be possible nor appropriate to survey here. A number of therapeutically oriented studies are summarized in the Appendix. It is impossible to do justice to the subtleties of these studies and their analyses in the short synopses given here, although it is


in precisely these subtleties that these studies make their major contribution. One feature that emerges is that the areas of applicability of CT or lack of it are quite different at different anatomic sites and, even, in different institutions where treatment policies and diagnostic capabilities may vary widely.

It is fair to say that the site-specific studies confirm in a broad outline the important role for CT suggested by the general studies presented in the previous section. As a more critically focused review of CT is made, the picture is filled in with greater detail, leaving the overall impression that CT is a very powerful tool for a broad range of management problems that span the broadest possible spectrum of anatomic sites and disease entities.

4. MONITORING OF TUMORS DURING TREATMENT AND IN POSTIREATMENT FOLLOW-UP

A tumor is not a static entity but will change with time. Knowledge of such changes can be most important in guiding therapy. CT is perhaps unique in its ability to accurately monitor gross changes in a tumor. This is because many other studies, while they may be very specific for detecting the presence or absence of tumor, usually do so inferentially and do not demonstrate the full three-dimensional extension of disease. Sequential CT studies are potentially of value for

1. Monitoring tumor regression with a view to tailoring the radiation therapy treatment fields accordingly (Le., reduced boost fields).

2. Monitoring response to an initial dose of radiation therapy in order to assess the resectability of a lesion after a preoperative course of treatment.

3. Monitoring the response of a tumor towards the end of planned therapy, with the idea of reducing or increasing the total dose based on tumor response.

4. Follow-up of patients after treatment to increase the chance of early detection of recurrence, with consequent improved salvage or palliation.

Interest has been widely expressed in all of the above uses of CT and some preliminary experience has been reported. However, it is fair to say that the assessment of the role and value of CT in these areas is still in its preliminary stages. These questions have an added importance in terms of the health planning and economic aspects of CT scanning since, if CT is to be used extensively in the follow-up of patients, it will increase manyfold

212 Michael Goitein

the number of scans performed per patient and, conversely, substantially decrease the patient population that can be supported by a CT scanner.

A case report of the use of CT to monitor regression of tumor during therapy was presented in the MGH study.(2) This was a patient with Hodgkin's disease whose lesion appeared to regress with conventional radiographic imaging and for whom a reduced field was planned. The rescan with CT demonstrated that although a significant bulk of tumor had indeed regressed, there was residual gross disease that would have been excluded from the reduced treatment fields. Such findings haye been reported by others in a number of situations. For example, Yu et al.(8)

looked at their patients with bladder cancer with a view to demonstrating residual and recurrent disease, which was observed in 10 out of 25 patients (40%). Hodson et al.(9) had similar findings in their series of bladder cases.

A limiting feature in some treatments is the degree of patient motion possible; variations can occur in setup from treatment to treatment and from simulation to treatment and from changes in anatomic relationships prior to and during the course of therapy. Clearly, CT scanning can materially contribute to information about these matters. For example, a patient has been treated at the MGH who had a perirenal tumor that was excised prior to therapy. Treatment featured, in part, a posterior proton beam designed to irradiate the tumor bed and extend a few millimeters into the kidney but spare the bulk of the kidney-a particularly important goal, since the patient's other kidney had been previously removed for unrelated reasons. There was some uncertainty as to whether postsurgical sequelae would cause the anatomic relationships between various tissues of interest to vary with time. The presence of both edematous and fibrotic tissues was thought to pose a potential problem. Partly to investigate this matter (as well as for dosimetric purposes), the patient was scanned on five separate occasions and the charged-particle path length was shown not to have changed by more than a few millimeters. Such a use of CT would certainly not be customary, but is a valuable capability in specific cases where very careful alignments and measurements are critical to the therapeutic strategy.

The role of CT in the assessment of tumors after therapy has been analyzed for a number of lesions. Howells et alYO) reviewed 53 CT scans on 38 patients with a diagnosis of cervical cancer. The CT scan was positive in 80% of the patients who had demonstrated recurrent disease. In 75% of the positive studies multiple tumor sites were detected. In 15% of the cases para-aortic adenopathy was discovered in addition to the primary disease, while distant recurrences were only observed in 20% of the cases. Husband et alYo analyzed CT for recurrent rectal tumors. Of ten patients with a mass on the initial baseline CT, only two had regression of their


tumors three to four months after treatment. The ten patients with no observable tumor on initial evaluation all remained negative, with a maximum follow-up of 18 months. Husband et al. concluded that CT can localize disease and offer follow-up for patients at high risk "in hope of detecting recurrence which would otherwise remain undetected." Brizel et al., (12) in their study of the radiotherapeutic applications of pelvic computed tomography, saw changes in tumor volume in 59% of 27 patients scanned after treatment and modified therapy for 8 (30%) of this group. Lee et ai.(13) analyzed the role of body tomography in the clinical follow-up of abdominal masses. They concluded the following:

our data show that CT is a useful non-invasive technique for following the progression or regression of masses in the abdomen or pelvis after therapeutic intervention. It can reveal clinically unsuspected progression of recurrence. Furthermore, it can provide reassurance of the original diagnosis of normalcy. Serial CT scans alone were helpful in elucidating the clinical status of the mass or confirming normalcy in 71 % of all cases. In another 16% CT added new information to the cases that other diagnostic studies had not. In only 6 cases (6%) was information obtained from CT either not helpful (2 %) due to technical difficulty or misleading (4%). Analysis ... shows serial CT scans to be helpful in 96% of patients with lymphoma, 90% of the patients with pancreatic or pelvic masses and 72 % of the patients with retroperitoneal disease.

The investigators who have followed patients with CT scans have generally been enthusiastic about its possible role in this area. However, it is in the very nature of a follow-up study that a long time must elapse before objective data are forthcoming and a critical assessment of the benefits of CT in this situation cannot yet be made.

5. TREATMENT OUTCOME

The studies discussed so far have analyzed the effect of CT scanning on the process of treatment. No study to date has measured the impact of CT scanning on the outcome of the treatment of malignant disease. It generally takes many years to establish the outcome of such treatment and the result is so susceptible to bias in the selection of patients that a controlled randomized prospective study is usually needed. Three questions are now addressed: (1) What kind of outcome is one interested in, (2) is a controlled randomized study feasible, and (3) what can be said about the likely impact on the outcome on the basis of the present data?

214 Michael Goitein

It is customary to analyze the outcome of treatment in terms of long-term (perhaps 5-year) survival of the patient. This is, of course, a highly desirable goal and also the most readily quantified one. A second measure is that of local control. Since the primary goal of localized therapy, such as radiation therapy or surgery, is to achieve local control of disease, and since metastatic spread of disease is a frequent occurrence in many cancers, the extent to which local control has been achieved can be a better measure of the success of the therapy. Local control of disease can be an important benefit to the patient, even if subsequent metastases lead to death. However, it is important to emphasize that these measures of outcome are highly restricted and do not address a number of very important matters all of which may be strongly affected by the data from CT. They include the quality of life after treatment, both for long-term survivors and for those destined to die from their disease, the preservation of function, and the avoidance of discomfort and pain. Thus the assessment of the complications of treatment is another important aspect of treatment outcome.

Can a randomized controlled study of the value of CT scanning be undertaken? The first question is whether a sufficient number of patients, sufficiently free of bias, can be studied to measure any improvement in outcome with adequate statistical significance. This depends entirely on the extent of the benefit offered by CT, a matter that is further discussed below. A second question is the extent to which such a study would be ethical. The execution of randomized control studies depend on the proposition that the treatment and control arms are, in the considered judgment of the investigators, likely to be of comparable benefit to the patient. To perform a CT study one would have to compare patients treated with the benefit of a CT scan to a randomly selected control group that did not receive a scan. Is it possible to assert that this latter group would receive equally good treatment? The high incidence of treatment changes based on CT that has already been documented makes it unlikely that this could be asserted with confidence. For such a proposition to be tenable, one would have to believe that there was a substantial possibility of an inferior treatment outcome from some cause that would offset the likely benefits already perceived for CT. In many comparisons of treatment protocols, such effects are readily imaginable. However, we are dealing here with the addition of a diagnostic tool in the assessment of a patient's condition and few would argue that the incremental information that would on the average be gained from this tool would be deleterious. In fact, the essence of the question that one wishes to answer is not whether CT scanning is of benefit to the patient, but whether that benefit is worth the cost of the additional procedure. One wishes to quantify the benefits in order to establish the cost-benefit relationship. This is a very different matter from the conventional controlled trial of competing therapeutic modalities. The question


is one of economics rather than one of medical management. If the cost of CT scanning were one-tenth or even one-fifth of what it now is, the question would be entirely without interest and no one would propose a randomized trial. Under these circumstances, it seems questionable whether a randomized trial is appropriate.

Given the difficulty of conducting a randomized trial, can one get any suggestion of the likely influence of CT scanning on treatment outcome in other ways? The MGH study has been analyzed(14) in an effort to extract an estimate of outcome from the data. The analysis was provoked by the quite startlingly high incidence of treatment changes due to the CT scans. In the MGH study inadequate tumor coverages occurred in 42% of the patients and, as discussed above, comparable numbers have been obtained by several others. At first blush, this number does not appear to be consonant with current clinical experience. Nowhere near 42% of patients treated with curative intent suffer a marginal recurrence. How is one to understand this and, more generally, how can one proceed from the measures of the impact of CT to an estimate of outcome?

Three factors will tend to demagnify the influence of inadequate tumor coverage on outcome. First, as suggested in Figure 1, not all cases of inadequate tumor coverage involve a complete loss of dose to a portion of the tumor. In fact, in the majority of the cases, in the absence of CT scan, the missed portion of the tumor (perhaps one quarter) would have been subject to only a fractional (on the order of perhaps 20%) dose deficit. When, without CT, part of a tumor would have had essentially no dose, a local treatment failure would almost certainly have resulted. However, when part of the tumor would have received only part of the dose, a treatment failure is not inevitable. This is analyzed further below. Second, even if adequate tumor coverage is in effect, one cannot guarantee local control, since it is rarely possible to deliver a sufficiently high dose to make this certain. Rather, adequate tumor coverage is likely to be associated with some probability, on the order of 50% in the MGH study, of achieving local control. Thus an improvement from inadequate to complete tumor coverage in a given patient cannot increase the probability of local control by 100% for that patient, but, at most, by a lesser amount, equal to the local control probability with adequate tumor coverage. Finally, even when long-term local control is achieved, there is a significant likelihood that the patient may develop metastatic disease, and this will reflect itself in a lesser improvement in the probability of long-term survival as compared to the improvement in the probability of local control. These three factors have been analyzed in detail in the MGH study, as will now be discussed.

Prior to the CT scan the radiotherapist was required to make a number of predictions about the patient's likely.outcome. These included an estimate of the expected local control frequency in the event that adequate tumor coverage was obtained and the likelihood that the patient would

216 Michael Goitein

develop metastases and die within five years of treatment even if local control were established. The physician's estimate was considered as being the most reliable measure of these quantities and was available for each patient, based upon the considered assessment of his or her individual circumstances.

Those cases, lying to the extreme right of Figure 1, in which part of the tumor would, without CT, have received essentially no dose, represent no problem for the analysis. One can presume with good confidence that they would have led to local failure. However, the majority of cases did not lie in this category. There is a dearth of clinical or animal data that would predict the outcome of treatment when a tumor is inhomogeneously irradiated. In order to assess this problem, a partial-tumor dose response

0

-10

-20

""' (! 0

)" -10 .... .....

...,j ...... -20 a:. ~ -30 C

~ -40 ...,j

c -50 ~ ~ 0 ~

~ c ~ -20 10.::

~

~ ~

-40

~ ~

-60

10 100

TCPo =50%

10

dose deficit /

/ 50%

TCPo =80%

2 3 5 7 10 20 30 50 70 100

PERCENT OF TUMOR VOLUME WITH DOS£ DEFICIT

Figure 3. Change in tumor control probability (TCP) as a function of the volume experiencing a dose deficit and the magnitude of that dose deficit. Results are given for full coverage TCPs (TCPo) of 20, 50, and 80%. From Goitein in ref. 14.

Applications of CT in Radiotherapy Treatment. Planning 217

model was developed. It was based upon an assessment of the control probability for a fraction of a tumor, which in turn was calculated on the basis of an assumed dose response relationship for the tumor as a whole and a presumed logarithmic dependence of tumor control probability on the number of donogenic cells. Armed with such a model, it is possible to calculate the change in tumor control probability when a variable fraction of a tumor volume is subject to any degree of dose deficit. The results obtained are reproduced in Figure 3 for three levels of tumor control probability (TCP). These predictions depend on the steepness of the dose response curve assumed for a given tumor. An analysis was made of the available clinical data and a slope such that the tumor control probability increased from 20 to 80% with a 40% increase in dose was assumed in the calculation. The sensitivity of the calculation as a whole was examined with respect to this parameter. Armed with this calculational technique, it is possible to assess for each patient what the likely change in tumor control probability would have been when the CT information was invoked to avoid what would otherwise have been a partial dose deficit to a partial tumor volume. This number is sensitive to the point on the dose response curve appropriate to the lesion, and this was assessed on a patient-by-patient basis, using for the estimate of tumor control probability the radiotherapist's estimate made prior to the CT scan. The net result of these calculations for the MGH study was that the average increase in tumor control probability attributable to CT for those patients for whom CT demonstrated inadequate tumor coverage would have been 14.5%-that is, a change from 37.5% without CT to 52% with CT. Averaged over all 77 patients in the study, the average increase in local control probability was 6 % . When the therapist's estimate of the probability of developing metastatic disease was considered, the average increase in five-year survival was estimated to be 3.5%. The distribution of improvements in tumor control probability is indicated in Figure 4. The sensitivity of the results to the assumed slope of the dose response curve was quite small, as summarized in Table 8.

Of this improvement, 3.4% of the local control improvement and 1.7% of survival improvement came from patients having 100% dose deficits in some fraction of their lesion. This contribution is essentially model independent. Patients having marginal tumor coverage contribute a further 0.4% to the local control figure and 0.2% to the 5-year survival figure.

The estimated impact on outcome seems, at first glance, to be rather modest. It is obviously a substantial decrease from the 42% figure that is the proportion of patients with inadequate tumor coverage. However, for those patients whose outcome was improved by CT, the improvement in tumor control was 14.5% out of a possible 52%. That is, for these patients

218

CHANGE IN TCP DUE TO CT (%)

tOO

20 t5 to 5 0

NO. PATIENTS

MAX POSSIBL E CHANGE IN TCP

•

50

Michael Goitein

• 100

ESTIMATED TCP WITH ADEQUATE

TUMOR COVERAGE(%)

Figure 4. Scatter plot indicating the distribution of TCP improvements resulting from CT versus the expected TCP after CT. The left-hand-side histogram shows the frequency distribution of the TCP improvement (it is the projection of the scattergram on the left-hand axis). From Goitein in ref. 14.

CT recovered at least one-quarter of the possible benefit of their therapy. Averaged over the entire population of 77 patients, CT recovered approximately 11 % of the possible benefit of their therapy.

These calculations are of interest for three reasons. First, the net effect is small enough that the improvement in outcome due to CT would be hard to observe and this suggests that exceptional care will be needed if clinical studies are designed. Second, they put the studies of the influence

Table 8. Sensitivity of the Estimate of Outcome to the Slope of the Dose Response Curve a

Slope of dose response curve ~80120 (%)

20 40 75

"From Goitein in ref. 14.

Average improvement resulting from CT in

local control probability

8.7 6.1 4.7

5-year survival

5.1 3.4 2.5


of CT on radiation therapy planning into perspective. The frequencies of change of treatment plan are startlingly high and somewhat at variance with one's impression of the adequacy of treatment technique. The 6% improvement in local control seems much more in line with experience. Finally, the gain due to CT suggested by the model, while hard to measure and hence to confirm, is substantial and worthwhile. This is now discussed further.

6. COST EFFECTIVENESS

Is CT worth the cost? This beguilingly simple question is very much on the collective public mind. To answer such a question definitively requires more data on the effects of CT scanning than presently exist. However, it is possible to calculate some order-of-magnitude numbers that bear on the question.

As discussed above, the potential benefits from the application of CT scanning in radiation therapy cover a broad range of possible sequelae. These include the possible influence of CT in avoiding some complications of treatment and thereby decreasing morbidity of treatment and improving the quality of life for treated patients, the improved palliation of terminal disease, and the potential for early detection of recurrent disease, thereby enhancing the possibility of salvage for some patients and improving palliation for others. To concentrate, as will now be done, on the possible improvements in local control and long-term survival is to underestimate the possible benefits, and it is important, since regulating agencies tend to focus exclusively on "cure," to realize and emphasize this point. In addition, improvements due to improved dose calculational capability and other influences of CT on the design of treatments are not included in the analysis. In the following discussion, the calculations of the previous section on the possible improvements in local control probability and 5-year survival as a result of CT scanning are accepted as reasonable estimates of the benefits from CT. It would be much better to use clinical data, but they are not available at this time. The restriction to these benefits are due to data being unavailable for the other benefits just mentioned. This analysis has been presented in the literature. (15)

One approach to such a question is to compare the cost of a CT scan with the total cost of therapy and juxtapose the ratio of these to the ratio of the improvement in outcome due to CT to the expected overall outcome of therapy. If the proportional cost of CT compared with that of therapy is less than the proportional improvement in outcome, one might conclude that CT is cost effective. The following analysis is based on current costs in the United States.

220 Michael Goitein

A body scan costs in the neighborhood of $250. (16) The cost of therapy, which is less well established, has been estimated to be $12,000.07 ,18) This is an average figure for patients whose initial therapy is successful. When there are complications of treatment, such as might accompany a failure of local control, the total cost of treatment was estimated as $36,000. To be conservative, the lower figure is used in the present calculation, leading to the estimate that a CT scan represents 250/12,000 = 2.1 % of the cost of treatment. The proportional benefit it bestows is the ratio of the improvement in local control due to CT (6%) to the probability of local control for overall therapy (estimated to be 52% in the MGH study)-that is, CT is estimated to improve the outcome of therapy by 6/52 = 11.5%. (The same figure would be reached if long-term survival, rather than local tumor control, were considered, since both the numerator and denominator would be multiplied by the same factor.) Thus, in this calculation, CT is cost effective, since its proportional benefit, 11.5%, exceeds its proportional cost, 2.1 %, by a ratio of 5.5. Put simply, if one pays $12,000 for a 52% chance of obtaining local tumor control, a further cost of $250 for a CT scan to increase the probability of tumor control by 6% looks like a bargain.

In this calculation the cost of CT has been equated with the charge for the procedure. This, of course, will tend to overestimate the cost. Furthermore, the estimate is more favorable to CT than the factor of 5.5 suggests, since one would normally expect that the costs of marginal improvements in therapy would be greater than the average cost of therapy.

Another way of looking at the problem is to estimate the cost savings the $250 expense of a CT scan may effect. The figures for the cost of therapy given above imply that the average cost of treating a patient whose initial treatment has failed is $24,000 more than the cost of a successful treatment. Thus, if a CT scan is not performed, a 6% chance that local disease will not be controlled will on the average lead to an additional cost of 6% x $24,000 = $1440. Offsetting this is the fact that there is a 2.5% chance (6%-3.5%) that, with CT, local control will be achieved but the patient will nevertheless die of disseminated disease. The additional cost of therapy in this case is probably less than the $24,000 figure above and, lacking data, is estimated here at half that amount, leading to an offsetting cost, with CT, of 2.5% x $12,000 = $300. Overall, then, the $250 CT scan is estimated to avert additional treatment costs of $1440 - $300 = $1140. This is a benefit ratio of a factor of 1140/250 = 4.6. This estimate does not include savings associated with regained productivity or estimates of the value of improved quality of life, both of which would further favor the positive benefit seen for CT. These calculations should be considered as very approximate. However, they do suggest that even if CT scanning leads only to benefits as small as those presented here, it is almost certainly a cost-effective procedure.


7. PATIENT POSITIONING

The use of the anatomic information provided by CT for the design of radiation treatments depends upon the patient being as nearly as possible in the same position for scanning as for treatment (or vice versa). The importance of this will vary from patient to patient. In some situations it is possible to relate the CT scan findings to stable anatomic landmarks such as bony structures and, in devising the treatment, work backwards from the fixed landmarks. This, for example, is the case for the brain, where the stability of the structures and the presence of good landmarks often makes it possible to use information obtained from scans taken under quite different conditions from those of treatment. However, for the majority of situations encountered in body scanning, such a translation significantly degrades the information obtained in the CT scan. It is therefore important that attention be given to making sure that scans are performed under known and appropriate conditions of patient positioning and alignment.

Undoubtedly the single most important factor in achieving appropriate patient positioning is the training and interest of the CT technician and the understanding of the radiologist supervising the scan. Without the active cooperation of the CT staff, it is impossible to realize the full benefits of CT scanning. It would be highly desirable that either the technicians have a background as radiation therapy technicians, or that a radiation therapy technician assist in studies of patients who are known to be likely to receive radiation therapy, or, at least, that the CT technician be exposed in a serious way to the techniques used for patient alignment in radiation therapy treatments.

There are substantial problems of communication and logistics in most hospitals. Often the CT scan is performed early in the course of the patient's work-up and before a decision has been made to institute radiation therapy-often before a radiation therapist has been involved in the patient's management. Thus, even though a malignancy may be suspected, the scans are often not thought of as potentially contributing to the planning of radiation therapy treatments. Under such circumstances, even though attention may customarily be paid to the problems of patient positioning for known candidates for radiation therapy, the proper steps will not be taken. Consequently, the scan may be of diminished value, or none at all, for purposes of planning the radiation therapy field, and it may be necessary to rescan the patient under the appropriate conditions, with the attendant increase in cost and unnecessary use of a valuable and limited resource. Attention needs to be given to this problem, both by those administering the use of CT scanners within the hospital and by regulatory agencies. There is no reason that the possible use of radiation therapy treatments

222 Michael Goitein

cannot be anticipated for the majority of patients being assessed for malignant disease. Hence, the need for accurate and reproducible patient alignment could be anticipated.

Flat table tops are in almost uniform use in radiation therapy. It is not clear whether the use of a flat table top achieves more stable patient positioning, probably not. However, it seems to offer greater reproducibility of patient position from treatment to treatment. Whatever the reason, the fact that the patient will be treated under these conditions requires the use of a flat surface for the support of the patient in the CT scanner. Figure 5, taken from the study of Hobday et al.,(3) shows the changes of anatomy that can occur between curved and flat patient support surfaces. Ideally, not only should the patient surface be flat, but it should have the same width as the treatment table top and use the same covering material. In this respect, some very good results have been obtained with the use of a thin vinyl-covered foam mattress.(19) This mattress is thin enough (6 mm) to ensure the stability of the patient, thick enough to be reasonably comfortable, has a washable nonslip surface, and is radiolucent.

Given that the scanner support is as similar as possible to the treatment table, the patient's position must be appropriate and properly documented. Such details as the position of the hands and feet should be recorded. Distinguishing the left from the right side of the patient is not always trivial. Certainly the single most helpful aid in relating the CT scan to the treatment situation is the use of skin surface markers in the CT scan, which can be used to identify skin marks in the scan situation. Such markers can be lengths of radio-opaque plastic catheter material (sufficiently narrow gauge, about # 5, that they are really "radio-gray" rather than radio-opaque) taped onto the skin surface or a thin bead of barium paste. Of course, if the patient has not had a therapy simulation with the addition of skin marks, there may be no skin marks to identify at the time of the scan. However, it is readily possible to invent some marks that will be useful in subsequent patient positioning, mark them on the skin surface, instruct the patient not to remove them, add the selected CT -visible material, and document what was done adequately. It is such details that require the thoughtful support of the CT staff.

CT -compatible immobilization casts can be helpful, particularly for treatments of the head and neck area and of the extremities. A perforated thermoplastict has been used that can be softened by immersion in 65°C water, then stretched over and formed to the patient's skin surface. It hardens in a few minutes. This material can be secured to a plywood board (covered with the previously mentioned mattress material), which can both be used in the CT scanner and on the therapy machine. Several such boards

t Available from Alimed, Inc., 68 Harrison Ave., Boston, Ma. 02111.


Figure 5. Effect of couch shape on body contour and internal anatomy. The top scan is of a patient on a dished platter and the bottom scan is of the same patient on a flat platter. From Hobday et al. in ref. 3.

could be available, some kept permanently at the CT scanner and the others at the appropriate treatment units. The advantage of perforated material is that the patient's skin surface can be seen through it so that correct registration of the patient and mask can be verified, skin marks can be seen, some skin sparing is possible, and the patient can breath and perspire freely (both during cast fabrication and subsequently).

224 Michael Goitein

Despite the best intentions, it is not always possible to achieve the desired patient position on the CT scanner. This happens, for example, when the normal treatment policy is to have the arms outstretched in some manner (akimbo, clasped over the head, or out to the side as is done in some chest wall irradiation). While this may sometimes be unavoidable, necessitating the careful analysis of the relative motion likely between the scanned and treatment situations, it is worth considering a modification in the treatment position. Owing to the generally modest diameter of the scanner tunnel, the CT scanner imposes more severe restraints than does the therapy machine. It may be preferable to adopt a treatment position that is compatible with the CT scanner than to accept different positions in treatment and scanning situations.

There are a number of scanning conditions that are discretionary. The questions discussed above fall into this category and little or no compromise of the utility of the scan for diagnostic purposes will result from paying attention to such alignment needs and related matters. However, there are factors that substantially affect the quality of the CT scan or its value for purposes of interpretation and these will now be discussed.

7.1. Breathing

One would never begrudge patients the benefits of respiration, but would that they would breathe either much more rapidly, or, preferably, very much more slowly, and in any event, with much less excursion of lung volume than they are wont to do. The question of what to do about breathing depends to a large extent on the speed of the CT scan. Machines whose scan times are in excess of a reasonable breath hold offer no practical alternative than to have the patient breath as shallowly as possible during the course of the scan. Machines whose scan times are on the order of 15-20 sec have only two realistic choices. Either the patient will be asked to maintain deep inspiration throughout the course of a scan or the patient can maintain shallow respiration. Fast scans, on the order of 10 sec or less, can be performed in any phase of suspended respiration but will be prone to artifact if scans are taken with continuous breathing. In the event that suspended respiration is not possible, a condition under which very many respiratory cycles are encompassed in the course of one scan is probably preferable to one in which only a few take place. Battista et al. (20) have suggested that the option for deliberate slow scanning should be readily available on CT systems.

The extent to which breathing represents a serious problem depends, of course, on the anatomic site. Clearly the thorax is likely to be significantly affected by respiration, as is the abdomen, where diaphragmatic motion can cause significant distortion of the organs. It is probably a lesser problem


in the pelvis, and relatively insignificant in the head and neck and extremities.

If one is able, as with a fast scanner, to scan under any desired condition of respiration, what would be the best respiratory condition to employ in scanning the patient? If it is only possible to take scans under one condition, one would pick the condition that would result in the best possible quality of scan, namely, a state of suspended respiration. With a fast scanner this should probably be with lung filling of the order of what is obtained in shallow respiration. With a 15-20 sec scanner one is obliged to work with deep inspiration. Usually, if attention is being paid to these matters, it should be possible to repeat at least a few critical sections in another state of respiration to assess the extent to which respiration causes significant change in the anatomic relationships. In such cases, in what phase of respiration should the second scan be taken? It is common to suggest, as Hobday et al. do, (3) that the second scan should be repeated with the patient breathing normally. I find myself in some disagreement with this position, which is primarily based on some notion of using "average" absorption coefficients for purposes of dose evaluation. I think it more valuable to perform the second scan also in suspended respiration, in a phase of the respiratory cycle taken so that the pair of scans bracket the expected treatment situation. This is because one is really less interested in the average position of organs than in their extreme excursions. That is, if one is to include a tumor mass in the treatment field, one wishes to have its margin within the field during all phases of respiration and one is much less interested in knowing that some average position is encompassed. Thus, if organ and tissue localization are of primary concern, one is better served by scanning once in inspiration and once in expiration, thereby defining the limits of possible motion.

Of course, very often, and particularly with fast scanners, these issues are academic and it is only occasionally that a significant effort is necessary to monitor the degree of motion during respiration. However, it is important to recognize that respiration can give rise to problems and to be prepared to make an analysis when necessary.

7.2. Contrast Media

Another ingredient of the CT recipe that raises itself in the dichotomy between diagnostic and treatment planning scans is the question of the use of contrast media. Oral contrast for intestinal delineation, intravenous contrast for kidney and bladder delineation, gas contrast for vaginal and bladder delineation, and various surgical clips for correlation with surgical findings are all routinely used. It is sometimes proposed that "treatment planning" scans be taken in the absence, where possible, of such contrast

226 Michael Goitein

agents. This notion, of course, arises from the desire to use the CT scan for dose calculation purposes. However, the delineation of tissues is, as repeatedly emphasized, extremely important for the design of the treatment planning field and should be de rigueur for "treatment planning" scans. It is relatively easy to take care of contrast media in the dose calculation process. Emphasis should be placed on making the treatment planning program capable of this rather than denying these enormously useful techniques in obtaining scans for treatment planning purposes.

It would be of great value if radio-gray rather than radio-opaque surgical clips were used. These would be made of a material sufficiently opaque to show with medium contrast on CT scans but not so dense that they would generate major artifacts in the event that they moved during scanning.

7.3. Bolus

Bolus distorts body shape, but it may be needed to avoid distortion of CT values near the body surface. On balance, it is better to accept the numerical error and to avoid anatomical distortion by eliminating bolus. This is partly because the diagnostic quality of scans is often very good, even when bolus is omitted, an observation supported by Hobday et al. (3)

As the dynamic range of scanner detectors increases in the more modern machines, use of bolus is likely to become redundant.

7.4. Unavoidable Discrepancies

There are circumstances under which one cannot avoid significant changes between diagnostic and treatment planning scans. This is illustrated by Hobday et al. in the case of a patient with carcinoma of the bladder. The two scans are shown in Figure 6, one (top) with a full bladder and the other (bottom) with an empty bladder. The former was necessary to delineate the tumor. However, it was elected to treat the patient with an empty bladder in order to minimize the volume of normal tissues included in the treatment. (3) Under such circumstances two sets of scans should be made.

7.5. From Scan to Treatment

Achieving reproducibility of position between the CT scan and treatment situation is not a new problem. It is one that has already been solved, at least at a certain generally acceptable level, in the process of achieving reproducibility of position between one treatment and the next and between a treatment simulation and the treatment itself. Quite careful and accurate


/

_. , '.- :k .• ~ b, .""',: ' ~ c.. __ ••• •

' .. , t :.

- f/' . .

-~\II'."'" .. -

Figure 6. Scans of a patient with carcinoma of the bladder. In the top scan, the bladder is full and the tumor (T) is clearly visible. In the bottom scan, same patient with an empty bladder. The tumor is no longer clearly delineated. B = bladder, R = rectum, b = bowel. From Hobday et al. in ref. 3.

techniques have been developed for these purposes and they often work very well. The reproducibility of the CT scan situation only requires the same attention to detail and perhaps needs to rest on a similar technology for its successful resolution.

7.6. Use of CT to Select Position for Treatment

To date, the long times necessary for CT reconstruction have made it virtually impossible to perform more than a limited number of scans. It has therefore not been possible to perform complete sequences of scans with the patient in a variety of positions to investigate whether any treatment advantage would be obtained by judicious choice of patient position. However, with the more recent scanners and, in particular, with the development of array processors and other special-purpose hardware that

228 Michael Goitein

enable very rapid scan reconstruction times, the situation has altered and it should now be possible to perform such analyses. Significant variations in anatomy take place when patient position changes. Sontag and Cunningham(21) have presented an example, which is reproduced in Figure 7, in which a patient was scanned in both the prone and supine position. Obviously, the mediastinal structures were significantly altered in this maneuver and one could imagine that the extent of lung sparing could be substantially affected. Gunderson, using conventional radiographic techniques, has shown the importance of patient position in avoiding irradiation of the small bowel(22) and that the influence of the degree of filling of the bladder on small bowel position is also substantial. As better technical capabilities in CT scanning become available, CT will greatly assist in the selection of the optimal treatment position, either for individual patients posing special problems or for general categories of disease.

8. DOSIMETRY

A major contribution to radiation therapy has been made by physicists in the development of techniques for the accurate calculation of dose distributions within the irradiated patient. The determination of both the absolute dose at some reference point and the relative dose distribution involve significant experimental and theoretical problems. If the level of accuracy were five times more relaxed than the 5% figure often mentioned as the goal of the dose calculation process, there would be less need for the assistance of the physicists. If the calculational accuracy were five times more stringent, the cost and difficulty of therapy would be substantially increased.

The ability of CT to measure the distribution of absorption coefficient throughout the body section opens up the possibility of improving the accuracy of the dose calculation. This is the case both for conventional external beam x- and gamma-ray therapy and for charged-particle therapy, both with electrons and with the more esoteric, heavier charged particles.

McCullough (23) has recently reviewed the application of CT to radiation therapy and the reader is referred to his excellent discussion for a more comprehensive treatment.

8.1. x- and Gamma-Ray External Beam Therapy

8.1.1. HOW MUCH OF A PROBLEM IS THERE?

Advocates of the use of CT for purposes of calculating dose exhibit a predilection for examples that feature treatments in the thorax. This is

AppHaatioDS of cr in Radiotherapy Treatment Planning

Medoot';num Elm Prone only ~ SYo,ne only

• Prone ond sup'ne

229

Figure 7. Two cr images of the same patient slice in which the patient was lying in a supine (top) and prone (middle) position. The superimposition of the two scans is shown in the bottom panel. The most dramatic change is seen in the shape and position of the mediastinum. The separation between points P and P' is an indication of the movement of tissue that has taken place. From Sontag and Cunningham in ref. 21.

230 Michael Goitein

no accIdent. For supervoltage energies (the diminishing importance of orthovoltage is perhaps fortunate for those who would calculate dose distributions, although this might well be an area in which CT would have a substantial impact) the most important inhomogeneity is that posed by air or air-containing cavities. Probably the greatest "inhomogeneity" is the external patient contour. It can, however, be well measured by much less costly techniques. However, the lungs and, to a lesser extent, other air cavities do have a major impact on the dose distribution. Jelden et al.,(24) Fullerton et al., (25) Battista et al., (26) and Sontag et al. (27) have presented examples of CT -guided treatment planning of patients with lesions in the thorax. Such treatments fall into three broad categories: (1) treatments of the mediastinum, such as are involved in mantle irradiation of lymphomas, (2) treatments of masses outside the lung, where treatment fields are likely to traverse lung tissue-esophageal and chest wall irradiations are examples, and (3) treatments of lung nodules. In the first situation the lung is largely shielded and the main calculational difficulty is due to the problem of estimating the scatter component of dose in an irregular field. Here, CT probably adds little to the dose estimation problem. In the second situation quite significant dose perturbations are experienced and the accurate mapping of the lung margins will significantly improve the calculational accuracy. However, one is left with a suspicion that the limiting problem in esophageal treatments is not so much dosimetry as it is a combination of biologic intransigence and imperfect tumor localization, and that the most significant problems in chest wall irradiation arise from geometric factors such as field abutment problems. Finally, in the irradiation of lung nodules there are significant questions about the scatter component of dose, which will benefit from CT data but for which the calculational techniques are as of yet inadequate.

Another site in which inhomogeneities may be relatively important is in the area of the paranasal sinuses, where quite complex treatment plans are often required and where the air-filled cavities are both difficult to assess without CT and pose a substantial dose perturbation.

Elsewhere in the body the main internal inhomogeneity is likely to be bone. Bone, however, is a significantly lesser problem. For example, a 3-cm-thick bone irradiated by a cobalt-60 beam is estimated to cause an 11 % dose deficit in its shadow.(28) When one considers that the treatment plan is likely to comprise at least two, and probably more, fields, it is clear that such a dose perturbation is relatively modest.

Given that there are inhomogeneities, the dose distribution within the target volume is likely to be inhomogeneous. Of course, some manipulation of beam angulation and weighting and the use of standard beam-modifying wedges can in some cases reduce this. This is particularly the case when the inhomogeneities are relatively large and smoothly varying. However,


in many cases the added information provided by CT will largely contribute to a documentation of a dose inhomogeneity that was already suspected. The use of CT makes little sense unless it is accompanied by a capability to compensate for such inhomogeneities by the fabrication of individually tailored compensating filters in those instances where major perturbations of dose exist.

As of yet there is no clear answer to the question posed in the title of this section. An answer will require a critical assessment of the influence of inhomogeneities on the dose distributions across a broad range of treatment sites and of the importance of proper compensation for dose inhomogeneities. Such a program would have to address the question of whether the calculational techniques are adequate. It is likely that there are a number of situations in which the detailed mapping of inhomogeneities by CT will provide a significant advantage, and it will be important to identify these situations.

8.1.2. HOW ACCURATE NEED THE DATA BE?

The necessary accuracy in the measurement of inhomogeneities directly dictates some specifications of the CT scanner, a matter addressed below. One must be concerned with both the spatial resolution and the absorption coefficient resolution of the scanner. There is a well-known correlation between these and the measurement accuracy, or dose.(29) This relationship, while important for understanding the operation of CT, has also given rise to a great deal of misunderstanding, particularly as regards the spatial resolution of which a CT scanner should be capable. In the present review, we will independently establish the spatial and absorption coefficient resolutions necessary for guiding dose calculations. These, then, will dictate the necessary scan dose.

The requirements on CT resolution have been discussed by Geise and McCullough (30) in some detail and their discussion should be consulted in this matter.

8.1.2.1. What Spatial Resolution is Required? (The spatial resolution is related to the pixel size used in the data reconstruction unless the pixels are smaller than the inherent resolution.) Consider, first, the uncertainty in position of the external contour. The line integral of absorption between the skin and a given point within the patient will have an uncertainty on the order of the spatial resolution times the average tissue absorption coefficient. In supervoltage therapy, the absorption of radiation is characteristically 5% per centimeter of tissue, or less, so that a 4-mm uncertainty in the position of the external contour would be at most responsible for a 2% dose uncertainty. Geise and McCullough(30) show that a 2% dose uncertainty is associated with a spatial resolution of 5 mm in the assessment

232 Michael Goitein

of a lO-cm thickness of lung and a 4.5-mm spatial resolution in the assessment of a 5-cm thickness of bone (electron density 1.8 times that of water). Since such thicknesses of bone would likely include some spongy bone of lesser density, this is probably an overestimate of the resolution necessary in the assessment of bony inhomogeneities. The spatial resolution of virtually all commercial CT scanners is significantly better than these requirements.

8.1.2.2. Random Error in Absorption Coefficient Measurement. What about absorption coefficient resolution? We will return to the question of the need to translate between the CT number and the estimate of the absorption coefficient appropriate to the therapy beam energy and, for the moment, assume that this translation can be made exactly. In this case, it is the uncertainty in the CT number that now concerns us. One should distinguish between the statistical fluctuations of CT number and systematic uncertainties. The former are usually dominated by statistical fluctuations in the number of transmitted photons detected. One wants to know what will be the uncertainty in the integrated path length between the skin surface and some point of interest as a result of such fluctuations in CT number. CT noise is highly correlated, so that, over short distances, it would be incorrect to add the uncertainties in CT number of each pixel in quadrature. The exact manner in which CT noise is added depends upon the shape of the region over which one is integrating. This question has been discussed by Chesler et ai.(31) Where a small, more or less circular, area is involved, the fractional uncertainty in the integrated (or average) CT number for a fixed dose will be N-3 / 4 times the fractional uncertainty in the CT number of a single pixel, where N is the number of pixels in the averaged area. If noise were uncorrelated and normally distributed, the uncertainties would add in quadrature and the fractional uncertainty in the average value would be given by N-1/ 2 times the fractional uncertainty in a single pixel. Thus, addition in quadrature overestimates the uncertainty in such an average. In contrast to an average over a circular region, an average over an elongated region reduces the effect of correlation between pixels, so that, in the limit in which integration is performed along a line that passes through the entire reconstructed region, the erors do, in fact, add in quadrature.(31l The situation with which we are concerned, namely, the integration between the skin and some point at depth, more nearly approximates the latter than the former situation. Therefore, the addition in quadrature of individual pixel uncertainties is both reasonably accurate and overestimates the fractional uncertainty in the line integral.

How accurately need the line integrals be known? An approximate answer to this question can be obtained by considering the radiation to be exponentially attenuated so that the intensity at depth, I, is related to the un attenuated intensity, 10 , by the relationship 1= Ioe -P, where p is the


line integral of J IL dx, IL is the absorption coefficient, and dx a unit of path length through it. By simple differentiation, one obtains the relationship

dp = _! dI p pI

If we consider a situation in which the point of interest is at a depth such that the dose is approximately 50% of the un attenuated dose (dI/ I = -0.5), the corresponding value of l/p is 1.44. The path length uncertainty would be l/p times the desired uncertainty in dose, so that a 2% dose uncertainty would result from a 2.9% uncertainty in path length. A more accurate calculation can be performed using tissue-air ratios and leads to the conclusion that a slightly greater uncertainty in path length can usually be accepted. For example, in a situation in which 10 cm x 10-cm cobalt-60 fields are used to irradiate a 20-cm-thick patient section, the change in dose at midplane is 3.6% when the electron density is varied by 7%, suggesting a multiplying factor of 1.9 rather than the figure of 1.44 quoted above.

Roughly speaking, then, a 2 % dose uncertainty would result from a 3.5% path length uncertainty and, if the path traverses 100 pixels and errors are added in quadrature as discussed above, the random uncertainty in each pixel could be as high as 35%. This, of course, is much larger an uncertainty than would be obtained in any CT scan of interest. The conclusion, then, is that statistical uncertainties in the CT number are not an important practical concern for the calculation of dose.

8.1.2.3. Systematic Error in Absorption Coefficient Measurement. In contrast to random errors, systematic uncertainties are a major problem. The above calculation can also be interpreted to mean that a systematic shift of 3.5% in the CT number will lead to a 2% error in the calculated dose. Such a systematic uncertainty is by no means improbable with conventional CT scanners. It can arise from inadequate dynamic ranges of the detectors, the inadequate size of the scan circle so that portions of the anatomy or supporting couch are omitted from the scan, beam-hardening artifacts, the sensitivity of the CT number to the eccentricity of the scanned object, and drifts in x-ray quality.

Major systematic uncertainties can arise from beam-hardening effects. The cupping artifact within the skull is a well-known example of this. Significant effects can occur both within and in the shadow of bone. (32) The most secure solution is the use of an iterative beam-hardening correction that would significantly improve the dose estimate when beam hardening occurred.(33,34) However, this is not likely to be necessary for x-ray dose calculations, since the conventional beam-hardening corrections used in CT scanners probably compensate adequately. This is a problem that needs further attention.

234 Michael Goitein

Contrast material within the patient can generate regions of apparent inhomogeneity. It would not be appropriate to always insist on scans used for dosimetry being performed without contrast medium. Rather, provision should be made in the treatment planning program to "reset" such a region to a presumed CT value (which might be that of some neighboring tissue or a value given by the user).

Another source of systematic error is the change in lung density that occurs with breathing. Of course, the total mass of lung tissue does not change, so that, when the integral of material from the exterior through the entire lung is involved, the reduction in density upon inspiration is likely to be at least partially compensated for by an increase in path length through the lung. Nevertheless, significant variations of the CT number, of as much as 150 Houndsfield units between quiet breathing and deep inspiration, can be experienced, (35) and, therefore, where dosimetry is strongly dependent on the path length through lung, such effects should be considered.

To assess and, perhaps, correct for drift in the CT number after calibration, it is wise to include routinely with each scan some materials of known CT value. Air, 50% ethanol, water, Lucite, nylon, and iodinated water solution are reasonable choices. These can be incorporated into the table top and the CT numbers within them can be checked for individual scans.

Nonrandom error in the CT number can result from reconstruction artifacts. These are usually caused by motion during the scan of a highcontrast object, i.e., a metal clip, air in the stomach or intestine affected by respiration or peristalsis, or the external patient contour (since the tissue-air or tissue-bolus interface is a high-contrast boundary). Such artifacts are both undesirable and inevitable in practice. They can give rise to significant errors if path lengths are calculated roughly parallel to the artifactual streaks. Path lengths calculated perpendicular to such streaks will decrease the significance of error, since they tend to have a "ringing" effect that will tend to preserve the value of line integrals along paths normal to the streaks. The minimization of such artifacts is an important goal in the design (and choice) of scanners. It should also be a standard feature of dose calculation software that artifacts can be identified and the region in which they occur assigned some reasonable average CT number to replace the erroneous values.

It has been customary to perform dose calculations as though the body were water equivalent. If CT is used to make a realistic estimate of the tissues in the beam, there may be a systematic shift in dosimetry. Typically, many tissues have linear absorption coefficients many percent higher than water, while fat tends to have CT values some 10% lower than water, so that, depending on the anatomic site, dose estimates may tend in a syste-


matic manner to be a few percent different from "conventional" dosimetry-even in the absence of significant air and bony inhomogeneities. Parker et at. have suggested that a 7% shift in the CT number (which is their estimate of the uncertainty in the CT number) would, for cobalt-60 and 8-MeV beams, change dose by less than 2% at depths on the order of 10 cm. The calculation discussed above suggests that such a variation in the CT number would alter the dose by about 3.5%. A systematic effect of this magnitude, while not a catastrophic change in dosimetry, should not be accepted without a conscious assessment of whether a small compensating change in the stated dosage aim would be appropriate.

In summary, CT scanners do exhibit quite substantial systematic errors in CT values and, in addition to interfering with the identification of tissues, these errors are a substantial limitation in the use of CT for calculating dose.

8.1.3. FROM CT NUMBER TO ELECTRON DENSITY

In estimating dose, one needs to know the absorption coefficient of tissues appropriate to the energy of the therapeutic radiation. Strictly speaking, one needs to know the spectrum of the radiation and the absorption coefficient at all energies present in the radiation. However, for all practical purposes, an effective energy and single absorption coefficient are adequate. For supervoltage energies, where Compton scattering is the dominant mode of interaction of the primary radiation, the absorption coefficient is determined by the electron density of the material, and in subsequent discussion we will use the electron density as the quantity of interest.

There are two ways in which a translation from CT number to electron density can be achieved: The first is the use of a look-up table that presumes a one-to-one correspondence between CT number and electron density; the second is the use of dual energy scanning.

A simple look-up table has been proposed by a number of investigators. Generally, the electron density relative to that of water is plotted against the CT number. An example of such a relationship is presented in Figure 8. The data are generally close to a straight line between the origin and electron densities near that of water, and there is then a break in slope for higher densities corresponding to regions containing a greater or lesser proportion of compact bone. Such a correspondence is strongly dependent on the conditions under which the CT scan was obtained (filtration, kVp ,

and other scanner characteristics) and the therapeutic-regime (if either very low or v,ery high energies are involved). Therefore it is essential that such relationships be determined for individual scanners and for all operating conditions normally employed. It would be unwise to accept any curve presented in the literature, even for a scanner of similar design.

236 Michael Goltein

Correlation of CT numbers to Electron Density in vivo

500

400 Fem"l 300

200 ,

100 I Breast prosthesis {IJI .1' } M d' f e lOS Inurn

0 I

-100 oi..! } Fat CT Seon I... 7 'b I

~ -200 I I

~ I f..-. -300

1\ '-> " Equotion (I)

-400 I 1

1 -500

1 I

I Compton Scon -600 1

I

-700 Lung

-800 1

1 - 900 I

I I

-1000 0 0.5 1.0 1.5 20 2.5

Electron density, relotive to woter

Figure 8. Relation of tissue CT numbers to electron density (electrons per cm3) in vivo. Patients were imaged by both CT and Compton scanning and corresponding tissue regions in the same patient were compared. For such regions, the average CT number is plotted against the average electron density, obtained by Compton scanning. The uncertainty bars shown correspond to standard deviations of individual tissue elements sampled and reflect the heterogeneity of tissue in vivo. The CT number scale used ranges from -1000 (air) to + 1000 (twice the attenuation of water). The CT scanner was operated at 135 kVp, with 3 mm of added aluminum filtration (half-value thickness of 7.2 mm aluminum). From Battista et al. in ref. 20.


The validity of a look-up table is based upon the applicability of a two-compartment hypothesis in which material is presumed either to be tissue equivalent (but of varying physical density) or to contain varying amounts of bone. The measurement of a CT number higher than that of water is then presumed to indicate an admixture of a proportion of bone (Le., calcium), the amount of which can be estimated from the CT number. Geise and McCullough,(30) Kijewski and Bjarngard,(36) Parker et ai.,(37)

Prasad et ai.(38) and Battista et ai.(20) have all discussed this question and presented relationships between the CT number and electron density appropriate to their scanners.

A theoretically more satisfying approach would be to make measurements of the absorption coefficient at two effective energies and deduce the contribution to the CT scans from the photoelectric process, thereby enabling a direct estimate of the electron density. Such a technique is of interest for tissue composition analysis and will be discussed below. Ki jewski and Bjarngard compared dual energy scanning to the straightforward look-up approach discussed above and concluded that the dual energy approach was not significantly more accurate for treatment planning purposes. (36)

Care should be taken if unusual materials are present within the scanned section. Figure 8 includes a datum point for a breast prosthesis that differed markedly in its electron density from that which would be presumed on the basis of the simple CT number versus electron density relationship appropriate for most tissues. The data in Figure 8 are also of interest in that the electron densities were established using a Comptonscatter scanner.

8.1.4. CALCULATIONAL TECHNIQUES

8.1.4.1. Introduction. Dose calculations that take inhomogeneities into account fall into two broad categories: (a) those that use pixel-by-pixel CT data and (b) those in which an average electron density is assigned to each of several large volumes of tissue.

The argument as to whether CT numbers should be transferred in toto from scanner to treatment planning computer hinges both on the question of whether they are necessary for the dose calculation and upon how the CT scan is to be used for tumor and normal tissue delineation. If the primary function supported by the treatment planning computer is that of dose calculation, it is unnecessary and, indeed, perhaps inadvisable to transfer the numerical arrays. Rather, the CT scan may be entered with the conventional digitizing hardware, with large regions being assigned a single effective electron density. One advantage of this method is that it is possible to synthesize information from other studies with the CT data

238 Michael Goitein

more easily if this is done on a sheet of paper or on a film of a CT scan. If, on the other hand, the CT data is to be used for tissue localization and for the design of the probably irregular field shape, then the transfer of CT numbers is necessary. However, there do not presently exist treatment planning programs that use CT information effectively for this application. A computer designed for this application must have both the hardware necessary for transfer of the data (probably a magnetic tape, possibly, but less satisfactorily, a floppy disk) and a graphics display device for reviewing CT sections in gray scale. These capabilities are expensive.

8.1.4.2. "Conventional Techniques". Several techniques for calculating the influence of inhomogeneities have been succinctly reviewed by Geise and McCullough.(30) The calculation of dose for external photon beams is best considered as the summation of a primary and secondary, or scatter, contribution. The primary dose at a given point is due to photons emanating from the therapy machine and directed towards that point, and is therefore governed by the material lying along a line between the source and the point of interest. The secondary component of dose is mainly due to interactions of photons that were originally directed away from the point of interest. Therefore, the integral of electron density between the source and point of interest is not the correct quantity to use in estimating the scatter contribution which, of course, can be a substantial proportion of the dose at depth within the patient. Nevertheless, many of the existing calculational techniques assume that the scatter contribution can be modified by factors depending on the path length of material between the source and point of interest. This is true of the effective source-to-surface distance (SSD) method and the Batho tissue-air ratio technique. (39) Thus these techniques will be reasonably accurate in situations in which the bulk of the inhomogeneity is a slab roughly perpendicular to the direction of beam incidence and will tend to have problems when a significant perturbation of scattered radiation is due to a distribution of inhomogeneity lateral to the beam direction. This latter situation, for example, would be the case if the presence of lung were not considered in estimating the mediastinal dose from an anterior-posterior field.

Similar techniques of calculation can be applied both to data in which inhomogeneities are characterized as large internally uniform regions and where fine-grained absorption coefficient information is provided on a pixel-by-pixel basis. In the former case the path length between a point of interest and the source is estimated on the basis of the distance that this line passes through each inhomogeneity multiplied by the effective density of that inhomogeneity; in the latter case the path length is calculated by a pixel-by-pixel integration.

The implementation of treatment planning in which large areas are assigned a uniform electron density could be facilitated by some minor


software development in the CT scanner. McCullough has suggested(40) that a display in which given ranges of CT number were assigned uniform and specific shades or colors would be helpful. Such a display would show separately regions corresponding primarily to air, lung, fat, soft tissue other than fat, low-density bone, and high-density bone. It would additionally be highly desirable to calculate in each of these regions the average attenuation coefficient of all included pixels. It would then be possible to transfer this average number, rather than the modal value, to the planning code. This, for example, would allow the correct average lung value to be used rather than some assumed value that might not be appropriate to the particular patient.

In transferring data through the medium of film, some attention must be given to the choice of window and level for the display, since significant errors in the location of boundaries can be made if this choice is improperly made.

8.1.4.3. Pixel-by-Pixel Calculations. A technique has been developed by Sontag and Cunningham(41) that is specifically intended to take advantage of CT data available on a pixel-by-pixel basis. This technique features the calculation of a correction factor to be applied to the dose calculated on the assumption of a uniform material. The correction factor is evaluated in terms of a primary component, which is computed on the basis of the radiologic path length through the material, as measured by integrating the CT numbers along a line between the source and point of interest, and a scatter component, which is evaluated in terms of that radiological path length and an effective field size. The effective field size is calculated on the basis of a summation over the CT numbers of tissues outside the direct path between the point of interest and source, and weighted by factors that, in essence, take into account the angular distribution of the scattering process. This calculation does not take into account the measured path length between the source and secondary scatter point or between the secondary scatter point and the point of calculation. The calculation has the unusual feature that the secondary scatter contribution in a homogeneous situation is not cylindrically symmetrical around the line between the source and point of interest. Experiments with this technique suggest that it does handle inhomogeneities quite well and does so in a reasonable amount of time. (42)

Without question, the most satisfying approach from a theoretical point of view is to integrate the scatter contribution of dose over the entire irradiated volume. Such an integral is most efficiently performed using a Monte Carlo sampling technique. Webb and Parker(43) have done this. They assert that it is "impracticable to contemplate performing Monte Carlo calculations for individual patients" and conclude that the technique is primarily of value in checking more approximate techniques for assessing

240 Michael Goitein

the effects of inhomogeneities. Webb and FOX(44) have concluded that in performing a Monte Carlo integration, a 5-mm pixel size is acceptable for purposes of dose calculation.

8.1.4.4. Comparison of "Homogeneous Inhomogeneities" and Pixelby-Pixel Techniques. It is important to know whether a pixel-by-pixel calculation technique is necessary to achieve an adequate accuracy of dose calculation. Such a technique is inevitably more expensive to implement than the "homogeneous inhomogeneity" approach that is already implemented on many commercially available treatment planning computers at this time. Prasad et al. (38) addressed this question. They compared the results of calculation with thermoluminescent device (TLD) measurements in a Rando phantom in a section that included lung and some minor amount of bone. The techniques that they compared were a pixel-by-pixel calculation of the radiologic depth of each point of calculation and, on the other hand, an effective SSD technique that essentially measured the radiologic depth as the sum of the depths through each of the large regions of inhomogeneity. They concluded that the two techniques were comparable in their accuracy, both showing variations from the measured values on the order of several percent. Parker et al. have presented a pixel-by-pixel technique. (37) They imply that its primary value is the practical one of simplifying and automating the data manipulation and do not suggest an important advantage in calculational accuracy for the technique.

It seems, then, that there is no clear advantage for the pixel-by-pixel technique in terms of calculational accuracy at this time. Any such advantages will probably come from improved methods of calculation such as the Monte Carlo,(43) the effective TAR method,(41) or some similar technique that accurately takes into account the complicated problem of the effect of inhomogeneities on the scatter component of dose.

8.1.5. CONTIGUOUS SCANS

It is, of course, quite unreasonable to think that one can ignore the inhomogeneities that occur in regions outside the plane of a given scan and only include the effects from tissues seen in the plane of calculation. This, however, is the assumption made in many programs that purport to take inhomogeneities into account. Of course, the influence of inhomogeneities on the primary component of radiation is correctly assessed by the data in the plane of calculation. However, the perturbation of the scatter component of dose will be affected just as much by inhomogeneities out of the plane of the scan as those in it. Battista et al. (26)

have convincingly shown this to be the case in their analysis of the dose distribution in the treatment of a carcinoma of the upper esophagus.


8.1.6. "THREE-DIMENSIONAL" TREATMENT PLANNING

The calculational techniques mentioned above are primarily twodimensional techniques that have easily accommodated CT scans precisely because they provide two-dimensional measurements. However, the patient, while not isotropic, has no strongly preferred axis from the point of view of dose calculation or treatment planning. Why, then, should calculations be restricted to beams that are coplanar and lie in the plane normal to the scan axis (which is generally parallel to the long axis of the patient)? In the past, treatments have tended to be constrained to such beams. To a great extent the use of more complex angulations of beam has been limited by a lack of security in the information necessary for their setup. The availability of high-quality three-dimensional data should now permit greater freedom, which, in selected cases, will result in improved treatments.

Perhaps the most important possibility that CT scans offer is that of viewing structures and volumes of interest from the beam's eye view. From such a viewpoint the beam angulation can be selected (using all degrees of freedom, i.e., not restricting the treatment to being coplanar) in such a manner as to exclude, to the extent possible, sensitive structures from the beam while including the target volume. The design of an irregular fieldshaping device- or the selection of a collimator setting and angulation-can then readily be made. This should be the first order of business for a treatment planning program, since the field shape determines which structures are "in" and which structures are "out" of the beam. This is a more important question than the calculation of the actual dose to each structure, and the capability of choosing the field aperture should have a higher priority than the capability for dose calculation. That this has not been the case in the past has largely been due to the lack of confidence in the spatial location of interesting structures and poor registration between serial sections where more than one cross section of the patient was, indeed, considered.

Treatment planning programs that provide a fully three-dimensional capability based on CT data are only now becoming available. McShan et al. have presented information on one such program that has been developed by their group. (45) This is an exciting development and further progress along these lines may be expected from several groups in the near future. The implementation of fully three-dimensional programs with their requirement for fairly sophisticated graphics capabilities is quite demanding of computer and peripheral hardware and is, therefore, quite expensive. Just how expensive this will prove to be and whether the additional costs will justify the benefits to be gained from these techniques remains to be seen.

242 Michael Goitein

8.1.7. INTEGRATION OF CT DATA WITH OTHER STUDIES

The availability of the detailed and extensive information offered by CT scans should not lead one to forget that many other studies of major diagnostic importance are often needed in the assessment of a patient's disease and should be used in the treatment planning process. The CT scan forms a very useful frame of reference within which to synthesize this information. Plain radiographs using special techniques, conventional polytomes and laminograms, contrast studies such as lymphangiograms, small bowel series, angiograms and pneumoencephalograms, scintigrams using a variety of radioactive tracers, ultrasonograms, and information from physical examination all need to be brought together. Traditionally, this has been done rather unsystematically and probably inaccurately, with the information usually being integrated on a sheet of paper from which data are entered into the treatment planning computer. In this sense, the presence of data in the computer is a backward step. Such data is often much less easily augmented or modified than are pencil marks on paper, where accessibility does not depend on software effort and where editing requires no hardware beyond an eraser and pencil. There is a major need to develop techniques that permit this synthesis of data to be incorporated in a rational manner into the treatment planning process.

8.2. Brachytherapy

The three-dimensional calculation of dose is, of course, just as feasible and interesting for brachytherapy as it is for external beam treatments. The benefits here are likely to be somewhat different, however. Inhomogeneities do not seem to play so large a role in brachytherapy, because the distances over which the high-dose region extends are generally sufficiently small so that there is little room for significant inhomogeneities. The benefits from CT scans are likely to be in their ability to assess the anatomical features covered in the high-dose volume and to pinpoint the exact location of particularly sensitive tissues so that the dose to them may be evaluated. Overlays of brachytherapy dose distributions onto CT slices demonstrate a remarkable wealth of information, which suggests that this capability will be very valuable.

8.3. Charged-Particle Therapy

8.3.1. HEAVY CHARGED PARTICLES

Inhomogeneities pose a very different and much more severe problem in charged-particle radiation therapy. This is because an inhomogeneity


affects the range, or penetration, of the charged particle as compared to the intensity of the radiation beam in the case of photons and other neutral radiation beams. This is illustrated in Figure 9. In this example a 3-cm thickness of bone exposed to a 60CO beam is shown to depress the dose in its shadow by some 11 %. In a treatment plan involving multiple fields, such a dose deficit would be unlikely to have major clinical consequences. By contrast, the charged-particle beam intensity is virtually unaffected in the immediate shadow of the bone-an advantage of charged-particle beams-but the dose in the distal 2 cm or so is reduced to essentially zero by the shadowing bone. If a tumor extended into the distal region, such a dose deficit could have extremely serious clinical consequences. This general picture is obtained for all charged-particle beams, including pi meson, heavy ion, proton, and electron beams.

It is therefore essential that the spatial distribution of inhomogeneities within the regions traversed by charged-particle beams be accurately known. It is providential that CT scanning became available at about the time that charged-particle radiation therapy began to be actively pursued.

Figure 9. The effect of a 3-cm slab of bone on the depth dose distribution of a 60CO beam compared with a charged-particle beam.

,--------Co 60 __ --,

BONE

,-----1--- CHARGED PARTICLES

--"I -+-'1 ----, 1 1 I : : I 1 1 I I I I

--Ii+- I 1 1 t+-~ 2cm 3cm I

DEPTH-

244 Michael Goitein

Indeed, so critical is the need to assess inhomogeneities that in the late 1960s Tobias had considered the use of charged-particle beams for threedimensional tissue-density reconstruction and, with Lyman, had made measurements on a phantom, using alpha particles, which were analyzed using a least-squares iterative reconstruction technique(29) before commercial x-ray CT was announced.

Heavy ions, protons, and pi mesons are similarly affected by inhomogeneities, although the multiple scattering of these particles sets a somewhat different scale for the effect. Three phenomena may be distinguished: (1) a range modification of the type represented schematically in Figure 9, (2) an edge scattering phenomenon, and (3) dose modification caused by thin slivers of material parallel to the beam direction and within interfaces roughly parallel to the beam direction. These are now discussed in turn.

The range modification is due to differences in linear stopping power among different materials. This is roughly, but not exactly, determined by the electron density of the material and therefore a transition scheme very similar to that needed to translate CT numbers for purposes of supervoltage therapy planning is required in charged-particle treatment planning. The conversion is between a water-equivalent path length-namely, that path length through water which would give the same range shortening as the thickness of test material-and CT number. Chen et al. have presented such a translation scheme.(46) Here, too, there is a question as to whether a simple one-to-one relationship between the CT number and waterequivalent path length is adequate or whether dual energy scanning may be necessary to determine the water-equivalent path length with sufficient accuracy. The latter technique has been described in its application to charged-particle therapy(47) and would certainly be sufficient for that purpose. It is likely, as Chen et al. have suggested, that such a refinement will not be necessary under normal circumstances and a simple one-to-one translation between the CT number and water-equivalent path length will be adequate.

The second effect is an edge scattering phenomenon, which is seen when an inhomogeneity does not cover the full beam area. In the shadow of the edge quite large intensity fluctuations can be observed. This phenomenon, which has long been known in connection with electron beam therapy, has recently been reviewed and a calculational method for assessing the extent of the dose perturbation has been developed.(48) The technique is in good agreement with measurements. (49) In general, dose perturbations of up to 50% can be seen under conditions of good beam parallelism. However, the dose perturbation is substantially reduced when there is significant angular divergence to the beam, a condition that is obtained when even modest amounts of overlying tissue are present. Thus, only very


superficial inhomogeneities turn out to pose a significant problem under most circumstances, and only air-tissue interfaces are then significant. Bone-tissue interfaces do not generally pose a clinical problem. Figure 10 shows a dose perturbation in the shadow of a Lucite inhomogeneity exposed to a near-parallel proton beam. The ±31 % dose perturbation is reduced to a ±12% effect by a thickness of overlying material only one-half that of the thickness of the inhomogeneity, as is seen in Figure 10(b). While the calculational technique presented in ref. 48 is quite accurate for single edges, it is not applicable in the general case of an arbitrarily inhomogeneous region. A full Monte Carlo calculation is necessary in such a situation and the necessary data could only come from CT. However, such a calculation is probably unnecessary, since, under most practical circumstances, such edge perturbations do not pose a significant clinical problem.

The effect of a long sliver on the dose distribution depends on its thickness. (50) An example of this is demonstrated in Figure 11. Multiple scattering of the beam wipes out the influence of the sliver if it is sufficiently thin. This is fortunate, since otherwise such inhomogeneities would have to be mapped with impossibly good spatial resolution. As it is, the spatial resolution of CT scanners is well matched to the requirements of heavy charged-particle radiation therapy. (47)

The calculation of dose in a charged-particle beam, and in particular the estimation of the perturbing effect of inhomogeneities, requires in principle a full-fledged Monte Carlo program. Such a program is time consuming. It is probably not necessary, at least for protons and heavy ions. An alternative technique has been described for calculating heavy charged-particle dose distributions using CT data. (51) This technique features an important concept, namely, the estimation of the calculational uncertainty. In conventional treatment planning, the uncertainty in the dose distribution is never explicitly presented. It is generally assumed that some overall normalizing uncertainty exists that is implicitly kept in mind. The uncertainties due to patient motion are reflected in a lack of realism in the representation of the dose distribution near the field margins. This, too, is implicitly recognized in developing a treatment plan, and fields are usually made larger than dictated by the tumor volume to allow for such effects. However, in charged-particle treatment planning quite significant uncertainties can exist throughout the treatment volume owing to uncertainty in the registration between an inhomogeneity and the beam (which is usually provided with a compensating filter to account for internal inhomogeneities). Calculational uncertainties in assessing the multiple scattering and edge scattering effects also occur. In some parts of the field these uncertainties will be negligible, and in other parts they may be very substantial. It is therefore recommended that treatment plans specifically reflect this, perhaps by presenting a pair of dose distributions, the first

~

~ ~ g ~ h: ~

~ cc

14

0

12

0

10

0

80

60

40

20

12

7 M

e V

Pro

tons

I I

I I

Im ~

1 A

ir

__

__

.. _

1.

Sc

an

Oln

9 ..

,.I

DIO

de

25

em

L

ucl

te

12

7 M

eV P

roto

ns

I I

I I

Lu

clte

1 3

em

i&kk?

4kfry

2 5 e

m

25cm

A

ir

1 ___

____

1..

-(S

can

nin

g D

iod

e

OLI~~ _

_ ~L-

__

L-_

_ J-_

_ -L

__

~ _

_ ~ _

__

_ L-_

_ L-_

_ L-_

_ ~

-1.0

-.

8

-.6

-.4

-.

2

0 .2

.4

.6

.8

1.

0 -1

.0

-.8

-.6

-.

4

-.2

o

.2

.4

.6

DIS

TA

NC

E (

em

) (a

) ( b

)

.8

1.0

Fig

ure

10.

Com

pari

sons

bet

wee

n m

easu

rem

ents

(poi

nts)

and

cal

cula

tion

s (l

ines

) for

127

-MeV

pro

tons

inci

dent

on

a 2

.5-c

m

Luc

ite-

air

inte

rfac

e. D

ose

pro

file

s ar

e m

easu

red

afte

r a fl

ight

pat

h in

air

of

25 c

m. (

a) S

impl

e L

ucit

e-ai

r ed

ge. (

b) L

ucit

e-ai

r ed

ge c

over

ed w

ith

a 1.

13-c

m-t

hick

Luc

ite

scat

tere

r pl

aced

jus

t up

stre

am.

Fro

m G

oite

in e

t al.

in r

ef.

49.

N

01:0. =-- ~ S- IlO ~ ~ t

Applications of CT in Radiotherapy Treatment Planning

20

REGION OF SLIVER n WITHOUT

SLIVER

PROTONS

~ ~ ~ 1

"1:,::: I

CENIRAL Al(IS

o L-______ ~ ______ _L ______ ~~~ ______ ~

o 5 10

OEPTH(cm)

247

Figure 11. A short (2.5 em long) Teflon sliver at a depth of 7.5 em in water, exposed to a realistic proton beam. Central-axis depth dose distributions are shown for sliver widths of 10,4, 2, and 1 mm. From Goitein and Sisterson in ref. 50.

showing the maximum possible dose and the second showing the minimum dose that might be delivered in a given treatment situation.

Treatment planning techniques have been presented for heavy ions(46) and pi mesons. (52) These rely on CT data for the measurement of the distribution of inhomogeneities.

8.3.2. ELECTRONS

Electrons pose the most difficult calculational problem of all radiations used in therapy. This is because, being light, they can scatter through quite large angles. They exhibit all the effects characteristic of charged particles mentioned above, but the spatial dimensions over which these effects occur are quite large and the calculational technique proposed for heavy chargedparticle therapy(51) is totally inappropriate for electrons. A satisfactory technique for calculating the dose distribution of electrons under arbitrary circumstances is not yet available. This is only tolerable because electron beams of practical value have a generally modest penetration and, therefore, relatively little opportunity to encounter significant thicknesses of inhomogeneity. Nevertheless, at the very least, one has to assess the path length of material through which electrons have to traverse to cover the

248 Michael Goitein

treatment volume-and thereby select the electron beam energy and perhaps design some relatively crude compensating bolus. CT scans are likely to be invaluable for this purpose. A correlation between CT number and water-equivalent path length as developed for charged particles(46) will be needed for electron beams. With the anatomic information of CT, it will at least be possible to ensure that the electron penetration is appropriate for the task. This is the most important requirement in electron treatments. However, the detailed distribution in the presence of fine-structured inhomogeneities will be something of a mystery until further calculational advances are made.

9. TISSUE CHARACTERIZATION

To say that tissues are imaged implies that some characteristic, which varies form tissue to tissue, is being measured. A single CT scan will measure the spatial distribution of a single quantity, the linear absorption coefficient, throughout the reconstructed section. If luck were with us, the tissues of interest would all have different absorption coefficients, diseased tissue would have an absorption coefficient different from its normal counterpart, variations from individual to individual would be less than the typical differences between tissue absorption coefficients, and the uncertainties in the measurements would be even smaller. Mankind is not that fortunate. Rather, the differences between many tissues of interest are quite modest and are at or below the limit of resolution of CT. Moreover, variations from individual to individual are substantial relative to these small differences. The lack of absolute accuracy and stability of CT numbers has, in part, contributed to the observed variability of CT numbers, and it may be that with more stable devices a more systematic pattern will emerge. Nevertheless, a single CT scan, not surprisingly, does not fully characterize the tissues it images. Therefore, a number of techniques are used to increase the information about tissues in an effort to identify pathognomonic factors.

The most straightforward of these techniques is the use of contrast media. Much of the early work was concerned with whether or not a given lesion did or did not "take up" observable quantities of iodinated contrast material. With time, this question has been refined, both in terms of quantitating the uptake of material relative to some control area and by looking at the time course of such uptake. As fast scans become available, the evaluation of the transit time for a bolus of contrast material becomes feasible, thus making possible estimates of tissue perfusion. As these techniques become refined, it will be necessary to provide the capability for comparative analysis of CT scans at the diagnostic console.


9.1. Tomochemistry

Since tissue absorption coefficients are energy dependent and their energy dependence is a function of elemental composition, it is possible to get information concerning elemental composition by analyzing CT scans taken at two different effective energies. In principle, if the absorption edges of the x-ray attenuation coefficients could be probed, it would be possible to obtain detailed elemental analyses of tissues. However, at the present time this is not possible. We await the advent of the tunable x-ray laser. In practice, only two quantities of interest can be determined. These may be thought of as the electron density and the effective atomic number of the material or, alternatively, the contribution from the Compton and photoelectric effects to the total absorption coefficient. High-atomicnumber elements, such as calcium in bone or iodine in contrast media, are primarily responsible for significant changes in effective atomic number. It may be that differences in hydrogen content can also be detected.

Two techniques for obtaining data have been proposed. The first is the use of two sequential scans taken with different x-ray tube potentials and, preferably, different filtrations.(35,53) The second is the use of a single scan in a scanner provided with split detectors, the first stage of which has a greater efficiency for detecting the lower-energy transmitted x-rays, and the second stage, being shielded by the first, senses the higher-energy components. (54)

The problem with tomochemistry is that the uncertainty in the atomic number determination is quite considerable. Figure 12 shows an example of an early effort to perform tomochemistry on a patient with a posterior lesion. The noise in the effective atomic number reconstruction is quite evident. To reduce it, it is important to achieve the greatest possible separation between the effective energies of the two scans. There is some controversy as to which measurement technique is superior. The split detector has the advantage that the two measurements are made simultaneously in the course of a single scan. This obviates problems of registration that arise when scans are taken one after the other, as is necessary in scans taken at different tube potentials. One commercial translate-rotate scanner was designed to permit a change of voltage at the end of each translation to minimize the delay between the two measurements and hence to avoid problems of registration. It is likely that the dual kVp technique will provide the greater sensitivity.(33) Moreover, the registration problem, which has been a concern to many investigators, is probably not very serious in many applications. This is precisely because the uncertainty in the determination of effective atomic number is so great. It is therefore likely that an average over a relatively large area of fairly uniform effective atomic number will be necessary to distinguish or identify tissues. In this case,

250 Michael Goitein

Figure 12. Reconstruction of (left) electron density and (right) effective atomic number from scans of a patient with a posterior lesion taken at 140 and 100 kVp. From Goitein (unpublished results).

small registration errors, such as are likely to be experienced in practice, will not cause significant errors in the measurement of effective atomicnumber. This is probably not the case, however, if the effective atomic

number of bone (i.e., bone mineral concentration) is to be determined in this manner, since quite sharp variations in CT number can be seen over relatively small distances.

Latchaw et al. (55) have presented a simplified technique for the analysis of dual energy scans and applied it in a study of 15 patients with brain

32r-----------------------------,

I 2 3 4 S 6 1 8 9 10 II 12 13 14 IS

Glial Tumors Meningiomas Metastases Miscellaneous

Cases

Figure 13. Distribution of the percentage change in tumor atomic number between scans taken with and without a contrast medium. From Latchaw et al. in ref. 55.


tumors of known histology. They used the dual energy technique to assess the change in effective atomic number when a contrast medium was used. They analyzed three quantities: the change in effective atomic number with and without contrast material, the change in absorption coefficient at the lower kVp setting with and without the contrast medium, and the shift in attenuation coefficient between two potentials with and without contrast media. They concluded that the fractional change in the effective atomic number between conditions of contrast and noncontrast media was the most useful parameter. With it, quite significant differences could be seen between groups of patients with glial tumors, meningiomas, and metastatic lesions. These data are shown in Figure 13. One of the first clinical studies was presented by Rutherford et al.,(53) who analyzed this technique extensively.

The potential of tomochemistry has excited many investigators. However, the determination of atomic number is relatively crude and the clinical value of the technique remains unclear.

9.2. Spatial Structure

Another feature that is potentially of value in distinguishing tissues of different histologies is the structure, or "texture," of the image. This feature is very familiar in ultrasonography, where some tissues can be distinguished on the basis of the spatial distribution of the reflected sound. Ritchings et

al.(56) have performed an autocorrelation analysis on CT scans of the liver and spleen in 53 patients in an effort to distinguish normal from cirrhotic livers. The autocorrelation functions did show significant differences, with the normal liver showing structure that was most significant at a distance corresponding to 10 pixels. Using discriminant analysis of the data, they were able to correctly identify 17 out of 18 cirrhotic patients and 10 out of 13 normal patients. If additional data from the spleen were included, the correct classification was obtained in all cases.

This kind of analysis is interesting and merits further attention and extension to other tissues. However, it is clear that the structural differences in CT scan are not as blatant as those seen in ultrasonography and it is likely, therefore, that this kind of information may be less helpful. However, the use of quantitative techniques of analysis could compensate for the greater subtlety of the effect.

9.3. Radiologic-Pathologic Correlation

There is a great need to correlate CT findings with pathologic determinations. Without such correlations the identification of abnormalities in the CT scan must remain indirect and insecure. In several studies mentioned

252 Michael Goitein

in the Appendix, this question has been addressed. These include studies of bladder disease, (8,9) bronchogenic carcinoma, (57) laryngeal carcinoma, (58) and lymphomas. (59) As one might expect, the correlations have shown quite good accuracy for the CT technique, but have suggested significant instances of false positive and negative findings. These studies have essentially addressed yes-or-no questions in which a pathologic condition was or was not diagnosed. One is interested in a more extensive and far more complex question, namely, when a CT scan suggests a volume of abnormality, does it correctly indicate its margins? It is often suggested that edematous or fibrotic changes adjacent to a malignancy may be mistaken for an extension of disease and, alternatively, that CT may sometimes underestimate disease spread. One would very much like to compare the spatial CT information with detailed pathologic mapping of disease. Such a program, while sorely needed, has not yet been reported.

10. SPECIFICATIONS FOR CT SCANNERS TO BE USED IN TREATMENT PLANNING

10.1. Introduction

The use of CT for planning treatment places a number of constraints on the CT scanner. These have been discussed by a subcommittee of the Committee on Radiation Oncology Studies (CROS)(17) and by Goitein.(60)

The CROS Subcommittee on CT Scanning and Radiation Therapy considered CT scanners in two settings. These were (1) CT scanners used primarily for diagnostic purposes, but with "therapy options," and (2) CT scanners dedicated to radiation therapy. The specifications for these two scanners were somewhat different. However, an increased awareness of the requirements of radiation therapy have made the more stringent requirements for the dedicated scanner highly desirable for both applications and the rapid advance in the state of the art of computed tomography has made them entirely feasible. Hence, there will be no distinction between these two classes of scanners in the following discussion.

The special needs of radiation therapy stem from

1. The necessity of fully delineating tumors and adjacent normal structures.

2. The necessity of scanning the patient in a position that is appropriate for, and can be reproduced during, treatment.

3. The desirability of using CT scans to assist in calculation of dose.

The specifications for a CT scanner that is to meet the needs of radiation treatment planning are now discussed in terms of these broad categories.


Of course, some features will be dictated by more than one of these concerns.

10.2. Three-Dimensional Delineation of Tumors and Adjacent Normal Structures

The best possible tissue differentiation is necessary if tumors and adjacent organs are to be well delineated. This, in turn, dictates that a CT scanner to be used in planning radiation therapy must have the best possible spatial and absorption coefficient resolutions. In addition, there is a major gain to be achieved if what I term "volumetric scanning" is feasible. This capability places special demands on the imaging and display capabilities of the system. The features necessary to support the calculation of dose have already been touched on in Section 8 and are further addressed in the present discussion.

10.2.1. SPATIAL RESOLUTION

The spatial resolution should be as good as possible. This is an area in which CT manufacturers are continually improving their equipment. Of course, improvements in spatial resolution are often accompanied by the increasing complexity of the scanner (primarily due to an increase in the number of detectors in single-motion machines) and there are no data relating the cost and benefit of this capability. Nevertheless, spatial resolution is a very important aspect of reconstruction quality that deserves heavy emphasis in the selection of equipment. An improvement in spatial resolution, by at least a factor of 2 or 3 and probably more, is potentially possible in the next decade and this development should be strongly encouraged.

The specification of the spatial resolution is not entirely straightforward. Analysis of the edge response(61) is reliable and reproducible but is not simple to do. In practice, with a sufficiently fine reconstruction and display grid, the analysis of a profile across the image of a thin high-contrast wire (point response) seems to be reasonable and is more easily done. Viewing bar patterns of varying spacings(62) is somewhat subjective and sensitive to the viewing conditions but can provide a reasonable and rapid way of assessing performance, especially relative performance. With such a test object, present CTscanners resolve up to ten line pairs per centimeter.

Of course, for a given patient dose, the absorption coefficient resolution will deteriorate at the highest level of spatial resolution. This does not dampen one's enthusiasm for improved spatial resolution, because many of the tissue interfaces of diagnostic interest are in fact of relatively high contrast. They include tissue-fat interfaces (which have approximately 10% contrast), tissue-air interfaces, and bone-tissue boundaries. The integrity

254 Michael Goitein

of such boundaries is often a major diagnostic clue as to whether and to what degree there may be extension of disease. The importance of tissue-fat boundaries can be inferred from the difficult diagnostic problem posed by patients with little fatty tissue.

The discussion of spatial resolution has so far been in terms of the resolution within the plane of the scan. It is impressive that the generally inferior resolution of CT scans normal to the plane of the scan does not seem to degrade the in-scan resolution as much as one might have expected. Nevertheless, the slice thickness needs to be matched to the in-scan resolution, and slice thicknesses on the order of 2 mm or so would be desirable in many cases. Such thin slices are also needed for volumetric scanning, as discussed below. Thin slices have implications for the x-ray tube characteristics and dose, which are also discussed below.

10.2.2. ABSORPTION COEFFICIENT RESOLUTION

The absorption coefficient resolution of most scanners is limited by statistical fluctuations in the number of detected photons. In order, then, to get the best possible absorption coefficient resolution for a given dose, it is important that the utilization of the x-rays striking the patient be as effective as possible. On the one hand, this implies that good pre-patient collimation is necessary. On the other hand, the detection of transmitted photons should be efficient both from the point of view of conversion efficiency of the photons and of the geometric efficiency of the detectors. Detectors with wide septa between detector elements will be inefficient in this regard.

The assessment of absorption coefficient resolution, or low-contrast resolution, is not easy and the comparison or specification of scanners in this respect is fraught with difficulties. Low-contrast "detectability" phantoms have been devised(63) and are of interest. It is important that the performance of a scanner be statistically limited. This can be determined by plotting the product of the variance in the CT numbers of a uniform field (such as a water bath) and some measure of the dose (either the response of a dose monitoring instrument or the mAs of the scanner) versus the dose. The dose per scan should be varied, if at all possible, by varying mA and not the scan time. This product should be constant over the normal range of operating conditions of the machine. Any deviations imply systematic effects (problems in the scanner operation and reconstruction algorithms) and detector and electronics noise limitations.

The absolute accuracy of the CT number is an important and often neglected feature of a scanner. Many of the early body scanners had quite poor accuracy due both to variations of CT number with scanning conditions and to drift in CT number with time. Absolute accuracy is needed for two

Applications of CT in Radiotherapy Treatment Planning 2SS

reasons: (1) for purposes of dosimetry, as discussed in Section 8.1.2.3, and (2) to assist in the identification of tissues. Many of the tissues of interest in the body differ by only a percent or less in CT number. Thus, CT is only on the verge of being able to distinguish many tissues. Lack of confidence in the absolute value of the CT number is in part responsible for the difficulty in identifying specific tissues and in distinguishing the normal from the abnormal situation.

An absolute accuracy of 2 % in CT number should be achievable and one could hope for even better: Many aspects of scanner design will affect the accuracy of the CT numbers. Some of these are the dynamic range of the detectors, the ability to scan the entire patient cross section, the stability of the x-ray source, the stability of the detectors, the filtration of the x-ray beam, and the nature of the beam-hardening correction. Beam-hardening artifacts are likely to playa significant role in distorting the CT number. (32) An iterative beam-hardening correction(33.34) is probably desirable, and may even be essential for this purpose. Implementation of an iterative beam-hardening correction is entirely feasible on the current generation of scanners. It is likely to increase the reconstruction time by approximately a factor of 3.

An even better tool for overcoming beam-hardening effects, which also has the advantage of facilitating the translation between CT number and electron density needed for dosimetry purposes, is the ability to perform dual energy scanning.(32.53) This capability is not needed for a standard scanner in a clinical setting. However, in a research application this capability would be of interest both for the reasons presented in this section and for purposes of tissue identification, as discussed in Section 9.

The incorporation of reference materials into the couch so they appear in each scan can significantly improve the accuracy of CT data. This can be vital for applications such as the determination of relative bone mineral concentration.

10.2.3. VOLUMETRIC SCANNING

The long reconstruction time of the early generations of CT scanners, together with other technical factors, made it difficult to take more than a modest number of slices, on the order of 16 or less, during a routine study. This necessitated sampling the volume of interest quite coarsely-typically, scans might be taken every 2 cm, with 13-mm-thick slices. While such scans provide a wealth of information, they are not fully adequate in the general case for planning radiation therapy treatments. This judgment has developed during the course of a clinical evaluation of CT(2) in which coarse sampling, such as just described, often proved unsatisfactory. The problem comes from two factors. First, such a coarse sampling makes it difficult to

256 Michael Goltein

define the cephalad and caudad extensions of disease as accurately as one would wish. Second, the coarseness of such scans makes it difficult to use contiguity of tissues for purposes of interpretation. For example, it is not infrequent in the pelvis to have some confusion as to whether a particular region corresponds to a loop of small bowel or a mass. This confusion can persist, even with the use of oral contrast media, either through incomplete filling of the intestine or through imaging an opacified loop of bowel that only partially enters the scan volume (partial volume effect). In such a situation the adjacent scan, 2 cm away, often fails to clarify which of these alternatives is correct.

Figure 14 is an effort to demonstrate the problem of coarse sampling in a schematic manner. Figure 14 (top) is a conventional CT section through the pelvis. In Figure 14 (bottom), a second section is shown whose resolution has been distorted to mimic the situation obtained along the long axis of the body when coarse sampling of the nature described above is employed. That is, alternate regions are blanked from view and, within the imaged volume, the data are uniformly smeared in one direction. It is clear from the comparison of these two images that the lateral definition of, for example, the seminal vesicles, or, indeed, any other structure of interest, is greatly compromised by this strategem. No sensible person would willingly accept such a degradation of resolution. Nevertheless, along the long axis of the patient, it is the norm for the majority of body scanners at this time.

For purposes of planning radiation therapy treatment, the entire volume of interest should ideally be scanned continuously and with quite good resolution along the long axis of the patient. If this is done, two benefits will accrue. First, the determination of the three-dimensional extension of demonstrable disease will be more accurate. Second, the "diagnostic" value of the scan will be greatly enhanced-that is, the identification of tissues and organs of interest and the determination of their normalcy will be improved.

Such "volumetric" scans can be viewed, of course, in any arbitrary plane, including sagittal and coronal views. The added value of such views has been recognized in head scanning, where, by suitable positioning of the patient, a direct coronal section can be obtained. (64) Its value is now being appreciated in body scans. An impressive series of such scans has been presented in the examination of the pancreas, often a difficult organ to evaluate in the direct transverse view.(65) A particularly elegant application of volumetric scanning has been developed by Glenn et ai.(66) in the evaluation of the lumbar spine.

Volumetric scanning involves taking multiple contiguous thin slices throughout the volume of interest. This volume, of course, will include the region of suspected disease and a margin above and below it. How finely

Applications of cr in Radiotherapy Treatment Planning 257

Figure 14. Comparison of two scans. (top) Conventional CT scan through bladder and seminal vesicles. (bottom) The same scan with coarse sampling and smearing of the resolution in the horizontal direction.

258 Michael Goitein

should the data be taken? The answer depends in part upon the nature of the disease and on the region of interest. The CT scanner at best delineates bulk disease and, as always, the therapist must determine final margins on the basis of predictions of microscopic extension. He will also wish to be guided by the location of nearby normal structures. Sampling theory suggests that the distance between scan slices should be one-half the desired spatial accuracy. In general, a reasonable goal would seem to be the definition of field margins with an accuracy of 1 cm, which then implies a slice spacing of 5 mm. Even finer sampling would be desirable for the more critical applications.

Such a scan sequence has immediate and stringent implications for the CT scanner. A typical volume of interest might be 20 cm in length and, if some margin is desired at either end, the inspection of a length of 30 cm would not be unusual. If this is to be sampled every 0.5 cm, it implies that the machine must be capable of taking 60 scans, each approximately 0.5 cm thick. These scans need to be taken in quite rapid succession if the entire procedure is not to take too long, which would compromise the ability to immobilize the patient and, perhaps, the possibility of doing the study at all if the scanner were needed for too long a time. The implications of this are discussed below in the context of scanner features necessary for accurate patient positioning.

The proposal that volumetric scanning should be the norm is met with surprising resistance. This seems to come from a number of factors: (1) those familiar with early generation scanners feel that it is not feasible; (2) it is suggested that, even if feasible, such a capability would be impossibly expensive to implement; (3) it is thought that volumetric scanning would be too time consuming; and (4) it is suggested that the dose to the patient from such a procedure would be excessive. Without question, volumetric scanning is possible. The only qualification is in terms of the density resolution that can be obtained in this mode and this question is discussed below. Such a capability is relatively inexpensive. The primary requirement, discussed below, is for increased mass storage or for data compression, and these may add about 3 % to the cost of the equipment. The duration of such a procedure depends on the repetition frequency with which scans can be made (discussed below) and on the reconstruction time. In current generation equipment the latter is on the order of 0.5 min, so that a 60-slice study would take half an hour to reconstruct. If this is too long a time, one can arrange to store the unreconstructed data and to batch process the reconstructions overnight. A "quick look" analysis of the data can be made either by reconstructing every second or third slice or by reconstructing on a coarse matrix. The dose to the patient is not, it turns out, too great, largely because of limitations on x-ray tube output. For a given x-ray tube output, the number of detected photons is roughly proportional to the slice


thickness. There is, in fact, some penalty to pay owing to the fact that not all x rays striking the patient are detected. In going from a 10-mm-thick slice to a 2-mm-thick slice there is probably a factor of 2 penalty to pay owing to this effect. If we ignore this effect, for a given mAs per scan, the dose delivered in four contiguous 2.5-mm slices would be the same as that delivered in one 10-mm thick slice. Of course, the noise in each scan would be worse by a factor of 2. If one wished to have the same quality of scan in the thinner slices, the mAs, and hence dose, would have to be increased by a factor of 4 per slice to compensate for the loss of flux. However, this is not possible owing to x-ray tube cooling problems (discussed below) and the mAs per slice would, if anything, have to be somewhat reduced. As a result, the absorption coefficient resolution in volumetric scanning cannot be as good as for multiple thick sections by something on the order of a factor of 2, and the dose is larger than that in thick section scanning by a small factor that depends on the slice thickness and the characteristics of the particular scanner. Where the statistical fluctuations in the data pose a problem in interpreting the image, adjacent slices can be averaged to reduce the noise. When viewing in planes other than the transverse plane, the noise reduction can be performed by averaging the data in planes contiguous to that of interest. For example, in viewing coronal sections, one would average several contiguous coronal sections so that the effective coronal section thickness would be of the order of 10 mm when low noise was required.

Volumetric scans are associated with relatively low numbers of detected photons, both because only modest x-ray exposures are possible and because small slice thicknesses are employed. It is therefore necessary that noise in the detector and its associated electronics be acceptably low in the face of low signals.

10.2.4. DISPLAY CAPABILITIES

The interpretation of CT scans is affected by the quality of their display. One of the significant innovations associated with CT technology was the introduction of contrast-enhancement techniques, without which it would be impossible to exhibit the full dynamic range of the data. Fortunately, the current technology of image displays has made good-resolution gray scale display the norm. Probably a 256 x 256 array would be the minimum acceptable display and a 320 x 320 or greater display would be more desirable. (For a given storage capacity, it would be more sensible to provide for a rectangular rather than square image. For example, a 256 x 400 matrix would be much more appropriate for both body sections and, rotated by 90°, for head sections than a square 320 x 320 image.) At least 64 full gray levels should be present. If rapid changes in window and level are to

260 Michael Goitein

be achieved, a deeper image memory, probably at least 10 bits, should be required. Multiformat cameras that offer gray scale images on transparent film have become the standard for hard copy. The use of color in the image display has not been widely accepted. Its major problem is that if the "palette" can be too easily varied, the images cannot be immediately appreciated and their interpretation can be highly subjective. A bright red "mass" can generally be placed almost anywhere one desires if sufficient control over the display characteristics is permitted. However, color has two important advantages that may yet make it attractive: (1) it increases the dynamic range of the display so that it is possible to unambiguously show lung structures, soft tissues, and bony densities in one picture and (2) it can be useful for coding, that is, for providing overlays-a particularly useful capability in dose distribution displays.

The linearity of the display and, particularly, of the hard copy device, can be important for purposes of dosimetry if the anatomic data is transferred to the treatment planning computer through the hard copy medium. The specification of linearity in these devices is therefore important. In addition, it should be possible to superimpose a grid (or tick marks) in two orthogonal directions on the image in order to establish the scale and check the linearity of the image.

Early scanners provided limited access to the quantitative CT information. The ability to perform at least a region of interest analysis, giving the mean and standard deviation of the CT numbers within a geometric or arbitrary region, is now available on most new scanners and should be considered a necessity. These features are of value both in the evaluation of the data for purposes of tissue identification and where absolute CT numbers are being used for dosimetry. The ability to view sagittal, coronal, and arbitrary sections is also valuable and, indeed, mandatory if volumetric scanning is performed.

10.3. Patient Positioning

Three general features are necessary: (1) appropriate scanner geometry, (2) the capability of completing the scan in a reasonably short period of time, and (3) the ability to document the patient position.

10.3.1. GEOMETRIC FACTORS

A flat scanner couch top is necessary (if the treatment couch is flat). If not provided with the scanner, a flat filler can always be made.

A large patient aperture is necessary if unacceptable distortions of the patient are to be avoided under certain circumstances. Probably a 60-cmdiameter aperture should be sought. Even this is inadequate for many arm


positions used in therapy. Probably it would be better to change the position used in therapy, if this is at all possible, in order to match the CT scan situation.

A large diameter of reconstruction is necessary so that the skin and any superficial localizing marks can be imaged and to achieve absolute accuracy of the CT number. The former is necessary both for registration of the patient using radio-gray skin marks and in order to determine the patient contour for dosimetry purposes. Accuracy of the CT number is necessary both for tissue identification and for dosimetry. A 50-cmdiameter reconstruction should be adequate in most instances. Finally, it should be possible to perform scans without the presence of bolus material. The use of a water bath in the original EMI scanner was an inspired solution to a range of problems. The effort to achieve comparable or superior scanning capability without a water bath or the use of bolus material has complicated the design of current machines. Nevertheless, this is an important capability if distortion of patient contour is to be avoided. To achieve it, good dynamic range of the detectors is necessary so they will respond linearly, even when exposed to the full beam. Also, a fast scan will minimize motion artifacts from the skin-air boundary.

10.3.2. PERFORMING A COMPLETE STUDY IN A REASONABLE PERIOD OF TIME

In order to achieve reproducible and reliable patient positioning, a study must be completed in an acceptably short period of time. The sequence of scans should not take longer than 10 or 15 min. (If studies both with and without a contrast medium are desired, an even shorter time per study would be desirable.) Such a time is comparable to the time in which patients are required to remain immobile during treatment, and, therefore, a shorter CT study time would not be a major advantage.

If volumetric scanning is desired, 60 scans need to be performed during this time, which implies a repetition frequency of approximately one scan every 10 sec. This requirement has numerous ramifications.

First, the x-ray tube must be capable of sustaining the heat load involved. With such a large number of scans it generally turns out that the anode cooling capacity is the limiting feature. Rotating anode x-ray tubes with anode cooling rates of 140,000 heat units per minute are available and such tubes are capable of performing up to 100 consecutive slices, one every 10 sec, with an output of 110 mAs per slice at 120 kVp • This is an adequate, though somewhat lower than optimal, x-ray tube output.

Current CT scanners are generally capable of scan times on the order of 2 to 3 sec, although additional time is generally required for start-up, slow-down, and, in some cases, rewind. Scan repetition rates of -1 per

161 Michael Goitein

5 sec are therefore currently possible from a mechanical point of view. Transfer of the measured data onto the mass storage medium (disk) can be time consuming and may limit the scan repetition rate for some scanners.

Since reconstruction times are generally longer than the 10-sec period of such repeated scans, and a significant proportion of that time is often required for the transfer of data between the computer memory and the mass storage device, it is not possible to reconstruct the scans as fast as they are taken. Rather, it is necessary to store the data until after the scan sequence is completed and then perform, or at least complete, the reconstructions. This places demands on the mass storage capacity that must be provided. This, in turn, depends on the number of measurements taken per slice, which can be as many as 1.5 million. Sixty scans then require the storage of 180 Mbytes of data. Considering that room will be required for storing completed images, it will generally be necessary either to augment the mass storage by an additional large disk or to provide data compression. With the latter, it is possible to achieve something like a 3 : 1 compaction of the data without loss of information. The cost of providing this capability is comparable to that of adding an additional disk and in either case is a modest increment on the cost of the scanner.

While deferring reconstruction until after the scan sequence is completed has the advantage of reducing the time the patient needs to spend in the scanner, it is nevertheless desirable to reconstruct the pictures as rapidly as possible, both so that the study can be reviewed before the patient leaves and to maintain the throughput of the unit. Reconstruction times of approximately 0.5 min per slice are widely available and highly desirable. A 60-scan volumetric scan could then be reconstructed in half an hour. It is possible to reconstruct on a coarser grid in a shorter time and to defer the full reconstruction until a later time, perhaps for overnight analysis. This, however, greatly increases the mass storage requirements.

Reconstruction times are a function of the number of measurements from which the reconstructions are made and the number of reconstructed pixels. Owing to oversampling, the former can be an order of magnitude more than the latter. The speed of reconstruction is partly a matter of cost, with the back projection process being a critical factor. Reconstruction of about 1 million measurements onto 100,000 pixels can easily be done in 30 sec, with an order of magnitude faster reconstruction being currently within the state of the art.

Obviously, automatic advance of the couch position between slices is necessary to achieve the repetition rate of 1 scan every 10 sec unless the operator is an Olympic class athlete.


10.3.3. DOCUMENTING THE PATIENT POSITION

Anyone who has used CT scans in planning radiation therapy has spent a considerable time trying, with varying degrees of success, to overlay the scan planes on AP and lateral films. The development of a plain film capability--called a scout view by one manufacturer-represents a major advance in documenting the CT scan situation and facilitating correlation of the scan with data from other studies. In this procedure the patient is moved past the source-detector system, which is held at a fixed angle, the slice thickness being tightly collimated to a thickness of 1 mm or so. This capability should undoubtedly be a requirement in CT scanners used for treatment planning purposes.

A slit-lamp indication of the scan plane and sidelights, probably offset by a well-defined distance from the plane of the scan, will assist in patient alignment and setup. The use of CT -compatible casts such as described in Section 7 is also valuable in establishing the patient position. Such devices should be provided with radio-gray markers for alignment purposes. A Polaroid camera in the scanner room to document the scan conditions would be a good investment.

10.4. Features Necessary for the Calculation of Dose

Most of the capabilities necessary to permit the calculation of dose have been discussed in preceding sections. These include absolute accuracy in the absorption coefficient measurements at the 2% level, availability of an iterative beam-hardening correction to assure this, a large region of reconstruction to avoid having portions of the patient outside the scan circle-a situation that makes the calculation of dose difficult for some beam angulations and that can lead to absolute CT number uncertainty, and the availability of a number of display options and good linearity in the hard copy device. In addition, it may be desirable to ensure that the quantitative data can be transferred to a treatment planning computer through an appropriate medium. Usually magnetic tape is the most versatile of the alternatives, although floppy disks have become quite popular. However, with volumetric scanning, floppy disks do not have adequate capacity. Various schemes for data transfer have been described.(1,37.67)

10.5. Summary of Specifications

Clearly, the specification of a CT scanner is a multifaceted problem. It would not be sensible to place inflexible requirements on all parameters, since compromises may be necessary, depending upon the available commercial equipment and the priorities of the individual user. To summarize

264 Michael Goitein

the above recommendations, the features of a cr scanner used in whole or in part for planning radiation therapy should include the following:

1. Excellent spatial and absorption coefficient resolutions. 2. The capability of scanning a large number (about 60) of relatively

thin (5 mm or less) sections continuously at a rate of 1 scan per 10 sec.

3. The capability of generating AP and lateral "plain film" views relative to which scans can be referenced.

4. A large patient aperture and a large diameter of reconstruction. 5. Scanning without bolus material. 6. High anode and x-ray tube housing cooling rates and high geometric

and photon conversion efficiency of the detectors. 7. Good display capabilities, including sagittal and coronal reconstruc

tion and display.

Table 9 presents some of these requirements in greater detail. These requirements are reasonable and there are commercially available scanners that can currently meet all of them.

10.6. Relationship between Diagnostician and Therapist

The relationship between the diagnostician and therapist deserves attention in three areas: (1) in the selection and purchase of a CT scanner, (2) in the day-to-day operation of the scanner, and (3) in their interaction over the interpretation and use of scans.

As discussed in earlier sections, CT has proven itself of substantial value in planning radiation therapy treatments. Moreover, patients with cancer form a significant proportion of the patients scanned on whole body scanners (70% in one study(68) were being assessed for a question of malignant disease) and about half of those patients with cancer will ultimately receive radiation therapy. Application to radiation therapy is often cited as part of the justification for purchasing a scanner. It is therefore reasonable to propose that therapy departments playa significant role in defining specifications for CT scanners bought by their institutions. All too often it is thought that providing a flat couch top, a large "hole," and, perhaps, an expensive treatment planning capability mounted on a remote diagnostic console will satisfy the needs of therapy. It is important that the therapy group play a more substantial and critical role in the selection of equipment.

In many centers with whole body scanners, despite the substantial proportion of patients with malignant disease being assessed, only a few

Tab

le 9

. Su

mm

ary

of C

T S

cann

er S

peci

fica

tion

s R

elev

ant t

o T

reat

men

t Pla

nnin

g A

ppli

cati

ons

>

'G

'C:I

Of

valu

e fo

r ~

110 5-

tiss

ue

pati

ent

dose

=

D

iscu

ssed

'" Q

Fea

ture

de

line

atio

n po

siti

onin

g ca

lcul

atio

n C

omm

ent

in s

ecti

on

- ~ .-j

Spa

tial

res

olut

ion

S·

in s

can

plan

e A

s go

od a

s po

ssib

le

10.2

.1

:= 110

out

of s

can

plan

e •

Min

imum

sca

n th

ickn

ess

.;;2

mm

10

.2.1

~

Abs

orpt

ion

coef

fici

ent

reso

luti

on

!i- II

rand

om

• A

s go

od a

s po

ssib

le

10.2

.2

.. 110

syst

emat

ic

Abs

olut

e ac

cura

cy <

2%

10

.2.2

'C

:I •

'<I!

"Pla

in f

ilm"

capa

bili

ty

• "S

cout

vie

w"

or

com

peti

tor'

s eq

uiva

lent

10

.3.3

.-

j .. II

Fla

t co

uch

top

• M

ake

it yo

urse

lf,

if ne

cess

ary

10.3

.1

i L

arge

pat

ient

ape

rtur

e ;;.

60 c

m in

dia

met

er

10.3

.1

II

Lar

ge r

econ

stru

ctio

n re

gion

•

• ;;

.50

cm in

dia

met

er

10.3

.1

a. Sc

an w

itho

ut b

olus

R

equi

res

larg

e dy

nam

ic r

ange

of d

etec

tors

10

.3.1

:s!

11

0

and

fast

sca

n =

=

N

umbe

r of s

lice

s con

tinu

ousl

y sc

anne

d ;;.

60

10.2

.3

S·

IIQ

Hig

h sc

an r

epet

itio

n ra

te

• 1

scan

per

10

s fo

r 60

cyc

les

10.3

.2

X-r

ay t

ube

cool

ing

Lar

ge a

node

and

hou

sing

coo

ling

rat

es

10.3

.2

Mas

s st

orag

e •

180

Mby

tes

extr

a, o

r da

ta c

ompr

essi

on

10.3

.2

Fas

t re

cons

truc

tion

tim

e •

• .;;

30 s

ec p

er s

lice

10

.3.2

A

utom

atic

cou

ch a

dvan

ce

• •

10.3

.2

Dis

play

and

har

d co

py

;;;'3

202

; ;;;

.64

gray

lev

els;

lin

ear

disp

lay;

10

.2.4

gr

id

and

/or

scal

e;

regi

on

of i

nter

est

anal

ysis

It

erat

ive

beam

-har

deni

ng c

orre

ctio

n •

10.2

.2

Dua

l en

ergy

sca

nnin

g (F

or r

esea

rch

appl

icat

ions

) E

ffic

ient

pho

ton

util

izat

ion

• 10

.2.2

L

ow e

lect

roni

c an

d d

etec

tor

nois

e 10

.2.3

N

•

CI\

til

266 Michael Goitein

"slots" are reserved for "therapy patients." In part, this is because of a failure to anticipate that a given patient may ultimately require radiation therapy. This is partly a function of the unfortunate distinction between "diagnostic" and "therapeutic" scans, which has left the impression that the latter relate to purely dosimetric needs. Greater attention needs to be paid to anticipating future needs when scanning patients in whom a malignancy is suspected. In addition, the CT staff needs to be aware of the needs of radiation therapy.

Nowhere is the misapprehension concerning the nature of therapeutic scans more serious than in the attitude towards the interpretation of scans. Obviously, if the full extent of disease is to be assessed in order to design a plan of treatment, the greatest possible expertise should be brought to bear on the analysis of the CT scans. This certainly should involve consultation between the diagnostician and therapist. This consultation should take place at a CT viewing console and not over a murky film. To beproductive, such a consultation will be time consuming and special recognition of this should be made.

10.7. Inexpensive Scanners

Both because of a desire to measure absorption coefficients at therapeutic beam energies and in recognition of the reduced spatial and absorption coefficient resolutions necessary to support dosimetric calculations (see Section 8), a number of designs have been proposed for scanners for what are termed "treatment planning purposes." A commercial option is available on the old Delft simulator. This uses the image-intensifying capability of the simulator in connection with special-purpose reconstruction hardware to provide tomographic sections of patients. Other simulator modifications have been proposed. These include the use of a multielement xenon-gas proportional detector,<69) the use of a fluorescent screen viewed by an Isocon camera, (70) and a film pack that is translated as the simulator rotates.(71) In addition, scans using high-energy sources such as cobalt-60 or cesium-137 have been proposed.(72,73)

The danger of such devices is twofold. First, efforts may be made to make interpretations of tissue densities that may not be supported by adequate data. Such scans tend to have rather poor resolution and will not provide anything like the capabilities of diagnostic CT scanners. Second, these units tend to be inexpensive only in prospect. In retrospect, and upon proper accounting of the time spent and effort made, they may be seen to be less so. For a department with limited resources it may well be that the purchase of the least expensive diagnostic CT scanner available together with the cheapest available simulator would be a better expenditure than


the purchase of a more elaborate simulator and the development or acquisition of a crude CT capability.

10.8. Treatment Planning Programs

There has been a recent burst of commercial activity in implementing a treatment planning capability for specific CT scanners. For the most part, manufacturers have been interested in using the computer and display capabilities provided for diagnostic CT viewing as the core of the treatment planning hardware. This effort has been in recognition of the important role of CT scanning in planning radiation therapy and, conversely, the important role that the needs of radiation therapy play in justifying the purchase of a CT scanner.

In general, the treatment planning programs use conventional treatment planning approaches and overlay them upon CT scans. In their present state of development, it is unclear as to whether they offer significant advantages over the use of conventional treatment planning computers in which the CT information is transferred using some form of hard copy such as film, Polaroid, or a printout. Ultimately, if three-dimensional treatment planning capability is implemented in which a beam's eye view of the treatment volume can be offered, allowing the user to define the field shape and angulation, these computer systems will be very useful. However, a number of points should be kept in mind: (1) at the present time these capabilities are not available and treatment planning may be an expensive option with a relatively modest technical gain, (2) the development of new code is a considerable expense that manufacturers are generally interested in minimizing, particularly since the CT market has quieted down, (3) access to the treatment planning capability is highly dependent on the extent and vigor of competition for use of the console (where the treatment planning code is mounted on an operator's console, one would expect it to be virtually inaccessible for treatment planning purposes; where a remote diagnostic console is involved, the availability of the code will depend on geopolitical factors), and (4) the development of clinically viable treatment planning codes has in the past involved person-decades of software development, and several strong commercial ventures have foundered in the process of developing such a capability. The current level of activity and commitment to a long-term developmental effort and systems support would be important matters to assess in deciding whether such a capability was worth acquiring.

The availability of a dose calculation should not dominate the assessment of whether a given scanner is appropriate for radiation therapy. In the present state of development of such programs, numerous other criteria,

268 Michael Goitein

discussed above, should be the primary factors in comparing the relative merits of scanners.

11. EPILOGUE

CT has made possible anatomic mapping and differentiation of tissues undreamt of a decade ago. These capabilities are of great value both for the detection and assessment of malignant disease for purposes of planning overall therapeutic strategy and in guiding radiation therapy. Sever~l studies concur in their evaluation of the potentially significant impact that CT has in planning treatments. There is every reason to expect that this impact will translate into worthwhile improvements in treatment outcome and that CT will be a cost-effective tool for treating malignant disease.

To realize the benefits of CT requires attention to the specification of the scanner and to its mode of use in daily practice.

CT has opened a new vista in the treatment of cancer and, in particular, in guiding the delivery of radiation. Patients will be the better for it.

ACKNOWLEDGMENTS

I am grateful to numerous colleagues for freely providing their data in advance of publication and for their advice. These include J. Ammon, J. J. Battista, H. E. Brizel, F. Casamissima, G. T. Y. Chen, N. J. Hodson, W. J. Hoogenraad, J. Jallamarque, M. T. Kademian, J. E. Munzenrider, R. P. Parker, M. V. Pilepich, B. Pullan, J. A. Purdy, B. A. Schlager, B. J. Smit, and J. W. Walsh. Dr. M. Angell kindly read the manuscript and Ms. Barbara Grzybek provided heroic assistance in its preparation.

APPENDIX: REVIEW OF SITE·SPECIFIC STUDIES ON THE IMPACT OF CT t

A.t. Prostatic Cancer

Treatment of prostatic cancer with radiation has become widespread since the studies of Bagshaw et al.(74) were reported. The treatment usually comprises a comprehensive pelvic field and is often delivered with a fourfield "box" technique, with a boost to the prostate and adjacent involved tissue. Quite sophisticated simulation procedures have been developed, using a variety of contrast media to define the margins of the prostate.

t This section is based on a review of the literature published before January 1980.


A number of studies have dealt with the role of CT in planning these treatments. Kademian et al. (75) evaluated 31 biopsy-proven cases of cancer of the prostate, primarily stages Band C. The CT scans followed the definition of the target volume by conventional techniques. These authors developed and initiated the use of a digital localizer, a probe that could be introduced into the rectum by the physician, who would place its tip at the most cephalad extent of palpable disease and determine the corresponding position on the anterior skin. Kademian et al. were then able to evaluate the adequacy of the treatment as judged by CT, both with the conventional techniques of localization and with the addition of the new localizer. In the 20 patients evaluated before the digital localizer was available, chan!Jes of at least 1 cm in one or more field margins were necessary in 9 (45%) of these patients. With the digital localizer, a change was necessary in only 1 of 11 patients (9%). Overall, changes were required in 10/31 (32%) of the cases. These authors emphasized the ability of CT to demonstrate the seminal vesicles. Ultrasound scans were obtained during the same study and were said to verify the CT findings to a large extent, but not to demonstrate the seminal vesicles as well. Three-quarters of the changes made were in the superior portion of the treatment field.

Pilepich et al.(76) evaluated 35 patients with prostatic cancer. They found that CT had limited efficacy in the definition of tumor size. However, they judged it to be useful in treatment planning. In their series, the field placement had to be changed on the basis of data gained from CT in 12/35 patients (23%.) and the field volume had to be increased in 4 of the patients (11%), for an overall incidence of changes in 16/35 (34%) of the cases. Pilepich et al. judged that changes which would make at least a 6% impact on the maximum tumor dose took place in 6 of the 35 patients (17%). For four of the patients these changes would have led to a dose change of at least 10%.

Hoogenraad and Bakker(77) have reviewed the impact of CT in 12 patients treated for prostatic carcinoma. They concurred that it was not possible to differentiate the tumor from adjacent normal tissue, but felt that CT was extremely helpful in delineating the seminal vesicles. In their study, CT scans led to changes in at least 4 of 11 patients (36%). Hoogenraad et al. made dose calculations with and without CT information about bony inhomogeneities and concluded that dose estimates in the absence of CT might underestimate or overestimate the tumor dose by 7-10%.

Brizel et al.(12) analyzed 14 patients with prostatic disease. They found that unexpected spread of tumor beyond the prostate was present in five cases (36%). They also found that CT led to a change in the goal of treatment for three patients and was responsible for increased accuracy of portal alignment based on the demonstration of normal and tumor anatomy

270 Michael Goitein

in six patients, leading to the conclusion that CT was of help in the treatment planning of nine patients (64 % ).

These studies are interesting from a number of points of view. They agree that the delineation of the tumor is not straightforward and that in many areas it is difficult to distinguish between malignant and nonmalignant tissue. On the other hand, it is clear that the prostate can be well demonstrated in most instances and that an abnormally enlarged gland, or one that is asymmetric, can readily be detected. The presence of enlarged seminal vesicles is an unexpected finding with cancer of the prostate, which these and other authors have noted. (!l2) Often such seminal vesicles lie outside the conventionally planned boost field. However, the importance of this is uncertain. Clearly, correlation between the CT findings and the pathologic evaluation is needed.

In all cases a substantial frequency of therapeutic changes was initiated by information gained from the CT scan. Averaged over the four studies, changes were made in 32% of the cases. The results are summarized in Table 10. It is interesting to note that different centers have found CT to change their standard practice in different ways. The Wisconsin group found that CT led largely to modifications of the superior margin of the field; the Utrecht group modified their standard technique primarily in the posterior margin of the field; and the MGH group found that unexpected lateral extension of the prostate (and sometimes seminal vesicles) has been one of the major surprises. Clearly, quite different conventional

Table 10. Summary of Studies of Prostate

Authors Comment

Kademian et al. a Conventional technique Digital localizer

Pilepich et al. b Dose changed ;;.6% Hoogenraad and

Bakerc

Brizel et al. d

"In ref. 75. 'In ref. 77. bIn ref. 76. dIn ref. 12.

Total number of patients

20 11

Overall: 31

35

11 14

Significant changes necessitated by CT

Number % of total of patients patients

9 45 1 9

10 32

6 17

4 36 9 64


practices were obtained at the different institutions and one should not assume that the detailed conclusions of anyone study will have universal application.

Two of these studies indicated that it may be possible to develop from the evaluation of the CT findings new treatment techniques that may obviate the continued need for CT scans. The Utrecht group felt that CT led them to modify the placement of their borders in a way that might be universally applicable to their patients. The Wisconsin group hopes that the digital localizer will be sufficiently effective in delineating the prostate to make it possible, once further CT evaluation has confirmed this, to proceed without routine CT scanning.

Clearly, the role of ultrasound in the evaluation of the prostate is an important one. These studies and others(78) have partially addressed the question of how important, but it remains unresolved. There is a preference for the CT scan relative to the ultrasound scan in these papers. It is clear, however, that one of the major problems of the CT scan is in the limited resolution in the cephalo-caudad dimension-a problem that may be substantially reduced with fast scanners having the capability to take and reconstruct a large number of views in a short period of time. However, the extent to which the preference for CT can be justified in face of its higher cost remains unclear.

A.2. Bladder Cancer

The role of CT in the staging and management of bladder cancer has been investigated by several groups. All of the studies indicate that CT has an important role to play in establishing the staging of the disease. Clinical staging is currently based upon techniques that include urography, cystography, bimanual examination under anesthesia, biopsy, and sometimes bilateral pedal lymphography. These techniques are often unable to determine whether extravesicle extension of disease is present. CT often provides this information. Yu et al.(8) have studied 60 consecutive patients with a histologic diagnosis of bladder cancer. Seventeen of these had surgical exploration and it was possible to correlate the staging based on CT and histologic findings in these cases. In 15 of the 17 cases the CT and surgical/histological staging agreed. There was a discrepancy in two cases: in one, CT demonstrated extension of disease unsuspected by clinical staging but confirmed in subsequent surgery; in the other, the CT scan failed to demonstrate a l-cm-diameter node. In the 43 patients for whom only clinical staging was available, CT upgraded the staging in 8 cases (19%). Six were upgraded from stage B to stage C, and two from stage C to D-that is, CT was important in finding clinically unsuspected extravesicular extension. Yu et al. reported that CT was not able to define

272 Michael Goitein

the extension of disease within the bladder wall, that is, could not differentiate among stages A, Bh and B2 •

Hodson et al.(9) have studied the role of CT in staging bladder cancer. Their data are shown in Table 11. Overall, the CT staging was different from the clinical staging in 28 out of 75 evaluable patients (37%). In 14 of these (19%) CT led to an upgrading of disease, primarily due to its demonstration of extravesicular extension of disease. In a further 14 cases the CT showed a lesser extent of disease than clinically observed. This was primarily due to the superior assessment of extra vesicular extension of disease by CT. Furthermore, it was possible to demonstrate extravesicular extension in 26 of 36 clinical stage III cases (72 %), thus enabling a division between stage T3A and T3B and above. Surgical-histological correlation was possible in 10 of the 75 patients. In four of these the CT demonstrated a residual mass that was histologically confirmed as cancer. In six, the CT found an increased bladder wall thickening, and in only one of these cases was the histology positive. Hodson et al. concluded that in irradiated bladder, localized thickening of the bladder wall cannot be regarded as a reliable sign of residual disease.

The impact of CT on the treatment technique was evaluated in two studies. Schlager et al. (79) found that treatment plans were changed in 6 out of 21 cases (29%) as a result of the CT scans. Hodson et al.(9) found that 9 out of 52 (17%) had a significant change of plan due to CT. In one case this was due to an altered dose estimate that would have been more than 5 % in error in the absence of CT information. In one case the irradiated volume could be reduced based on CT information, and in the remaining seven cases the target volume coverage was judged to be inadequate on the basis of the CT scans. They concluded that the CT findings must cast doubt on the assumption that geometric miss is a relatively unimportant factor in failure to control bladder cancer. Yu et al.(8) also concluded that

Table 11. Staging Comparison in Bladder Carcinoma a

CT stage

T2, T3A, T3B, T4A, T4B, Clinical Number of superficial deep perivesicle adjacent pelvic

stage patients muscle muscle involvement organ wall

T2 14 4 7 3 T3A,B 36 10 23 3 T4A 12 8 3 1 T4B 13 6 7

Total 75

"From Hodson et al. in ref. 9.

AppUeations of CT In Radiotherapy Treatment Planning 273

with CT, "we can plan more accurately and confidently, avoiding 'geographic misses,' excessive 'safety margins' and protecting other organs."

CT was also thought to be valuable in the assessment of the response to treatment. Yu et al.(S) found that 10 out of 25 (40%) patients followed with CT had residual disease. Hodson et al.(9) found that 50% of their patients receiving only radiotherapy had residual disease at a mean time of 20 weeks after treatment. Yu et al. suggest that with such follow-up, "it may be possible to undertake appropriate surgery at an earlier stage of recurrence in those patients who were treated with definitive irradiation and decrease the incidence of feared complications often attending later intervention. "

Schlager et al. (79) analyzed the distribution of the dimensions of the bladder in their patients and found wide variability in the numbers. Their data indicated that a 10 cm x 10 cm field "will encompass the bladder adequately in most patients." Another interpretation would be that such a field would be unduly generous in a substantial proportion of patients and that CT scanning would permit a careful tailoring of the field to the individual circumstances-an interpretation supported by the other studies. Schlager et al. also analyzed the extent to which the bladder volume changes after voiding. They found a smaller volume reduction than anticipated, (69 ± 10)%. Yu et al. often performed scans both in the prone and supine positions and with and without negative contrast medium (C02 ) in the bladder. These maneuvers were primarily to define disease rather than to choose an appropriate situation for therapy.

The value of CT in assessing nodal disease was addressed by Hodson et al.(SO) Of 11 patients shown to have positive nodes by lymphography, only 4 were demonstrated by CT scan. All six cases missed by CT proved to have nodes smaller than 1.5 cm in diameter, which was the criterion for a positive node on CT scan. In 25 patients who had negative nodes on lymphography, CT demonstrated 1 case with a positive node that was proven upon subsequent surgery.

Brizel et al. (12) analyzed 17 patients with bladder carcinoma. In eight cases (47%) CT demonstrated a larger tumor than previously suspected. Five patients (27%) were upstaged owing to muscle involvement or extravesicular extension not appreciated in endoscopy and biopsy. Brizel et al. judged CT to be helpful in the radiation therapy treatment planning of 12 of the 17 cases (71%).

Results of these studies are summarized in Table 12. Husband and Hodson were able to make a valuable radiologic

pathologic correlation by scanning 14 patients just prior to cystectomy.(Sl) The CT gave one false positive and one false negative in this series and, in addition, five patients had abnormal thickening of the bladder wall after their initial course of radiation therapy. None of these patients were found

Tabl

e 12

. Su

mm

ary

of S

tudi

es o

f Bla

dder

Yu

eta

l.b

Bri

zel

et a

l. d

N

Sch

lage

r et

al.

a H

odso

n et

al.

C

...:a ""

Sta

ging

G

ave

addi

tion

al i

nfor

ma-

Surg

ical

ly e

xplo

red

grou

p C

T s

tagi

ng d

isag

reed

wit

h 5

/17

(29

%)

upst

aged

bec

ause

of

tion

on

"unu

sual

con

-1

5/1

7 c

ases

CT

sta

ging

cl

inic

al s

tagi

ng i

n 2

8/7

5

mus

cle

invo

lvem

ent

or

tou

r" a

nd "

poss

ible

ag

reed

. (3

7%)

of c

ases

: up

grad

ed

extr

aves

icle

ext

ensi

on n

ot

exte

nsio

n of

tum

or

Cli

nica

lly

stag

ed

14

/75

, do

wng

rade

d 14

/75.

ap

prec

iate

d in

end

osco

py a

nd

outs

ide

the

vesi

cula

r C

T u

pgra

ded

stag

ing

in 8

/43

C

T g

ood

for

dete

ctin

g bi

opsy

. 0

/17

dow

nsta

ged.

w

all.

" (1

9%)

case

s-pr

imar

ily

by

extr

aves

icul

ar e

xten

sion

but

de

mon

stra

ting

uns

uspe

cted

po

or a

t de

mon

stra

ting

ex

trav

esic

ular

dis

ease

. in

volv

emen

t of

adj

acen

t or

gans

. S

urgi

cal/

hist

olog

ical

co

rrel

atio

n in

ten

pati

en

ts-

five

conf

irm

ed,

five

wit

h bl

ad-

der

wal

l th

icke

ning

see

n in

CT

w

ere

hist

olog

ical

ly n

egat

ive.

Tre

atm

ent

6/2

1 (

29%

) ch

ange

in

Bel

ieve

tha

t w

ith C

T,

"we

can

9/5

2 p

atie

nts

had

sign

ific

ant

CT

was

of

help

in

trea

tmen

t te

chni

que

plan

due

to

CT

. pl

an m

ore

accu

rate

ly a

nd

plan

cha

nge

due

to C

T.

plan

ning

: ch

ange

in

Rx

obje

c-co

nfid

entl

y, a

void

ing

'geo

-ti

ve o

r po

rtal

con

figu

rati

on,

6;

grap

hic

mis

ses,

' exc

essi

ve

chan

ge i

n po

rtal

ali

gnm

ent,

6;

'saf

ety

mar

gins

' an

d pr

otec

t-T

otal

, 1

2/1

7 =

71%

. in

g ot

her

orga

ns."

Fol

low

-up

Det

ecte

d re

sidu

al d

isea

se i

n 50

% o

f pa

tien

ts r

ecei

ving

3

/8 p

atie

nts

had

dem

onst

rate

d 1

0/2

5 (

40%

) of

pat

ient

s.

radi

othe

rapy

onl

y ha

d re

spon

se t

o R

x: o

ne p

atie

nt

resi

dual

dis

ease

at

a m

ean

chan

ged

from

pal

liat

ion

to

tim

e of

20

wee

ks.

cura

tive

sur

gery

du

e to

CT

.

Com

men

ts

Ana

lyze

d va

riab

ilit

y of

R

evie

wed

60

cons

ecut

ive

case

s.

Par

t of

a r

evie

w o

f 72

con

secu

-f.

blad

der

volu

me.

ti

ve c

ases

of

pelv

ic d

isea

se.

:r

to !!. ~

"In

ref.

79.

<I

n re

f. 9

. =

bI

n re

f. 8

. dI

n re

f. 1

2.

~


to have a mass on cystectomy, but two had evidence of microscopic tumor within the bladder wall. These data are consistent with those of Seidelmann et al., (83) who reported an accuracy with CT of 81 % in 21 patients in whom surgical confirmation was available, and Kellet et al., (84) who found agreement between CT and pathologic findings in 12 of 15 patients (80%).

A.3. Other Pelvic and Abdominal Diseases

Brizel et al. (12) reviewed 72 consecutive patients with pelvic malignancies who had a total of 104 CT examinations. Their results as they relate to the prostate and bladder have already been referred to above. Overall, they found that CT demonstrated a tumor volume larger than originally suspected in 40% of the 72 patients, and unsuspected extension of tumor in 43 % of the cases. The CT findings led to an upward revision of staging in 25% of the patients. They found that treatment planning was aided in 44 of 72 cases (61%). "In 22 patients, CT findings led to significant adjustments in total dose, fractionation, portal configuration, or treatment objective." In the remaining 22 patients CT was reponsible for increased accuracy of portal alignment.

Many of the studies evaluating the role of CT in radiation therapy treatment planning have been confined to patients being treated with curative intent. Ammon et al., (85) in discussing the role of CT in the staging and treatment of urologic cancer, emphasized that CT has "undoubted value in treatment planning of urologic tumors where palliation only is the aim." Brizel et al., since they reviewed consecutive cases, also considered patients who were being planned for palliative treatment. They reviewed their cases to determine whether a preselection of those who had benefited from CT would be possible. They stated: "No criteria could be established. Some patients thought to be candidates for palliation were treated for cure either initially or after follow-up scans. Others, initially thought to be candidates for radical radiotherapy were treated palliatively." Clearly, the discovery that a patient's disease is such as to enable a potentially curative treatment to be attempted is a major benefit and it will be interesting to see this aspect of CT further documented. Moreover, the effectiveness of palliation also depends very much on the technical details of treatment delivery. The ability to offer good palliation of disease is one of the major capabilities of modern cancer therapy. Thus, it is quite appropriate that CT scanning be evaluated for the role it could play in the palliative treatment of malignancies. Such an evaluation has not, apparently, yet been made.

Brizel et al. concluded that "computed tomography proved to be of such value in staging, treatment planning and follow-up that it is recommended as an adjunct in the workup and management of patients with pelvic malignancies referred for radiation therapy."

276 Michael Goltein

Yeh(86) has compared ultrasonography and computed tomography in the assessment of carcinoma of the gallbladder. Seven patients were assessed by both modalities. Yeh concluded that details of the structures in and around the gallbladder were shown better on ultrasonograms. He identified the deficiencies in CT scanning as being caused by artifacts due to motion from gas and/or contrast in the gastrointestinal tract, which might be decreased with faster scan times than those of the 2.5-min scanner used in this study.

Lee et al. (87) have evaluated computed tomography as a technique for the assessment of malignant gastric neoplasms. They scanned 25 patients with abnormal upper gastrointestinal studies. Eleven of these patients were subsequently proven to have cancer and 14 did not. In the work-up of these patients, special care was taken in the administration of contrast agent and scans were performed with the patient both prone and supine and, where necessary, in the lateral decubitus position. The CT image of the stomach wall was good, probably owing to the use of a fast (4.8-sec) CT scanner. While the investigators concluded that CT was not the primary diagnostic modality for detecting malignant lesions (barium studies and endoscopy playing this role), they believe that "CT may provide valuable additional information regarding intra-abdominal extent." Such information is useful for staging the lesion, assessing the surgical resectability, and evaluating the tumor's response to therapy. They concluded that CT was crucial in avoiding laparotomy in two patients in their study, provided significant information for planning surgery in one patient with lymphoma by demonstrating a large ulcerating gastric lesion perforating the pancreas, was useful for documenting results of chemotherapy and radiotherapy in another patient with lymphoma, detected a recurrence in a patient who had had surgery, and helped shorten the clinical work-up in some patients with suspected gastric tumor by demonstrating pancreatic carcinoma invading the stomach in one case and enlarged spleens causing abnormal barium studies in others. The CT findings were "correct" in the 14 patients without gastric cancer.

A significant component of the studies of other sites within the abdomen and pelvis has been devoted to the use of CT in the evaluations of the response to therapy and of suspected recurrent disease. These studies are discussed in more detail in Section 4 above.

A.4. Thorax and Chest Wall

The thorax and chest wall present unusually severe problems of dosimetry owing to the wide range of tissue densities present. This is


discussed in Section 8. The localization of tumors in these areas has been discussed by a number of authors. Emami et al. (88) reviewed the effectiveness of CT scanning in planning radiation therapy in 32 patients with bronchogenic carcinoma. Tumor was seen in all but one of these patients. The treatment volume was changed in a total of 12 patients (38%), being increased in 10 and decreased in 2. Areas of tumor involvement not suggested by other data were seen in the CT scan in 21 patients. In 11 patients, direct extension of the tumor into previously unsuspected areas was seen; unsuspected nodes were found in nine patients, and other areas of tumor involvement were seen in six patients. In two patients, areas of involvement suggested by the conventional radiographs were found to be normal on subsequent CT scan.

Staging of disease was affected by CT. The T stage changed in 14 patients (44%), being raised in 13 and lowered in 1. The assessment of nodal stage was changed in nine of the patients (28%).

It is of interest that CT changed the treatment intent for five patients (16%). In four cases the original curative plan was changed to palliative treatment; in one case a patient planned originally for palliative treatment received curative treatment as a result of the CT scan.

CT was judged to be essential if it led to a significant change in volume or if part of a tumor would have been missed without it. This was the case for 17 of the 32 patients (53%). CT was judged to be helpful in a further 14 patients, and not useful in the remaining patient. The authors conclude that with CT

local treatment failure as a result of inadequate definition of tumor and nodal extension decreased. More confident selection of patients for radical radiotherapy should be possible, while for others with uncontrollable bone or chest wall invasion, palliative techniques should be more appropriately offered.

An interesting complementary study has been reported by Underwood et al.(57) These authors correlated the CT findings in 18 patients with proven bronchogenic carcinoma with the histologic evidence from staging mediastinoscopy or thoracotomy. That is, they evaluated the question of extension of disease to the mediastinum from proven bronchogenic carcinoma. Their findings are summarized in Table 13. The accuracy of CTwas 12/18 (67%). The specificity was 93%, and sensitivity 56%. Their findings led them to conclude that CT "should not replace surgical exploration of the mediastinum to evaluate the extent of spread (stage) of bronchogenic carcinoma." Had the CT evaluation been relied upon, one-third of the patients would have been inappropriately staged. However, CT did provide useful information concerning the mediastinum not obtained with conventional

278 Michael Goitein

Table 13. Correlation of CT Scan with Mediastinoscopy a

Result

True positive (positive CT, positive mediastinoscopy)

True negative (negative CT, negative mediastinoscopy)

False positive (positive CT, negative mediastinoscopy)

False negative (negative CT, positive mediastinoscopy)

Total

"From Underwood et al. in ref. 57.

Number of patients

4

8

5

18

Percentage of total patients

22

44

6

28

100

roentgenograms in 3 of the 18 patients (17%). These workers concluded that

CT scanning of the thorax should not be used as a definitive staging technique in lieu of the current accepted operative staging procedures; however, scanning of the thorax is certainly useful in the evaluation of selected patients with bronchogenic carcinoma.

This study provides valuable pathologic correlation and could serve as a model for much needed studies of CT findings.

Casamassima et al. (89) have reported on their experience with CT in localization of Hodgkin's disease in the mediastinum. They do not report specific numbers, but conclude that CT is valuable in the assessment of thoracic disease to (a) define extension of disease to the thoracic wall, (b) define the extent of contiguous involvement to the pulmonary parenchyma and ilo-mediastinal disease, (c) define residual disease after therapy, and (d) distinguish postirradiation fibrotic changes from relapsing disease. Others, it should be noted, have had difficulty in distinguishing fibrosis from residual or recurrent tumor.

Mintzer et al.(90) have conducted a prospective evaluation of 100 patients who proved to have chest malignancies, comparing the results of conventional chest tomography and CT. They concluded that conventional tomography was more useful than CT in evaluating the hilus, but that CT was the superior technique in assessing the mediastinum. As a result, in the later cases in their study, a CT finding of mediastinal disease was used as an indication for avoiding thoracotomy for a curative resection. However, while the CT was accurate in 12 of 16 patients with proven mediastinal involvement, it was falsely negative in the remaining 4 patients. In 42 patients with metastases to the chest, CT scans of the lung parenchyma


were more sensitive than whole lung tomography but had little additional impact on patient treatment, although in 18 of these patients the results of computed tomography or conventional tomography directly affected patient therapy.

Munzenrider et al. have assessed the role of CT in the treatment planning of the internal mammary nodes in 46 postmastectomy patients. (91) They found that the "internal mammary nodes were not routinely visualized, but abnormalities consistent with enlarged nodes were seen in 6 patients and gross recurrence in another." Their main emphasis was on the measurement of the chest wall thickness, which they evaluated at several points, both laterally and along the long axis of the body. Munzenrider et al. found a significant range of thicknesses-making a "standard" specification of the nodal target volume inappropriate. They estimated that from 18 to 42% and from 54 to 74% of measured interspace chest wall thicknesses on the operated and nonoperated sides, respectively, were thicker than 31 mm at 3 cm from midline. They also found that the measured chest wall thickness correlated positively with the patient weight. They conceded that ultrasonography may be a competitive technique for the measurement of chest wall thicknesses and recommend a comparison with CT scanning for this purpose. The use of standard thicknesses led to a variation from the prescribed dose for cobalt-60 en face irradiation of -7 to + 11 %. For electron beam irradiation a more significant error would be introduced by the use of standard thicknesses, and in a further study(92) Munzenrider et al. concluded that the

risk of internal mammary (node) underdosage with direct [en face] portals was estimated to be 5 to 10% with cobalt-60 and 15 and 20 MeV electrons, and 34% with 12 MeV electrons,

and 75% with tangential cobalt-60 portals. They concluded, therefore, that

routine determination of chest wall thickness [on a patient by patient basis] is recommended.

It is interesting to compare this observation with those of Kademian et al.(75) and Hoogenraad and Bakker,(77) who, in analyzing the CT scans of the prostate, concluded that the CT analysis could lead to standardized treatment that would.obviate the necessity of a patient-by-patient evaluation of disease. Such an experience is not necessarily universal. Two propositions may hold: (1) CT scans may make it possible to establish the range of variation of abnormal as well as normal anatomy to the point that standardized treatment techniques can be devised that would be adequate for an acceptable proportion of patients; and (2) where a significant range

280 Michael Goitein

of tissue disposition exists, individualized CT scans would permit the sparing of normal tissues that would otherwise be irradiated to high dose by the standardized techniques. The extent of the benefit provided by the second possibility will, of course, vary from site to site and depend on the range of abnormal and normal anatomy, as well as on the nature of the tissues that could potentially be spared by individually scanning patients. The benefit will probably be found to be significant in many instances. Certainly one should not accept a "standardized" treatment technique without documenting the potential tissue sparing that could be achieved if individualized scanning were performed.

A discussion of the role of CT in the evaluation of breast cancer is not complete without mention of the special-purpose CT scanner developed for the diagnosis of breast disease. A study of 1000 patients scanned by this machine has recently been reported by Chang et al. (93) They concluded that the special-purpose CT unit, using contrast medium enhancement, is superior to mammography, thermography, and physical examination for diagnosing both benign and malignant breast disease. They reported that CT has unique advantages in (a) the detection and diagnosis of early breast lesions less than 5 mm in maximum diameter, which are practically impossible to detect on xeroradiographs or film mammograms (two cases were detected in this category), (b) the diagnosis of cancer in breasts so large that mammograms are not satisfactory (9 of 44 proven breast cancers fell into this category), (c) the correct diagnosis of benign fibrocystic disease when it mimics cancer on mammograms, (d) assisting in the choice of biopsy sites when multiple lesions are seen in mammography, (e) the detection of residual cancer in the face of postoperative breast changes, (f) the differentiation of recurrent disease from fibrosis following therapy, and (g) the preoperative detection and localization of potentially precancerous lesions of the breast. The diagnostic accuracy of CT in 44 proven cancers was 93%, as compared to 80% for mammography, 55% for tomography, and 64% for physical examination. In 70 benign lesions CT was accurate in 84 %, compared with 50% for mammography and 53 % for physical examination. It is ironic that, in spite of these impressive results, this machine may not be commercially developed.

A.S. Head and Neck Tumors

The value of CT scanning in the head and neck has not received the attention given to other sites. In part, this may be due to some technical difficulties in scanning this area. There are rapid variations of density due to both bone and air spaces. The cross section is small, but is only accessible to body scanners that are optimized for large body sections-so detector overranging and spatial resolution can be a problem. Bolus material is


often needed, but may distort anatomy. However, the quality of scans is rapidly improving with the more modern equipment that can scan rapidly, take thin sections, and have an adequate dynamic detector range. With such scanners we can expect to see a much more satisfactory evaluation of the head and neck.

Hodson and co-workers have analyzed the role of computed tomography in the management and treatment of tumors of the paranasal sinuses. (94)

In the evaluation of 38 patients (14 with recurrent disease), CT demonstrated "significantly greater tumor spread [than conventional studies] in nearly half the patients (17/38). The additional information in nearly all cases concerned extension in the posterior and superior directions." CT was considered to be especially valuable in defining the extent of disease to adjacent organs, especially soft tumors of the pterygoid region, parapharyngeal space, and orbit. However, the absorption coefficient value could not distinguish between tumor, benign disease, or fluid. Contrast agents were helpful in the evaluation of intracranial extent. Hodson et al. felt that treatment evaluation after a course of radiation therapy allowed better assessment of the extension of disease into adjacent cavities, since the radiation frequently caused reduction in tumor size, allowing drainage to be reestablished and the distinction made between blocked fluid-filled sinuses and tumor-filled sinuses. Positive neck nodes were observed, but they had all been found clinically. A coronal scan was judged to be necessary in order to evaluate the orbital floor and cribriform plate. CTwas considered to be valuable in assessing the operability of lesions and the extent of surgery that would be required. In planning the radiation therapy of 12 patients, the CT scanning demonstrated inadequate tumor coverage, leading to a change in treatment plan in 3 of 12 (25%) cases.

The value of CT in the assessment of tumors of the brain has been widely accepted. There is an immense literature dealing with this subject, primarily from the point of view of establishing the diagnosis. Perhaps because the role of CT seems self-evident for these tumors, there have been no studies specifically assessing the impact of CT on planning their treatment. Assessments have been made of the use of CT scanning in the follow-up of patients with primary brain tumors, (95-97) and clinical pathologic correlations have been made in some cases.

Mancuso et al. (98) have discussed the role of computed tomography in the assessment of tumors of the parotid gland. They compared CT sialography to conventional sialography and concluded that "CT permitted a better appreciation of the location and direction of growth of tumor in 6 of the 7 patients." Only one of the seven masses proved to be malignant and Mancuso et al. stated that "the single malignant parotid tumor was so obvious on the CT sialogram that a preoperative diagnosis of malignancy could be made with confidence." They also concluded that the CT

282 Michael Goitein

sialogram, performed prior to biopsy, offers great promise in assisting the choice of surgical approach in a situation in which damage to the facial nerves is a major problem. These authors co-opted the patient into the contrast injection process, prevailing upon him to "continue to inject contrast material ... until he actually feels pain; at this point he stops the injection and waves at us to start the scan." No discussion is given of possible alternative reasons for the patient waving.

Miller and Norman(99) have reviewed ten cases in which neck masses were studied with CT as well as with conventional imaging techniques. They concluded that the CT was extremely useful in defining both the osseous and soft tissue extent of the lesion. The relationship of lesions to the spinal canal was established, and in one case the CT scan demonstrated unsuspected thoracic extensior., thus avoiding a possible geographic miss by the radiotherapist.

Mancuso and Hanafee have conducted a prospective study of 66 patients to evaluate the relative merits of CT and laryngography in the assessment of laryngeal disorders. (58) One might think that with the excellent visualization of the larynx that can be obtained upon direct laryngoscopy and with the excellent quality of conventional radiographic techniques, CT would have little application to the evaluation of the larynx. Indeed, this did seem to be the initial impression in the early days of CT. However, Mancuso and Hanafee have concluded that the CT examination has an important role to play in the assessment of the larynx. They found that CT complemented direct laryngoscopy and biopsy for treatment planning and was superior to both modalities for showing deep infiltration, invasion of cartilage, and extension to the soft tissues of the neck. In 13 patients they were able to obtain a radiologic-pathologic correlation between the CT findings and the pathologic .examination of tissue samples. CT was accurate in all 13 of these cases, while laryngography had 1 false positive and 2 false negative findings in the 6 patients who were available for this correlation. Among the cases in which a less precise radiologic-pathologic correlation was available, CT correctly diagnosed 36 of 38 cases of paralaryngeal involvement and was in error in the other 2 cases. Overall, CT was superior in 10 patients and equal to laryngography in 24. In four patients the laryngogram offered more information. Xerography or widelatitude soft tissue techniques were routinely employed early in the study. However, the accuracy of CT was such that these examinations were discontinued for the later patients. Forty-two patients were evaluated for lymph node metastases. Clinically positive nodes were verified by CT in eight patients. In two of these the nodes could not be palpated but were confirmed pathologically. These workers concluded that CT was equal or superior to laryngography on laryngeal function and fine mucosal detail but revealed deep extension into areas that are invisible in direct laryngos-


copy or only indirectly suggested by laryngographic findings. The examination of the larynx is greatly improved with the use of a fast (5 sec or less) CT scan and would be expected to improve with the use of thin (thinner than the 10- to 13-mm-thick sections used in this study) closely spaced sections.

A.6. Other Sites

A.6.1. SARCOMAS

Munzenrider et al. have evaluated the role of CT in the treatment of liposarcomas. Ten patients, seven pretreatment and three posttreatment, were studied. In the seven patients who were scanned prior to therapy, CT provided unique information on at least one treatment margin in five cases (71 %). CT was compared with arteriography for four patients and, in one it was judged that arteriography would have underestimated the superior tumor extension. CT was compared with echography for two patients. Tumor was seen in both cases by both techniques, but echography was judged to underestimate the superior extension in one patient and failed to show the relationship of the bowel and intestine to the lesion in the other. In the follow-up of five patients, CT detected one recurrence and failed to detect a second.

Computed tomography can be expected to have an important role in planning the radiation treatment of soft tissue sarcoma of the extremities. This is because such lesions often occupy a major fraction of the cross section of the extremity. In such a circumstance, successful uncomplicated treatment depends to a large extent on the ability to spare the largest possible amount of the uninvolved tissues. The assessment of the target volume, then, is quite critical and CT seems to have a major role in assisting in this evaluation.

Levine et al. (100) have reported the results of a prospective study of 50 unselected patients with suspected musculoskeletal tumors, consisting of 29 with bone tumors, 15 with soft tissue tumors, and 6 who did not have tumors. (The tumors were ultimately biopsy or surgically proven.) In 13 (26%) of the patients, computed tomography correctly suggested a specific histologic diagnosis or established the presence or absence of a lesion when other imaging techniques gave equivocal results. CT gave a better indication of the tumor extent and its relationships with adjacent tissues in 27 patients (54%). CT had a significant effect on the management of 33 of the patients (66%). In six patients this was because CT was negative, obviating the necessity for biopsy and further management. In the majority of patients, the influence of CT resulted either from delineation

284 Michael Goltein

of a soft tissue mass not shown by other methods or from superior demonstration of the tumor. For some patients a more complete resection of the tumor or a different radiation therapy field became possible as a result of CT information. The patients in this study also had conventional radiographic studies, 99mTc-pyrophosphate bone scans (all patients), 67Ga_ citrate scans (12 patients), sonography (23 patients), and angiography (22 patients). The relative merits of these various imaging modalities were evaluated in detail and an algorithm for the work-up of patients with suspected musculoskeletal tumors was developed. This featured initial evaluation with plain films and mammography followed by CT scanning for primary soft tissue tumors, technetium-99m scanning for malignant bone tumors, and either biopsy or observation for benign bone tumors. CT scanning was advocated for patients for whom a single bone tumor was detected in the technetium scan. These investigators concluded that CT is "very valuable in assessing the extent of muscoloskeletal tumors and often has a profound influence on their management."

A.6.2. TESTICULAR TUMORS

Computed tomography has been assessed in the staging and management of testicular neoplasms. Lee et at. reviewed the CT findings in 26 patients with primary testicular neoplasms. (101) The CT evaluation was correct in 18 of the 20 patients who had either surgical (10 patients), lymphographic (9 patients), or adequate clinical (6 patients) confirmation. The study confirmed that CT is useful and accurate in detecting retroperitoneal and pelvic nodal metastases in patients with testicular tumors. The investigators emphasized that CT "can aid in the planning of radiation therapy portals, and may be used to follow tumor response after therapy."

Husband et aly02) analyzed CT scans in 65 patients with malignant teratoma. They analyzed the role of CT scanning in detecting metastases in the abdomen, mediastinum, and lung and compared CT with lymphography and whole lung tomography. In the lungs they found that CT revealed lung metastases not seen by whole body tomography in one-third of the patients (CT was positive for lung nodules when whole lung tomography was negative in 5 patients and gave a greater number of metastases than whole lung tomography in 15 patients). In the abdomen, they considered that lymphography, while exact, failed to demonstrate the full tumor extent that could be revealed by CT and felt CT and lymphography to be complementary techniques. They stated that CTwas important in demonstrating the size and extent of tumor in the retroperitoneum, which is important for staging, selection and monitoring of therapy, radiation therapy treatment planning, and choice of surgical approach. CT revealed unsuspected lymphadenopathy in 4 of 65 patients. They considered CT to be a more


accurate method of detecting mediastinal tumors. In summary, they considered that CT is "making a significant contribution to the management of malignant teratoma."

Husband et al. (103) analyzed the influence of CT on treatment policy in 100 patients with testicular teratoma. Changes were made in 31 % of these patients. Treatment was withheld for ten patients, radiotherapy initiated for one patient, surgery initiated for one patient and contraindicated for two, and chemotherapy introduced for nine patients, continued for five patients, stopped for one patient, and the regimen changed for two patients. This analysis highlights the important role that CT can play in guiding the management of disease-a role often underestimated in studies directed primarily at the role of CT in radiation therapy treatment planning.

A.6.3. LYMPHOMAS

Redman et al. have considered the role of CT as an adjunct in the staging of Hodgkin's disease and non-Hodgkin's lymphomasy04) They reviewed the role of CT for 25 patients being staged for histologically documented untreated disease. They concluded that "CT scanning cannot completely replace lymphography in the evaluation of lymph nodes. [However], the CT scan often demonstrated a much greater extent of nodal morphology than was expected from the lymphogram."

Lee et al.(S9) have studied 24 patients with Hodgkin's disease and 56 patients with non-Hodgkin's lymphoma. Thirty-one of these had an exploratory laparotomy or an autopsy making a pathologic correlation possible. The CT findings were 90% accurate, with 3 of 31 scans being false negatives. Thirty-seven patients had both CT and lymphangiography but no surgical correlation. Both modalities were in full agreement in 31 cases (84%). Twelve patients had only CT diagnosis but substantial clinical follow-up and in these the CT findings were supported by the clinical evidence. An overall accuracy of 90%, with a true positive rate of 86%, was achieved. The major discrepancy in interpretation occurred when CT scans were made at coarse (4-cm) intervals. This led to an altered strategy in which an initial sequence of coarsely spaced scans was taken. If these scans were positive, scanning was terminated; if negative, additional scans were interlaced with the first series. CT was also advocated for following response to treatment and was considered superior to lymphangiography for this purpose.

Pilepich et al.(1OS) retrospectively reviewed 38 patients with Hodgkin's disease and 59 with non-Hodgkin's lymphoma who had received CT scans. They obtained valuable information from abdominal scans, but they particularly emphasized the role of CT in evaluating anterior mediastinal masses. Of 28 patients with such masses, 15 had extension along the anterior

286 Michael Goitein

chest wall, and appreciation of this resulted in significant alteration of treatment technique in 9 patients (32%).

Alcorn et aly06) studied 33 patients with malignant lymphoma, 16 of whom had stage I or II disease and underwent staging laparotomy. For these patients, for whom a CT -pathologic correlation was possible, there was complete agreement in nine cases. In three patients the CT did not demonstrate the abnormality of the spleen found in laparotomy. The CT scan missed two cases with lymph node involvement and was false positive for abnormal nodes in three cases. Correlations were also obtained with 67 Ga-citrate scans. CT was judged to be sufficiently reliable to replace lymphangiography and 67 Ga imaging in patients with advanced disease.

Schaner et al. (107) evaluated 30 patients with histologically proven lymphoma within one week of bipedal lymphangiography to assess the relative sensitivities of CT and lymphangiography in defining the location and extent of lymphomatous involvement. They concluded that the CT scan was superior in its ability to define the size and position of involved nodes. In particular, lymphangiography underestimated the superior extent of nodal involvement in 13 patients. It was unable to demonstrate tumor extension into the renal and/or splenic hila in 11 patients and invasion into the root of the mesentery in 3. However, lymphangiograms were positive in two patients for whom the CT scans were interpreted as negative. These investigators also examined 26 patients with non lymphomatous intra-abdominal malignancy. They found CT to be of value in eight patients with either seminoma or teratocarcinoma of the testis. Five of these patients had gross periaortic adenopathy, which was significantly underestimated by the lymphangiogram but demonstrated by CT. Ten patients with ovarian carcinoma were studied, but the CT findings were not helpful, even though two patients had positive nodes upon lymphangiography. None of the eight patients with carcinoma of the pancreas showed abnormalities on either CT scanning or lymphangiography.

Breiman et aly08) have reviewed CT scans in 27 patients with Hodgkin's disease or non-Hodgkin's lymphomas and were able to get pathologic correlations in 18 cases. Their results are reproduced in Table 14. The accuracy of CT is seen to be 91 % in 68 areas of interest.

Desforges et al., in a recent review,(109) considered the role of CT in the staging and management of Hodgkin's disease and stated that

CT scanning, in addition to detecting tumor in areas difficult to assess by other noninvasive tests, has special value for the radiotherapists in giving cross-sectional dimensions of a mediastinal tumor, allowing for accurate mapping of treatment fields.


Table 14. Correlation of CT Interpretation with Staging Laparotomy a

Region

Para-aortic lymph nodes

Mesenteric lymph nodes

Spleen Liver

"From Breiman et al. in ref. 108.

CT positive

Pathology Pathology positive negative

6 2

2 10 0

1 0

CT negative

Pathology Pathology Not positive negative biopsied

0 8 2

12 2 7 0

16 0

There seems to be broad agreement among these studies about the role of CT in the evaluation of the patient with lymphoma. CT cannot, it seems, replace lymphangiography, since the evaluation of small nodes with abnormal architecture is not feasible with CT scanning and is the forte of lymphangiography. However, the evaluation of nodes not normally or reliably opacified is a complementary strength of CT scanning. Moreover, there seems to be a consensus that the extension of nodal disease is often significantly underestimated by lymphangiography and can be assessed well with CT.

REFERENCES

1. J. E. Munzenrider, M. Pilepich, J. B. Rene-Ferrero, I. Tchakarova, and B. L. Carter, Use of body scanner in radiotherapy treatment planning, Cancer 40, 170-179 (1977).

2. M. Goitein, J. Wittenberg, M. Mendiondo, J. Doucette, C. Friedberg, J. Ferrucci, L. Gunderson, R. Linggod, W. U. Shipley, and H. V. Fineberg, The value of CT scanning in radiation therapy treatment planning: a prospective study, Int. J. Radiat. Oneol. Bioi. Phys. 5, 1787-1798 (1979).

3. P. Hobday, N. J. Hodson, J. Husband, R. P. Parker, and J. S. Macdonald, Computed tomography applied to radiotherapy treatment planing: techniques and results, Radiology 133,477-482 (1979).

4. D. P. Ragan and C. A. Perez, Efficacy of CT-assisted two-dimensional treatment planning: analysis of 45 patients, Am. J. Roentgenol. 131,75-79 (1978).

5. B. A. Schlager, S. O. Asbell, A. Baker, S. G. Mann, and D. M. Sklaroff, The application of computerized tomography in radiation treatment planning, private communication (1979).

6. H. L. Abrams and B. J. McNeil, Medical implications of computed tomography ("CAT scanning"), New Eng. J. Med. 298, 255-261 (1978).

7. H. L. Abrams and B. J. McNeil, Medical implications of computed tomography ("CAT scanning"), New Eng. J. Med. 298, 310-318 (1978).

288 Michael Goitein

8. S. W. Yu, R. H. Sagerman, G. A. King, C. T. Chung, and T. W. Yu, The value of computed tomography in the management of bladder cancer, Int. l. Radiat. Oncol. Bioi. Phys. S, 135-142 (1979).

9. N. J. Hodson, J. E. Husband, and P. A. Hobday, The role of CT in the management and therapy planning of carcinoma of the bladder, in Proceedings of the European Association of Radiology Workshop an the Use of Computerized Tomographic Scanners in Radiotherapy in Europe, 28-30 March 1979, Geneva, British Institute of Radiology Special Series, Supplement IS, 78-80 (1981).

10. R. Howells, J. W. Walsh, T. A. Hazra, M. Amendola, and R. Be1grad, Assistance of computerized tomography in the optimization of the management (diagnostic and therapy) of patients with cervical cancer, in Proceedings of the European Association of Radiology Workshop on the Use of Computerized Tomographic Scanners in Radiotherapy in Europe, 28-30 March 1979, Geneva, British Institute of Radiology Special Series, Supplement 15,62-64 (1981).

11. J. E. Husband, N. J. Hodson, and C. A. Parsons, Use of computed tomography in recurrent rectal tumors, Radiology, 134, 677-682 (1980).

12. H. E. Brizel, P. A. Livingston, and E. V. Grayson, Radiotherapeutic applications of pelvic computed tomography, l. Comput. Assist. Tomogr. 4, 453-466 (1979).

13. J. K. T. Lee, R. G. Levitt, R. J. Stanley, and S. S. Sagel, Utility of body computed tomography in the clinical follow-up of abdominal masses, l. Comput. Assist. Tomogr. 2,607-611 (1978).

14. M. Goitein, The utility of CT in radiation therapy: an estimate of outcome, Int. l. Radiat. Oncol. Bioi. Phys. S, 1799-1807, 1979.

15. M. Goitein, Benefits and cost of computed tomography in radiation therapy, lAMA, 244,1347-1350 (1980).

16. R. G. Evens and R. G. Jost, Utilization of body computed tomography units, Radiology 131,695-698 (1979).

17. J. R. Stewart, J. A. Hicks, M. L. M. Boone, and L. D. Simpson, Computed tomography in radiation therapy, Int. l. Radiat. Oncol. Bioi. Phys. 4, 313-324 (1978).

18. Committee for Radiation Oncology Studies (CROS), Proposal for a Program in Particle Beam Radiation Therapy in the United States, Cancer c/in. Trials I, 153-208 (1978).

19. M. Goitein, Mattress for treatment couch, Radiologic Technology, to be published. 20. J. J. Battista, W. D. Rider, and J. Van Dyk, Computed tomography for radiotherapy

planning, Int. l. Radiat. Oncol. Bioi. Phys., 6, 99-108 (1980). 21. M. R. Sontag and J. R. Cunningham, Clinical application of a CT-based treatment

planning system, Comput. Tomogr. 2, 117-130 (1978). 22. L. L. Gunderson, Radiation Oncology and New Modalities, in Alimentary Tract

Radiology, Vol. III (A. R. Margulis and H. J. Burhenne, eds., pp. 593-619, C. V. Mosby, St. Louis (1979).

23. E. C. McCullough, Computed tomography in radiation therapy treatment planning, in RadialogyoftheSkullandBrain, Vol. 5 (T. H. NewtonandG. Potts,eds.), C. V. Mosby,St. Louis (1971).

24. G. L. Jelden, E. S. Chernak, A. Rodriguez-Antunez, J. R. Haaga, P. S. Lavik, and R. S. Dhaliwal, Further progress in CT scanning and computerized radiation therapy treatment planning, Am. l. Roentgenol. 127, 179-185 (1976).

25. G. D. Fullerton, W. Sewchand, J. T. Payne, and S. H. Levitt, CT determination of parameters for inhomogeneity corrections in radiation therapy of the esophagus, Radiology 126, 167-171 (1978).

26. J. J. Battista, J. Van Dyk, and W. D. Rider, CT-aided radiotherapy planning with emphasis on dose calculations in three dimensions, in Proceedings of the European Association of Radiology Workshop on the Use of Computerized Tomographic Scanners in Radiotherapy in Europe, 28-30 March 1979, Geneva, British Institute of Radiology Special Series, Supplement IS, 147-151 (1981).


27. M. R. Sontag, J. J. Battista, M. J. Bronskill, and J. R. Cunningham, Implications of computed tomography for inhomogeneity corrections in photon beam dose calculations, Radiology 124, 143-149 (1977).

28. H. E. Johns and J. R. Cunningham, The Physics of Radiology, 3rd Ed., p. 446, Charles C. Thomas, Springfield, Ill. (1969).

29. M. Goitein, Three dimensional density reconstruction from a series of two dimensional projections, Nucl. Instrum. Methods 101,509-518 (1972).

30. R. A. Geise and E. C. McCullough, The use of CT scanners in megavoltage photon beam therapy planning, Radiology 124, 133-141 (1977).

31. D. Chesler, S. J. Riederer, and N. J. Pelc, Noise due to photon counting statistics in computed x-ray tomography, J. Comput. Assist. Tomogr. 1, 64-74 (1977).

32. R. E. Alvarez and A. Macovski, Energy-selective reconstructions in x-ray computerized tomography, Phys. Med. Bioi. 21, 733-744 (1976).

33. P. J. Joseph and R. D. Spital, A method for correcting bone induced artifacts in computed tomography scanners, J. Comput. Assist. Tomogr. 2, 100-108 (1978).

34. O. Nalcioglu and R. V. Lou, Post-reconstruction method for beam hardening in computerized tomography, Phys. Med. BioI. 24, 330-340 (1979).

35. P. J. Robinson and L. Kreel, Pulmonary tissue attenuation with computed tomography: comparison of inspiration and expiration scans, J. Comput. Assist. Tomogr. 3, 740-748 (1979).

36. P. K. Kijewski and B. E. Bjarngard, The use of computed tomography data for radiotherapy dose calculations, Int. J. Radiat. Oncol. Bioi. Phys. 4, 429-435 (1978).

37. R. P. Parker, P. A. Hobday, and K. J. Cassell, The direct use of CT numbers in radiotherapy dosage calculations for inhomogeneous media, Phys. Med. Bioi. 24, 802-809 (1979).

38. S. C. Prasad, G. P. Glasgow, and J. A. Purdy, Dosimetric evaluation of a computed tomography treatment system, Radiology 130,777-781 (1979).

39. ICRU, Report 24, Determination of Absorbed Dose in a Patient Irradiated by Beams of X or Gamma Rays in Radiotherapy Procedures, International Commission on Radiation Units and Measurements, Washington, D.C. (1976).

40. E. C. McCullough, Potentials of computed tomography in radiation therapy treatment planning, Radiology 129, 765-768 (1978).

41. M. R. Sontag and J. R. Cunningham, The equivalent tissue-air ratio method for making absorbed dose calculations in a heterogeneous medium, Radiology 129, 787-794 (1978).

42. M. R. Sontag and J. R. Cunningham, Corrections to absorbed dose calculations for tissue inhomogeneities, Med. Phys. 4,431-436 (1977).

43. S. Webb and R. P. Parker, A Monte Carlo study of the interaction of external beam x-radiation with inhomogeneous media, Phys. Med. Bioi. 23, 1043-1059 (1978).

44. S. Webb and R. A. Fox, The direct use of CT data for inhomogeneity corrections in radiotherapy planning, in Proceedings of the European Association of Radiology Work-28-30 March 1979, Geneva, British Institute of Radiology Special Series, Supplement 15,98-102 (1981).

45. D. L. McShan, A. Silverman, D. M. Lanza, L. E. Reinstein, and A. S. Glicksman, A computerized three-dimensional treatment planning system utilizing interactive colour graphics, British Journal of Radiology 52, 478-481 (1979).

46. G. T. Y. Chen, R. P. Singh, J. R. Castro, J. T. Lyman, and J. M. Quivey, Treatment planning for heavy ion radiotherapy, Int. J. Radiat. Oncol. Bioi. Phys. 5, 1809-1819 (1979).

47. M. Goitein, The measurement of tissue heterodensity to guide charged particle radiotherapy, Int. J. Radiat. Oncol. Bioi. Phys. 3, 27-33 (1977).

48. M. Goitein, A technique for calculating the influence of thin inhomogeneities on charged particle beams, Med. Phys. 5, 258-264 (1978).

290 Michael Goitein

49. M. Goitein, G. T. Y. Chen, J. Y. Ting, R. J. Schneider, andJ. M. Sisterson, Measurements and calculations of the influence of thin inhomogeneities on charged particle beams, Med. Phys. 4, 265-273 (1978).

50. M. Goitein and J. M. Sisterson, The influence of thick inhomogeneities on charged particle beams, Radiat. Res. 74, 217-230 (1978).

51. M. Goitein, Compensation for inhomogeneities in charged particle radiotherapy using computed tomography, Int. J. Radiat. Oneol. Bioi. Phys. 4, 499-508 (1978).

52. K. R. Hogstrom, A. R. Smith, S. L. Simon, J. W. Somers, R. G. Lane, I. I. Rosen, C. A. Kelsey, C. F. Von Essen, M. M. Kligerman, P. A. Berardo, and S. M. Zink, Static pion beam treatment planning of deep seated tumors using computerized tomographic scans, Int. J. Radiat. Oneol. Bioi. Phys. 5, 875-886 (1979).

53. R. A. Rutherford, B. R. Pullan, and I. Isherwood, Measurement of effective atomic number and electron density using an EMI scanner, Neuroradiology 11, 15-21 (1976).

54. G. DiChiro, R. A. Brooks, R. M. Kessler, G. S. Johnston, A. E. Jones, J. R. Herdt, and, W. T. Sheridan, Tissue signatures with dual-energy computed tomography, Radiology 131,521-523 (1979).

55. R. E. Latchaw, J. T. Payne, and L. H. A. Gold, Effective atomic number and electron density as measured with a computed tomography scanner: computation and correlation with brain tumor histology, J. Comput. Assist. Tomogr. 2, 199-208 (1978).

56. R. T. Ritchings, B. R. Pullan, S. B. Lucas, R. A. Fawcitt, J. J. K. Best, I. Isherwood, and A. I. Morris, An analysis of the spatial distribution of attenuation values in computed tomographic scans of liver and spleen, J. Comput. Assist. Tomogr. 3, 36-39 (1979).

57. G. H. Underwood, R. G. Hooper, S. P. Axelbaum, and D. W. Goodwin, Computed tomographic scanning of the thorax in the staging of bronchogenic carcinoma, New Eng. J. Med. 300, 777-778 (1978).

58. A. A. Mancuso and W. N. Hanafee, A comparative evaluation of computed tomography and laryngography, Radiology 133, 131-138 (1979).

59. J. K. T. Lee, R. J. Stanley, S. S. Sagel, and R. G. Levitt, Accuracy of computed tomography in detecting intraabdominal and pelvic adenopathy in lymphoma, Am. J. Roentgenol. 131, 311-315 (1978).

60. M. Goitein, Computed tomography in planning radiation therapy (editorial), Int. J. Radiat. Oneol. BioI. Phys. 5, 445-447 (1979).

61. P. F. Judy, The line spread function and modulation transfer function of a computed tomographic scanner, Med. Phys. 3, 233-236 (1976).

62. AAPM Task Force on CT Scanner Phantoms, Report Number 1: Phantoms for Performance Evaluation andOuality AssuraneeofCTSeanners, AAPM, Chicago, III. (1977).

63. G. Cohen and F. A. DiBianca, The use of contrast-detail-dose evaluation of image quality in a computed tomographic scanner, J. Comput. Assist. Tomogr. 3, 189-195 (1979).

64. R. Tadmor and P. F. J. New, Computed tomography of the orbit with special emphasis on coronal sections. II. Pathological anatomy, J. Comput. Assist. Tomogr. 2, 34-44 (1978).

65. W. D. Foley, T. L. Lawson, and F. Quiroz, Sagittal and coronal image reconstruction: application in pancreatic computed tomography, J. Comput. Assist. Tomogr. 3, 717-721 (1979).

66. W. V. Glenn, M. L. Rhodes, W. M. Altschuler, L. L. Wiltse, C. Kostanek, and Y. M. Kuo, Multiplanar display computerized body tomography applications in the lumbar spine, Spine 4, 262-352 (1979).

67. R. N. Abel, L. A. Buhler, W. T. Chu, J. M. Slater, I. R. Neilsen, and D. C. Zimmerman, Direct data link between a CT scanner and a radiation treatment planning system, Med. Phys. 6, 452-453 (1979).


68. R. J. Alfidi, A. R. Antunez, and J. R. Haaga, Computed body tomography and cancer (editorial), Am. J. Roentgenol. 127,1061 (1976).

69. D. Boyd, J. Coonrod, J. Dehnert, D. Chu, C. Lim, B. Macdonald, and V. Perez-Mendez, A high pressure xenon proportional chamber for x-ray laminographic reconstruction using fan beam geometry, IEEE Trans. Nucl. Sci. 21, 184-187 (1973).

70. R. M. Harrison and F. T. Farmer, The determination of anatomical cross sections using a radiotherapy simulator, Br. J. Radiol. 51, 448-453 (1978).

71. S. Webb, S. C. Lillicrap, H. Steere, and R. D. Speller, Application of a radiotherapy simulator for imaging of body cross sections, Br. J. Radiol. 50, 152-153 (1977).

72. G. A. Thieme, W. R. Hendee, G. S. Ibbott, P. L. Carson, and D. L. Kirsh, Cross-sectional anatomic images by gamma ray transmission scanning, Acta Radiol.14, 81-112 (1975).

73. J. M. Galvin, B. E. Bjarngard, and A. Cormack, In vivo determination of attenuation coefficients for radiation therapy, Proc. Soc. Photo-Opt. Instrum. Eng. 47, 27-39 (1975).

74. M. A. Bagshaw, G. R. Ray, and D. A. Pistenma, External beam radiation therapy of primary carcinoma of the prostate, Cancer 36,723-729 (1975).

75. M. T. Kademian, W. L. Caldwell, J. A. Zagzebski, and J. Glassner, Prostate boost field localization with computerized tomographic scanning, in Proceedings of the European Association of Radiology Workshop on the Use of Computerized Tomographic Scanners in Radiotherapy in Europe, 28-30 March 1979, Geneva, British 'Institute of Radiology Special Series, Supplement 15,71-74 (1981).

76. M. V. Pilepich, C. A. Perez, and S. Prasad, Computed tomography in definitive radiotherapy of prostatic carcinoma, Int. J. Radiat. Oncol. Bioi. Phys. 6, 923-926 (1980).

77. W. J. Hoogenraad and C. J. G. Bakker, CT scanning in radiation therapy of the prostate, in Proceedings of the European Association of Radiology Workshop on the use of Computerized Tomographic Scanners in Radiotherapy in Europe, 28-30 March 1979, Geneva, British Institute of Radiology Special Series, Supplement 15,65-67 (1981).

78. F. R. Paquette, A. S. Ahuja, P. L. Carson, L. A. Mack, G. S. Ibbott, and M. L. Johnson, A comparative study of computerized tomography and ultrasound imaging for treatment planning of prostatic carcinoma, Int. J. Radiat. Oncol. Bioi. Phys. 5, 289-294 (1979).

79. B. Schlager, S. O. Asbell, A. S. Baker, D. M. Sklarofi, H. G. Seydel, and B. J. Ostrum, The use of computerized tomography scanning in treatment planning for bladder carcinoma, Int. J. Radiat. Oncol. Bioi. Phys. 5, 99-103 (1979).

80. N. J. Hodson, J. E. Husband, and J. S. Macdonald, The role of computed tomography in the staging of bladder cancer, Clin. Radiol. 30, 389-395 (1979).

81. J. E. Husband and N. J. Hodson, Computerized axial tomography for staging and assessing response to treatment, private communication, 1979.

82. F. E. Seidelmann, N. E. Reisch, W. N. Cohen, J. R. Haaga, P. J. Bryan, and T. R. Havrilla, Computed tomography of the seminal vesicles and seminal vesicle angle, Comput. Axial Tomogr. 1, 281-284 (1977).

83. F. E. Seidelmann, W. N. Cohen, P. J. Bryan, S. P. Temes, D. Kraus, and G. Schoenrock, Accuracy of CT staging of bladder neoplasms using the gas-filled method: report of twenty-one patients with surgical confirmation, Am. J. Roentgenol. 130, 735-739 (1978).

84. M. J. Kellett, R. T. D. Oliver, J. E. Husband, C. D. Kelsey, and I. Fry, CT scanning as an adjunct to bimanual examination for staging bladder tumors, Br. J. Urology, 52 (2), 101-106, April 1980.

85. J. Ammon, W. Frik, J. H. Karstens, H. Rubben, and J. Schofiers, Total body computed tomography of the urogenital system, Urol. Ausg. A 18, 1-13 (1979).

86. H. C. Yeh, Ultrasonography and computed tomography of carcinoma of the gall bladder, Radiology 133, 167-173 (1979).

292 Michael Goitein

87. K. R. Lee, E. Levine, R. E. Moffat, L. R. Bigongiari, and A. S. Hermreck, Computed tomographic staging of malignant gastric neoplasms, Radiology 133, 151-155 (1979).

88. B. Emami, A. Melo, B. L. Carter, J. E. Munzenrider, and A. J. Piro, Value of computed tomography in radiotherapy of lung. cancer, Am. 1. Roentgenol. 131, 63-67 (1978).

89. F. Casamassima, L. Forzini, and L. Cionini, The CT scanner in the study of the mediastinal localization of Hodgkin's disease, in Proceedings of the European Association of Radiology Workshop on the Use of Computerized Tomographic Scanners in Radiotherapy in Europe, 28-30 March 1979, Geneva, British Institute of Radiology Special Series, Supplement 15, 37-41 (1981).

90. R. A. Mintzer, S. R. Malave, H. L. Neiman, L. L. Michaelis, R. M. Vanecko, and J. H. Sanders, Computed vs. conventional tomography in evaluation of primary and secondary pulmonary neoplasms, Radiology 132, 653-659 (1979).

91. J. E. Munzenrider, I. Tchakarova, M. Castro, and B. Carter, Computerized body tomography in breast cancer. I. Internal mammary nodes and radiation treatment planning, Cancer 43, 137-150 (1979).

92. J. E. Munzenrider, M. Castro, and I. Tchakarova, Computerized body tomography in breast cancer. III. Internal mammary dosimetIy, private communication.

93. C. H. J. Chang, J. L. Sibala, S. L. Fritz, S. J. Dwyer III, and A. W. Templeton, Specific value of computed tomographic breast scanner (CT/M) in the diagnosis of breast diseases, Radiology 132,647-652 (1979).

94. N. J. Hodson, C. A. Parsons, and P. A. Hobday, The role of CT in the management and therapy planning of tumors of the paranasal sinuses, in Proceedings of the European Association of Radiology Workshop on the Use of Computerized Tomographic Scanners in Radiotherapy in Europe, 28-30 March 1979, Geneva, British Institute of Radiology Special Series, Supplement 15,29-31 (1981).

95. J. E. Marks and M. Gado, Serial computed tomography of primary brain tumors following surgery, irradiation and chemotherapy, Radiology 125, 119-125 (1977).

96. M. A. Mikhael, Radiation necrosis of the brain: correlation between computed tomography, pathology and dose distribution, 1. Comput. Assist. Tomogr. 2, 71-80 (1978).

97. R. J. Carella, N. Pay, J. Newall, A. T. Farina, I. I. Kricheff, and J. S. Cooper, Computerized (axial) tomography in the serial study of cerebral tumors treated by radiation (a preliminary report), Cancer 37, 2719-2728 (1976).

98. A. Mancuso, D. Rice, and W. Hanafee, Computed tomography of the parotid gland during contrast sialography, Radiology 132,211-213 (1979).

99. E. M. Miller and D. Norman, The role of computed tomography in the evaluation of neck masses, Radiology 133, 145-149 (1979).

100. E. Levine, K. R. Lee, F. R. Neff, N. F. Maklad, R. G. Robinson, and D. F. Preston, Comparison of computed tomography and other imaging modalities in the evaluation of musculoskeletal tumors, Radiology 131,431-437 (1979).

101. J. K. T. Lee, B. L. McClennan, R. J. Stanley, and S. S. Sagel, Computed tomography in the staging of testicular neoplasms, Radiology 130, 387-390 (1979).

102. J. E. Husband, M. J. Peckham, J. S. Macdonald, and W. F. Hendry, The role of computed tomography in the management of testicular teratoma, Clin. Radiol. 30, 243-252 (1979).

103. J. E. Husband, A. Barrett, and M. J. Peckham, Evaluation of computed tomography in the management of testicular teratoma, private communication, 1979.

104. H. C. Redman, E. Glatstein, R. A. CasteIIino, and W. A. Federal, Computed tomography as an adjunct in the staging of Hodgkin's disease and non-Hodgkin's lymphomas, Radiology 124,381-385 (1977).

105. M. V. Pilepich, J. B. Rene, J. E. Munzenrider, and B. L. Carter, Contribution of computed tomography to the treatment of lymphomas, Am. 1. Roentgenol. 131,69-73 (1978).

Applieations of CT in Radiotherapy Treatment Planning 293

106. F. S. Alcorn, V. C. Mategrano, J. P. Petasnick, and J. W. Clark, Contributions of computed tomography in the staging and management of malignant lymphoma, Radiology 125, 717-723 (1977).

107. E. G. Schaner, G. L. Head, J. L. Doppman, and R. C. Young, Computed tomography in the diagnosis, staging and management of abdominal lymphoma, J. Comput. Assist. Tomogr. 1, 176-180 (1977).

108. R. S. Breiman, R. A. Castellino, G. S. Harell, W. H. Marshall, E. G1atstein, and H. S. Kaplan, CT -pathologic correlations in Hodgkin's disease and non-Hodgkin's lymphoma, Radiology 126, 159-166 (1978).

109. J. F. Desforges, C. J. Rutherford, and A. Piro, Hodgkin's disease, New Eng. J. Med. 30,1212-1222 (1979).

5

Positron Imaging: Some Practical Considerations

ROBERT J. NICKLES and S. JOHN GATLEY

1. INTRODUCTION

1.1. Scope of the Problem

Nuclear imaging techniques seek to unravel the pharmacokinetics governing the distribution of labeled compounds. These images detail how the body operates on the radiotracer and hold the promise of distinguishing the diseased state in a noninvasive clinical test. At the same time, these procedures offer the basic scientist a unique opportunity to study physiological phenomena in the intact species. This duality of diagnostic utility and research potential tends to stretch positron tomography today, as it will do to nuclear magnetic resonance (NMR) imaging tomorrow. What is being displayed for sectional examination is something more subtle than radiographic opacity and requires a period of fundamental study to precede the establishment of clinical protocols. Both clinical and basic nuclear imaging studies, however, share the goal of determining organ function, revealed in the dynamics of tracer localization.

Living organisms are composed of compounds whose major elemental constituents are hydrogen, carbon, nitrogen, and oxygen. In addition, halogenated analogs of some physiological substrates perform creditably through the initial stages of a metabolic process, but steric effects and

ROBERT J. NICKLES and S. JOHN GATLEY • Medical Physics Section, Department of Radiology, 3321 Sterling Hall, 475 N. Charter St., University of Wisconsin, Madison, Wisconsin 53706.

295

296 Robert J. Nickles and S. John Gatley

Table 1. Radioisotopes of Hydrogen, Carbon, Nitrogen, Oxygen, and Fluorine Suitable as Biological Tracers

Representative Decay mode Typical production biological

RadionucIide tl/2 {3± (Ey keY) reaction reference

3H 12 yr (3- 6Li(n, a) 3H Many lle 20 min (3+ 14N(p, a) lle 1 14e 5700 yr {3- 14N (n, p) 14e Many 13N 10 min (3+ 160 (p, a) 13N 2 140 71 sec (3+(2314) 14N(p, n) 140 3 ISO 123 sec (3+ 14N(d, n) ISO 4 190 29 sec (3-(198, 0 0 0) 180(d, p) 190 5 17F 66 sec (3+ 160 (d, n) 17F 6 18F 110 min (3+ 2oNe(d, a) 18F 7

decreasing binding energies generally restrict this substituent to fluorine. Table 1 lists a few unstable isotopes of these five elements that have been employed as physiological tracers, along with their decay characteristics and a typical nuclear reaction leading to each radionuclide; these are the candidates, the starting material for the synthesis of a labeled compound capable of tracing the pathways of the authentic counterpart. The list breaks down into essentially two classes: long-lived 3H and 14C suitable for invasive studies and short-lived positron emitters permitting activity measurements with external detectors sensitive to penetrating annihilation radiation. These natural circumstances force us to consider positron imaging if we hope to trace most compounds through an intact, living system.

1.2. Linkage

The triangular relationship(8) linking the radiopharmaceutical, the instrument, and the biomedical problem pervades the discipline of nuclear medicine. This interplay becomes even stronger in positron imaging procedures. Many of the excellent reviews of the field(9-11) have examined the practical consequences that unfold from the simple physical characteristics of radiation as sketched in Table 2.

In recent years, positron imaging has entered clinical service, and a division of attitude has become apparent between the groups that strive for optimal image quality ("pretty pictures") and those who regard the image merely as an intermediate in the quantitation of regional metabolic rate constants. These divergent attitudes often mark the difference between the clinician relying on experience to visually identify the fingerprint of organ disfunction and the investigator seeking a more rigorous relationship,

Positron Imaging: Some Practical Considerations 297

such as the effect of therapy on tracer uptake. The prevailing sentiment will influence what type of instrument evolves, what labeled agents are developed, and what studies are performed with the full tomographic capability available.

1.3. Historical Perspective

Positron coincidence localization predates(12.13) the 99mTc-Anger camera coalition(14.15) that has dominated nuclear medicine for the past decade. The increasing availability of both positron-emitting tracers and minicomputers capable of image reconstruction has revived interest in positron tomography. Most imaging instruments have developed along one of two distinct directions: (a) laminographic systems based on two opposed planar detectors(16-21) or (b) transverse systems employing a closed ring array of detectors. (22-31) This original dichotomy is simplistically represented in Figure 1, which is an attempt to group the major imaging instruments by their functional similarities. In fact, this initial orientation is purely historical, as planar devices now rotate and ring arrays are made to scan in an effort to offer both transverse section and area imaging from the same positron camera. (9) The impact of transmission CT scanning on radiography has been so great, however, that the term "positron tomography" has become understood to mean transverse section imaging, with the

Table 2. Nuclear Imaging Consequences of the Physical Decay Properties of Positron Emitters

Physical characteristics

1. Short lifetime

2. Energetic radiations

3. Two-photon decay

Consequences

Primary

Couples production and use

Rapid decay

Fast syntheses

High specific activity

High dose to workers

Penetrates detector shielding

Coincidence imaging opens tomography

Secondary

Shared facilities, tight logistics

Repeatable studies, low patient dose

Limits complexity of labeled compounds

Ideal tracers for receptor sites

Remote handling, massive shielding

Heavy collimation, limited resolution

Leads to quantitative in vivo autoradiography, improved diagnosis, radiopharmaceuticals


LONGITUDINAL TRANSVERSE

STATIONARY

CONTINUOUS

No!

16,17 18,19 20,21 22-25 26

Figure 1. A simplistic representation of the decisions facing the radionuclide imaging clinician intending to enter the field of positron emission tomography. The references are intended to be neither exclusive nor exhaustive.

term "laminography" being reserved for the focused imaging of longitudinal planes.

The static form of Figure 1 tends to hide the impressive advances that have been made in the 20-year history of positron tomography. Only recently, data rate limitations have been lifted to allow image data to be acquired in seconds, with reconstruction times reduced to a similar overhead.(31) Spatial resolution has been pushed to roughly 1 cm (full width at half maximum) in most systems, capable of revealing most functional defects in the high-contrast sectional studies. With typical fields of view of SO cm, body sections can be imaged with uniformity and known sensitivity. (32) Finally, the advent of multislice machines marks a major advance in the effort to maintain constant physiology throughout a study through the simultaneous acquisition of image data.

1.4. A Modest Alternative

Several of the prototype positron cameras have been engineered into costly commercial devices. (33) This, in turn, fuels the cost-benefit polemic, and the polarized public attitude is not unexpected. In fact, doing fast


coincidence is an instrumental challenge; scaling it up to large detector arrays is necessarily a high-stakes venture. Include the compact cyclotron facility with a shielded laboratory, and the hospital administrator is faced with a several million dollar capital expenditure. On the other hand, dependence on extramural support further alienates the clinician and tends to make it so that emission tomography will always be the plaything of a few major centers of excellence.

Against this backdrop, we offer a modest alternative based on two premises:

1. That the availability of positron-emitting precursors can be broadened through access to nearby research accelerators, reactors, and generator systems.

2. That low-cost laminographic imaging is possible through a modular approach, building on the existing in-house equipment.

2. POSITRON IMAGING

2.1. Physical Background

2.1.1. ORIGIN OF THE RADIATION

A proton-rich nuclide will decay to its neighboring stable isobar. Positron emission becomes allowed as the available energy exceeds the 2mc2 = 1.022 MeV rest mass deficit. At higher energies, positron decay competes even more favorably against K capture owing to the greater slice of phase space available to its three-body final state. This breakup insures a continuous distribution of positron energies, with average kinetic energies considerably below 1 Me V in radio nuclides of medical interest. These positrons are slowed to thermal energies on a picosecond time scale after passing through millimeter-long path lengths of condensed material. The most probable fate for the thermalized positron that of joining an atomic electron to form positronium. Radiative transitions to the hydro genic ground state (L = 0) are rapid, leaving the positronium in either the singlet (antiparallel) or triplet (parallel) spin state. Their respective lifetimes (.25 and 150 nsec) and decay modes (2y and 3y) facilitate identification. Further, the quenching of the triplet state acts as a unique probe of the chemical properties of the stopping substrate. (34) This quenching assures that the great majority of positrons will decay with the emission of two colinear photons, each with the characteristic energy of the electron rest mass mc 2 = 511 keV.


2.1.2. CONSEQUENCES FOR IMAGING POSITRON DISTRIBUTIONS IN VIVO

2.1.2.1. Single-Photon Detection. The 511-keV photon energy of positron annihilation reminds the nuclear medical clinician that his imaging instruments have been increasingly tailored to the 140-keV emission of 99mTc. A superficial comparison of attenuation coefficients suggests that at the higher energy, the photon will suffer less attenuation in the patient but will require thicker collimator septums and NaI detectors for equivalent full-energy peak efficiency. This is verified in phantom studies, where both spatial resolutions and sensitivities are disappointing. In fact, these phantom studies do not prepare one for the simple fact that 511-keV photons make very respectable images in the common clinical situation, where a lowcontrast radiotracer distribution usually underlies a substantial layer of scattering tissue.

The improved rejection of Compton-scattered events is the result of the improved fractional energy resolution and the decreased scattering cross section as the incident energy increases. The scattering angle (J

depends on the scattered energy E as

mc 2 Eo -E 1 cos (J = 1---- = 1--(Eo-E) (1)

Eo E aE

where a = Eo/mc 2 measures the initial energy Eo in units of the electron rest mass. The lower-energy threshold E' is typically set at the lower edge of the full energy peak, which determines the smallest scattering angle rejected by the pulse height analysis. Figure 2 shows this relationship with the half-angle (J defining the system cone of blindness, tightened up by the

T

~~------~

99mTc

~ Eo #~ OE.~E. NlEt ....... ) ... ~M \

~E Eo Ey •

140 511keV

15"10 8"10

450 160

Figure 2. A diagrammatic representation of a Compton scattering event through an angle 8, resulting in an energy degradation Eo - Es barely perceptible to a gamma spectrometer of energy resolution jJ..Eo/ Eo. As the fractional energy resolution improves, the cone of uncertainty rapidly closes.


improved fractional energy resolution. To estimate the effect of this restricted angular acceptance, we employ the differential collision cross section dO' for the scattering(3S) of photons into the incremental solid angle d n, related to Eo, E, e, and the classical electron radius ro:

(2)

Integration of this collision cross section between the limits of Eo (no scatter) and the lower threshold E' gives the scattering probability P in terms of the fractional scattered photon energy y == E' / Eo as

P = P(y, a) = r~ [(~ _ 1) In(1 _ y) + Y _ y2 + -;(_1_ - 1)J a a 2 a 1-y

(3)

Inspection of P(y, a) emphasizes two points:

1. A Ge(Li) detector with an energy threshold set 1 ke V below the 99mTc photopeak barely equals the scatter rejection available to NaI at 511 keY.

2. The temptation is strong to lower the threshold E' to recapture those events lost through Compton scatter in the detector. The excellent scatter rejection achieved with 511-ke V photopeak counting is quickly sacrificed by this expedient.

In summary, single-photon imaging of positron emitters with heavily collimated gamma cameras can be surprisingly good in the clinical situation. The absence of scattered radiation in the narrow energy window removes low-frequency noise from the image, although the improved instrumental resolution is marred by the overlaid collimator pattern. A serious Moire problem arises if these images are subsequently digitized at a conversion gain that is insufficient to retain this pattern. This artifact has been remedied(36) by rotating an offset tungsten collimator, with a dramatic improvement in image quality. The remaining problem of low inherent sensitivities is frequently solved by increasing the administered dose, exploiting the "forgiving" nature of short-lived positron emitters.

2.1.2.2. Coincidence Detection of Annihilation Radiation. The colinearity in the two-photon decay has long been recognized for its potential to localize the origin of an annihilation event. Several basic properties act to set ultimate limitations on the spatial resolution achievable through backprojection. The range of the positron separates the site of annihilation from the desired site of positron emission. Somewhat more serious for large planar systems is the deviation from colinearity imparted on the photon pair by the residual momentum of the electron-positron


pair prior to annihilation. These effects have been well studied(37,38) and their contributions (in quadrature) are not significant in current instruments.

Another basic effect can be expected to operate in laminographic devices based on Anger cameras. The two annihilation photons can carry away no angular momentum, since they originate from the singlet state of positronium. This constrains their respective linear polarizations to lie in orthogonal planes. (39) Furthermore, in an impending Compton scattering event within the NaI detector material, the scattered photon tends to be ejected normal to the electric vector of the incident radiation. (40) These multiple Compton-photoelectric interactions dominate the detection process in thick camera crystals at high energies and impart a sizeable component to the spatial resolution. (41) What has previously been neglected is that the combination of the orthogonal polarization and the Compton polarimeter nature of Anger cameras acts to correlate that component of position uncertainty in the opposing camera heads. Furthermore, it does so about an angle (90°) that could not be better chosen to minimize this contribution to the overall system resolution. Besides entertaining those with a penchant for second-order effects, these considerations caution one from adding all contributions in quadrature.

2.2. A Modular Positron Camera

Our goal has been to develop a positron imaging device subject to. the criteria in Table 3. Existing equipment should be used: retired gamma cameras, NIM electronics, and the nuclear medicine section computer with a CAMAC interface. After the optimal electronic parameters have been established, the NIM electronics should be repatriated, to be replaced with home-built, single-purpose components dedicated to the task.

Table 3. Positron Camera Design Criteria

Needs to have

high sensitivity stationary detectors open modular design low cost

Not intended to be

stand-alone, dedicated large field of view

In order to

match a limited production capability permit fast sampling of tracer kinetics bench-test positron imaging ideas act as tomographic "training wheels"

Since

the computer must serve all nuclear imaging needs fields will be restricted to known lesions


Table 4. Detector Characteristics

Parameters

Supplier Graded filter(42) Number of phototubes Origin of energy signal 511-ke V photofraction 511-keV timing (vs. Pilot U)

Detector A Detector B

GE RAD II Searle Positron III [Pb (3 mm) Sn (1 mm) Cu (1 mm)] 19 19 summed X±, y± common tenth dynode 18% 46% 14 nsec 13.6 nsec

2.2.1. COINCIDENCE DETERMINATION

303

The detectors are an opposed pair of Anger camera heads, characterized in Table 4, subscripted (A, B) to designate the origin of the gamma event. Fast coincidence between photo events is determined by constant fraction timing discrimination (CFTD), overlapped with a 40-nsec timing aperture. Time-to-amplitude conversion (T AC) shows a measured system time resolution of 22 nsec, in good agreement with the quadrature sum of the individual detectors when opposed by a fast plastic scintillator, as shown above. This time resolution results in a real-to-accidental ratio of 20/1 in a typical cerebral imaging situation. The close interplay between the coincidence timing and dynamic range of the accepted gamma energies is revealed by "imaging" such signal pairs as (TAC, Bsurn = BA + BB) and (BA, BB), as though they were the (X, Y) signals from a conventional gamma camera.

The electronics for these measurements are shown in the two circuits of Figure 3. For simplicity, we shall consider the various source configurations viewed by a pair of discrete detectors (5 cm x 5 cm NaI). Figure 4 shows such a set of TAC-vs.-Bsurn and BA-vs.-BB images in the simplest case of a point source centered between the two detectors in air. The effect of spreading the activity out into a plane, or introducing realistic scatter, results in the dominance of Compton-Compton events with the real coincidence slice. Figure 5 shows the altered topography as the point source is shifted normal to the center line out of the sensitive volume, with the photoevent sum peaks giving way to interdetector scatter. The time resolution can be seen by the 12-nsec TAC width of Figure 5, taken at the Bsurn = 1.022 Me V photopeak slice. This time spread can be seen to increase at lower summed photon energies, where the flume pattern characterizes a "walk" misalignment on the stop-side timing discriminator.

These effects of scatter, off-axis activity, and widened apertures in energy and time are most clearly revealed in a cinema playback of a dynamic


r z

t------(/ TAC 0 Y L.....;;;T-=-'

1-+-----11---0 Z

LH'lOOnSEC

t------(!---O X L..::.:::":"':...... ESUII

y

Figure 3. Block diagram of the electronics used for two-parameter analysis of the characteristic signal pairs (TAC, Bsum = BA + BB) and (BA, BB)' The detectors are 5 cm x 5 cm NaI scintillators, viewing the positron point source at 10 em spacing. Constant fraction timing discrimination (CFTD) triggers the time-to-amplitude convertor (TAC) at 50 keY on each branch. The stop side has been slightly misaligned in its walk adjustment. Analysis exploits the dual ADC "front end" of an imaging computer expecting the (X, Y) signal pair from an Anger camera (see text).

study taken as the parameter in question (e.g., source position) is steadily advanced. The two-parameter analysis demonstrates that distributed activity and successive layers of overlying scattering material quickly mute the sharp features of the energy-time domain. Windows in both time and energy need to be carefully optimized to achieve the best performance in each particular type of study.

2.2.2. POSITION SIGNAL PROCESSING

Two approaches to laminography are implemented, focusing either single or multiple planes.

Fig

ure

4.

Isom

etri

c an

d i

nten

sity

-cod

ed d

ispl

ays

of

the

dete

ctor

pai

r vi

ewin

g a

cent

ered

poi

nt s

ourc

e of

88G

e. T

he (

E. u

rn, T

AC

) sp

ectr

um o

n t

he

left

an

d (

EA

, E

B)

on t

he r

ight

are

dis

play

ed a

s (X

, Y

) in

th

e im

ages

sho

wn.

Th

e ve

rtic

al Z

ax

is h

as b

een

soft

ened

fo

r di

spla

y by

th

e sq

uare

-roo

t op

erat

ion

. The

refo

re (

X,

Y, Z

) re

pres

ents

[E

. urn

, T

AC

, (c

ount

s)1/

2J.

"1:1 Ii ! a i· ~

51 It ~ ,. ::. !. ~ =

=

~

Co ! =

11 ~ =

til

Fig

ure

5. D

ispl

ays

sim

ilar

to

Fig

ure

4, w

ith t

he m

idpl

ane

posi

tron

pO

int s

ourc

e di

spla

ced

norm

al t

o th

e de

tect

or's

com

mon

cen

ter

line

be

yond

the

sen

siti

ve v

olum

e. T

he

topo

grap

hy o

f th

e (5

0-fo

ld r

educ

ed)

coun

ting

rat

e sh

ows

peak

s at

pho

ton

ener

gies

cor

resp

ondi

ng to

in

terd

etec

tor

scat

ter.

The

car

eful

ana

lysi

s of

suc

h so

urce

sca

tter

dis

trib

utio

ns is

a n

eces

sary

pre

lude

to

syst

em d

esig

n.

fM

~ " [ !"

' ao; ~ i ~ ~ c:'l ~

~


CAMERA#I

-y

-y+

307

TO GAMMA CAMERA CONSOLE

- X (DIFFERENCING, ... )

t------- x+ AND COMPUTER

Figure 6. Electronic block diagram depicting the analog processing to focus a single-plane image. By gating the summed signals with the coincidence criteria, the difference and shaping circuits in the console are shielded from the 2 x 105 c/sec singles traffic. The resulting (X, Y) signal pairs are then digitized by the dual ADCs of the conventional gamma camera interface, under the control of standard imaging software.

2.2.2.1. Analog Processing of Single-Plane Images. In an effort to preserve the familiar identity of the camera-computert system, Figure 6 shows the simple interjection of the coincidence and analog signal mixing circuitry between the gamma camera heads and the camera console. The four position signals from each camera head are weighted, summed, delayed, and gated by the coincidence criteria. This relieves the slower circuits downstream from the overwhelming (-2 x 105 c/sec) singles traffic. In this way, such key processing node points as signal differencing and stretching face only modest (-3 x 103 c/sec) coincidence rates and a potential bottleneck is averted. (13)

The focal plane position is determined a priori by the weighting factor, while the depth of focus is set by the detector spacing, nominally 50 cm. The speed of analog signal processing and the simplicity of the system are attractive but paid for by the need to anticipate the desired focal plane.

2.2.2.2. Software Processing of Multiple Images. Clearly ail operations of weighted summation, normalization, and even energy selection can be done digitally a posteriori if the two sets of four position signals are individually digitized. The CAMAC(43.44) interface provides a generalized

t Digital Equipment Corp, PDP 11/40, Gamma 11.

308

INVERT DELAY

Robert J. Nickles and S. John Gatley

CAMACRESIDENT a-FOLD 11- BIT ADC

CAMAC DATAWAY

TO COMPUTER

STROBE

Figure 7. Block diagram showing the simplified electronics for multiplane laminography in software. The eight position signals are shaped and delayed for presentation to the gated inputs of an octal ADC. This module converts the signals to eight ll-bit addresses, passed to core through the CAMAC data way by the crate controller. The weighted summation, differencing, normalization, and backprojection are performed by an assembly language code on an event-by-event basis.

"front end" to almost any imaging computer. An octal ADCt simultaneously digitizes the eight position signals to ll-bit accuracy in 70 fJ-sec when accompanied by a coincidence strobe. The resulting look-at-me (LAM) request for service through the CAMAC crate controller transfers the eight-address set to core and initiates the computation of the event coordinates (X, Y) backprojected onto four planes as shown in Figure 7. The placement, spacing, and number of focused planes are readily selected in the acquisition program,(45) but core limited to four 64 x 64 (word) images or two 128 x 128 (byte) images.

The computation time approaches 1 msec/event for the 300 assembly language instructions and imparts nonparalyzable losses saturating at 1 kHz. Losses through all node points are monitored by a pulser pair injected(46) into test inputs of the camera heads and skimmed out of the data stream after conversion. By hardwiring the backprojection algorithms currently

t EG&G Ortec Model AD811, 100 Midland Rd, Oak Ridge, Tenn.


performed in software, the processing time could be made to approach the irreducible overhead of the CAMAC ADCs.

2.2.3. MULTIPLANE DYNAMIC IMAGING

Frame-mode dynamic image acquisition is handled by a FORTRAN

program passing four plane sets directly into Gamma 11 disk files. The selectable frame rate is keyed to the real-time clock and disk storage limits the study to 400 frames (100 time intervals). The same program interrogates a bank of CAMAC scalers, forming a convenient sideband for the parallel monitoring of

1. Counting rates from probes over such fields as heart, lungs, and exhaled gas to widen the compartmental perspective.

2. Frequency-converted physiological transducers. 3. Data losses scaled by logging test signals throughout the study.

This latter channel is important for the correction of observed regional activity curves. These losses could have been monitored by inclusion of a fixed positron source in the field view, but there is precious little field or count rate ability to spare.

2.2.4. IMAGE CORRECTIONS

Regardless of the imaging mode, two image corrections must be addressed. The first is to recognize that accidental coincidence events overlay an axially symmetric distribution I(X, Y) = I(r, 8) = I(r) on top of the real events in the observed image. The single-photon image seen by each uncollimated detector is essentially flat within the circular field X2 + y 2 "., R2 of radius R. The uncorrelated origins of signals A and B insure that the midplane random image will be the pointwise convolution overlapping these two unit circles (r "., R = 1) given by the mensuration formula(47)

I(r) = 7TR2 - 2[r(R2 - r2)1/2 + R2 sin-l(~)] (4)

Actually, the random images recorded in successive focal planes are the result of convolving circles of increasingly dissimilar radii. This image contamination is minimized by keeping the real! accidental ratio greater than 20/1 in clinical imaging situations.

More important is the severe distortion impressed on the real image of a uniform plane source. This geometrical "flood-field correction" follows directly from the analytic form(48) of the point source response function,

310 Robert J. Nickles and S. John Gadey

involving expressions similar to equation (4). This correction is axially symmetric, but strongly dependent on both radius and position along the line between detectors. This nonuniform response discourages scintiphotography and quenched enthusiasm for positron cameras in the decade prior to computers with image division capability. Million-count floods acquired at each focused plane suffice to correct their respective images.

2.2.5. PERFORMANCE

The performance of the system is summarized in Table 5·. The laminographic capability can be qualitatively seen in the four-plane image of an I-shaped phantom placed on the uppermost focal plane, shown in Figure 8. In summary, the camera is a curious mixture of old ideas and current hardware. The computer makes it work, as the flood correction removes the crippling distortion. Furthermore, the NIM-CAMAC approach makes it easy to operate, achieving the goal of effective positron imaging with sufficient sensitivity to be served by a small research accelerator.

3. RADIOPHARMACEUTICAL SYNTHESIS

3.1. General Considerations

Because of the short half-lives of the radioisotopes under consideration, synthesis times and yields must be optimized. Apart from the obvious financial need to conserve isotopes, radiation exposure to the staff is usually minimized when a synthesis can be done rapidly and in high yield. An important principle for the design of synthetic schemes is that the radiolabel should be incorporated as late as possible into the reaction sequence. Purification of the radiopharmaceutical must also be rapid, making flow-

Table 5. Positron Camera Performance

Spatial resolution in focal plane Depth of focus at 50 cm detector separation Sensitivity in air (680 a)

Centered point source Centered disk phantom

Photopeak counting rate with disk ~ource Singles A Singles B True coincidences

"Disk = 20 em (diam.) x 5 cm (thick).

llmm 30mm

64 c/I'Ci-sec 37,000 c/I'Ci-sec

700 c/I'Ci-sec 580 c/sec/I'Ci

31 c/sec/I'Ci


Figure 8. The multiplane program focuses four images positioned at k (top left), it; (top right), -& (bottom left), and -k (bottom right) of the nominal 50-em interdetector spacing. An I-shaped phantom 14 cm long is placed in the uppermost plane and sharply delineated in the lower right image. The lateral resolution is 11 mm (full width at half maximum) in the focused plane. This resolution is two-fold defocused in a plane displaced by 3 cm, defining the depth of focus.

through processes preferable to batch processes, both because of speed and radiation safety.

The stringency of the above requirements varies with the half-life of the isotope being used. Thus, only hot-atom and flow-through reactions have been used with 150, while the ttO-min half-life of 18F allows batch reactions and multistep syntheses, as long as yields are reasonably high. For llC, synthetic precursors such as CO2, CH20, CH3I, and HCN are produced in as continuous a fashion as possible, then used in the terminal steps of reactions.

High-pressure liquid chromatography (HPLC) is now the method of choice for purifying of compounds labeled with short-lived radionuclides, especially 13N and llC. A great difference between "orthodox" and radiopharmaceutical chemistry is that the latter is done on as small a scale (as near carrier free) as possible. This poses many problems, but has the advantage that purification can often be done with analytical HPLC columns.

The specific radioactivity required of a radiopharmaceutical varies according to its proposed use. Generally, specific activities should be high enough to avoid possible toxic pharmacological effects. Several groups are currently striving towards the production of radiopharrnaceuticals designed


to map populations of hormones or neurotransmitter receptors. Since the number of molecules of the compound administered will have to be of the same order as the number of receptor sites, this goal requires very high specific activities (1 Ci/ #Lmol).(49) The theoretical maximum specific activity of l1C is roughly 5000 Ci/ #Lmol.

3.2. Review of Recent Developments

Several collections of reviews and research articles dealing with radiopharmaceutical chemistry have been published in the last few years. (50-53) These all include chapters on positron-emitting radiocompounds. The Proceedings of the Pirst(54) and Second(55) International Symposia on Radiopharmaceutical Chemistry, publjshed in the Journal of Labelled Compounds, also contains many interesting papers. The monograph Short-Lived Radioactive Gases for Clinical Use,(56) published in 1975, gives useful details on handling accelerator-produced nuclides. The book Isotopic Carbon,(57) published in 1949, summarizes the impressive early work with 11e.

Recently, Prench scientists at Orsay have managed, by careful treatment of reagents and target gases, to increase the specific activity of l1C compounds made via l1CH3I and l1CH20 to a level of 1 Ci/ #Lmol.(58) The compounds made include imipramine and chlorpromazine. This group has increased the specific activity of l1C compounds by almost two orders of magnitude since 1976.(59) Another such increase ought to allow receptor mapping.

The last few years have also seen the development of l1C amino acid synthesis by the high-temperature Bucherer-Strecker reaction introduced by the Oak Ridge group. (60) Several physiological and nonphysiological amino acids have been reported and are under study for use as pancreas-(61) and tumor-localizing agents. (62) Although chemical syntheses of amino acids invariably result in the optically inactive dl compounds, recent developments in the use of "chiral eluants,,(63) should now allow easy separation of optical isomers.

There have been several important recent developments in the field of 18p radiopharmaceuticals. The Brookhaven group has pioneered the use of 18P2. This reagent is produced by deuteron bombardment of neon containing a trace of P2 gas. (64) Although some 18p sticks to the target walls and some reacts with impurities to produce organic 18p gases, a reasonable fraction of the 18p activity can be recovered as 18P2. To date, this reagent has been used to produce several compounds, including 18p_5_ f1uorouracil(65) and 18P-2-deoxy-2-f1uoroglucose.(66) The latter has generated considerable interest for brain, (67) heart, (68) and tumor(69) studies.

The St. Louis group has made several very important contributions to the use of the 18p fluoride ion.(70) By trapping H 18p from a neon target


on cesium hydroxide, they generate a 18F of very high specific activity. They were thus able to obtain 18F-3-deoxy-3-fluoroglucose only an order of magnitude away from being carrier free. The importance of this to those wishing to make receptor-binding radiopharmaceuticals is tremendous. This method also allowed Tewson and Welch to study the biodistribution of 18F-2-fluoroethanol,(71) a compound whose extreme toxicity demands a very high specific activity. Finally, the St. Louis group has been able to produce aromatic fluoro compounds with no carrier added by a modification of the method of Wallach,(72·73) involving thermal decomposition of diazopiperidine compounds. One of the agents they have made is the tranquilizer 18F-haloperidol.(74) It seems very probable that, with modification, the techniques of Tewson and Welch will allow useful amounts of 18F radiopharmaceuticals to be made from the more accessible reactorproduced 18F-. (75)

Although there has been a resurgence of interest in fundamental biochemical and physiological studies with 13N,(76-79) relatively few new compounds of radiopharmaceutical interest have been reported in the last couple of years. The synthesis of 13N nitrous oxide(80) and octylamine(81) deserve mention for their potential to map cerebral blood flow through their ability to freely diffuse across the blood-brain barrier.

Gallium-68, which is available from generators. from the 275-day parent 68Ge, has also been the subject of recent work. The 68Ga-EDT A complex is excluded by the blood-brain barrier, 68Ga phosphates and phosphonates localize in bone, (82) 68Ga human serum albumin, (83) and red blood cells are confined to the vasculature. Generators have now been made that yield 68Ga oxine(84) and ionic 68Ga,(85) rather than the EDTA chelate. Platelets and white blood cells(86) have been labeled with these precursors. The bulk of today's nuclear medicine is done with scanning agents whose localization is independent of the actual radionuclide. There seems little doubt that with 68Ga, positron-emitting equivalents of most of the 99mTc radiopharmaceuticals can be made. This versatility insures that the power of positron tomography can be brought to bear on the problems currently dominating routine scanning procedures.

4. RESULTS

4.1. Practical Aspects in a Clinical Setting

In the tolerant research setting, positron imaging with short-lived tracers is taxing; in the clinical setting, it is baroque. The careful orchestration of balky accelerators, syntheses, and imaging instruments crumbles as the patient is unable to find a parking place. The bulk of frayed relations can


be avoided if a seemingly extravagant attention is paid to communications. The tandem vaUlt, radiochemistry lab, imaging suite, and computer room are on a party-line intercom. Further, TV monitors at all stations depict at a glance both the incandescent beam on target and the well-being of the subject beneath the positron camera. The patient is never alone, and every effort is made to assuage his very legitimate concerns amid the pandemonium that surrounds these studies.

The nuclear medicine technologist and radiopharmacist are quick to learn the rhythm imposed by the rapid physical decay. As an extreme example, a current research protocol requires investigation of the change in tumor blood flow throughout a course of radiotherapy. The washout of H2 150, introduced as a bolus through an indwelling arterial catheter implanted for chemotherapy, is followed by conventional dynamic imaging. The synthesis of the water is continuous,(88) saturating to an ample 2S-mCi dose. The water is made isotonic, sterile by millipore filtration, and injected via the catheter in less than 20 sec. The study is abbreviated by the 2-min decay, with multiple views permitted by the low absorbed dose to the patient.

Positron imaging breeds a form of discipline not normally required in a nuclear medicine department. Many gamma cameras are grossly undershielded, so that a 99mTc procedure in progress can be affected by the proximity of several millicuries of annihilation radiation in a nearby syringe or passing patient. This becomes more troublesome in the case of a positronemitting gas, prone to leak around mouthpieces. Unshielded and airborne, the influence of the source is unpredictable by virtue of the several-meter positron range. Fans and vacuum cleaners insure a rapid turnover of fresh air in the vicinity of the patient and camera.

Positron tomography is a quantitative imaging mode, because the tissue attenuation can be corrected by a previously measured transmission image. This opportunity to ascertain the fractional dose taken up by the organ in question carries with it the need for ancillary measurements, such as activity in blood or urine, inclusion of a known aliquot in the field of view, and the simple recording of time. For example, there is an experimental protocol that involves investigating the localization of 18F-S-fluorouracil(65) in patients undergoing tumor biopsy. (89) A digital stopwatch is started at end-bombardment, accompanies the dose through its chemical synthesis, follows the patient through injection, imaging, and biopsy, to return with

) Figure 9. (top) The lateral cerebral distributions (clockwise from upper left) of 68Ga-EDTA, 13NH3, 18p_3FDG, and 18p_2PDG focused at midplane in the same normal sUbject. The 400-kcount images are acquired in 15 min, after equilibrating with a typical intravenous dose of 300 ~Ci. Prame-mode dynamic imaging is augmented by CAMAC-scaling the counting rates observed by probes over other organs. (bottom) Brain and heart uptake are shown here for 18p_2FDG and 18p- 3FDG.

Positron Imaging: Some Practical Consideration

'"F-2FDG

. ~ I (\ .. '_-..... - I

I

.J BRAIN

1000 r'-- ---....~-.'.----.

U . ~ ... ~ .. / I '-,- .,..--

~ 100 ." I z . HEART

5 I U ......

I~s~ I I

315

'"F-3FDG


the tissue specimen for in vitro counting. At each benchmark, the time and activity are recorded, with the physical presence of the clock impressing a temporal order on the conduct of the study.

4.2. Representative Images

Figure 9 (top) shows a collage of lateral views of the brain midplane of a normal volunteer sequentially imaged with 68Ga-EDTA, 13NH3, 18F_ 2FDG, and 18F-3FDG. The dynamics of the 18F activity are included [Figure 9 (bottom)] for the two fluorosugars in the subject's heart and brain.

The 68Ga-EDT A is eluted directly from a commercial 68Ge_ 68Ga generator. The 68Ga complex is excluded from the normal brain volume by the blood-brain barrier, making it suitable for tumor detection. The wide range(83-87) and ready availability of 68Ga-Iabeled pharmaceuticals insures a workhorse role in positron tomography. Ammonia, labeled with 13N, is reduced from 13NO;-,(90) formed by proton irradiation of water. The localization of 13N ammonia has been shown(91.92) to depend on a number of physiological factors, with a nonlinear dependence on regional perfusion. This complexity prevents the use of a 13NH3 scan as a quantitative map of brain flow, but does not preclude its use as a qualitative image to assess the relative effects of intervention or stress.

The 18F-2-fluoro-2-deoxyglucose is prepared by the method of Ido,(66) starting from 18F2 formed by deuteron irradiation of a Ne-1 % F2 mixture.(64.93) The 8% (decay-corrected) yield from the initial 15 mCi of fluorinated triacetal glucal results in sufficient 18F-2FDG to image two subjects with the positron camera. Probes over the brain, heart, and bladder monitor 18F activities in the scaler side channels of the dynamic study, revealing the 70% cerebral extraction fraction in the first-pass kinetics. Similar studies with 18F-3-fluoro-3-deoxyglucose, prepared by the method of Tewson,(70) modified to start with reactor-produced 18F-, (75) show a diminished cerebral contrast. Neither the images nor kinetics of 18F-3FDG show any discernible carrier effects in five repeated studies in the same subject, studied with 10 J.'mol of added carrier and no carrier added. The ability of the fluorosugars to "freeze in" the distribution of metabolic activity epitomizes the radio nuclide image as a map of organ function.

5. CONCLUSIONS

Radionuclide imaging procedures can see past morphology to reveal organ function. This unique perspective has been extended to the biochemical level by the use of short-lived positron emitters to label tracers that are ideal in their authenticity. With specific activity being pushed to


its theoretical limit, such features as hormonal receptor sites may soon be mapped without perturbing the delicate balance under investigation.

The fact that these tracers also lead naturally to emission tomography seems, at first, fortuitous. On second thought, however, it is largely the promise of heightened contrast through tomographic imaging that sparks the quest for more subtle mechanisms than the "mechanical" localization of 99=rC compounds. Positron emission tomography allows the scientist to follow a radiopharmaceutical through the intact subject with a quantitative assay that equals autoradiography in its regional isolation. This same ability will give rise to more sensitive and specific diagnostic tests when translated into the wider clinical arena.

ACKNOWLEDGMENTS

The authors are greatly indebted to M. T. Madsen, R. D. Hichwa, D. J. Simpkin, J. L. Martin, W. J. Shaughnessy, and M. E. Daube for their contributions to all aspects of this work. The U.W. tandem group generously allocated beam time, and support was provided by NIH-NCI through the multidisciplinary grant P01-CA-19278-04. We would also like to acknowledge NIH l-K04-HL00213 for its support through a research career development award (RJN) , and the Wellcome Trust for a travel award (SJG).

REFERENCES

1. A. P. Wolf and C. S. Redvanly, Carbon-II and radiopharmaceuticals, Int. 1. Appl. Radiat. Isot. 28, 29-48 (1977).

2. M. G. Straatman, A look at 13N and 150 in radiopharmaceuticals, Int. 1. Appl. Radiat. Isot. 28, 13-20 (1977).

3. G. A. Russ, R. E. Bigler, J. R. Dahl, J. Kostic, J. M. McDonald, R. S. Tilbury, and J. S. Laughlin, Regional imaging with oxygen-14, IEEE Trans. Nucl. Sci. 23, 573-576 (1976).

4. T. Jones, D. L. Levene, and R. Green, Use of carbon dioxide for inhalation radiocardiograms and measurements of myocardial perfusion. Dynamic Studies with Radioisotopes in Medicine, 751-764, IAEA, Vienna (1971).

5. R. J. Nickles and Y. F. Au, The oxygen c1ock-a dual-tracer physiological timer, Phys. Med. Bioi. 20, 54-66 (1975).

6. R. J. Nickles, S. J. Gatley, M. T. Madsen, R. D. Hichwa, J. L. Martin, D. J. Simpkin, and J. R. Halama, High-yield synthesis of the sequence 13N2, 13N20, llC2 H2, H2150, H2 140 and 17F- for rCBF studies, 1. Labelled Compd. 16, 90-91 (1979).

7. G. D. Robinson, Biologically active 18F-ftuoroorganic compounds: rapid, high specific activity synthesis of 18F-ftuorocarboxylates and derivatives, in Radiopharmaceuticais and Labeled Compounds, Vol. I, pp. 423-432, IAEA, Copenhagen (1974).


8. H. N. Wagner, Principles of Nuclear Medicine, p. 1, W. B. Saunders, Pa. (1968). 9. G. L. Brownell, J. A. Correia, and R. G. Zamenhof, in Recent Advances in Nuclear

Medicine, Vol. 5 (J. H. Lawrence and T. F. Budinger, eds.), pp. 1-49, Grune and Stratton, New York (1979).

10. M. E. Phelps, Emission computed tomography, Semin. Nucl. Med. 7, 337-365 (1977). 11. R. M. Lambrecht, in The Chemistry of Radiopharmaceuticals (N. D. Heindel, H. D.

Burns, T. Houda, and L. W. Brady, eds.), pp. 87-107, Masson, New York (1978). 12. F. W. Wrenn, M. L. Good, and P. Handler, The use of positron-emitting radioisotopes

for the localization of brain tumors, Science 113, 525-527 (1951). 13. G. L. Brownell and W. H. Sweet, Localization of brain tumors with positron emitters,

Nucleonics 11, 40-45 (1953). 14. P. V. Harper, K. A. Lathrop, and A. Gottschalk, in Radioactive Pharmaceuticals (G. A.

Andrews, J. W. Kniseley, and H. N. Wagner, eds.), U.S. Atomic Energy Commission (1966).

15. H. O. Anger, Scintillation camera, Rev. Sci. Instrum. 29, 27-33 (1958). 16. H. O. Anger, Sensitivity, resolution and linearity of the scintillation camera, IEEE Trans.

Nucl. Sci. 13, 380-392 (1966). 17. G. Muehllehner, Positron camera with extended counting rate capability, J. Nucl. Med.

16,653-657 (1975). 18. C. A. Burnham and G. L. Brownell, A multicrystal positron camera, IEEE Trans. Nucl.

Sci. 19 (3), 201-205 (1972). 19. L. R. Carrol, Design and performance characteristics of a production model positron

imaging system, IEEE Trans. Nucl. Sci. 25, 606-614 (1978). 20. C. B. Lim, D. Chu, L. Kaufman, V. Perez-Mendez, R. Hattner, and D. C. Price, Initial

characterization of a multi-wire proportional counter positron camera, IEEE Trans. Nucl. Sci. 22, 388-394 (1975).

21. M. K. Kopp and C. J. Borkowski, Positron-sensitive proportional counters using resistance-capacitance position encoding, Oak Ridge Nat. Lab. Rep. ORNL-TM-5083.

22. J. S. Robertson, R. B. Marr, M. Rosenblum, V. Radeka, and Y. L. Yamamoto, in Tomographic Imaging in Nuclear Medicine (G. S. Friedman, ed.), pp. 142-153, Society of Nuclear Medicine, New York (1973).

23. S. E. Derenzo, T. F. Budinger, J. L. Cahoon, R. H. Huesman, and H. G. Jackson, High resolution computed tomography of positron emitters, IEEE Scintillator and Semiconductor Counter Symposium, New Orleans, 1976, IEEE Trans. Nucl. Sci. 24, 544-558 (1977).

24. Y. L. Yamamoto, C. J. Thompson, and E. Meyer, Dynamic positron emission tomography for the study of cerebral dynamics in a cross section of head using positron emitting 68Ga-EDTA and 77Kr, J. Comput. Assist. Tomogr. 1 (I), 43-56 (1977).

25. Z. H. Cho, J. Chan, and L. Eriksson, Circular ring transaxial positron camera for transverse section reconstruction, IEEE Trans. Nuc/. Sci. 23 (1), 613-617 (1976).

26. Therascan 3128, Atomic Energy of Canada Limited, P.O. Box 6300, Ottawa, Canada. 27. L. R. Carroll, G. O. Hendry, and J. D. Currin, Design criteria for multislice emission

computed tomography detector system, Cyclotron Corporation, Berkeley, Ca. 28. M. M. Ter-Pogossian, M. E. Phelps, E. J. Hoffman, and N. A. Mullani, A positron

emission tomograph for nuclear imaging (PETT), Radiology 114, 89-98 (1975). 29. EG&G Ortec Incorporated ECAT, Model 4832/01, 100 Midland Road, Oak Ridge, Ten. 30. N. A. Mullani, C. S. Higgins, J. T. Hood, and C. M. Currie, PETT IV; Design analysis

and performance characteristics, IEEE Trans. Nucl. Sci. 25, 180-183 (1978). 31. M. M. Ter-Pogossian, N. A. Mullani, C. S. Higgins, J. T. Hood, and N. J. Caston, The

performance of PETT V; A positron emission tomograph for fast dynamic studies, J. Nucl. Med. 20, 628 (1979).


32. E. J. Hoffman, S. C. Huang, and M. E. Phelps, Effect of object size in quantitative positron computed tomography, J. Nucl. Med. 19, 683 (1978).

33. Comments by Drs. Oldendorf and Phelps, in Non-Invasive Brain Imaging (H. J. DeBlanc and J. A. Sorenson, eds.), p. 200, Society of Nuclear Medicine, New York (1975).

34. H. J. Ache, B. Djermouni, and E. D. Handel, Positron annihilation as a nuclear probe for the study of biophysical properties, presented at the 178th Meeting of the American Chemical Society, Washington, D.C. (1979).

35. O. Klein and Y. Nishina, tiber die Streuung von Strahlung durch freie Elektronen nach der neuen relativistischen Quantendynamik von Dirac, Z. Phys. 52, 853-868 (1929).

36. B. Brunsden, P. V. Harper, and R. N. Beck, Elimination of collimator hole pattern by double displacement of a hexagonal array, 1. Nucl. Med. 16, 517 (1975).

37. M. E. Phelps, E. J. Hoffman, S. C. Huang, and M. M. Ter-Pogossian, Effect of positron range on spatial resolution, 1. Nucl. Med. 16,649-652 (1975).

38. G. Muehllehner, Resolution limit of positron cameras, 1. Nucl. Med. 17, 757 (1976). 39. C. S. Wu and I. Shaknov, The angular correlation of scattered annihilation radiation,

Phys. Rev. 77, 137 (1950). 40. R. D. Evans, The Atomic Nucleus, p. 681, McGraw-Hill, New York (1955). 41. H. O. Anger and D. H. Davis, Gamma ray detection efficiency and image resolution in

sodium iodide, Rev. Sci. Instrum. 35, 693-697 (1964). 42. G. Muehllehner, R. J. Jaszczak, and R. N. Beck, The reduction of coincidence loss in

radionuclide imaging cameras through the use of composite filters, Phys. Med. Bioi. 19, 504-510 (1974).

43. CAMAC-A modular instrumentation system for data handling, European Atomic Energy Community-EURATOM, EUR-4100E, March (1969).

44. J. E. Holden, R. J. Nickles, A. J. Kiuru, and R. E. Polcyn, The use of a generalized modular data handling system in nuclear medicine, Radiology 122,411-416 (1977).

45. M. T. Madsen, Ph.D. thesis, University of Wisconsin, 1979. 46. R. J. Nickles, J. E. Holden, A. J. Kiuru, and R. E. Polcyn, Inclusion of physiological

data into computerized nuclear medicine dynamic studies, 1. Nucl. Med. 17, 297-300 (1976).

47. CRC Handbook of Chemistry and Physics, 48th Ed., p. A-257, Cleveland, Ohio (1967). 48. A. M. J. Paans, R. J. Nickles, E. J. de Graaf, W. Vaalburg, S. Reiffers, A. Steenhoek,

and M. G. Woldring, A dual-head Anger camera system for the imaging of positronemitting radionuclides, Int. 1. Nucl. Med. Bioi. 5, 266-273 (1978).

49. R. C. Reba, R. E. Gibson, W. J. Rzeszotarski, F. Vieras, J. K. Mazaitis, and B. Francis, Receptor-binding radiotracers: a class of potential radiopharmaceuticals, 1. Nucl. Med. 20,350-357 (1979).

50. Radiopharmaceuticals and Labelled Compounds, Vols. 1 and 2, International Atomic Energy Authority, Vienna (1973).

51. G. Subramanian, B. A. Rhodes, J. F. Cooper, and V. J. Sodd (eds.), Radiopharmaceuticals, Society of Nuclear Medicine, New York (1975).

52. M. J. Welch (ed.), Radiopharmaceuticals and Other Compounds Labelled with Short-Lived Radionuclides, Int. 1. Appl. Radiat. Isot. 28, 1-233 (1977).

53. N. D. Heindel, H. D. Burns, T. Honda, and L. W. Brady (eds.), The Chemistry of Radiopharmaceuticals, Masson, New York (1978).

54. Abstracts of the First International Symposium on Radiopharmaceutical Chemistry, J. Lab. Compo Radiopharmaceut.13, 155-290 (1976).

55. Abstracts of the Second International Symposium on Radiopharmaceutical Chemistry, 1. Lab. Compo Radiopharmaceut. 16, 1-233 (1978).

56. J. C. Clark and P. D. Buckingham, Short-Lived Radioactive Gases for Clinical Use, Butterworths, London (1975).


57. M. Calvin, C. Heidelberger, J. C. Reid, B. M. Tolbert, and P. F. Yankwich, Isotopic Carbon, Wiley, New York (1949).

58. G. Berger, M. Maziere, R. Knipper, C. Prenant, and D. Comar, Automated synthesis of llC-labelled radiopharmaceuticals; imipramine, chlorpromazine, nicotine and methionine, Int. I. Appl. Radiat. Isot. 30, 393-399 (1979).

59. C. Marazano, M. Maziere, G. Berger, and D. Comar, Synthesis of IIC_methyl iodide and ltC-formaldehyde, Int. I. Appl. Radiat. Isot. 28,49-52 (1977).

60. R. L. Mayes, L. C. Washburn, and B. W. Wieland, Synthesis and purification of llC_ carboxyl labelled amino acids, Int. I. Appl. Radiat. Isot. 29, 186-187 (1978).

61. M. Zalutsky,J. Wu, T. Wickland, P. V. Harper, and A. Moosa, ltC-Tryptophan: synthesis and purification using high pressure liquid chromatography, I. Nucl. Med. 20, 663 (1979).

62. B. Schmall, J. R. Dahl, and R. E. Bigler, Synthesis of ltC-labelled a-aminoisobutyric acid as a potential agent for tumor visualization, Abstract NUCL 083, 178th Annual Meeting of the American Chemical Society (1979).

63. P. E. Hare and E. Gil-Av, Separation of D and L amino acids by liquid chromatography: Use of chiral eluants, Science 204, 1226-1228 (1979).

64. R. M. Lambrecht, R. Neirinckx, and A. P. Wolf, Cyclotron isotopes and radiopharmaceuticals XXIII. Novel anhydrous fluorinating intermediates, Int. I. Appl. Radiat. Isot. 29, 175-183 (1978).

65. J. S. Fowler, R. D. Finn, R. M. Lambrecht, and A. P. Wolf, The synthesis of lsF_5_ fluorouracil, I. Nucl. Med. 14, 63-64 (1973).

66. T. Ido, C. N. Wan, V. Casella, J. S. Fowler, and A. P. Wolf, Labeled 2-deoxY-D-glucose analogs. lsF-labeled 2-deoxy-2-f1uoro-D-glucose, lsF-2-deoxy-2-f1uoro-D-mannose and 14C-2-deoxy-2-f1uoro-D-glucose, I. Lab. Compo Radiopharmaceut. 14, 175-183 (1978).

67. M. Reivich, D. E. Kuhl, A. P. Wolf, J. Greenberg, M. E. Phelps, T. Ido, V. Casella, J. S. Fowler, B. M. Gallagher, E. Hoffman, A. Avari, and L. Sokoloff, in Cerebral Function, Metabolism and Circulation (D. H. Ingvar and N. A. Lassen, eds.), pp. 190-191, Munksgaard, Copenhagen (1977).

68. B. M. Gallagher, A. Ansari, H. Atkins, V. Casella, D. R. Christman, J. S. Fowler, T. Ido, R. R. MacGregor, P. Som, C. N. Wan, A. P. Wolf, D. E. Kuhl, and M. Reivich, lsF-labelled 2-deoxy-2-f1uoro-D-glucose as a radiopharmaceutical for measuring regional myocardial glucose metabolism in vivo: tissue distribution and imaging studies in animals, I. Nucl. Med. 18, 990-996 (1977).

69. P. Som, H. L. Atkins, D. Bandyopadhyay, J. S. Fowler, R. R. MacGregor, K. Matsui, Z. H. Oster, D. F. Sacker, C. Y. Shiue, and A. P. Wolf, Early detection of neoplasms with radiolabeled sugar analog, I. Nucl. Med. 20, 662 (1979).

70. T. J. Tewson, M. J. Welch, and M. E. Raichle, lsF-labelled 3-deoxy-3-fiuoro-D-glucose synthesis and preliminary biodistribution data, I. Nucl. Med. 19, 1339-1345 (1978).

71. T. J. Tewson and M. J. Welch, Synthesis of carrier-free ISF-f1uoroethanol, I. Nucl. Med. 20,663 (1979).

72. O. Wallach, Uber einen Weg zur leichten Gewinnung organischer F1uorverbindungen, Ann. Chern. 235, 255-271 (1886).

73. O. Wallach and F. Heusler, Uber organische F1uorverbindungen, Ann. Chern. 246, 219-244 (1888).

74. T. J. Tewson and M. J. WeIch, A synthesis of carrier-free lsF-labelled aromatic compounds and its application to the preparation of the lsF-labelled neuroleptic haloperidol, I. Nucl. Med. 20, 671 (1979).

75. S. J. Gatley and W. J. Shaughnessy, Synthesis of lsF-3-deoxy-3-f1uoro-D-glucose from reactor-produced f1uorine-18, Int. I. Appl. Radiat. Isot. 31, 339-341 (1980).

76. J. P. Witter, S. J. Gatley, and E. Balish, Distribution of nitrogen-13 from labelled nitrate (13N03") in humans and rats, Science 204, 411-413 (1979).


77. J. P. Witter, E. Balish, and S. J. Gatley, Distribution of nitrogen-13 from labelled nitrate e3N03) and nitrite (13N0 2") in germ-free and conventional flora rates, Appl. Environ. Microbial. 38, 870-878 (1979).

78. T. M. Tiedje, R. B. Firestone, M. K. Firestone, M. R. Betlach, M. S. Smith, and W. H. Caskey, Methods for the production and use of 13N in studies of denitrification, Soil. Sci. Soe. Amer. Proe. 43, 709-716 (1979).

79. A. J. L. Cooper, J. M. MacDonald, A. S. Gelbard, R. F. Gledhill, and T. E. Duffy, The metabolic fate of 13N-labelled ammonium in rat brain, J. Bioi. Chem. 254, 4982-4992 (1979).

80. J. Thomas, J. C. Meeks, C. P. Wold, P. W. Shaffer, S. M. Austin, and W. S. Schieu, Formation of glutamine from 13N-ammonia, 13N-dinitrogen and 14C-glutamate by heterocysts isolated from Anabaena eylindrica, J. Bacterial. 129, 1545-1555 (1979).

81. R. J. Nickles, S. J. Gatley, R. D. Hichwa, D. J. Simpkin, and J. L. Martin, The synthesis of nitrogen-13 labelled nitrous oxide, Int. J. Appl. Radiat. Isot. 29, 225-227 (1978).

82. R. E. Lade, M. Castiglioni, and A. P. Wolf, Synthesis and metabolic studies of ultra-high activity 13N-labelled amines, Abstract NUCL 100, 178th Annual Meeting of the American Chemical Society (1979).

83. M. K. Dewanjee and D. J. Hnatowich, R. Beh, New 68Ga-labeled skeletal imaging agents for positron scintigraphy, J. Nucl. Med. 17, 1003-1007 (1976).

84. S. J. Wagner and M. J. Welch, Gallium-68 labeling of albumin and albumin microspheres, J. Nucl. Med. 20, 428-433 (1979).

85. G. J. Ehrhardt and M. J. Welch, A new germanium-68/gallium-68 generator, J. Nucl. Med. 19,925-929 (1978).

86. R. D. Neirinckx and M. A. Davis, Development of a generator for ionic gallium-68, J. Nucl. Med. 20, 681-682 (1979).

87. M. J. Welch, M. L. Thakur, R. E. Coleman, M. Patel, B. A. Siegel, and M. M. Ter-Pogossian, Gallium-68 labeled red cells and platelets: new agents for positron tomography, J. Nucl. Med. 18, 558-562 (1977).

88. R. J. Nickles, A. J. Kiuru, S. M. Schuster, and J. E. Holden, A catalytic generator for the production of H2 140 and H2 150, Prog. Nucl. Med. 4,72-79 (1978).

89. L. M. Lieberman, B. W. Wessels, A. L. Wiley, S. J. Gatley, R. J. Nickles, D. Young, W. H. Wolberg, and A. E. Bogden, 18F-5-ftuorouracil studies in humans and animals, Int. J. Radiat. Oneal. 6, 505-509 (1980).

90. W. Vaalburg, J. A. A. Kamphuis, H. D. Beerling-van-der-Molen, S. Reiffers, A. Rijskamp, and M. G. Woldring, Int. J. Appl. Radiat. Isot. 26, 316-318 (1975).

91. M. E. Raichle, Metabolic and biochemical studies of the brain with short-lived positronemitting radionuciides, Abstract NUCL 14, 178th Annual Meeting of the American Chemical Society (1979).

92. M. E. Phelps, S. C. Huang, D. E. Kuhl, C. Selin, and E. J. Hoffman, Can N-13 ammonia be used as a cerebral blood flow tracer?, J. Nucl. Med. 20, 611 (1979).

93. W. J. Shaughnessy, S. J. Gatley, R. D. Hichwa, L. M. Lieberman, and R. J. Nickles, Aspects of the production of 18F-2FDG via 18F2 with a tandem Van de Graaf accelerator, Int. J. Appl. Radiat. Isot. 32, 23-29 (1981).

6

Modern Optical Methods for the Storage of Radiographs

U. KILLAT

1. INTRODUCTION

Eighty-five years after the detection of the principles of radiographic diagnosis by W. C. Roentgen in 1895, the world's annual production amounts to some 2 billion radiographs. Scaled to the environment of a medium-size hospital, this means that the x-ray examinations will amass about 200,000 films per annum, which are stored for several years.

The reasons for storing old radiographs are the following.

1. Their status as legal records: depending on the legislation of the individual country, radiographs have to be stored for up to 30 years, thus providing a medico-legal record for future contingencies.

2. Their use in comparison views: radiologists and other medical staff require earlier x-ray studies to evaluate, for instance, a patient's progress or the temporal development of an anomaly.

3. Their use as a teaching file: the saving of radiographs provides a teaching file that may also be used for retrospective research.

As the radiograph is one of the most important data sources in modern medicine, effective means for storage, retrieval, and transfer of this information are of utmost importance.

U. KILLAT • Philips GmbH Forschungslaboratorium Hamburg, Vogt-Kolln-Str. 30, D 2000 Hamburg 54.

323

324 U. Killat

In today's x-ray files the original radiographs are stored, and this "simple" solution results in one of the greatest problems of most x-ray departments. Filing, retrieval, and transport of normal-size films (typically 14 in. x 17 in.) is space and time consuming, tedious, and expensive. Owing to the great size of these documents, their handling is costly to automate. Among the complaints put forward, lack of space is certainly the most prominent. In the example given above, about 1500 m3 of storage room are required for 30 years of filing. This is particularly annoying, if one bears in mind that the retrieval request rate drops below 2 % after two years,o)

Besides its bulkiness, today's x-ray store exhibits functional problems such as the lack of integrity and the cumbersome loan service organization. The files on active patients are in constant demand for consultation, reevaluation, review, and teaching purposes. The resulting extensive handling of envelopes containing several radiographs frequently results in temporary or even permanent loss of radiographs. The search for a lost radiograph is quite ineffective in the present system and strongly reminds one of looking for a needle in a haystack. Owing to the more rigid requirements recently developed in several countries, (2) the lack of integrity of a patient's file can probably no longer be tolerated. This, in turn, could inhibit the transfer of information between departments or hospitals-an idea that can by no means be desirable.

Obviously, the classical concept for storing radiographs has its drawbacks and a more sophisticated solution to the problem has to be developed. This new storage system should incorporate, in a flexible manner, data obtained from novel imaging techniques such as x-ray video, computer tomography, nuclear imaging, ultrasound imaging, and infrared imaging.

Furthermore, it should incorporate the following required features:

1. Small size, without any noticeable information loss. 2. Permanent integrity. 3. Simple maintenance and retrieval. 4. Simple techniques for producing copies for loan service activities. 5. Compatibility with other data processing techniques used in the

hospital. 6. Acceptable cost.

As a first attempt in this direction, photographic minification systems have been developed and introduced in hospitals. Section 2 is entirely devoted to minification systems. It describes the systems available on the market, the experiences gained so far, and the physical limitations to which they are bound.

Modern Optical Methods for the Storage of Radiographs 325

Section 3 gives a brief outline on notions that are related to the information capacity of radiographs. This section provides some analytical tools for use in the following two sections.

Section 4 describes a method of optical information storage that became very popular when lasers of sufficient power and coherence length became available: holography. The storage of pictures is one of the applications of holography, which is fully understood in its principles but lacks the elaboration of certain applications and the evaluation of suitable recording materials. The state of the art with respect to the storage of gray tone images is documented by simple analytical models and some instructive experiments related to the demands of the storage of radiographs.

In contrast to the analog optical methods described in Sections 2 and 4, Section 5 is concerned with picture-handling techniques in a digital environment. The key issue is the development of optical mass memories which solve the problem of memory capacity, but which have to be interfaced to the types of documents that must be kept safe in a hospital.

Finally, Section 6 traces the lines along which the latter approach could merge into an integrated communication system for all kinds of information storage and transfer for inter- and intra-hospital communication.

2. MINIFICATION SYSTEMS

Microfiches have successfully been used to file patients' case reports. The stored information can be made available by using a special reader or by producing a full-size hard copy print.

A common means of storage of radiographs and case reports in a common location would obviously simplify a hospital's file management and could solve the crucial space problem. To this end a variety of formats have been proposed for the storage of radiographs(3) and a discussion has evolved as to how "mini" "mini" should be. It seems that basically two formats have been evolved, which have found their entrance to the hospital:

100-mm film. The advantage of this size is the adequacy of the image for direct diagnosis without a magnifying viewer. The saving in storage area is about 90% and it therefore seems to be economically attractive. In addition, the format corresponds nicely to that of images gained from image-intensifier tubes or the Polaroid images of computed tomography (CT) scanners. Finally, this size is well suited for distributing copies by mail.

35-mm film. The argument for this format is simple but decisive: the improved saving of space (about 99%). Another advantage could be that it is well adapted to computer-controlled image densitometry.

326 U. KiDat

The 100 mm x 100 mm format is found in the Delcomat system (De Oude Delft(4), whereas the 35-mm film is mainly used by LogEtronics.(5) A third system that was successfully tested in hospitals(6) but that has been withdrawn from the market was the Retnar system by Kodak, which used copies of 43 mm x 50 mm.

In the following we describe only the Delcomat and LogEtronics systems.

2.1. The Delcomat System

The Delcomat system comprises three parts:

1. The Delcomat copying machine. The Delcomat copying machine allows the user to make copies (100mm x 100 mm) from radiographs up to a size of 350 mm x 430 mm and can be operated under daylight conditions. The unit has a copy-repeat facility for making any number of copies of one original.

2. The Odelcamatic processor. The Odelcamatic processor provides fully automatic processing of Delcomat copies at a rate of 150 copies/hr.

3. The Odelcard filing system. The Odelcard filing system exists only in some pilot installations(7,8) and has not been introduced on the market. Copied films and case reports are kept together in a special plastic envelope. This envelope is punch coded to leave circular holes in the top, corresponding to date of birth, date of arrival in the hospital, diagnosis, etc. The envelopes are stored in a cabinet from which they can be retrieved by a mechanical system addressing the punched holes by a couple of pointed rods.

For the copies, the Delcopex RP film (Agfa-Gevaert) is recommended. The film is positive working and hence the copies are of the same polarity as the original radiographs.

2.2. The LogEtronics System

The key components of the LogEtronics system are the minification camera and the projection viewer (an automatic processor is provided as well):

1. The minification camera is suited for copying radiographs up to 380 mm x 460 mm (transmission mode) and case reports (reflection mode). In the transmission mode, the camera has the unique feature of automatic dynamic range reduction by means of an electronically


driven unsharp-masking method. The intensity of the cathode-ray tube (CRT) spot used to illuminate the microfilm is controlled by the average density in the vicinity of the point under consideration. The benefit of this method is twofold: • All parts of the picture can be examined at the same illumination

level, because they are all recorded in the same linear portion of the characteristic curve,

• The detail contrast is enhanced as compared to an unmasked reproduction.

2. For a 35-mm copy, a projection reviewer is indispensable and has been developed by LogEtronics for viewing the minified radiographs at original size.

The image produced on the record film (Agfa Copex Pan Rapid or Ilford HR 31) is of reversed polarity as compared to the original radiograph; it is therefore referred to as a "black-bone" master, in contrast to the "white-bone" original.

Two ways to organize files have been evolved:

1. Producing black-bone copies immediately after the evaluation of the original radiograph; only one file is necessary.

2. From each black-bone master a white-bone copy is made. Two files are kept: the black-bone master resides in a security file; in the second file are kept the white-bone prints, which are made available for reevaluation or loan service activities.

2.3. Experience with Minification Systems

The performance of minification systems should be evaluated in terms of the requirements 1-6 listed in Section 1, with emphasis on the first requirement, i.e., small size and preservation of information. However, it seems that in the literature major attention has been focused on cost and simplicity of handling (items 3, 4, and 6). As a general comment on the studies performed with both minification systems, Delcomat and LogEtronics, it must be noted that the results obtained by different authors are by no means unequivocal.

2.3.1. THE DELCOMAT SYSTEM

The Delcomat system has been qualified as being economically attractive and easy to handle by many of its users. (8-12) The contrary has also been stated.(13) Deininger et alY4) performed a study in which radiologists

328 U. Killat

first judged the copy and then the original. As a result, it was found that the diagnosis from the copy was false in 3 % of the cases and more difficult to arrive at in 13% of the cases, which covered 15 types of radiographic investigations.

Twice the number of false negative diagnoses as compared to full-size radiographs has been reported by Taenzer and Schneider.(1S)

2.3.2. THE LOGETRONICS SYSTEM

From the viewpoint of maintenance and cost some authors feel content with the LogEtronics system,(16,17) whereas others complain about cumbersome filing procedures(12) and high cost. (13,15) Ruygrok(18) noted a marked edge-enhancement effect that may lead to errors of diagnosis in investigations of the thorax and intestines. An overexposure at the rim of the radiograph causes information losses that are particularly annoying in mammography. According to Taenzer and Schneider,(1S) the reduction in diagnostic quality is comparable to that found with the Delcomat system.

ROSS(19) comments on some of the criticism and recommends only the use of a black-bone file (see Section 2.2). The missing second copying step will considerably improve image quality. In addition, space and time for running the second file is saved. It must be noted, however, that this advice also relinquishes the major advantage of the double-file policy: absolute integrity in the safety file.

2.4. Physical Limits for Minification Systems

Any storage system can be considered as an information channel for transporting information from a source, i.e., the original information, to a destination, i.e., the reproduced information.

With respect to the storage of radiographs, the information loss may be described in terms of false diagnoses, as it has been done in Section 2.3 for both the Delcomat and LogEtronics systems.

non- Ideal 0- channel -0 --t

Figure 1. Visualization of signal impairments in photographic processes. (a) Amplitude limitation; (b) addition of noise; (c) damping of high-frequency components.

Modem Optical Methods for the Storage of Radiographs 329

However, for the systems designer it seems to be more helpful to classify information losses in terms of the model of a non-ideal transmission channel. The signal transfer for an analog input is shown in Figure 1 for a one-dimensional example. The comparison of output and input signals reveals three types of impairments relevant to photographic minification processes:

(a) Amplitude limitation or nonlinear behavior, (b) Addition of noise, (c) Damping of high-frequency components

A brief discussion of these three aspects will reveal the physical properties to which minification systems must adhere to in order to make information losses tolerable.

2.4.1. LARGE AREA SIGNAL TRANSFER

The linear or nonlinear behavior of a film can be read from the characteristic curve relating the final density to the exposure (which, in turn, is inversely proportional to the negative of the density of the original transparency) .

2 3 ----- log exposure

input signal

Figure 2. Signal transfer at the characteristic curve for Delcopex RP film. The characteristic curve has been taken from Bollen in ref. 20.

330 U.KiDat

For the positive working Delcopex RP film, this curve is shown in Figure 2. The derivation of an output signal from the input signal from this curve is known as "large area signal transfer." This model considers only details that are large enough to not influence each other. If the curve of Figure 2 were a straight line with slope -1, then each density value of the original could be transferred to the same value in the final print. The S-shaped behavior of the curve, however, results in a clipping of the extreme density amplitudes of the original. As the density range examined in radiographs covers at least three density units, a linear portion of this range would be desirable. Nevertheless, the majority of radiographs exhibit a lower density range of, say, two density units, and for these purposes the Delcopex RP film is a reasonable choice.

In the LogEtronics system, the problem of nonlinearity is circumvented by mapping great density differences to a small, linear portion of the characteristic curve. This is, however, accompanied by a loss of the lowfrequency content of the image, owing to the process of harmonization.

2.4.2. NOISE

A homogeneously illuminated film exhibits density fluctuations (granularity) resulting from the statistical properties of the silver grains in the emulsion. Storage of a picture involves combining two noisy processes: The noise in the original radiograph is added to the noise of the storage medium. As both noise sources are independent of each other, their variances are additive and therefore

2 2 S2 2 CT tot = CT prim + CT star (1)

where CT is the rms noise and the subscripts prim and stor stand for primary noise and storage film, respectively. The factor S2 results from the fact that the storage film, and hence its noise patterns, have to be magnified by a linear factor of S in order to get a full-size reproduction. Equation (1) is somewhat simplified in the sense that it neglects the transformation of the primary noise by the nonlinear characteristic curve and the modulation transfer function of the storage film.(21) However, it provides some evidence that the granularity of the storage film (measured in some appropriate way, e.g., granularity parameter G or Selwyn coefficient(21» should be lower than the granularity of the original by at least a factor of 4S2. This constraint can barely be satisfied by the Delcomat and LogEtronics minification systems. As an example, we mention that for Delcopex RP film, G is an order of 0.5 f,Lm2,(20) which is one to two orders of magnitude lower than the values found with radiographs (see Section 3).


2.4.3. RESOLUTION

The damping of high-frequency components causes a reduction in spatial resolution. To reproduce all details of the original, the resolution capabilities of the storage system should exceed those of the original by a factor of two or so.

Ruygrok(18) investigated the minification systems with respect to resolution and obtained the following results. The LogEtronics system shows a resolution of 5 Ip/mm at the margin and 7 Ip/mm in the center of the picture. For comparison, the values for the Delcomat system are 9 Ip/mm and 10 Ip/mm, respectively. In view of the resolution of about 5 Ip/mm found with frequently used film-foil combinations, Ruygrok therefore deduced a preference for the Delcomat system. However, the determination of the tail of the modulation transfer function (MTF) must not be overestimated, because most frequently in medical radiography the detection of small details is not MTF limited, but noise limited. (22)

2.5. Conclusion

The storage and retrieval of x-ray images poses a serious problem for most hospitals. A first attempt to reduce the problems of space, time, and cost has been the development of minification systems, two of which have been marketed in Europe and the United States: the Delcomat system and the LogEtronics system. They seem to be tailored to different views regarding the purpose of the stored copy, and there seems to be no consensus among radiologists in this respect:

1. If the radiograph should remain suitable for a renewed diagnostic interpretation of all details (including those that were not relevant to the original examination), then no observable loss of pictorial information is tolerable. Any image processing before storage is not allowed. Delcomat then seems to be the best compromise.

2. If the radiograph is considered as additional but not indispensable morphological information to the x-ray report to confirm a diagnosis, then the loss of irrelevant information and the enhancement of relevant details is seen as an advantage. Consequently, the LogEtronics system will be preferred.

A replacement storage system must be of the same or better quality than the system it replaces. In Section 1 we have set up some criteria for judging storage systems and we will reconsider them with respect to the two systems discussed so far.

332 V.KiDat

1. The space savings of 90% for the Delcomat system and of 99% for the LogEtronics system, are a clear advantage over direct archiving. False diagnoses on the order of a few percent do occur with both systems when evaluating the final prints. It is a question of personal view whether to consider this to be tolerable rather than unacceptable. The physical reasons for the occurrence of information loss are mainly the lack of a linear characteristic curve for the recording material and, to a lesser extent, the addition of noise and limited resolution.

2. None of the manufacturers provide an efficient retrieval system. 3. Once the 100 mm x 100 mm format is accepted for direct viewing,

the Delcomat system provides a simple solution for producing copies.

4. Permanent integrity can be achieved by keeping two files, a security file and a working file. Owing to its lower space requirements, the LogEtronics system is favored as a solution.

S. Integration of case reports and radiographs is possible with both systems but is more convenient with the LogEtronics system. A more comprehensive integration in other information systems has not been considered by any of the manufacturers.

6. The estimate of the total cost, including investments, material cost, and personnel for copying work, varies from author to author. The increasing prices of silver and the possibility of silver recovery is, however, in the long run a striking argument for minification systems.

In summary, the performances of present-day minification systems are not fully satisfying. Indeed, their use in hospitals is very limited. Enquiries in the United States concerning the reasons thereof showed that practical considerations such as the cost and the difficulty of viewing predominate over legal status or even the preservation of diagnostic information. (23) For the Federal Republic of Germany, reference lists report a delivery of 66 LogEtronics systems and 88 Delcomat systems.

Minification systems are a means for solving the space problems of hospitals, but they do not meet the remaining requirements for a modern storage and retrieval system.

Therefore other, perhaps less obvious, solutions have been investigated on a research basis. The problem seems to be well defined, yet intricate enough to be intellectually challenging. Two different approaches are discussed in Sections 4 and 5. In the following section, we introduce some basic concepts and notions to facilitate the understanding of later sections.


3. STORAGE OF RADIOGRAPHS: THE INFORMATION THEORY APPROACH

Before discussing new methods for storing radiographs, it would be desirable to define the properties to which storing devices must adhere in order to do their job. With respect to the crucial point of image quality preservation, the question arises as to what amount of information stored in a radiograph has to be consequently stored in a new storing device.

Obviously, the answer to this question depends on the specific radiograph chosen and must be answered individually for each radiograph. Therefore, it is perhaps more reasonable to ask for an upper bound of information content: What is the information capacity of a radiograph?

The first analysis of information capacity was given by Shannon,(24) who developed some powerful theorems within the context of electrical communication channels. The adaptation of this work to the photographic recording of information is challenging and not a straightforward task; all models proposed so far lack either mathematical accuracy or applicability to easily measurable quantities. We shall adopt here an intuitive model,(25) rather than a rigorous theory, to derive some instructive estimations. For a more thorough discussion of the subject the reader is referred to some further references. (22,26,27)

3.1. Information Capacity of Radiographs

The reasoning given here assumes the following model of a photographic layer as an information storage device. The unit area of the film is divided into n picture elements of size (.1X)2. Such a picture element is distinguishable from its neighboring elements in that its density exceeds that of its surroundings by a certain value !:t..D. Within a given maximum density range D2 - Db m steps may be distinguished on the basis of a criterion given below, and by analogy to Shannon's approach the information capacity C is defined as

C = n log2 m (2)

A one-dimensional illustration of this model is shown in Figure 3 for n = 6 and m = 4. Obviously each level is subject to uncertainty because of the noise superimposed on the true signal. Noise can mimic a signal, and the problem of level separation may be formulated in this way: To what extent must the signal exceed the noise in order that the probability be sufficiently high for the registered signal to be a true one and not simulated by some random process? Thus the problem has to do with the very nature of noise and its statistical properties, and it may become very intricate in its details.

334 U. Killst

D i t kG'

T "!::"x ..

___ x

Figure 3. One-dimensional illustration of the cell model for elementary determination of information capacity.

The general answer, as developed by Rose, (28,29) can, however, be stated in a simple form: The smallest signal difference flDmin by which a detail of size A can be detected from the background must exceed the rms fluctuations in A by a factor of k:

flDmin = ku A (3)

where U A denotes the standard deviation of density fluctuations evaluated in A, and k ensures some level of probability P for the signal to be true. Assuming Gaussian noise, one has a probability P > 1-10-6 for k = 5 and this is a frequently used value. (29)

For the number of gray tone levels dm resolvable in an interval dD we have

(4)

The number of gray tone levels assigned to a picture element of size A in a radiograph with a density range D2 - Dl is straightforwardly obtained:

fD 2

m = 1+k- l u-;/(D)dD D,

(5)

Since n = A -t, equation (2) can be rewritten as

C = A -llog2 (1 + k-l fD 2 uAl (D) dD) D,

(6)

For radiographs, the maximum signal excursion D2 - Dl cannot be considered to be independent of the cell size A: The linear dimensions of


small objects shown in a radiograph are of approximately the same size in the direction of the x-ray beam and perpendicular to it. As the observed density difference is proportional to the thickness in the beam direction and to the difference 1:J.p, of linear attenuation coefficients, one obtains

(7)

where 'Y is the slope of the characteristic curve in the region of interest. Investigation of equation (6) reveals that the 1/ A term dominates the logarithmic term: Binary coding gives the highest information capacity. (TA(D) can be determined experimentally. It would, however, be desirable to have a measure of noise that is independent of the aperture A. Such a convenient measure of noise properties is the Wiener spectrum W describing the noise power spectrum. (22) (T A is related to W by virtue of equation (8):

(T~ = J t: W(u, v)IT(u, vW du dv (8)

where T is the optical transfer function of the aperture of size A. A determination of C from experimental Wiener spectra is given in the next section.

3.2. Information Capacity of Radiographs: Example

In the following, equations (4) and (6) are evaluated using some experimental data of a typical film-foil combination (Curix RP1-Cawo Uni).

A Wiener spectrum for a certain density level is shown in Figure 4. The frequency dependence has been understood since the work of Rossmann: (30,31) The low-frequency maximum is determined by the fluctuations of the absorbed x-ray quanta obeying Poisson statistics. This white noise input is low-pass filtered by a diffusion effect in the screen. The result is a bell-shaped curve with a base constituted from the grain noise of the film emulsion. To a first approximation this spectral dependence is the same at all density values. Under this assumption the D dependence of W can be derived from the low-frequency maximum Wo alone, for which a simple relation holds:

w: 2 _-1 0- 'Y q (9)

where q is the average number of absorbed quanta and 'Y is the slope of the characteristic curve at the respective exposure. With the help of equation (9), Wo is related to the exposure and-via the characteristic curve-to the final density. Figure 5 shows such a plot. The influence of film grain noise has been included in the full curve.

336 U. Killat

I 10 -+-----------------WO

5

2 3 4 5 ____ f Ilp/mmJ

Figure 4. Wiener spectrum of film-foil combination (Curix RP l/Cawo Uni) for D = 0.4.

20

----------------------°O~------------------~--------------------~2-------

-0

Figure 5. Calculated Wiener spectrum at origin. Woo as a function of density for the film-foil combination Curix RP l/Cawo Uni (fuil curve). The dashed curve gives the contribution of the film granularity of Curix RP 1 and relies on measurements of Bol\en in ref. 20.

Modern Optical Methods for the Storage of Radiographs

100

dm dO

1 50

0~--------~--------~2~--------*3------

- °specular

337

Figure 6. Differential number of gray tone levels calculated from a noisy image [see equation (4)].

Thus uA(D) is fully determined by equation (8) and Figures 4 and 5. Taking, in equation (7), !l.1L = 5 mm -1 (corresponding to the bone-tissue contrast at 50 ke V), C has been evaluated for our example and yields, for binary coding, a capacity C = A -1 of about 600 bit/mm2 •

In view of the limited geometric resolution of the radiographic imaging process and the linear dimensions of radiographic objects, binary coding does not seem to be a realistic model. A more conservative cell size of, say, 125 ILm x 125 ILm is more appropriate. With this value an information capacity of C = 210 bit/mm2 is derived and we end up with roughly 2 x 107

bits for a normal-size radiograph (30 cm x 30 cm). This value will be taken as a guideline in Section 5.

For the study of gray tone preservation, an evaluation of dm/ dD before and after the storage process is useful. For use in Section 3, a plot of the differential number of gray tone levels evaluated for an aperture of size A = 1 mm2 is shown in Figure 6. This distribution will be referred to as the "primary distribution of gray tone levels" for the Curix RPI/Cawo Uni system. Density values refer to the specular density, because this quantity is relevant to the holographic storage process.

3.3. Information Content, Information Capacity, and Image Quality

The information capacity has been introduced as an upper limit to the information stored in a recording meqium. The actual information content

338 U. Killat

will be lower: a continuous signal, with each sample value depending upon neighboring sample values, is expected to have a lower information content than a random signal. On the other hand, our cell model of the photographic layer is not fully tailored to the way in which information is coded in a radiograph. Consequently, a real storage device that behaves as our model would spoil image quality, even though it has a sufficient information capacity. Since some readers may find this confusing, we want to shed some light on this problem because it is important for the scanning devices and hard copy units dealt with in Section 5.

From the calculation of information capacity in Section 3.2 we can show that a picture element (pixel) of 125 I-'m should be coded with 4 bits. One may then argue that a radiograph coded in this way should be reproduced without any notice~ble information loss.

This is not true! In the cell model all density values are equally probable at any pixel; no correlation between neighboring pixels is assumed. In radiographs these correlations exist because the typical objects in radiographs are larger than the resolution limit. For these extended objects, noise is averaged out and the number of resolvable gray tone levels increases [see equation (5)]-though the information content per unit area is not increased!

The point we want to make is best described with reference to Figure 7. Figure 7 shows a trace along several resolution elements together with an attempt to quantize this signal. As in Figure 3, the number of quantization steps is deduced from the noise evaluated in a single cell. However, owing to the strong correlation of the picture elements, the smooth (dashed) line becomes recognizable. Conversely, the staircase approximation of the smooth curve is felt to be a severe falsification of the true signal. The same unwanted effect of pseudocontours is shown in Figure 8, where an actual radiograph is coded with 4 bits (a) and 7 bits (b) per pixel.(32l

This discussion stresses the importance of coding for all problems of information storage, because there is in general no fundamental correlation between information capacity and image quality.(33l Thus the number of

o

I __ resolution cells

Figure 7. Effect of quantization noise. Quantization steps are deduced from the rms values of noise within a resolution cell.


0) b)

Figure 8. Reproduction of a radiograph from a 2048 x 2048 matrix: (a) 4 bit/pixel; (b) 7 bit/pixel. From M. Schlindwein in ref. 32.

bits per pixel necessary for the digitization of radiographs as described in Section 5 can best be determined experimentally with test pictures as shown in Figure 8. Coding techniques can then be used to eliminate redundancy and to reduce the apparent number of bits per pixel to values on the order of or even less than those derived in Section 3.2. To this end, source coding techniques are used for the digital storage of radiographs.

Another aspect of the term "information content" should also be considered: The information of a radiograph is in general evaluated by a human observer. The diagnostically relevant information content is therefore limited to what the radiologist actually sees. From the viewpoint of information theory the criterion for the detectability of a signal is that it exceed the rms noise to some degree. Since the work of Rose, it is known that the eye approaches this model of an ideal pickup device, which is only limited by the noise of the scene it is looking at. This statement is no longer true when very faint contrasts are approached: The approximate contrast threshold of the eye is W = 0.01. Thus there is a difference between visibility and detectability. It has been pointed out by Motz and Danos(34) that there exist density differences in conventional radiographs that meet the detectability criterion but that are obscured to the human being and therefore lost for diagnosis. Here is a possibility to extend the present diagnostic capabilities by contrast-enhancement techniques. With respect to the storage of radiographs, it would be desirable that these "invisible" signals be preserved. The testing of storage methods by visual inspection of test pictures is therefore not always a sufficient criterion for judging future-oriented systems.

340 U. Killat

Figure 9. Photograph (left) and fine stlUcture (right) of a hologram.

4. HOLOGRAPHIC STORAGE

The bottleneck of minification systems is the linear reproduction of the gray scale range of the original radiograph (see Section 2).

Therefore, the key issue in the holographic recording of radiographs is that the recording material is exposed to an almost homogenous exposure: Figure 9 (left) shows a photograph of a hologram plate. The fine structure of the hologram is revealed by a microscope [Figure 9 (right)] and may contain only small signal amplitudes even if there are great brightness differences in the reconstructed scene. Thus the problem of objects producing huge exposure differences seems to be circumvented by this technique.

Another advantage-as compared to conventional microfiches-is that holograms have proven to be insensitive to scratches or dust in the reconstruction process.

Baily et al. (35.36) introduced the holographic method for the storage of x-ray images. An example of their work, a reconstruction of a radiograph, is shown in Figure 10. More sophisticated methods for hologram recording and reconstruction were believed to further improve image quality. Holographic techniques were likewise proposed for data storage. and a remarkable state of the art has been developed in this field. (37) Less attention has been paid to a thorough evaluation of the gray scale performance of the holographic method, which is of utmost importance for the application considered here. In the following sections we will investigate the holographic method with respect to its use in a storage system of radiographs. (38,39) The principles of holography are documented in several books. (40-42) Nevertheless, on the assumption that there are readers unfamiliar with the field, a concise review is offered in Section 4.1.


Figure 10. Real holographic image of a radiograph of a human hand; diameter ofthe hologram: 1 mm. From N. A. Baily et al. in ref. 36.

4.1. Description of the Holographic Method

Holography is an interference method of recording light waves that have been scattered or transmitted by an object illuminated with coherent light. The principle of holographic recording is shown in Figure 11.

A monochromatic wave a illuminates the object transparency and travels to the recording plane. In the recording plane it is superimposed on a so-called reference wave R that has been derived from the same laser source as the object wave O. Owing to the high degree of coherence of the two waves, an interference pattern will form in the volume of space where the two beams overlap. The photographic record of this pattern, the hologram, contains sufficient information about the object wave to permit its reconstruction. The hologram can be considered as a sophisticated

342

..---..-..-t I

x

1 o

..--...---- 1 L

! SCREEN

I

d

U. Killat

.. z

Figure 11. Recording and reconstruction of a hologram. 0, object beam; R, reference beam; R', reconstruction beam; T, object transparency; H, hologram; I and L are the linear dimensions of the hologram and the transparency, respectively. The dashed lines denote the reconstruction process.

grating with some remarkable properties: Upon illumination by a wave R' making an angle (+0) with the positive z axis, a real image of the object is displayed at a distance d from the hologram. In addition, other waves are created, in particular a broad zeroth order that can be separated from the desired reconstructed image provided that the recording angle 0 is sufficiently high.

The maximum angle between a ray from the edge of the object transparency and the reference beam also determines the maximum spatial frequency F contained in the interference pattern. It turns out that a factor of 2 has to be applied to the space-bandwidth product of the object to obtain that of the recorded hologram. Thus F is related to the maximum spatial frequency f contained in the object transparency by

F=2Sf (10)

where S = L/ I is the linear size-reduction factor. Thus the required resolution derived from our application is on the order of 100lp/mm (f =

2Ip/mm, S = 10), which is far below the resolution limits of typical recording materials (see Section 4.3.2).

For a more detailed description of recording geometries, their properties, and their different applications the reader is referred to the literature quoted above.

4.2. Concept of a Holographic Storage System

A recording geometry, which is particularly suited for a holographic storage system of radiographs, is displayed in Figure 12: The radiograph

Modem Optical Methods for the Storage of Radiographs

al

bl

opaque screen

radiograph

reconstructed real image

recording medium

--8----7-/.7

// //

// //

// //

// // /

R'

343

Figure 12. Holographic storage system: (a) Recording process; (b) reconstruction process.

to be stored is placed behind the lens L and illuminated by the object beam, which can form different angles with the optical axis of the system. This object beam is then superimposed on the reference beam in the storage plane. The storage plan in Figure 12 is a rotating disk containing many holograms. Retrieval is accomplished by illuminating the processed hologram by the beam R' [Figure 12(b)]. An individual hologram is addressed by rotating the disk and controlling the orientation of the mirror M. The real image is always displayed at the same place and can be observed on a screen or by means of a TV chain.

Using a juke-box-type arrangement of several disks, a mass memory with simple and rapid retrieval of information can be built provided that

1. A storage material suitable for simple in situ handling is available.

344 U.KiDat

2. The image quality in the reconstructed image is sufficient at sizereduction ratios at least comparable to those found with minification systems.

It is the second, more fundamental, aspect that will be extensively discussed in the next section. The implications of this analysis with respect to recording materials are examined in Section 4.3.2. In this context, the in situ handling of the storage media (concern (1» will also be discussed.

4.3. Image Quality in the Holographic Recording of Radiographs

An example showing the good gray tone capabilities of holographic reconstructions is given in Figure 13. Figure 13(a) shows the original transparency: a radiograph of a human orbit, with an accompanying gray tone tablet. Figure 13(b) gives the reproduction of the same object from its image on a high-resolution film. Figure 13(c) shows the holographically reconstructed image of the original in Figure 13(a). The same high-resolution plate has been used to record both the photograph of Figure 13(b) and the hologram of Figure 13(c). In both cases the storage area equals that of the original transparency. The reproduction of gray tone levels in Figure 13(c) is clearly better than that in Figure 13(b).

Although the result of Figure 13(c) is very encouraging, up to what size reduction ratios used in the holographic storage process can such an image quality be preserved? Indeed, a degradation of image quality upon reproduction from small holograms is observed, as is demonstrated in Figure 13(d). However, such experimental evidence is only of limited value; it is necessary to evaluate the physical constraints that determine image quality and to compare these models with the experimental findings. As mentioned above, the resolution requirements for the holographic medium do not pose problems. Thus, as in the case of direct minification, the crucial point is the preservation of gray tone levels. In the following, we make an attempt to relate physical properties of the storage material to gray scale performance in the final reproduction.

4.3.1. GRAY SCALE REPRODUCTION

For faithful gray tone reproduction a storage system must

1. Exhibit a linear characteristic of large area signal transfer (see Section 2.4.1)

2. Add negligible noise so that faint signals are not lost in noise.


(a) (b)

(e) (d)

Figure 13. Reproduction of a radiograph (human orbit) and a gray scale. (a) original transparency; (b) reproduction after intermediate storage on a high-resolution film; (c) reproduction after holographic storage on the same high-resolution film; (d) reconstruction of the radiograph in (a) from a hologram at S = 20, with nom,lifiuse illumination of the object.

346 U. Killat

To understand these requirements in terms of holographic storage, we must examine the holographic reconstruction process in greater detail.

For faithful wave front reconstruction the hologram must contain a true replica of the original wave front in terms of an amplitude transmittance modulation. (For simplicity we neglect phase modulation.) This implies that the recording material must exhibit a linear transmittance exposure characteristic. As in the direct photographic process, we are concerned with a signal transfer problem, and again linear transfer curves are in general not available. To avoid the resulting image distortion, (40) linear recording is achieved by restriction to small signal amplitudes. Small signal behavior is ensured when the intensity of the reference beam is far greater than that of the object beam. The price to be paid for the resulting good behavior is a reduced diffraction efficiency, which is disastrous with respect to the noise problem considered next. For the ensuing discussion linear recording is assumed.

4

--- Dsppcular

Figure 14. Gray tone reproduction of holographic recording: The diagram shows the reconstruction intensity I as a function of specular density. 0: beam ratio 20: 1, 11 = 0.22%, <ll = 6.5 X 1010 (Ip/mm)-2; /:::": beam ratio 200: 1, 11 = 0.026%, <ll = 5.3 X 1010 (ip/mm)-2. The shaded areas show the range of the results of equations (11) and (12) within experimental error.


The reconstructing beam R' illuminates the hologram and generates an image of the original on a screen. The light distribution corresponding to an ideal noise-free reproduction may be labeled Ii. In passing through the hologram the light is also affected by the holographic medium itself, i.e., its microscopic structure. At the screen one therefore observes an additional intensity In originating from the random nature of the storage medium. (For a film emulsion these noise sources are primarily the film grains.) This noise light In will degrade the gray tone performance of the holographic system with respect to both requirements (1) and (2).

4.3.1.1. Large Area Signal Transfer. Large area signal transfer has been studied (38,39) using a gray tone wedge ranging from 0.1 to 4 units of specular density. The real image reconstructed from the hologram was scanned by a photosensitive detector. As a result, Figure 14 shows the negative logarithm of the relative brightness of the reconstructed gray steps plotted against the original gray scale. In such a plot, an ideal reconstruction would yield a straight line of slope 1 passing through the origin.

In Figure 14 two results with different reference-to-object beam ratios are shown. For a beam ratio of 20: 1 the dots lie on an S-shaped curve rather than on a straight line. The deviations for low densities (high intensities in the reconstruction field) are due to nonlinear behavior in the hologram, the dynamic range of the interference pattern still being too large. This interpretation is strengthened by the fact that for a beam ratio of 200: 1 (triangles) the low-density deviations are completely suppressed.

However, this suppression is accompanied by increasing saturation effects for higher density values.

The explanation is straightforward: With increasing beam ratios, the diffraction efficiency decreases and the signal light intensity Ii is reduced, whereas the noise light level In remains constant. For Ii «In only the constant value In will be measured.

For a quantitative explanation of the experimental results we have to calculate the average light intensity of an image point in a reconstructed gray step. The average light intensity I has two contributions: the average signal intensity and the average noise intensity:(40)

(11)

where the bars denote spatial averages performed within the region of one gray step of the gray tone wedge. For faithful gray scale reproduction, noise light intensity must be small compared to signal light intensity.

Low noise light levels are observed when

1. The overall noise power scattered from the holographic medium (and expressed in terms of the scattered flux spectrum <1» is small.

348 U. Killst

2. The fraction of noise light falling into the aperture L/ d of the (reconstructed) radiograph is small.

High signal intensities result from

(i) High diffraction efficiency T/ of the hologram. (ii) High relative transmittance T/Tav of the considered image point as

compared to the average transmittance of the object transparency.

The exact expression for In/Lin a resolution element of the reconstructed object of size L 2 has been derived for light of wavelength A (38):

- 2 In ~L Tav -=-=---2-Ii T/(Ad) T

(12)

Equation (12) is adequately interpreted by items 1,2, (i) and (ii) that which we have intuitively introduced above. Note that there is no dependence on hologram size.

Equations (11) and (12) have also been evaluated quantitatively using experimental values for T/ and~. Owing to the measurement inaccuracies in T/ and ~, these results are given as shaded areas lying slightly below the experimental values. It is seen, however, that the theory agrees quite well with the experimental results. This confirms our contention that it is the noise contributed from the holographic medium that considerably limits the density range achievable in the reconstructed image.

An extension of the density range-particularly by avoiding the lowdensity losses due to nonlinear behavior-can be accomplished using diffuse illumination; (35,36) a diffuser inserted between the lens and the transparency in Figure 11 leads to a nearly uniform intensity in the holographic plane. The improvement obtained under these conditions becomes obvious from a comparison of the upper curve (beam ratio 20: 1) in Figure 14 and the corresponding result obtained using a ground glass, as shown in Figure 15. The latter curve exhibits a much longer straight portion.

A diffuser, however, inevitably introduces speckle with serious consequences for the gray level resolution. We discuss these consequences in Section 4.3.3. Nevertheless, the result of Figure 15 might be relevant to more sophisticated forms of illumination such as random phase masks, (43,44)

These devices avoid speckle formation while still preserving the advantages of dynamic range reduction and redundant recording. Unfortunately, they have not yet been realized for large image areas.

Summarizing the results of this section, we conclude that holographic storage methods can be expected to exhibit a linear gray tone transfer in


-- Dspe,ular

Figure 15. Gray tone reproduction of holographic recording, with diffuse illumination of the object: The diagram shows the reconstruction intensity I as a function of specular density. 0: beam ratio 20: 1, 11 = 1.1 %, <I> = 8.3 X 1010 (Ip/mm)-2. The full curve gives the results of an evaluation of equations (11) and (12).

a dynamic range of roughly three units of specular density. In the next section we restrict our attention to this linearly reproduced density range and investigate signal detection in the presence of noise.

4.3.1.2. Signal Detection and Number of Gray Tone Levels. In Section 4.3.1.1 we have avoided the term "signal-to-noise ratio" (SNR) for the quantity l;/ In. The reason has to do with the coherence of light. The noise light not only lowers the contrast in the reconstruction but coherently adds to the signal light. The resulting interference pattern has a grainy appearance owing to the random phases of the noise signals. An approximate measure of SNR is (Ii In)1/2. The influence of noise is shown in Figure 16. Figure 16(a) shows an enlarged portion (lOx) of a neutral-density filter (Kodak Wratten) what was part of the coherently illuminated scene. In Figure 16(b), the reconstruction from a hologram recorded at the reduction ratio S = 20 is given. The aperture sub tended by the hologram was the same as that of the optical system used to record Figure 16(a). The reconstructed image exhibits a grainy appearance, which makes it difficult

350 U. Killst

a) b) c)

Figure 16. Reproduction of a neutral-density filter (Kodak Wratten) with a linear enlargement of 10. (a) Original filter, coherently illuminated; (b) reproduction from a hologram at S = 20; (c) reproduction from a hologram at S = 20 recorded under diffuse illumination.

to recognize small differences in gray tone levels in a picture or determine their boundaries. In order to avoid misleading the reader, we repeat that the noise in Figure 16(b) is contributed by the storage material and must not be confused with speckle noise [see Section 4.3.3 and Figure 16(c)].

The problem of signal detection from a noisy background has been introduced in Section 3.1. Noise has been found to limit the number of gray tone levels that can be distinguished in an area of size A. For the reader's convenience we rewrite the differential number of gray tone levels:

dm 1 -=--dD kO"A

(13)

This equation can be applied to the holographic reconstruction process, provided that (T A includes both the primary noise of the radiograph and the noise introduced by the holographic storage medium. The analysis(38) leads to the following expression for the differential number of gray levels found in an area of size A (sufficiently large to contain many resolution elements):

dm (A)1/2

dD = k[Wo(D) + S2w(D)]1/2 (14)

where Wo is defined as in Section 3.2 and

2 cI>(1 O-D)av w(D) = 2(loglO e) 71 lO- D (15)

Evaluation of equation (14) before (w = 0) and after the storage process (w ¥- 0) gives a measure of the information loss occurring in holographic storage. This is demonstrated in Figure 17: The dashed curve is replotted


from Figure 6 and shows the primary distribution of gray tone levels determined from the Curix RP1/Cawo Uni combination. The full curve shows the final distribution of gray levels calculated from equation (14). One sees that the initial distribution is drastically changed and dn/ dD after the storage process is essentially governed by holographically introduced noise, its influence increasing with increasing density values. Evaluating Jedm/ dD)dD for the dashed and solid curves, we find that about 30% of the information content has been lost in the storage process. If one reduces the requirement for the linear size-reduction ratio (S = 10), the full curve will get closer to the dashed curve. A deviation of the two curves of 10% at most can be obtained for S = 2. However, such S values cannot be of great practical interest!

For radiographs taken without intensifying screens, the situation is worse; they exhibit less primary noise and the effect of the holographically introduced noise will be even more pronounced.

An improvement over the unpleasant situation described in this section can be expected from materials with 11/$ values, which considerably exceed those of conventional silver halide emulsions, on which the results of Figure 17 are based. This leads us to an overview on holographic recording materials presented in the next section.

100

-----------

°O~--------~---------*2----------~3

~ Dspocular

Figure 17. Number of gray tone levels calculated from a noisy original. Dashed curve: primary distribution taken from Figure 6. Full curve: distribution after holographic storage calculated from the dashed curve and equation (14) with Tav = 0.1,1) = 1.1 %, <I> = 6 X 10-10 (Ip/mm)-2, S = 10, and A = 1 mm2 •

352 U. Killat

4.3.2. HOLOGRAPHIC RECORDING MATERIALS

A hologram must either change the amplitude of the incident light field (absorption hologram) or its phase (phase hologram). Absorption holograms, most frequently obtained from silver halide emulsions, are limited to diffraction efficiencies below 6.25%. Phase holograms, in which the optical path length varies spatially owing to a surface relief or a refractive index modulation, achieve maximum diffraction efficiencies of 33.9% (thin hologram) and 100% (volume hologram). Volume holograms differ from "ordinary" thin holograms by their finite thickness and their capability of storing a great number of holograms in the same volume.

With respect to the problem of storing radiographs, the following requirements become obvious from our discussion:

1. No limitation of resolution below 100 Ip/mm (Section 4.1). 2. 'TI/4> values on the order of 5 x 108 (lp/mm)2 [equations (14) and

(15)]. 3. In situ handling.

The last concern contains two aspects:

1. Unexposed areas of the storage medium must not be affected by the recording process at some other place of the disk.

2. Rapid in situ processing must take place without affecting neighboring parts of the storage disk.

The severest constraint set by our requirements concerns the scattered flux spectrum 4>. Unfortunately, quantitative data are only available for a few materials. People frequently measure signal and noise from test objects made up of a ground glass and a central opaque spot. The ratio r of reconstructed intensities in the two areas is a convenient (but arbitrary) measure of signal and noise irradiances. (45)

In Table 1 we have collected data for some materials that could be of interest for our purpose. If not otherwise stated, the data are taken from a comprehensive review by Biedermann. (46) Some care must be taken with respect to diffraction efficiency values. To avoid nonlinear recording, practical values are always limited to a few percent. In view of the requirements above, it is readily seen that none of the materials actually available are suitable for our purpose, though all of them have sufficient resolution. The best candidates with respect to noise performance are dichromated gelatin and photopolymers. From the viewpoint of handling, photothermoplastics are attractive: They allow the selective addressing of storage areas by a charging process and rapid in situ development. (47,48) Unfortunately, image

Mat

eria

l

Silv

er h

alid

e em

ulsi

on

(Agf

a 8

E 7

0)

Ble

ache

d em

ulsi

on

(Kod

ak 6

49F

) D

ichr

omat

ed g

elat

in

Pho

toth

erm

opla

stic

P

hoto

poly

mer

s E

lect

ro-o

ptic

cry

stal

s (m

etal

-dop

ed L

iNb

0 3)

Typ

e

Abs

orpt

ion,

thin

Pha

se,

thin

/vol

ume

Pha

se, t

hin

Pha

se, v

olum

e P

hase

, thi

n P

hase

, vol

ume

Pha

se, v

olum

e

Tab

le 1

. H

olog

raph

ic R

ecor

ding

Mat

eria

ls a

Upp

er f

requ

ency

lim

it

(lp/

mm

)

>3

00

0C

1000

30

00

4000

' 15

00

10,0

00h

Max

imum

dif

frac

tion

ef

fici

ency

for

si

ne g

rati

ng

(%) 5

46

b

30

90

30g

85

92h

·Unl

ess

othe

rwis

e st

ated

, all

data

from

Bie

derm

ann

in r

ef.

46.

"Fro

m C

olbu

rn e

t al.

in r

ef.

45.

'Fro

m V

ilkom

erso

n in

ref

. 52.

'F

rom

Cre

deIl

e an

d Sp

ong

in r

ef.

53.

"Fro

m K

illat

and

Ter

reIl

in r

ef.

47.

"Fro

m O

rlow

ski

in r

ef.

54.

'Fro

m B

usch

man

n in

ref

. 50

. dF

rom

Bie

derm

ann

in r

ef.

51.

Sca

tter

ed fl

ux

at 8

00-1

000

lp/m

m

[(lp

/mm

)-2]

9 X

10

-10

d

9 X

1

0-8

•

8 X

10-

10•

rb

(dB

)

18-2

1

23-2

7

17

24-2

8

In s

itu

hand

ling

No

No

No

Yes

Y

es

Yes

~ i- ii! 0 'C

I i I 1 GO i ~

fa ff

CI iI , CI - f C -; 'C

I if'

1M

!It

1M

354 U. Killat

contrasts have been found to be limited to a value of about 100 and the image quality of reconstructed radiographs is not satisfactory.(49) The field of grain-free volume-type materials, such as photopolymers and electrooptic crystals, has not yet been thoroughly investigated with respect to noise properties and ·gray scale performance. There is some hope that these materials will improve the gray scale performance of a holographic store to a satisfactory level. We would like to encourage research in this area.

4.3.3. SPECKLE EFFECT

Thus far we have assumed that our object wave contains no phase fluctuations. This assumption will be invalid

1. If diffuse illumination of the object transparency is used and/or 2. If the object transparency itself, or its support, contains phase

irregularities.

Diffuse illumination of the transparency is often used to provide a nearly uniform intensity in the hologram plane, thus reducing the problem of nonlinearities (see Section 4.3.1.1) and making the hologram insensitive to dust and scratches. However, the increase of redundancy is gained at the expense of the grainy appearance ("speckle") of the reconstructed image.

As an example, Figure 16(c) shows the image of the transparency of Figure 16(a) obtained from a hologram, taken in the same geometry as in Figure 16(b) but formed using diffuse illumination.

The same trouble with speckle will be encountered in case 2, which is typical for radiographs. Radiographs strongly scatter light and thus inherently act as diffusers. This is demonstrated in Figure 13(d), where the radiograph of Figure 13 (a) is reconstructed from a small hologram (S = 20). The difference in quality compared with Figure 13(c) (S = 1) is solely that in Figure 13 (c) the speckles are so fine that they are no longer objectionable. The only way to reduce speckle while preserving (willingly or not) diffuse illumination is to enlarge the aperture (i.e., the ratio 1/ d), because the speckle pattern spot size is diffraction limited. (55) As this method increases the bandwidth of the signal to be recorded by the hologram, the method is limited by the resolution of the storage material.

However, the point we want to make in this context is that this method reduces speckle only at the expense of increased noise contributed by the storage device. This becomes obvious from our discussion of equation (12), in which we showed that L/ d = S(// d) should be as small as possible to attain as Iowa noise contribution as possible. Obviously, there exists a


trade-off between the methods to fight speckle noise and the noise contributed by the storage material.

This situation could be improved if it were possible to produce a coherently illuminated nonscattering image of the radiograph in an intermediate storage device.

Such an intermediate image can be produced by simply copying the original radiograph using a conventional duplicating film (for example, Kodak X-OMAT). However, for a storage system used in the hospital this would not be a practical solution; moreover, a rapidly erasable storage device is required for this purpose. In principle, such devices have been developed, but the gray tone capabilities of these incoherent-coherent interfaces, such as Phototitus, (56) are not adequate in their present form.

4.4. The Bottlenecks of Holographic Image Storage

A great density range can be preserved in the reconstructed image even though the storage medium is modulated only with small signals. This is the key issue for the holographic storage of gray tone images.

However, in order to take full advantage of this method for the storage of radiographs, a number of problems have to be solved that have been identified in this section:

1. To avoid nonlinear recording and in order to take advantage of redundant recording, sophisticated illumination techniques (e.g., use of random-phase masks) are necessary. These devices are feasible but do not yet exist in sufficient size and quality.

2. Due to the coherent addition of signal and noise light from the storage medium, the noise imposed by the storage medium is a much more severe restriction than for non coherent methods. Materials with adequate noise performance are not yet available.

3. Radiographs tend to produce speckle patterns upon holographic reconstruction. Therefore a nonscattering replica (in its simplest form from a duplicate film) has to be provided to avoid speckle formation. Erasable incoherent-coherent transducers do not yet have acceptable quality.

In conclusion, we feel that, presently, holography cannot satisfactorily solve the problems of radiograph storage.

From the system's point of view, a holographic store could provide rapid access to the stored information (see Figure 12). However, radiologists would in general not go to this holographic store situated somewhere in the hospital and look on a screen. Therefore, the holographic storage

356 U. Killat

system must be coupled to some units providing a hard copy, or a "soft copy" on a CRT display. Thus, from the system's point of view, it is reasonable to decouple the functions of storage and information output. It is this issue that is rigorously discussed in the next section.

S. DIGITAL STORAGE OF RADIOGRAPHS

Novel high-capacity mass memories for computer applications have been proposed in a variety of concepts in the last ten years or so. (57) One approach, whiCh combines the well-known disk store principle with the submicron resolution of optical methods, is already marketed as video long play (VLP) for consumer applications. In addition, high-performance versions for computer peripherals and other applications are being developed. (58.59) Today, a direct-read-after-write (DRAW) recorder with a capacity of more than 10lD bits per side of the disk is available. "Juke box" systems with a final capacity of more than 1014 bits and an access time of less than 3 sec are presently being studied. (58) Such a unit would occupy about 5 m2 of floor space and would meet the requirement for 30 years of filing in a medium-size hospital.

Thus, at a first glance, the solution to the filing problem is almost available. However, the optical disk stores bits and the problem we are concerned with is radiographs! Thus the memory has to be interfaced to real objects and these interfaces will play an important role in this section. Nevertheless, the optical disk is the pacemaker for a novel concept of systems in all kinds of filing problems.

5.1. Description of the Method

The building blocks for a system to be developed for the storage of radiographs are shown in Figure 18.

1. At the input terminal the radiographs are scanned by an image pickup device (represented by the TV camera in Figure 18). The density of each picture element (pixel) is evaluated and then converted to a digital format. In this format the data are transmitted to the storage system.

2. Conversely, at the retrieval terminals, the data received from the store are converted to analog signals that drive the output unit (hard copy and/or display).

3. Before entering/leaving the optical disk store, the data pass a coding/decoding unit that reduces the redundancy in the data and thus saves memory space.


picture picture Input retrieval

terminal terminal

coder! decoder

optical disk store

transmission system

Figure 18. Digital storage of radiographs.

4. At the optical disk store, the data are stored on the disk together with some address information and coded in a way to ensure error-free recording.

These different units are discussed in detail in the following sections.

5.2. Picture Input Devices

The devices at the picture input terminal are optoelectronic converters that convey pictorial information to a sensor element and then amplify and process the output signal of the sensor.

The conversion of optical information into an electronic counterpart can be performed on the basis of the whole image (w mode), on a per-line basis (l mode), or on a per-pixel basis (p mode). Typical representations of the different types of picture input devices are

1. A TV camera equipped with a Vidicon tube (w mode). 2. A charge-coupled device (CCD) line scanner (l mode). 3. A microdensitometer (p mode).

As a rule of thumb, one may state that image quality and the recording

358

Type

Example

Spatial resolution

Image area at the device

Dynamic range

Recording time per pixel

Table 2. Image Input Devices

w mode I mode

Vidicon tube CCD linear image sensor (Fairchild CCD143)

1400 lines 2048 pixels

13 x 17mm2 27 mm x 13 /Lm

<40 dB -60dB

60 nsec 500 nsec

U. Killat

pmode

Microdensitometer (Optronix P 1000)

25 /Lm x 25 /Lm

Original size, up to 14 in. x 17 in.

256 levels within 0-3D

35 /Lsec

time per pixel increase in this order. Some typical features are shown in Table 2.

None of the picture input devices presently available are fully satisfactory. The drawback of microdensitometers, which provide sufficient image quality, is their slowness and their price, which ranges up to -$100,000. Sophisticated mechanics guaranteeing a high degree of accuracy are the main reason for these drawbacks and considerable improvements cannot be expected. A better option on a future high-quality scanner at a reasonable price is given by the CCO linear image sensor, which will be discussed in greater detail.

5.2.1. CCO LINE SCANNER

A CCO linear image sensor(60) consists of a linear array of photodetectors (MOS capacitors) and an analog CCO shift register [Figure 19(a)]. Once an electric field is applied across the capacitors, incident photons will generate photocurrents, which discharge the capacitors to a degree proportional to the intensity and the integration time during which photons are incident. After a fixed integration time (-25 msec), the residual charges of the photodetectors are transferred in parallel into a shift register (duration -2 msec). When the image signals have been transferred into the shift registers, integration of the next set of image points and serial readout from the shift registers are simultaneously executed.

With commercially available CCO lines, a dynamic range of about 60 dB can be obtained. The main problem-as with most array-type devices-lies in the homogeneity of the response of different elements of the line. Today, typical values are in the range of - 5 %. (61) For twodimensional scanning of a radiograph various geometries can be chosen. One example, in which both the radiograph and the CCO line are fixed, is given in Figure 19(b). The linewise deflection is obtained by means of

Modem Optical Methods for the Storale of Radiolf8Phs

a)

b)

readout CCD

0000000000

photosensitive area

--CGD

line sensor

line deflection

image

Figure 19. CCD imager. (a) CCD line sensor; (b) two-dimensional image scanner.

359

a rotating prism. The analog output signal from the image sensor is fed to a lO-bit AID converter that completes the interfacing to digital equipment. It seems that the rapidly evolving technology of CCD linear and area image sensors will provide for a satisfactory image input device in the near future.

5.3. Picture Output Devices

To maintain evidence of diagnosis, a hard copy unit is indispensable. In addition, a display ("soft copy") is helpful for a rapid overview, for the selection of images, or the proper detail of an image (blowup view). Similarly, a display will be used at the input terminal for control purposes.

5.3.1. DISPLAY UNITS

The oldest display unit that has found widespread use is the cathode ray tube (CRT) display. Though other display techniques are being developed, such as plasma-discharge displays, (62) liquid crystal displays, (62)

electroluminescent displays, (62) or magneto-optic displays, (63) the image quality of a high-resolution CRT has not yet been achieved by any other

360 U. Klllat

technique. Moreover, the resolution of 1000 to 2000 lines and of 50 to 100 gray levels found with high-quality CRTs is more than sufficient for most applications. In many cases, standard TV displays will meet the requirements. Means can be provided to extend the dynamic range by applying a variable-contrast window (common practice in x-ray computer tomography) or to virtually increase spatial resolution by a zoom facility.

5.3.2. HARD COPY UNITS

The description of hard copy units must be divided into two categories: silver and non-silver techniques. Although everybody wishes to get rid of the expensive silver technology, it is a fact that, for highly professional and scientific applications requiring the best hard copies available, photographic materials are used.

On the other hand, copying automats, computer output devices, and facsimile equipment use non-silver techniques that are being developed to deliver high-quality gray tone and color performance.

5.3.2.1. Silver Techniques. In considering silver techniques, we can distinguish the properties of the photographic material finally used as the hard copy from the optical device that exposes the silver emulsion: the recorder.

The films can be classified as

1. Wet processed media. 2. Dry processed media. 3. Instantly processed media.

Table 3 gives an example for each of these types together with some figures of merit. We may summarize that awkwardness of handling goes along with the quality of the final hard copy.

Table 3. Photographic Materials

Wet processed Dry processed Instantly processed Type medium medium medium

Example AGFAPAN400 3M 7869 Polaroid Land film Type 808

Resolution 55lp/mm 3001p/mm 7-9Ip/mm

Density range 0.1-2.8 0.2-3.4 0.2-2.5

Maximum slope of 0.7 2.4 2 characteristic curve

Sensitivity 400 ASA 100 erg/cm2 0.1 erg/cm2

Modem Optical Metbods for tbe Storage of Radiograpbs 361

For recorders, similar remarks apply to those made in connection with picture input devices, because both techniques are similar to some extent: With the input devices, a fixed exposure is applied and the attenuation of light is measured; with recording devices, the exposure is controlled in order to obtain a certain attenuation value in the final print. Similar to our discussion of input devices, we distinguish recorders according to their mode of operation: w mode, 1 mode, and p mode.

1. For operation in the w mode, a high-resolution CRT display is used and imaged onto a camera system by conventional optics. The advantages of this method are • No moving parts. • Inertialess movement of .spot light. • Sealed device. • Well-understood support electronics. • Unaffected by shock and vibration. The disadvantage is the low power density that is caused by the inefficiency of the phosphorus screen. In addition, the lack of uniformity of the phosphorus and errors of the lens system will reduce image quality. Gray tone resolution is limited to about 50 distinguishable gray tone levels.

2. For operating in the 1 mode, two different methods are worth mentioning: • A line CRT with a fiber-optic faceplate is used to convey the

light to the photosensitive medium. Exposure is achieved by moving the recording material in front of the faceplate. The major advantage of this device over the aforementioned conventional CRT is an increased light efficiency due to the fiber-optic coupling.

• A linear light source is focused on a linear array of light valves, which is optically mapped on the recording material. For the light valves, polarization switches are chosen, which make use of electro-optic(64) or magneto-optic effects. (65) As these approaches are open to a development towards low-cost integrated solid-state devices, the magneto-optic printer is described in more detail in the next subsection.

3. Finally, in the p mode, point sources are used to illuminate the recording material. Again, as with microdensitometers, an accurate mechanical design and a highly stable light source ensure high quality at the price of low speed. For the point sources, two satisfactory solutions are being marketed: • A light-emitting diode (LED) is moved by a mechanical carriage

to subsequently address the pixels of aline. Gray tones are achieved by pulse-duration modulation of the LED.

Tab

le 4

. H

ard

Cop

y U

nits

for

Rec

ordi

ng o

n Si

lver

Em

ulsi

ons

Typ

e w

mod

e I

mod

e I

mod

e p

mod

e

Exa

mpl

e D

un

n I

nstr

umen

ts

Tek

tron

ix 4

643

(use

s P

hili

ps m

agne

to-o

ptic

O

ptro

nix

P15

00

Col

or C

amer

a 63

1 C

RT

wit

h op

tic

prin

ter

(LE

D r

ecor

der)

(u

ses

CR

T)

face

plat

e)

Res

olut

ion

Nom

inal

ly 1

400

line

s N

omin

ally

100

0 lin

es

60 p

.m x

60

p.m

25

p.m

x 2

5 p.

m

(CR

T)

Fil

m d

imen

sion

s 8

in.

x 10

in.

Rol

l fil

m,

wid

th:

Dev

ice

cont

ains

256

22

in.

x 17

in.

2

2cm

pi

xels

/lin

e R

epro

duci

ble

gray

N

o d

ata

avai

labl

e 12

16

(64

) 64

le

vels

E

xpos

ure

tim

e (e

)/

r =

2 m

in

r =

12

sec

e =

200

sec

(20

sec

) fo

r e

=1

40

se

c

repe

titi

on t

ime

(r)

20

48

-x-2

04

8 m

atri

x R

emar

ks

Incl

udes

pap

er

Lab

orat

ory

mod

el;

proc

esso

r nu

mbe

rs i

n br

acke

ts

are

expe

cted

to

be

achi

eved

by

the

end

of 1

982

p m

ode

SE

P V

izir

(la

ser

reco

rder

)

12.5

f.Lm

x 1

2.5

f.Lm

40.5

cm

x 4

0.5

cm

64

e =

60

sec

1M ~

N ~ i


• A laser beam is directed to the individual pixels of a line by means of a galvanometer mirror. The intensity of the laser beam is modulated by an acousto-optic modulator.

The features of the different techniques for producing hard copies are summarized in Table 4 for some prominent examples.

Magneto-Optic Printer. Linear arrays of magneto-optic polarization switches can be used for the construction of a hard copy unit. (65) The principle is shown in Figure 20. The array of light valves (pixel size: 60 x 60 ~m2) is imaged onto the recording material. In order to produce gray levels, exposure can be controlled by a pulse-length modulation.

In magneto-optic light valves, use is made of the Faraday effect, which is quite strong in certain epitaxial garnet layers possessing an internal magnetization either parallel or anti parallel to the normal of the film (Figure 21). When looking along the normal of the film, the plane of polarization of linearly polarized light transmitted by the film is rotated either to the left or to the right, depending on the direction of the magnetization present in the pixel cell (Faraday effect). Light with one of the possible polarization planes is blocked at the output by a polarization foil, whereas light of the second type of polarization is transmitted to a certain degree. Switching

cylinder len5 - --,Fr

information input

(5hift regi5ter) - ~~~~1~~;~~7;~ magneto-optic light valve5 -----/-::....--h/.W"?JJf7f,,r,,{

objective -----+

recording material

Figure 20. Hard copy unit using magneto-optic light valves.

364 u. Killat

memory cell

Figure 21. Viewing of magnetic domains using the Faraday effect and polarization optics. From B. Hill and K. P. Schmidt in ref. 63.

from one state of magnetization to the other, and hence from the on state to the off state of the light valve (or vice versa), is accompanied by simultaneous application of a magnetic field to the whole film and selective heating of individual pixel elements via a thin film structure. The whole device (a 256-pixelline), including its driving electronics, has dimensions of 25 mm x 25 mm. Several of these elements can be put together to achieve lines of, for example, 2048 pixels.

The cycle time for one line is presently limited to 1 msec by the heat produced in the coils generating the magnetic field. Redesigning will reduce this time to 100 JJ.sec. For gray tone recording, the exposure range must be enlarged. Thus, for three decades of exposure, the cycle time has to be set to 10 msec. The exposure for a picture of 2000 lines will thus last 20 sec. A laboratory model that is presently being developed is designed for 64 gray levels. Thus a compact solid-state technique is evolving for hard copy units that might be considered as a counterpart to the CCD linear imagers for image scanning. The image quality achievable with an older magnetooptic printer (256 pixels, 16 gray levels) may be judged from Figure 22.

5.3.2.2. Non-Silver Techniques. Electrophotography is presently the only non-silver technique that has found widespread use. However, with respect to gray tone capability, other methods seem to be more promising. We mention ink jet recording,(67) magnetic stylus recording,(68) the elec-

Modem Optical Methods for the Storage of Radiographs

Figure 22. Image obtained with a magneto-optic printer. From B. Hill in ref. 66.

365

trosensitive technique (Bell and Howell), and the electrophoretic technique, (69) which is an attractive new method for continuous tone recording.

Electrophoretic Recording. Charged solid particles suspended in a solvent migrate under the influence of an electric field. This phenomenon of electrophoresis is used to guide toner particles onto a metal drum. The electric field is applied between that drum and a recording electrode and its strength controls the amount of toner particles deposited on the drum (see Figure 23).

co rrI I'r

Image drum

supply nozzle

charged loner parhcles

removol n0221(>

050~m

~'200V

UCa" " " 0 V

Figure 23. Principle of electrophoretic hard copy technique. From H. D. Hinz et at. in ref. 69.

366 U. Killst

Figure 24. Electrophoretic recording; reproduction of a detail of Figure 13a. From Hinz et al. in ref. 70.

For image recording, the drum is rotating at a constant rate while the recording electrode scans the drum parallel to its axis. After the entire image has been recorded on the drum, a transfer of toner particles onto the final paper is accomplished by means of a transfer roller. A resolution of 2 lp/mm (50% value of the modulation transfer function) has been obtained. Density values are limited only by the recording time per pixel.

Advanced concepts of electrophoretic recording investigate comblike recording electrodes that will reduce the recording time for A4-sized images to less than 1 min. This is comparable to photographic techniques when exposure and development steps are added. An example of the image quality obtained so far with the electrophoretic method is given in Figure 24, which shows a detail of the picture of Figure 13.

5.3.2.3. A Comment on Hard Copy Units. Hard copy units are certainly the most critical unit of the whole system. In addition, they can be considered as primary information sources for CT pictures. From the development described in this section it seems that a hard copy unit satisfactory in all respects (spatial and gray level resolution, price, and speed) is not yet available. On the other hand, it may be acceptable to have a somewhat reduced scale of density levels and/or resolution, as they will be offered in low-price devices such as the garnet and the electrophoretic printer: This is because the digital format will allow any form of picture manipulation. Instead of one image that is a true replica of the


original, the radiologist may be even more satisfied with an overview (at reduced quality), an enlarged detail (of sufficient quality), and some examples of processed images (e.g., low-pass filtering).

5.4. Source Coding of Radiographs

It has been mentioned in Section 3.3 that a matrix representation of x-ray images is highly redundant. The properties of the imaged objects introduce a high correlation of neighboring pixels. Suitable coding schemes can be used to reduce the effective number of bits per pixel necessary for faithful storage in the digital store.

Several coding schemes, such as predictive coding, run-length coding, or transform coding, have been investigated and compression factors of up to ten have been obtained. We will not discuss these schemes at length but refer the reader to the literature.(71-73) As an example, we present a radiograph coded with 1 bit/pixel in Figure 25. Its quality can be checked against that of Figure 8(a) using 4 bit/pixel but no redundance reduction. For the implementation of these techniques, a trade-off has to be made: High compression factors go along with sophisticated electronics and high sensitivity against bit errors. Therefore, the implementation of a coder/decoder will always represent a compromise.

Figure 25. :o>ource coding with SH transform: 1 bit/pixel. From P. Lux in ref. 73.

368

dala signa l facus error

U. Killat

polariz ing b .. a msplil1Pr

laser

fEt::~3:::::::;~~=====:!~r-_semitrnnsparent mirror

..,p--- trock - error detector

Figure 26. Recording and reading from an optical disk.

5.5. Optical Disk Store

In the optical disk system, digital data are represented in the form of small holes burnt into a storage layer by means of a focused laser beam. Figure 26 illustrates the construction of the storage plate and the optical system that provides the direct-read-after-write (DRAW) feature.(58.59)

The light of an AIGaAs laser source is deflected by a polarizing beamsplitter, passes a quarter-wave retardation plate, and is then focused

Figure 27. Optical disk recorder. From K. Bulthuis et al. in ref. 59.

Modem Optical Methods for tbe Storage of Radiographs

Figure 28. Data recorded as holes of diameter 0.9 ~m. From K. Bulthuis et al. in ref. 59.

369

onto the storage layer for recording and/or reading purposes. The light reflected from the storage layer passes the retardation plate twice and is therefore able to pass the polarizing beamsplitter unaffected. This reflected light is modulated by the succession of areas of different reflectivitynamely, areas containing a hole or areas in their virgin state. Therefore, the reflected light contains enough information to confirm the correctness of the written data and to derive servo signals for focusing and tracking from sector photodiode detection units. An assembled recorder is shown in Figure 27.

The storage disk itself, in fact, consists of two disks sealed to each other with spacers in between. Thus the tellurium-based storage layers are kept in a miniature clean room; dust or scratches on the outside surface of the disk do not prevent proper recording or reading. The holes containing the information are placed in a continuous spiral groove with a pitch of 1.67 ~m, and the diameter of a hole is 0.9 ~m (Figure 28). The data are cut into blocks of 1 Kbit; each block has a heading for addressing and synchronization purposes. The raw data are coded by adding some redundant bits in a way that single-bit errors can be corrected upon the decoding procedure. If large dropouts occur, they are detected in the reading process immediately after the writing procedure. In such a case, the respective sector is invalidated and rewritten. In this wayan almost error-free digital recording is possible.

The most crucial points in the design of an optical disk recorder are: 1. to attain as high an information density as possible, and 2. to construct a simple and efficient mechanics for the exact positioning of the reading/writing unit. To elaborate on these two points:

1. The spot size of the holes in the storage layer is limited by the cutoff frequency of the carefully designed focusing objective when

370 V.KiDat

used with the laser source of wavelength 0.82 ILm. The resolution obtained is of the same size as the wavelength of the laser source. Another parameter that determines the information capacity of a disk is the track spacing. Its minimum value is determined by the amount of cross talk that can be tolerated between adjacent tracks. This spacing is also determined by the size of the light spot. For a track spacing of 1.5 ILm, the cross talk reduces to 30 dB, which is sufficient for most applications. Under these conditions the information capacity of a 30-cm disk can amount to some 1010 bits per side.

2. To reduce the use of high-precision mechanics in the recorder, the concept of a pre grooved disk was developed. Such a disk contains a continuous spiral groove (pitch: 1.67 ILm) already in its virgin state. The light reflected from the groove provides tracking information for the servo systems that keep the writing/reading unit in the proper radial position. In this way, accuracy requirements for the mechanics are tied to the accuracy of the pre grooved spiral on the disk. This in turn can be guaranteed by a high-precision mastering machine used in the production of pre grooved disks.

Though the storage method described in this section uses optical technology, it is perfectly matched to its electronic surroundings. This has already been proven with the video disk system, upon which most of the technology for the digital store is based. We therefore recommend the extensive literature on the VLP system for further reading, especially a series of articles in the July 1978 issue of Applied Optics.

6. TOWARDS FUTURE HOSPITAL COMMUNICATION SYSTEMS

In the preceding sections, different technologies have been proposed for an adequate solution to the problem of efficient radiograph storage.

Photographic miniaturizing systems and holographic methods are representatives of analog methods that must compete with the more recently developed digital approach. With considerable research efforts, probably each of the solutions that "in principle work" could be refined to a degree that a practicable system would evolve. However, the rapid development in semiconductor technology and the flexibility of digital data processing will certainly lead to a concentration of efforts in favor of digital solutions. Such a situation is not unique but, rather, typical for many fields (e.g., multichannel audio recorders, transmission and switching of telephone


calls). In particular, the situation in the telecommunication field will be of utmost importance for information storage and retrieval in hospitals.

The current trend in telecommunications aims at networks transparent to all types of services: transmission of speech, text, diagrams, pictures, and data. A storage system in a hospital should not be restricted to the storage of x-ray images but should preferably include

1. Speech (e.g., comments of the physician using a dictaphone). 2. Text (e.g., case reports). 3. Diagrams (e.g., ECG, EEG). 4. Pictures (e.g., radiographs, thermographs, etc.).

Referring back to the illustrative example of Roentgen diagnosis, it would be desirable to do some processing on the raw data available from the store. Image processing equipment could back the storage/retrieval system in allowing for

1. Image manipulation, such as spatial filtering, contrast enhancement, subtraction of images.

2. Automated image analysis, such as pattern recognition or the quantitative evaluation of the size of diagnostically relevant areas.

Thus the system of Figure 18 has to be supplemented by input/output terminals for other services and a picture processing system, which may be at the same site or may be a remote computer connected to the hospital's information system via some sort of network for the exchange of data. Similarly, transmission channels should be available to interconnect hospitals, scientific centers, and outside practioners for rapid exchange of information. Broadband channels have to be included because of the high data volumes (> 107 bits) related to a single radiograph. A sketch of the hospital communication system of the future is given in Figure 29. For some of the devices shown in this figure, the state of the art has been described in Section 5. We would like to point out that the technological trends already discernible on the horizon do favor the approach given here:

1. For the input/output of pictorial information the greatest push is· expected from the Telefax service. As this service is still evolving, attention should be paid to the different areas needing devices with good gray tone performance; besides medical radiographs, these are nondestructive testing in industry, photogrammetry, weather forecasting, crime investigation, and the publishing industries.

372 U.KW.t

text pIcture termInal phone

Intrahospltal transmIssIon network

optIcal dIsk store

Figure 29. Hospital communication system.

2. The "soft copy" will profit from the development of a high fidelity TV with increased resolution.

3. Decreasing cost and space and increasing the speed of processors and memories will facilitate the development of picture processing units.

4. Fiber-optic cables, which are already in operation in field trials, (74,7S)

provide channels of up to I-Gbit/sec capacity with repeateriess transmission over several kilometers. (76)

The problem of a future hospital storage system will not only lie in the components, for which rather optimistic perspectives have been outlined; the problem will likewise lie in the architecture and the data management of the integrated storage, processing, and retrieval system handling large amounts of data. Fortunately, these problems have been recognized and the first developments for an integrated picture storage and processing system are presently being installed and will be tested in German hospitals. (77,78)

They represent a first step towards a network with immediate access to a picture archive as well as to the hospital's data base, with alphanumeric input and output and graphic interaction, access to processing power for picture and data, and picture-phone communication between the physicians. This type of communication system will improve the efficiency of medical care by making diagnoses faster, more comprehensive, and more reliable.


ACKNOWLEDGMENTS

The author would like to thank Drs. W. Dallas, H. Dammann, and H. J. Schmitt for many helpful comments. He is particularly indebted to J. Coumans, Dr. Meyer-Ebrecht, and R. Weissel, who made their research results relevant to Sections 5 and 6 available to him prior to publication.

REFERENCES

1. U. Schneider and V. Taenzer, Probleme der Informationsiibertragung bei der Mikroverfilmung von Rontgenaufnahmen, Roentgen BI. 25, 26-30 (1972).

2. K. H. Bernhardt, Die Mikro-Dokumentation von Rontgenaufnahmen mit dem "LinearSystem," Roentgen BI. 32, 329-334 (1979).

3. H. von Prosch, Formatreduzierte Krankenblatt- und Rontgenbildarchivierung, Roentgen Ber. 5, 205-211 (1976).

4. N. V. Optische Industrie "De Oude Delft", Delft, The Netherlands. 5. LogEtronics Inc., Springfield, Virginia. 6. E. Lehtinen and I. Tamminen, Microfiling and Automatic Filing of Radiographs, Roentgen

BI. 31, 300-303 (1978). 7. S. Laser, Archivierung von Rontgenfilmkopien verbunden mit der Moglichkeit einer

einfachen Datenverarbeitung, Radiol. Clin. 39, 252-255 (1970). 8. C. John, The reported film storage problem solved? 100 mm film miniaturization using

the Oldelft, De1copex and Ode1card systems, Radiography 19, 191-196 (1978). 9. H. Arvidsson, in Planning of Radiological Departments (M. Kormano and E. Stieve,

eds.), pp. 122-123, Georg Thieme Publishers, Stuttgart (1974). 10. I. Rossi and P. Virtama, in Planning of Radiological Departments (M. Kormano and E.

Stieve, eds.), pp. 123-125, Georg Thieme Publishers, Stuttgart (1974). 11. E. Kolihova, D. Obenbergerova, and M. A. Ansari, Unsere Erfahrungen mit dem

De1copex-System, Raddiagn. 19, 702-705 (1978). 12. L. Steinhart, Vorstellungen iiber eine Systemlosung der Archivierung von Rontgenauf

nahmen, Raddiagn. 19, 706-708 (1978). 13. F. Bossard and M. E1cke, Zur Anwendung von Rontgenbild-Verkleinerungssystemen in

einer zentralen Universitiits-Rontgendiagnostikabteilung unter Beriicksichtigung der Betriebsstrukturen, Roentgenpraxis (Stuttgart) 30, 16-27 (1977).

14. H. Deininger, K. Goergens, H. Heitzeberg, F. Heuck, W. Hoeffken, J. Kahlstorf, G. Meier, L. Mitrovics, G. Miiller, H. Oeser, Ch. F. Pahlow, R. Sorensen, H.-S. Stender, V. Taenzer, H. Weissleder, and E. Zeitler, Untersuchungen iiber die praktische Verwendbarkeit eines Rontgenfilm-Verkleinerungs-Systems, Fortschr. Roentgenstr. 120, 353-356 (1974).

15. V. Taenzer and U. Schneider, Rontgenaufnahmen-Mikroarchivierung: Eine Bilanz, Fortschr. Roentgenstr. 120,616-620 (1974).

16. G. P. van Dijk, Rontgenarchivierung mit Mikrofilm, Chirurg 47, 62-65 (1976). 17. L. W. Busching, Implementing a new x-ray film management system, Radiol. Technol.

49,603-610 (1978). 18. C. Th. Ruygrok, Ein Vergleichsversuch der modernen Kopiersysteme, presented in

Symposium iiber Fragen der Archivierung und des Informationsaustausches in der Rontgendiagnostik, Munich, 15-16 November, 1971.

19. E. Ross, Der logetronisierte Mikrofilm in der Rontgendiagnostik bei integrierter Verfilmung und Archivierung, Fortschr. Roentgenstr. 124, 585-594 (1976).

374 U. Killst

20. R. Bollen, private communication. 21. J. C. Dainty and R. Shaw, Image Science, Academic Press, New York (1974). 22. R. F. Wagner, Noise equivalent parameters in general medical radiography: the present

picture and future pictures, Photogr. Sci. Eng. 21, 252-262 (1977). 23. K. G. Vosburgh, Storage and retrieval of radiographic images, Radiology 123,619-624

(1977). 24. C. E. Shannon, A mathematical theory of communication, Bell Syst. Tech. J. 27, 379-423,

623-656 (1948). 25. J. H. Altmann and H. J. Zweig, Effect of spread function on the storage of information

on photographic emulsions, Photogr. Sci. Eng. 7, 173-177 (1963). 26. R. F. Wagner, D. G. Brown, and M. S. Pastel, Application of information theory to the

assessment of computed tomography, Med. Phys. 6, 83-94 (1979). 27. H. Kanamori and S. Murashima, Information capacity of radiographic images for the

random signal and continuous objects, Jpn. J. Appl. Phys. 9, 182-194 (1970). 28. A. Rose, The sensitivity performance of the human eye on an absolute scale, J. Opt. Soc.

Am. 38, 196-208 (1948). 29. A. Rose, Vision: Human and Electronic, Plenum Press, New York (1973). 30. K. Rossmann, Modulation transfer function of radiographic systems using fluorescent

screens, J. Opt. Soc. Am. 52, 774-777 (1962). 31. K. Rossmann, Recording of x-ray quantum fluctuations in radiographs, J. Opt. Soc. Am.

52, 1162-1164 (1962). 32. M. Schlindwein, private communication. 33. H. J. Metz, S. Ruchti, and K. Seidel, Comparison of Image Quality and Information

Capacity for Different Model Imaging Systems, J. Photogr. Sci. 26, 229-233 (1978). 34. J. W. Motz and M. Danos, Image information content and patient exposure, Med. Phys.

5 (1), 8-22 (1978). 35. N. A. Baily, A. L. Pearlman, and E. C. Lasser, Holographic storage of Roentgenogram

images, Radiology 99, 190-191 (1971). 36. N. A. Baily, R. L. Crepeau, and E. C. Lasser, Further developments in the use of

holographic methods for the storage of Roentgenographic images, Invest. Radiol. 7 (2), 118-123 (1972).

37. B. HilI, in Advances in Holography, Vol. 3 (N. H. Farkat, ed.), Marcel Dekker, New York (1976).

38. U. KilIat, Holographic microfiche of picture-like information, Opt. Acta 24, 453-462 (1977).

39. C. Clausen and U. KilIat, Holographically stored x-ray images: gray tone reproduction, Med. Phys. 5,181-187 (1978).

40. R. J. Collier, C. B. Burckhardt, and L. H. Lin, Optical Holography, Academic Press, New York (1971).

41. G. Groh, Holographie, Berliner Union and Kohlhammer, Stuttgart (1973). 42. M. Francon, Holography, Academic Press, New York (1974). 43. Y. Tsunoda and Y. Takeda, High density image-storage holograms by a random phase

sampling method, Appl. Opt. 13, 2046-2051 (1974). 44. A. Iwamoto, Artificial diffuser for Fourier transform hologram recording, Appl. Opt. 19,

215-221 (1980). 45. W. S. Colburn, R. G. Zech, and L. M. Ralston, Holographic Optical Elements, Technical

Report AFAL-TR-72-409, Harris Electro-Optics Center of Radiation, Ann Arbor (1973). 46. K. Biedermann, Information storage materials for holography and optical data processing,

Opt. Acta 22, 103-124 (1975). 47. U. Killat and D. R. Terrell, Performance and limitations of photothermoplastic devices,

Opt. Acta 24,441-452 (1977).


48. U. Killat and D. R. Terrell, Conventional and Self-Developing Phototherrnoplastic Devices: Sensitivity and Rapid Development, J. Photogr. Sci. 26, 183-188 (1978).

49. U. Killat, unpublished results. 50. H. T. Buschmann, in Optical and Acoustical Holography (E. Cametini, ed.), Plenum

Publishing, New York (1970). 51. K. Biedermann, The scattered flux spectrum of photographic materials for holography,

Optik (Stuttgart) 31, 367-389 (1970). 52. D. H. R. Vilkomerson, Measurements of the noise spectral power density of photosensitive

materials at high spatial frequencies, Appl. Opt. 9, 2080-2087 (1970). 53. T. L. Credelle and F. W. Spong, Thermoplastic Media for Holographic Recording, RCA

Rev. 33, 206-226 (1972). 54. R. Orlowski, private communication. 55. J. C. Dainty, in Laser Speckle and Related Phenomena (J. D. Dainty, ed.), Springer-Verlag,

Berlin (1975). 56. M. Grenot, J. Pergrale, J. Donjon, and G. Marie, New electro-optic light valve device

for image storage and processing, Appl. Phys. Lett. 21, 83-85 (1975). 57. B. Hill, U. J. Schmidt, and H. J. Schmitt, Optical Memories, J. Appl. Sci. Eng. A 1,

39-48 (1975). 58. G. C. Kenney, D. Y. K. Lou, R. McFarlane, A. Y. Chan, J. S. Nadan, T. R. Kohler,

J. G. Wagner, and F. Zernike, An optical disk replaces 25 mag tapes, IEEE Spectrum 16, February, 33-38 (1979).

59. K. Bulthuis, M. G. Carasso, J. P. J. Heemskerk, P. J. Kivits, W. J. Kleuters, and P. Zalm, Ten billion bits on a disk, IEEE Spectrum 16, August, 26-33 (1979).

60. D. V. McCaughan and B. R. Holeman, in Charge-Coupled Devices and Systems (M. J. Howes and D. V. Morgan, eds.), pp. 241-295, Wiley and Sons, New York (1979).

61. H. D. Bohring, H. Busch, J. Christiansen, J. Coumans, R. Grewer, H. D. Hinz, H. LObi, D. Meyer-Ebrecht, R. Pape, U. Rothgordt, and R. Weissel, Elektronische Bildbank, Report BMFT-FB 081 2074 (1980).

62. T. Okoshi, in Electronic Imaging (T. P. McLean and P. Schagen, eds.), pp. 103-144, Academic Press, New York (1979).

63. B. Hill and K. P. Schmidt, Fast Switchable Magneto-Optic Memory-Display Components, Philips J. Res. 33, 211-225 (1978).

64. S. Jyomura, I. Matsuyama, and A. Kumada, A high contrast image display device using a PLZT ceramic, Ferroeleetries 15,51-59 (1977).

65. B. Hill and K. P. Schmidt, Integrierte Lichtmodulationsmatrizen aus magneto-optischem Eisengranat fur neuartige Datensichtgeriite und optische Drucker, NTG Fachber. 67, 106-116.

66. B. Hill, private communication. 67. C. H. Hertz and A. Mansson, Electronic ink jet device, Proc. IFIP Congr. 74, 185-189

(1974). 68. M. Nomura, Direct imaging process by magnetic ink powder transfer, presented at the

2nd International Conference on Business Graphics, Washington, November 1979. 69. H. D. Hinz, H. LObi, and U. Rothgordt, Electrophoretic recording of continuous-tone

images, J. Appl. Photogr. Eng. 6, 69-72 (1980). 70. H. D. Hinz, H. Lobi, and U. Rothgordt, private communication. 71. M. Kunt, Source coding ofx-ray pictures, IEEE Trans. Biomed. Eng. 26,121-138 (1978). 72. P. Lux, Redundancy reduction in radiographic pictures, Opt. Acta 24, 349-366 (1977). 73. P. Lux, A novel set of closed orthogonal functions for picture coding, AEU 31, 267-274

(1977). 74. E. G. Rawson and R. M. Metcalfe, Fibernet: multimode optical fibers for local computer

networks, IEEE Trans. Commun. 26,983-990 (1978).

376 U. Killat

75. J. E. Midwinter and J. R. Stern, Propagation studies of 40 km of graded-index fiber installed in cable in an operational duct route, IEEE Trans. Commun. 26, 1015-1020 (1978).

76. C. Baack, G. Elze, and G. Wolf, Feasibility of optical gigabit transmission systems, Proc. Suppl. ECOC 1978,44-46 (1978).

77. D. Meyer-Ebrecht and R. Weissel, The integrated picture processing/picture base system PICASSO, NTG Fachber. 67, 245-253 (1979).

78. D. Meyer-Ebrecht, The management and processing of medical pictures-an architecture for systems and processing devices, presented at the Proceedings of the 2nd Workshop of Picture Data Description and Management, Asilomar, 27-28 August 1980.

Index

A150, 161, 186 Absorbed dose, 4, 6, 17,23

distribution, 61, 62, 68, 73 gamma-ray, 35 neutron, 40, 41 neutron distribution, 34 values, 82

Absorption coefficient, 228, 231, 232, 248, 249 coefficient resolution, 231, 232, 253, 254,

265 energy coefficients, 142, 147 hologram, 352 photoelectric, 142, 155

Acids, amino, 312 Acrylics, 137, 150, 159, 172 Activation detectors, 42, 43 Activities, specific, 311, 312 Additives, 158

corrective, 158, 159 Adipose tissue, 136, 137, 138, 141, 142,

143, 144, 146, 161, 164, 167, 186 Adiprene, 136 Agar, 136, 160 Agents, gelling, 160 Aluminum, 135

oxide, 159 Amplitude limitation, 329 Analysis, in vivo neutron activation, I, 15,

78 Angular dependence, 21, 31 Angular distribution of neutrons, 10 Angular mass scattering powers, 144, 147,

153

377

Annihilation positron, 300 radiation, 30 I

Anthropomorphic phantoms, 78, 133, 137, 171, 179

AP/3, 161, 186 AP 16, 161, 186 AP 1L2, 161, 186 AP ISFI, 161, 186 Apparatus

vacuum-mixing, 172, 173 vacuum-pouring, 176

Atomic number effective (Z) method, 148, 149

Attenuation coefficient, 170 effective method, 104, 108 linear, 106, 170, 335 mass, 142, 143, 147, 149 photoelectric, 149

Average breast, 143, 164, 165, 169, 188 tissue, 139, 141, 165

BIIO, 162, 187 Backprojection, 30 I, 308 Bakelite, 135 Base materials, 158 Basic data method, 150 Batho TAR technique, 238 Beams

external, 228 neutron, 7 proton, 180

378

Beams (cont.)

quality, 68 reactor neutron, 15

Beam-shaping devices, 73 Beef steak, 170 Beryllium target, 7, 10, II Binary coding, 335 Biological dosimeters, 7 Biological relative effectiveness, I Biological tracers, 296 BL 8 thixotrophic gel, 136, 161, 186 BL/L2, 161, 186 Black-bone

copies, 327 file, 328

Bladder cancer, 271 Blood, 138, 141, 161, 164, 167, 186 Body

organs, 139, 141, 165 phantoms, 133, 179

Bolus, 136, 226 alba, 135, 136 Lincolnshire, 137

Bone, 104, 128, 230 inner, 139, 141, 143, 162, 165, 168, 187 compact, 138 cortical, 137, 138, 139, 141, 142, 143,

144, 146, 162, 165, 168, 187 substitute, 135, 180

Boost fields, 202, 203 BR/12, 164, 188 BR/13, 164, 188 BR/14, 164, 188 BR, AV./L2, 164, 188 Brachytherapy, 242 Bragg-Gray, 18

cavity, 38 Bragg peak, 171, 180 Brain, 137, 139, 141, 163, 165, 169, 187

tissue, 170 Breast, 137

average, 143, 164, 165, 169, 188 tissue, 139, 141

Breathing, 224, 234 Bremsstrahlung, 143 BRN IL6, 163, 187 BRISR2, 163, 187 Buildup, 56, 57, 58, 71, 128

Calcium carbonate, 159 orthophosphate, 159

Calculations dose, 265 Monte Carlo, 126, 239, 245 pixel-by-pixel, 239 techniques, 237 transport, 69

Californium-252 neutron source, 35 Calorimeters, 27, 28

tissue-equivalent, 66 Camera

Anger, 302, 303 gamma, 303 minification, 326 modular positron, 302 positron performance, 310

Cameron wax, 135 Cancer

bladder, 271 prostatic, 268

Canonical components, 156 Captive processes, 145 Capture therapy, I

boron, neutrons, 15 Carbonate, calcium, 159 Carcinogenic properties, 158 Casts, immobilization, 222 CCD line scanner, 357, 358 Cell

anoxic, 2 cycle, I, 2

Chamber ionization, 17

Index

tissue-equivalent ionization, 10, 17, 18, 19, 23, 29, 30, 37, 65

Charged-particle, 143 equilibrium, 56, 59 therapy, 242, 244 treatment, 245

Chemistry, radiopharmaceutical, 312 Chest wall, 276 Class

A,166 B,166 C,166

Clinical stage, 196 Coding source, 367 Coefficient, 106, 135

absorption, 228, 231, 232, 248, 249,253 absorption, resolution, 253, 254, 265 attenuation, 170 effective attenuation, 104, 108 energy absorption, 142, 147

Index

Coefficient (cant.) linear attenuation, 106, 170, 335 mass attenuation, 142, 143, 147, 149 mass-energy-transfer, 145 Selwyn, 330

Coherent, 142 scattering, 155

Coincidence detection, 30 I determination, 303

Cold-cure polymers, 158 Cold-cure resins, 176 Collision, 143

stopping powers, 153 Compact bone, 138 Complementary, 155 Components

a, 2 p, 2 canonical, 156 gamma-ray, 10 linear, 2 quadratic, 2

Composition, elemental, 138, 140, 141, 186, 188,249

Compton scattering, 300 Computed tomography, 147, 181, 182, 195,

325 Contamination, 40 Contrast

media, 225, 248 medium, 234, 250

Control, local, 214, 217 Conversion, TAC (time-to-amplitude), 303 Copies

black-bone, 327 white-bone, 327

Correction flood-field, 309 image, 309

Corrective additives, 158, 159 Cortical bone, 137, 138, 139, 141, 142, 143,

144, 146, 162, 165, 168, 187 Cost effectiveness, 219 Cross sections

elemental, 142 neutron, 147

CT (computed tomography), 73,117,118, 119

impact, 268 number, 235, 237, 239, 244, 254 scanner, 252

CT (computed tomography) (cant.)

scanner specification, 265 See also ECT

Curix RPI-Cawo Uni, 335, 336 Curve

characteristic, 329, 335 dose response, 218

Cycle, cell, I, 2 Cyclotron, 3, 4, 7, 8

Do, 2 d + Be, 3, 8, 9 d(l3.5) + Be, 70, 71 d(l6) + Be, 65

spectra, neutrons, 45 d(50) + Be, 69, 72 d + T, 3, 69, 70, 72

neutron beams, 44, 65 neutron generators, 7 neutron sources, 13 spectra, 44

Damage, sublethal, I, 2 Damping, 329, 331 Data, basic method, 150 Decrement line method, 69 Delcomat system, 326, 327, 332 Delcopex RP film, 326, 329, 330 Densitometry, 325 Density

determinations, 178

379

electron, 142, 154, 235, 236, 237, 250 mass, 138, 141, 157, 178

Dental phantom, 181, 182 Dental radiology, 181 Detection, 196

coincidence, 30 I signal,349 single-photon, 300

Detectors activation, 42, 43 characteristics, 303 low-contrast, 254 moderating spheres, 42, 43 silicon diodes, 42, 43 track, 42, 43

Deuteron, 3, 7, 8 Diagnostic radiology, 134, 181 Differential scatter-air ratios, 121 Diffraction efficiency, 353 Diffuse illumination, 354 Digital, 356, 357

380

Displacement correction factors, 25, 26, 27 Display

capabilities, 259 units, 359

Dose absorbed, 4, 6, 17,23,34,40,41,73,82 calculation, 265 curve, 218 depth, 171 depth characteristics, 70 equivalent, 5 neutron distribution, 31 perturbation, 244 rates, 14 response, 5,218 total, 201

Dosimeters biological, 7 Iyoluminescent, 30 neutron, 15 solid-state, 30 thermoluminescent, 30, 33

Dosimetry, 228 fast neutron, 4, 76 interface, 56, 61 ionization, 22 neutron, 1,4,6, 16, 17, 18 neutron intercomparisons, 16,79,80,81,

82, 83, 85, 86 protocols, 65 radiobiological, 54 radiotherapy, 65

Dynamic range, 358

ECT (emission computed tomography), 134 Effective atomic number, 149

(Z) method, 148 Effective attenuation coefficient method,

104,108 Efficiency diffraction, 353 Electron, 140, 143, 144, 166, 167, 247

density, 142, 154, 235, 236, 237, 250 stopping powers, 144, 147

Electronic equilibrium, 106, 127 Electrophoretic recording, 365, 366 Electrophotography, 364 Elemental composition, 138, 140, 141, 186,

188,249 Elemental cross sections, 142 Elemental equivalence, 156 Elementally-equivalent materials, 135

Elements, chiral, 312 Emission, tomography, 183 Emitters, positron, 297 Emulsions, photographic, 30, 34 Energy

absorption coefficients, 142, 147 deposition, 4 dose-average lineal, 48 frequency-average lineal, 48

Index

lineal (y) spectrum, 6, 43, 50, 51 linear transfer (LET) spectrum, 6, 43 mass transfer coefficient, 145 neutron, 6, 7, 35, 43 photon, 6 saturation-correction dose-average lineal,

48,53 Epithermal neutrons, I Epoxy

CB2, 161, 186 resin (systems), 137, 159, 172

Equilibrium, electronic, 106, 127 Equivalent, elementally, 135, 166 Equivalent method, 160 Equivalent tissue, 135, 148, 166 Equivalent tissue-air ratio method, 113,

114, 116 Ethanol, 159

ethoxy, 136 Ethoxyethanol, 136, 161, 186 Ethylene glycol, 159 Exposure time, 362 Extended Y method, 153, 160 External beam therapy, 228

F flouride ion, 312 Failure, marginal, 210 Faraday effect, 363, 364 Fast neutrons: see Neutrons Fat, 136, 137, 139, 159 Fields

boost, 202, 203 radiation, 20 I

Filters flattening, 73 polyethylene, 13 wedge, 73

Fission neutrons, 15 Fitting ratio, 151, 154 Flattening filters, 73 Fluence

measurements, 29

Index

Fluence (cont.) neutron, 16, 79 spectrum, 29

Flux, scattered, 353 Foaming agent, 159 Fragmentation, 145 Frequency, upper limit, 353 Frigeno

gel, 162, 186 liquid, 161, 186

Function modulation transfer, 331 optical transfer, 335

Gallium-68, 313 Gamma camera, 303 Gamma-ray, 6, 7, 9, 10, 15

absorbed dose, 35 component, 10 contribution, 9 energy spectra, 44

Gases, radioactive, 312 Geiger-Muller (GM) counters, 34, 37, 38, 39 Gel, 164, 175

agents, 160 Frigerio, 161 substitutes, 137 thixotropic BL8, 136

Gelatin, 136, 159 Generators

d + T neutron, 7 neutron, 3, 4, 14

Geometry phantom regular, 180 Glass, 135 Glucose, 159 Glycerol, 156, 159 Glycol ethylene, 159 Granularity, 330 Gray scale reproduction, 344 Gray tone levels, 337, 349, 351 Gray tone reproduction, 346, 349 Griffith urethane, 161, 186 Guar,160

Hard copy units, 360, 362, 363, 366 Harris wax, 135 HB/P2, 162, 187 HB/SR4, 162, 187 Head and neck: see Tumors

Heterogeneity, 180 phantoms, 106 tests, 178

Histology, 196 Holograms, 340, 342

absorption, 352 phase, 352 thin, 352 volume, 352

Holographic image, 341, 355 Holographic method, 341

381

Holographic recording, 346, 349 Holographic recording materials, 352, 353 Holographic storage, 340, 350, 351, 355 Holographic storage system, 342, 343 Hospital communication systems, 370, 372 HPLC (high-pressure liquid chromato-

graphy), 311 Human muscle, 170

IB/1, 162, 187 IB/LI, 162, 187 IB/SRI, 162, 187 Illumination, diffuse, 354 Image

holographic, 341 holographic storage, 355 input devices, 358 quality, 337, 344

Imaging corrections, 309 multiplane dynamic, 309 polaroid, 325 positron, 295, 314

Immobilization casts, 222 Incoherent scattering, 142, ISS Inelastic scattering, 145 Information

capacity, 334, 335, 337 content, 337, 339 loss, 350

Inhomogeneities, 107, 230, 237, 238, 240, 242,244

corrections, 103 medium, 121 tissue, 6, 103, 104, 105

Inner bone, 139, 141, 143, 162, 165, 168, 187 Integration volume, 106 Interface effects, 127 In vivo

fast neutron dosimetry, 76, 78

382

In vivo (conI.)

neutron activation analysis, I, 15, 78 Ion, F fluoride, 312 Ionization

chamber, 17 dosimetry, 22 heavy (ions), 244, 247 tissue-equivalent chamber, 10, 17, 18, 19,

22, 23, 29, 30, 65 Isodose

distributions, 69, 72, 73 shift, 106, 108, III

Isotopic sources, 7

k u,40 KD/LI, 163, 187 Kerma,16

factors, 4, 30, 66, 145, 146, 147 K,145 neutron, 29, 82 neutron-rate, 10 rate, 7, 8 ratio, 25, 27 relative values, 55

Kidney, 139, 141, 163, 165, 169, 187

Laminography, multiplane, 308 Laparotomy staging, 287 Laponite, 136 LET (linear energy transfer), I, 43

restricted, 43 Lincolnshire bolus, 137 Lineal dose-average energy, 48 Lineal energy (y) spectrum, 6,43,47, 50, 51 Lineal frequency-average energy, 48 Lineal saturation-corrected dose-average

energy, 48, 53 Linear attenuation, 106 Linear attenuation coefficients, 170, 335 Linear components IX and /3, 2 Linear energy transfer, I, 6 Liquid, 175

Frigerio, 161 high-pressure chromatography, 311 Witt, 162

Lithium, 12 Liver, 139, 141, 143, 163, 16~, 169, 187 LN II, 163, 187 LN 19, 163, 187 LN ISR4, 163, 187

LogEtronics, 326, 327 system, 328, 332

Index

Lung, 104, 105, 137, 139, 141, 163, 165, 169, 187, 230

burdens, 184 substitute, 135, 159, 174, 180 tissue, 114

LV/LI, 163, 187 Lymphomas, 285 Lyoluminescence, 31, 32

Magneto-optic printer, 363 Mammographic phantoms, 177, 181 Mammographic test strip, 177 Markite, 135 Marrow

red, 139, 141, 143, 146, 163, 165, 169, 187 yellow, 139

Mass angular scattering powers, 144 attenuation coefficients, 142, 143, 147,

149 density, 138, 141, 157, 178 electron stopping, 147 energy-transfer coefficient, 145 photoelectric attenuation coefficient, 149 proton stopping, 147

Materials holographic recording, 352, 353 quasiequivalent, 165

Media, contrast, 225, 248 Medicine, nuclear, 183 Methods

basic data, 150 effective attenuation coefficient,

104, 108 effective atomic number (i), 148 effective SSD, 109 equivalent tissue-air ratio, 113, 116 extended Y, 153, 160 holographic, 341 isodose-shift, III power law tissue-air ratio, 112 quasiequivalent, 157

Microdosimeter, 357 Microdosimetric measurements, 51 Microdosimetric methods, 37, 47, 50 Microdosimetric techniques, 47 Microfiches, 325 Microfilm, 327 Microspheres, phenolic, 157, 159

Index

Minifications camera, 326 systems, 325, 327, 328

Miss marginal, 20 I Mix D, 136 Moderating spheres, 42, 43 Modification range, 244 Monte Carlo, 106, 127

calculations, 126, 239, 245 MSj IS, 161, 186 MSj20, 161, 186 MSj22, 161, 186 MSjGl, 161, 186 MSjG2, 161, 186 MSjLl, 161, 186 MSjSR4, 161, 186 MTF (modulation transfer function), 331 Multiplane dynamic imaging, 309 Multiplane laminography, 308 Muscle, 136, 138, 141, 143, 144, 146,

161, 164, 167, 186 human, 170 striated, 138 substitutes, 137

Neutron, 166, 167 absorbed dose, 40, 41 activation analysis, I, 15, 78 angular distribution, 10 beams, 7, IS, 73 boron capture therapy, 15 californium-252 source, 35 cross sections, 147 dose distribution, 31, 34 dosimeters, 16 dosimetry, 1,4,6, 16, 17, 18,76,78 dosimetry intercomparisons, 16, 79, 80,

81, 82, 83, 85, 86 energy, 6, 7, 35, 43, 44 epithermal, I fast, 1,2,3,4,65,71,78, 171, 180 fission, 15 fluence, 16, 79 generators, 3, 4, 14 generators, d + T, 7 interactions, 145 kerma (rates), 10, 16, 29, 82 output, 8 protocols, 65 radiotherapy, I, 14, 65 reactor therapy, 15

Neutron (conI.)

relative sensitivity, 38 sources, 6 sources, d + T, 13, 14 spectra, 10, 16,44 spectral distribution, II therapy, 8, 65, 71 thermal, I, 15,78 thermal fluence, 15 yield,7

Noise, 329, 330, 333, 349, 350 power spectrum, 335 quantization, 338 RMS, 330 signal-to-noise ratio, 349

Nonlinear behavior, 329 Nuclear medicine, 134, 183 Nuclear track detectors, 30 Nylon-6, 137, 158

Odelcamatic processor, 326 Odelcard filing system, 326 OER (oxygen enhancement ratio), 2 Optical disk, 368, 369 Optical disk recorder, 368 Optical disk store, 368 Optical transfer function, 335 Organs, 139, 141, 159, 165

body, 139, 141, 165 brain, 139, 141, 165 kidney, 139, 141, 165 liver, 139, 141, 165 lung, 139, 141, 165 thyroid, 139, 141, 165

Orthophosphate, calcium, 159 Output

devices, 359 neutron, 8

Oxygen enhancement ratio, I, 2

p + Be, 3, 8, 13 Pair production, 142, 155 Particles,

charged, 143 charged, equilibrium, 56, 59

Pathology tumor, 196 Patient

motion, 212 positioning, 221, 260, 263, 265

Penumbra, 68, 70, 73

383

384

Permissible variation, 154 Phantoms, 176

anthropomorphic, 78, 133, 137, 171, 179 body, 133, 179 dental, 181, 182 heterogeneous, 106 low-contrast detectability, 254 mammographic, 177, 181 regular geometry phantom, 180 thorax, 184

Phenolic microspheres, 157, 159 Photoelectric

absorption, 142, 155 mass attenuation coefficient, 149

Photon, 140, 142, 165, 167 contribution, 74 energy, 6 single detection, 300

Photopeak counting rate, 310 Physics, health, 183 Physiological tracers, 296 Pi mesons, 244, 247 Pixel

-by-Pixel calculations, 239 -by-Pixel techniques, 240 size, 231

Polaroid images, 325 Polycarbonate, 150 Polyethylene, 137, 158, 159

filter, 13 Polyesters, 172 Polymers, 135, 176

based substitutes, 174 cold-cure, 158

Polystyrene, 150 Polyurethane, 136, 137, 159, 172 Pork, 170 Position

patient, 221, 260, 263 signal processing, 304

Positron annihilation, 300 camera performance, 310 distribution, 300 emitters, 297 emitting radiocompounds, 312 imaging, 295, 298, 314 tomography, 297, 298, 314

Positronium, 299 Powders, 175 Power, 135

electron stopping, 144

Power (coni.)

mass angular scattering, 144 proton stopping, 144 stopping, 143

Index

Power law tissue-air ratio method, 112 Presswood, 135 Printer, magneto-optic, 363 Processes

captive, 145 radiation, 143

Processing, position signal, 304 Production, pair, 142, 155 Properties, carcinogenic, 158 Proportional counters, 50

TE, 35, 49, 51 Prostate cancer, 268 Protein, 159

contents, 138 substitutes, 159

Proton, 8, 140, 143, 144, 166, 167, 171,244 beam, 180 stopping power, 144, 147

PVC, 159

Quadratic components, ex and {3, 2 Quality

control, 171, 178 image, 337, 344 radiation, 4, 6, 43, 201

Quasiequivalent materials, 165 method, 157, 160 substitutes, 166

Radiation annihilation, 301 characteristics, 140, 165, 167, 170 fields, 201 processes, 143 quality, 4, 6, 43, 201

Radioactive gases, 312 Radiocompounds, positron-emitting, 312 Radiograph

information capacity, 333, 334, 335 storage, 323, 333 transmission, 104

Radiographic films, 75, 76 Radiology

dental, 181 diagnositc, 134, 181

Index

Radiopharmaceutical, 310, 311, 312 chemistry, 312 synthesis, 310

Radiotherapy, 134, 179 dosimetry, 65 fast neutron, I, 2, 3, 65 neutron, 14

Rando, 135 Rate

dose, 14 kerma, 7, 8 neutron kerma, 10 photopeak counting, 310 TE gas flow, 20

Ratio fitting, 151, 154 kerma, 25, 27 oxygen enhancement, I, 2 scatter/air, 69, 121 signal-to-noise, 349 tissue/air, 74, 106, 109

RBE (relative biological effectiveness), I, 2, 7, 10, 43, 70

Reactor, 7 neutron beams, 15 neutron therapy, 15

Reactors fast neutron, 78 thermal neutron, 78

Real tissue, 137 Recording

electrophoretic, 365, 366 holographic, 346, 349, 352, 353 optical disk, 368

Recurrence, marginal, 215 Red marrow, 139, 141, 143, 146, 163,

165, 169, 187 Reference man, 138, 139 Resins, 158

cold-cure, 176 Resolution, 331, 338, 362

absorption coefficient, 231, 232, 253, 254, 265

spatial, 231, 253, 265, 310, 331, 358 time, 303

Resonance regions, 146 Retnar system, 326 Rice, 136

sodium bicarbonate mixtures, 136 RM/GI, 163, 187 RM/L3, 163, 187 RM/SR4, 163, 187

S phase, 2 Sarcomas, 283 SB/3, 162, 187 SB/4, 162, 187 SBj7, 162, 187 Scanner, CCD line, 357, 358 Scanning

Compton, 236 CT,252 CT specification, 265 dual energy, 265 volumetric, 253, 255

Scatter-air ratio, 69 differential, 121

Scatter flux, 353 Scattering

coherent, 155 Compton, 300 edge phenomenon, 244 elastic, 145 incoherent, 142, 155 inelastic, 145 (mass) angular, powers, 144, 153 rejection, 30 I

Selwyn coefficient, 330 Sensitivity, 310 Shielding, 8, 15

blocks, 76 Shonka plastics, 135, 174 Siemens' wax, 135 Signal

detection, 349 large area transfer, 347 to-noise ratio, 349 transfer, 329

Silicon diodes, 30, 31, 42, 43, 78 Silver recovery, 332 Simulation procedures, 147 SKI I, 162, 187 SK/2, 162, 187 Skeleton, 143, 165

tissues, 139, 141 total, 164, 165, 169, 188

385

Skin, 137, 138, 141, 162, 164, 168, 187 sparing effect, 56

Sodium azide, 160 Solids, 172 Solubility, 157 Sources

californium-252 neutron, 35 coding, 367 isotopic, 7

386

Sources (cont.)

neutron, 6 Spallation, 145 Spatial resolution, 231, 253, 265, 310,

331,358 Spatial structure, 251 Speckle effect, 354 Spectra

d(16) + Be neutrons, 45 energy, 7 fluence, 10, 29 gamma-ray energy, 44 lineal energy, 50, 51 linear energy transfer, 43 neutron, 10, 16, 29 neutron distributions, II neutron energy, 6, 7, 35, 43, 44 photon energy, 6 time-of-flight, 10

Spectrometry: see Spectra Spectrum

noise power, 335 Wiener, 335, 336 See also Spectra

SSD (source-to-surface distance), 74, 106 effective method, 109, 238

Staging clinical, 196 laparotomy, 287

Stopping powers, 143 collision, 153 electron, 144, 147 proton, 144, 147

Storage digital, 356, 357 holographic image, 341, 355 holographic system, 342, 343 optical disk, 368, 369 radiographs, 323, 333

Striated muscle, 138 Sublethal damage, I, 2 Substitute, 135

bone, 135, 180 elementally equivalent, 166 gel, 137 lung, 135, 159, 174, 180 polymer-based, 174 protein, 159 quasiequivalent, 166 tissue, 135, 136, 158, 160, 161, 162, 163,

164, 186, 188 Sucrose, 156

Synchrotron, 8 Synthesis, radiopharmaceutical, 310

Delcomat, 326, 327, 332 holographic storage, 342, 343 hospital communication, 370, 372 LogEtronics, 328, 332 minification, 325, 327, 328 Odelcard filing, 326 Retnar, 326

Index

T AC (time-to-amplitude conversion), 303 Talc, 136 TARs (tissue-air ratios), 109

Batho technique, 238 equivalent, 106 power law, 106 ratio, 109

Target, 8, 12 beryllium, 7, 10, 11 lifetime, 14

TCP (tumor control probability), 216, 217 TE

calorimeter, 66 gas, 17,20,65 ion chamber, 18, 19,22,23,29,30,35,37,

65,87 liquid phantom, 66 plastic, 17 proportional counter, 35, 49, 51 tissue equivalent, 10, 17, 87

Technique Batho TAR, 238 calculational, 237 thera peutic, 196

Teflon, 159 Temex, 135

latex, 136 Test

homogeneity, 178 mammographic strip, 177

Testicular tumors, 284 TH/L2, 164, 188 Therapeutic technique, 196 Therapy

boron neutron capture, 15 brachytherapy, 242 capture, I charged-particle, 242, 244 external beam, 228 fast neutron radio-, I, 2, 3, 71 neutron, 8 reactor neutron, 15

Index

Thermal neutron fluence, 15 neutrons, I, 15, 78

Thixotropic gel, BL8, 136 Thorax, 276

phantom, 184 Three-dimensional treatment planning, 241 Thyroid, 139, 141, 164, 165, 169, 188 Time

exposure, 362 recording, 358 resolution, 303

TOF (time-of-flight), 37 spectra, 10 technique, 39,46

Tissue adipose, 136, 137, 138, 141, 142, 143,

146, 161, 164, 167, 186 average, 139, 141, 165 brain, 170 breast, 139 characterization, 248 delineation, 265 differentiation, 253 equivalent, 135, 146, 148, 166 equivalent ionization chamber, 10, 17 inhomogeneities, 6, 103, 104, 105 lung, 114 real, 137 simulation, 133, 137 skeletal,.l39, 141, 165 soft, 138, 141, 143, 156, 164 substitutes, 135, 136, 158, 160, 162, 163,

164, 186, 188 total soft, 143, 164, 165, 169, 188 See also Tissue-air

Tiss ue-air, 74 equivalent ratio method, 113, 114, 116 power law ratio method, 112

TL: see TLD TLD

dosimeter, 33, 104 response, 78 thermoluminescent devices, 32

Tomochemistry, 249, 251 Tomography

computed, 147, 181, 182, 195,325 emission, 183 positron, 297, 298, 314

Total dose, 201 Total skeleton, 164, 165, 169, 188 Total soft tissue, 143, 164, 165, 169, 188

387

TPX (methylpentene polymers), 137, 158 Tracer

biological, 296 elements, 159 physiological, 296

Track detectors, 42, 43 nuclear, 30

Transmission radiography, 104 Transport calculations, 69 Treatment

charged-particle, 245 modality, 209 outcome, 213 plan, 75, 76, 77, 103, 199, 230, 252, 265,

267 three-dimensional, 241, 267

Triolein, 136 TSK/Ll, 164, 188 TSK/SF3, 164, 188 TST/LJ, 164, 188 Tumor

coverage, 202, 203, 204, 207, 215, 217 head and neck, 280 pathology, 196 regression, 211 testicular, 284

Uncertainties, 23 Units

display, 359 hard copy, 360, 362

Urea, 156 Urethane, Griffith, 161, 186

Vacuum mixing apparatus, 172, 173 pouring apparatus, 176

Van der Graaff, 3 Vidicon tube, 357 Volume

holograms, 352 integration, 106, 121

Volumetric scanning, 253, 255

W,4,17 WN, 22, 24 Wall

effect, 20 thickness, 20

Water, 135, 156, 159, 162, 186

388

Wax Cameron, 135 Harris, 135 Siemens', 135

Weber powder, 137 Wedge filters, 73

White-bone copies, 327 Wiener spectrum, 335, 336 Witt liquid, 162, 187

Yellow marrow, 139

Index

Date post:	12-Oct-2020
Category:	Documents
Upload:	others
View:	0 times
Download:	0 times

Progress In Medical Radiation...

Documents