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Development of improved radiation therapy techniques using narrow scanned photon beams Björn Andreassen
Development of improved radiationtherapy techniques using narrow
scanned photon beams
Björn Andreassen
c© Björn Andreassen, Stockholm 2010
ISBN 978-91-7447-189-2
Printed in Sweden by Universitets service US-AB, Stockholm 2010
Distributor: Department of Physics, Stockholm University
Puff..
Abstract
The present thesis is focused on the development and application of narrowscanned high energy photon beams for radiation therapy. The introduction ofphysically and biologically optimized intensity modulated radiation therapy(IMRT) requires a flexible and accurate dose delivery method to maximizethe treatment outcome. Narrow scanned photon beams is a fast option forIMRT since it is not dependent on mechanically moving heavy collimatorleafs and is largely independent of the complexity of the desired dose distri-bution. Scanned photon beams can be produced by scanning an electron beamof low emittance, incident on a thin bremsstrahlung target of low atomic num-ber. The large fraction of high energy electrons that are transmitted throughthe target has to be removed by a strong purging magnet.
In the thesis a strong purging magnet, coupled with a magnetic scanningmagnet, is presented for an intrinsic electron energy of 50 - 75 MeV and asource to isocenter distance of 75 cm. The available scan area at isocentercan be as large as 43 x 40 cm2 for an incident electron energy of 50 MeVand 28 x 40 cm2 at 75 MeV. By modifying the existing treatment head of theracetrack microtron MM50, it was possible to experimentally produce rele-vant dose distributions with interesting properties from 50 MV scanned nar-row photon beams while deflecting the transmitted electrons onto a simplifiedelectron stopper. The deflection of the transmitted electrons was studied bothexperimentally and by the Monte Carlo method.
With high energy photons, treatment verification is possible through PET-CT imaging of the positron annihilations induced by photonuclear reactions inthe patient. Narrow scanned high energy photon beams is the ideal beam qual-ity since the activation efficiency and the effective photon energy will be moreuniform than the filtered photon beam from a full range bremsstrahlung target.The 11C activity induced by a 50 MV scanned narrow photon beam was mea-sured using PET-CT imaging and compared with Monte Carlo simulations.The combination of fast flexible dose delivery with treatment verification us-ing PET-CT imaging makes narrow high energy scanned photon beams a veryinteresting treatment modality for biologically optimized adaptive radiationtherapy.
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List of Papers
This thesis is based on the following papers, which are referred to in the textby their Roman numerals.
I Brahme, A., Gudowska, I., Larsson, S., Andreassen, B., Holm-berg, R., Svensson, R. (2006) Application of Geant4 in the devel-opment of new radiation therapy treatment methods. Proc. 9thConf on astroparticle, Particle and space physics, detectors andmedical physics applications, 451-461
II Andreassen, B., Svensson, R., Holmberg, R., Danared, H.,Brahme, A. (2009) Development of an efficient scanning andpurging magnet system for IMRT with high energy photonbeams. Nucl. Instr. and Meth. A, 612:201-208
III Andreassen, B., Strååt, S., Holmberg, R., Näfstadius, P., Brahme,A. (2010) Fast IMRT with narrow high energy scanned photonbeams. Submitted to Med. Phys.
IV Svensson, R., Andreassen, B., Lind B, Brahme, A. (2010) De-velopment of a diagnostic treatment unit combining fast radio-biologically optimized adaptive IMRT with in vivo 3D dose de-livery verification and tumor responsivness by PET-CT imaging.Manuscript
Reprints were made with permission from the publishers.
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Contents
Abstract . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . ivList of papers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . vTable of contents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . vii1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12 Radiation therapy optimization . . . . . . . . . . . . . . . . . . . . . . . . . . . . 33 Methods of delivering IMRT . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5
3.1 Dynamic MultiLeaf Collimation (DMLC) . . . . . . . . . . . . . . . . . . . . . . . . . . 53.2 Tomotherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 63.3 Single Arc Therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 63.4 Scanned Beams . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7
4 Scanning and purging magnets for narrow scanned photon beams . . 84.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 84.2 Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 94.3 Magnet design . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10
5 Dose distributions, measurements and simulations . . . . . . . . . . . . . . 125.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 125.2 Modification of the treatment unit . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 125.3 Beam measurements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 135.4 Spot matrix . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 145.5 Monte Carlo simulations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16
6 Positron emission tomography measurements . . . . . . . . . . . . . . . . . 186.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 186.2 Positron activation kernels of high energy photons beams in graphite . . . . . . 206.3 Treatment unit and phantom . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 206.4 PET-CT camera . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 216.5 Monte Carlo . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 216.6 Experimental results and simulations . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22
7 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 28Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 30Bibliography . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 31
1. Introduction
Today, about one third of the population in Sweden will at some point in theirlife develop cancer (SBU, 2003). Since most types of cancers develop late inlife and the fact that the average life span is increasing means that this figurewill continue to increase. Many of the most common types of cancer, such asnon-small lung cancer and breast cancer, are primarily treated with surgeryin combination with chemotherapy and/or radiation therapy. More than halfof the patients diagnosed with cancer in Sweden will, during their treatment,receive radiation therapy either as a curative or palliative (alleviate pain orsymptoms) treatment. Although a lot of research has gone into gene therapyand molecular biology, an effective alternative to surgery and radiation therapyis not realistic in the near future. The continued development and improve-ment of radiation therapy is thus of great importance, for both the survivaland quality of life of cancer patients.
Shortly after the discovery of X-rays in the late 19th century by WilhelmConrad Röntgen, the first cancer patient was treated with external radiationtherapy. The outcome and possibility to treat deep seated tumors with radia-tion therapy has since then been improved by technical innovations such as the60Co unit, betatron, traveling and standing wave linear accelerators and the mi-crotron, increasing the energy and quality of the photon and electron beams.The main goal in curative radiation therapy is to eradicate the tumor cellswhile sparing the healthy surrounding tissue. To do this effectively, detailedknowledge is required about the shape and position of the tumor and organsat risk, coupled with methods to accurately calculate the 3D dose delivery.The introduction of computer tomography (CT) in the 70’s, largely replacing2D X-ray film, made it possible to image the tumor in detail in 3D. Simulta-neously, delivery methods to conform the radiation dose to the actual shapeof the tumor, conformal radiation therapy (3DCRT) (Takahashi, 1965) wasdeveloped. The next milestone in radiation oncology was the introduction ofintensity modulated radiation therapy (IMRT) in the early 80’s (Brahme et al.,1982), which today in the clinic is a widely used technique. Later advancesin diagnostics, such as magnetic resonance imaging (MRI) and positron emis-sion tomography, most recently combined with computer tomography (PET-CT), has further advanced the accuracy in distinguishing the tumor from nor-mal tissue topograhy. The radiation delivery methods based on IMRT havebeen continuously improved, using multileaf collimation (MLC) or tomother-apy/arc therapy to shape and modulate the radiation beam. Equally important
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Chapter 1 Introduction
as optimizing the dose delivery to physical objectives is to include the biologi-cal responses of both tumor and normal tissues. In the BIO-ART (biologicallyoptimized predictive assay based radiation therapy) approach (Brahme, 2003,2009), PET-CT imaging and radiopharmaceuticals such as FDG are used tostudy the tumor response between fractions and with PET-CT dose deliveryimaging, used to make adaptive corrections to the treatments.
Today, in conventional photon therapy units, an accelerator delivers anelectron beam (4 - 20 MeV) to a 360 deg rotary gantry where a full rangebremsstrahlung target is combined with a flattening filter, beam blocksand multileaf collimators to create, shape and direct the desired beam. Byinstead magnetically scanning a high energy electron beam (50 - 75 MeV),incident on a thin bremstrahlung target, the treatment can be improved (PaperI and III). By not only relying on block or multileaf collimators the speedis mainly limited by the magnetic scan speed. Sharper dose gradients areeasily achieved at the edges of the MLC (Paper III) and treatment verificationis possible through photonuclear induced positron activity measurable byPET-CT imaging (Paper IV). The development of the thin target scaningbeam technique is the focus of the present thesis.
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2. Radiation therapy optimization
Therapy optimization tries to solve the problem of how to best deliver a hightumor lethal dose without damaging surrounding normal tissues. One of themost useful degrees of freedom in dose delivery is to intensity modulate theincoming beams, to avoid severe side effects to normal tissues. The radiationbeam, when using intensity modulated radiation therapy (IMRT), is not onlyshaped to fit the projection of the target volume (conformal radiation ther-apy), the photon fluence is also varied across the delivered beam. When treat-ing tumors of complex shape the delivered dose and treatment outcome canbe substantially improved by having multiple intensity modulated beams fromdifferent directions. Intensity modulated radiation therapy has been under con-stant development since the early 1980’s when the first paper on the subjectwas published (Brahme et al., 1982). Since then IMRT has gained wide ac-ceptance in the clinic and several investigations indicate the clinical benefitsof IMRT (Westmark et al., 2002; Veldeman et al., 2008). With physical opti-mization using IMRT, the delivered dose distribution is better confined to theinternal target volume while sparing sensitive surrounding tissue. This methodrepresents an improvement in the treatment outcome compared to conformalradiation therapy, when treating complex shaped tumors. To take full advan-tage of IMRT, eradicating the tumor with minimal normal tissue toxicity, thereis a need to develop effective optimization methods to not only optimize thephysical dose delivery but also to utilize biologically optimized radiation ther-apy (Brahme, 1995, 2001, 2004). By including tissue and tumor response datain the treatment planning process, the beam profiles and number of beams canbe biologically optimized to significantly improve the treatment outcome. Thedose to the normal tissue is generally decreased, while the the dose to the tu-mor can be locally increased since part of the tumor may be hypoxic and thusmore radioresistant. Biologically optimized IMRT can improve the treatmentoutcome by delivering a higher dose to the hypoxic regions and reducing itslightly to well oxygenated tumor regions and normal tissue.
Modern biologically based intensity modulated treatment techniques canincrease the treatment outcome by as much as 20-30% for advanced resistanttumors (Brahme, 2005). Such improvements can only be achieved when all theimportant links in the radiotherapy chain are well integrated and of high qual-ity. This has to work all the way from tumor diagnostic and predictive assayof radiation sensitivity using PET-CT imaging to advanced adaptive radiobi-ologically based treatment optimization and high accuracy radiation therapy
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Chapter 2 Radiation therapy optimization
dose delivery and treatment follow up, such as using the BIO-ART approach(Brahme, 2003). The present thesis is focused on methods for rapidly deliver-ing intensity modulated dose distributions by scanning narrow photon beamsover the tumor volume.
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3. Methods of delivering IMRT
Intensity modulated radiation therapy can be delivered in a number of ways.The first and technically simplest way is to use prefabricated metal compen-sators, this method is time consuming and not very common today. Anotherway is to use the MLC, either in a static setting or to dynamically modulate thefield. Tomotherapy does not have a classical MLC but delivers a fan beam ofradiation through a binary collimator system while rotating around the patient.Today a form of simplified fast rotational therapy is also generally availableusing a single arc rotation. Originally designed for brain treatments, a narrowdynamic robotic linac is also an option for IMRT. And finally IMRT can bedelivered by rapidly scanning a narrow photon, electron or ion pencil beam.
3.1 Dynamic MultiLeaf Collimation (DMLC)In conformal radiation therapy the introduction of MLC’s greatly improvedthe possibility of shaping the beam and thus better conform the delivered doseto the target volume (Brahme, 1988a,b; Källman et al., 1988). A typical MLCconsists of 20-80 tungsten leaves, on each side, that can be moved individu-ally to create field openings with complex shapes. Today, the most commonway to do an IMRT treatment is by using a multileaf collimator (Yu et al.,2008) in a step and shot mode or by moving the leaves during an irradia-tion. In the multiple static field technique (step and shoot), introduced in 1991(Boyer et al., 1991), the leafs are moved between the different fields but notduring irradiation. The delivered photon fluence distribution, from multiplefields, to the target volume is then intensity modulated as the different fieldsare superimposed. The other more advanced option is to move the leafs duringan irradiation. In 1978 it was shown that it is possible to create a wedge shapeddose distribution by moving one collimator towards an opposite stationarycollimator (Kijewski et al., 1978). A general technique to create the desireddose distributions by dynamic multileaf collimation was developed in the mideighties (Källman et al., 1988), generally independant of how the delivery wasperformed. The sliding window approach to a DMLC treatment was proposedin 1992 (Convery and Rosenbloom, 1992), where the leaves move from oneside to the other with a variable gap/window between opposite leaves. Theone dimensional fluence intensity at any point is then decided by the time itis under the window. Since the speed of individual leafs can be altered, it ispossible to create arbitrary dose distributions. The analytical solution was first
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Chapter 3 Methods of delivering IMRT
given 1993 (Svensson et al., 1993) and the technique was developed further(Spirou and Chui, 1994; Stein et al., 1994) using the analytical solution to theproblem and is still in use in the clinic today. There are a number of differentcommercially available DMLC systems with both step and shoot and dynam-ical motion of the multileaf collimators. Although the use of metal compen-sators allows dose distributions similar to a DMLC, the preparation time ismuch higher in producing the compensators. The speed of an IMRT treatmentwith a DMLC depends on the complexity of the desired shape as the leafshave to travel across the field. Today treatment times from 10-20 minutes areachievable, depending on the complexity of the treatment and on the vendor(Butler et al., 1994; Bortfeld, 2006).
3.2 TomotherapyTomotherapy was developed in the early 90’s (Mackie et al., 1993; Carol,1995) with the first patient trials in 1994 (Mackie et al., 1993). The first com-mercially available system was NOMOS MIMiC Peacock tomotherapy sys-tem while today the spiral system emanating from the University of Wisconsinis dominant. In tomotherapy the target volume is divided into narrow slices,which are irradiated separately by a fan beam while rotating the gantry aroundthe patient. While the gantry rotates at a constant speed the fan beam is colli-mated by a narrow binary collimator, which can be individually opened/closedduring each segment to modulate the intensity in each of the 64 vanes. Thetreatment is planned for 51 fixed directions covering a 360 degree rotation,with the total treatment time determined by the number of arcs required tocover the entire target area. Several investigators have found that tomother-apy has a dosimetric, at least in the planned target volume (PTV), advantageover other IMRT methods (Cao et al., 2005; Oliver et al., 2009) while it gen-erally requires a longer treatment planning time and has considerable longertreatment time then single arc therapy (Oliver et al., 2009).
3.3 Single Arc TherapyAs an alternative to tomotherapy the idea to use multiple overlapping arcs witha conventional MLC doing the beam shaping for each angle in each arc wasproposed in 1995 (Yu, 1995). For each beam angle the intensity modulation isachieved by superimposing two dimensional uniform distributions from sev-eral overlapping arcs. Although this technique has been proven to producehighly conformal dose distributions (Cao et al., 2005), it has not been appliedmuch in clinical practice (Yu et al., 2008). This can be partly credited to longtotal treatment time for large tumors as each arc takes approximately 1 minute(Tang et al., 2010) and as a result research and development has gone into
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3.4 Scanned Beams
single arc therapy, sometimes referred to as arc modulated radiation therapy(AMRT) (Cameron, 2005; Otto, 2007). There are today commercial productsbased on single arc IMRT such as Varian’s RapidArc and Electa’s VMAT. Thebasic principle behind single arc therapy is that the delivered dose distributionis very similar if all the homogeneous beams from one beam angle, for themultiple arc variant, is delivered as a single portal (Tang et al., 2010). Duringthe rotation the dose rate can be varied, but there are serious restrictions on theleaf translation movement to avoid excessive treatment times. This often leadsto a reduced treatment accuracy as compared to Tomotherapy or biologicallyoptimized few field techniques.
3.4 Scanned BeamsIn the above mentioned IMRT delivery methods a broad homogeneousphoton beam, created by full range bremsstrahlung targets and flatteningfilters, is used for the intensity modulation. For photon therapy, IMRT byscanning the photon beam was proposed in the 80’s (Brahme, 1987, 1992;Svensson et al., 1998b; Svensson and Brahme, 1996) and some experimentalwork has been made (Blomquivist et al., 1998; Svensson et al., 1998b;Rosenqvist, 1990; Lind and Källman, 1990). A conventional photon therapyunit has a full range bremsstrahlung target where no primary electrons aretransported through the target while a treatment unit designed for IMRT byscanning the beam would, to produce a narrow photon beam, have a thintarget. Scanning a broad photon beam from a full range bremsstrahlung targetallows much less Intensity modulation and as such could only complement aMLC, while a sufficiently narrow beam could, as has been shown in paper III,in some cases modulate the beam without the use of the MLC. Using a thintarget to produce bremsstrahlung beams, a large part of the electrons incidenton the target are transported through and have to be adequately purged fromthe photon beam. Since scanning the beam requires no mechanical movementthe IMRT delivery time can be reduced, compared to using a broad beamand a DMLC, and the photon energy fluence and mean energy is morehomogeneous and thus better suited for treatment verification by PET-CT(Janek et al., 2006; Brahme, 2003, 2009). By scanning the beam it is possibleto deliver IMRT beams in fractions of a second (for a 10 x 10 mm2 area) andcould therefore be used during Single Arc Therapy, to create more optimalIMRT fields from every angle and thus improve the dose delivery. For Iontherapy, beam scanning is done in clinical practice. However, ions represent aspecial case since the source is by necessity narrow and broadening it withcompensators would result in unwanted fragmentation of the nucleus andunwanted energy dispersion. As such scanning the beam across the targetvolume is a good choice for ion therapy.
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Chapter 4 Scanning and purging magnets for narrow scanned photon beams
4. Scanning and purging magnets fornarrow scanned photon beams
4.1 IntroductionTo maximize the clinical possibilities with a treatment unit based on scannednarrow photon beam IMRT, the photon beam should be as narrow as possible.Reducing the bremsstrahlung target thickness, down to a few mm of a lowZ material, results in a narrow bremsstrahlung beam but also a large amountof transmitted high energy electrons. Increasing the intrinsic electron energydecreases the width of the bremsstrahlung beam but also reduces the deflec-tion capabilities of the scanning system and the purging magnet. And finally,reducing the effective source to tumor distance decreases the width of thephoton beam and increases the accuracy in dose delivery. At the same timethe scanning angle needs to be increased to allow the treatment of also largetumor volumes.
d)
a)b)
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Figure 4.1: Magnetic flux density in a plane through a complete, gantry based, scaningand purging magnetic system for an incident electron energy of 75 MeV. a) 97 degbending magnet, b) first scanning magnet, c) last scanning magnet, d) purging magnet.
The stigmatic magnetic scan system used today in the racetrack microtronMM50 is capable of scanning a 50 MV photon beam, from a full rangebremsstrahlung target, to a maximum field size of 40 x 32 cm2 at a sourceto isocenter distance(SID) of 100 cm. When using narrow scanned photonbeams only a small amount of the delivered photon fluence is incident on thecollimators and the requirements on the collimation system is thus greatly
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4.2 Methods
reduced. By only using a multileaf collimator, no wedge filter or blockcollimator are needed, the SID can be shortened to 75 cm with the samespace available for the patient. A simulation of the magnetic flux densityof the complete magnetic scaning and purging system is shown in Fig 4.1.For the new system the critical components are the last scanning magnetand the purging magnet. The first scanning magnet and the bending magnethave parallel poles and increasing the strength of these magnets is relativelyeasy. In paper II we examined the potential of a new design for a scaning andpurging system for an incident electron energy of 50 - 75 MeV, a SID of 75cm and a bremsstrahlung target of 3 mm beryllium.
4.2 MethodsThe geometry of the yokes for both the scanning and purging magnet wasdeveoped using the CAD program Solid Edge, the geometry was then im-ported into the magnetic field simulation software Opera3D/Tosca. Opera3Dsolves the differential equations governing the magnetic field by a finite ele-ment method. The B-H characteristics of a commercially available steel typewas used. An example of a simulation of flux density in a plane through thegap of both magnets is shown in Fig 4.2 along with the the mesh in that plane.The size of the mesh is smallest in the gap region to increase the accuracy inthe region with both the highest gradient and the highest flux density.
Figure 4.2: Magnetic flux density in a plane through the gap of the scanning andpurging magnet with parallel deflection angles. The mesh grid in the plane is alsoshown.
Opera3D allows tracking of monoenergetic charged particles through thesimulated magnet field. By tracking electrons with intrinsic energy of 50 - 75MeV down to the exit of the bremsstrahlung target and there continuing thetransport with the mean energy of the transmitted electrons (ICRU, 1984a,b)the deflection capability of the purging magnet was examined. Similarly,
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Chapter 4 Scanning and purging magnets for narrow scanned photon beams
the photon scan was determined by tracking the intrinsic electrons downto the target and there, at the effective bremsstrahlung production depth(Svensson and Brahme, 1996), projecting the position and direction of thephoton down to the isocenter level. Analytical expressions for the angularand latteral spread of the transmitted electron beam was used to estimate thelevel of electron contamination in the scanned photon field.
4.3 Magnet designSince the purging magnet is stationary the shape of the divergent gap is de-termined by the scanning limits of the photon beam. The new SID of 75 cmincreases the divergent opening angles of both the scanning and purging mag-nets and thus decreases the strength of the magnets. To improve the integralmagnetic field of the purging magnet the length, in the beam direction, is in-creased and it is placed closer to the last scan magnet to reduce the gap at itsentrance. Furthermore the dimensions of the purging magnet are increased,the coil structure is leaning downwards and its size is more than doubled. Thestrength of the scanning magnet is increased by altering the shape of the yoke,increasing the power of the coils and the length of the magnet in the beamdirection. In Paper II (Figure 2) the scanning magnet and the purging magnetare shown either in a parallel or in a orthogonal arrangement. The orthogo-nal placement of the purging magnet resulted in a weaker magnetic field inthe purging magnet, due to increased leakage field into the gap part of thescanning magnet, and leaves less space for an eletron stopper. In Fig 4.3 thescanning and purging magnets are shown operating in the parallel arrange-ment with a divergent gap corresponding to an available photon scan of ±20cm at a SID of 75 cm. Many models with different scan area and distance be-tween the magnets were tested, the most promising ones with a scan of ±15and ±20 cm, are presented in Paper II.
The magnetic scanning and purging system is designed to be used with twostationary electron stoppers, one for each scan direction as described in paperIV. These will have an entrance channel to which the transmitted electronbeam is directed irregardless of the scan of the photon beam. The proximityof the two magnets results in an overlapping of the magnetic fields. Since thescanning of the photon beam and the deflection of the transmitted electronbeam can not be treated separately, the magnetic system was put through anextensive iterative process to find the scan limit of the photon beam whileallays deflecting the transmitted electron beam to the same area on the electronstopper.
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4.3 Magnet design
Figure 4.3: Scanning and purging magnets in a parallel arrangement and a rudimentaryelectron stopper.
The available scan area at the isocenter can be as large as 43 x 40 cm2 for anincident electron energy of 50 MeV and 28 x 40 cm2 at 75 MeV. Simple calcu-lations on the dose contribution from the transmitted electron beam comparedto the photon dose show that it is possible to use a thin bremsstrahlung targetto produce narrow scanned photon beams and get satisfactory deflection of thetransmitted electron beam using a stationary purging magnet system. This isbased on the assumption that it is possible to safely collect all the transmittedelectrons in an effective electron collector placed downstream of the purgingmagnet.
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Chapter 5 Dose distributions, measurements and simulations
5. Dose distributions, measurementsand simulations
5.1 IntroductionThe main challenge when using narrow scanned photon beams for IMRT isto effectively deflect the high energy transmitted electron beam. This problemhas been treated analytically (Svensson et al., 1998b) and by particle tracking,through the simulated magnetic field, in combination with analytical transportequations, presented in paper II. A more useful approach, is to incorporatethe magnetic fields of the scanning and purging magnets (Andreassen et al.,2009) into a Monte Carlo simulation, taking into account the most relevantcomponents of the treatment head. This allows simulations of both the directelectron contamination of the useful photon beam, the scattered radiation andbremsstrahlung contamination produced in the dedicated electron stoppers,placed on each side down stream of the purging magnet.
By modifying the treatment head of the MM50 and using the existing purg-ing and scanning system it was possible to create relevant dose distributionswith the scanned photon beam while the transmitted electron beam was de-flected onto one of the block collimators with the multileaf collimator under-neath. This allowed us to examine the scanning system, the deliverable dosedistributions as well as the treatment time and the impact of the electron andphoton contamination on the delivered dose distribution. In paper III, the ex-perimental geometry were introduced into a Geant4 simulation incorporatingthe magnetic field of the existing purging magnet, simulated with Opera3D.This allowed detailed analysis of the experimental results and provides a ba-sis for further simulations of a dedicated electron stopper with a scaning andpurging system designed for narrow photon beam scanning.
5.2 Modification of the treatment unitThe racetrack microtron MM50, from Scanditronix AB, has a stigmatic scan-ning system that creates a homogeneous photon dose distribution, by mag-netically scanning the incident electron beam in a circular pattern on thebremsstrahlung target. For the experimental studies presented in paper III, thefull range target of wolfram and copper was replaced with a thin target of 3 -6 mm beryllium normally used to reduce the energy of electron beams in 1.7and 3.3 MeV steps. The purging magnet, placed directly after the last scan-
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5.3 Beam measurements
ning magnet, was designed to remove photon produced leptons at the end ofthe bremsstrahlung target and was only able to produce a static magnetic field,and not dynamically alter the magnetic field strength depending on the scan ofthe photon beam. The part of the quasi-gaussian distributed transmitted elec-tron beam, that missed the top of the block collimator, therefore varied withthe scan of the photon beam. The segmented ion chamber that is normallyused for quality assurance and to have a steady dose rate during treatments,was removed. This was done to not damage the ion chamber but also so thatthe transmitted electron beam would be as narrow as possible. One of theblock collimators was fully extended and served as an electron stopper withthe multileaf collimator covering the same area.
5.3 Beam measurementsTo have the maximum amount of separation between the bremsstrahlung beamand the transmitted electron beam, the electrons where deflected in the oppo-site direction to the scan of the photon beam.
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Figure 5.1: Film measurement at downstream end of the purging magnet and the onecm thick wolfram collimator. The photon kernel, from 50 MeV electrons incident ona bremsstrahlung target of 3 mm beryllium, is scanned 130 mm at isocenter. a) Thepurging magnet is turned off so the transmitted electron beam follows in the directionof the photon beam. b) the purging magnet is turned on, separating the photon kernelfrom the transmitted electron beam.
In Fig 5.1, film measurements taken at the downstream end of the purgingmagnet and 1 cm of wolfram collimator, no buildup material was used in frontof the film and the treatment head as filled with air under normal pressure.The photon beam, generated by an electron beam of 50 MeV incident on 3mm of beryllium, was scanned to 130 mm at isocenter, 1 m below the target.In Fig 5.1a the transmitted electron beam follows the direction of the photonbeam, since the purging magnet is turned off. A background of scattered elec-
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Chapter 5 Dose distributions, measurements and simulations
trons is also seen in the figure. In Fig 5.1b the purging magnet is turned on andthe electron beam is deflected in the opposite direction of the photon beam.The energy dispersion of the transmitted electron beam is clearly seen, withpart of the electron beam deflected into the wolfram collimator.
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Figure 5.2: a) Bremsstrahlung kernel from a 3 mm beryllium target, measured withEDR2 film at SID 100 cm. b) The measurement setup, the film is tightly packed be-tween 8 cm of PMMA.
5.4 Spot matrixThe photon kernel was measured in a plane parallel to the beam with EDR2type film as shown in Fig 5.2a and in orthogonal planes shown in Fig 5.3. Ananalytical fit of the measured kernel was used as input for an inverse planningiterative deconvolution scheme (Lind, 1990). The resulting optimal dose dis-tribution minimized the over dosage needed at the edges and corners, to makesure that the there is no under dosage in the target area. This algorithm alsominimizes the delivered dose distribution outside of the target area.
14
5.4 Spot matrix
/ cm
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l. D
ose
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ose
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Figure 5.3: Film measurement, at isocenter below 8 cm of PMMA, of 50 MV pho-ton kernel generated by a 50 MeV electron beam incident on 3 and 6 mm berylliumbremsstrahlung targets. The relative dose is normalized to the forward photon yieldfrom a 6 mm beryllium target. a) 3 mm beryllium bremsstrahlung target, b) 6 mmberyllium bremsstralung target.
For the leaning cardioid dose distribution presented in paper III:figure 7the spot distribution is shown in Fig 5.4a. The iterative deconvolution schemedelivers a quasi continuous pencil beam intensity distribution across the entirefield, which is not practical from a scanning time point of view. Therefore thespot matrix was discretized to a number of equal intensity pencil beams spotsusing a Levenberg-Marquard optimizer with an analytical approximation ofthe pencil beam kernel and the analytical derivatives of the kernel. For thecardioid shape distribution a spot matrix, shown in Fig 5.4b, of 250 spots wasused which, with a pulse repetition of 200 Hz, takes little more then a secondto produce. The number of complete scans required to achieve the desireddose level depends on the size of the field but also the energy of the incidentelectron beam and bremsstrahlung target as discussed in paper III.
15
Chapter 5 Dose distributions, measurements and simulations
/ cm
/ cm
Nr
Sp
ots
a) b)
Figure 5.4: Spot distribution to create a leaning cardioid shape dose distributionwith photon kernels from an 50 MeV electron beam incident on a 3 mm berylliumbremsstrahlung target. a) Spot distribution with continuous intensity from the inverseplanning software. b) Spot distribution in discrete steps.
5.5 Monte Carlo simulationsTo compare with the experimental results, Monte Carlo simulations wererun taking the magnetic field into account. The Monte Carlo simulationswere made using GEANT4 version 9.2.3 (released 19 December 2009)with G4EmStandardPhysics. This version includes a more accuratemultiple scattering model, G4UrbanMscModel93, that has been tuned to fitexperimental results of the angular distribution from 13 MeV and 20 MeVelectrons scattered in thin foils (Faddegon et al., 2009). The root mean squarescattering angle
√θ̄ 2(z0) and root mean square radial spread
√r̄2(z0) were
estimated to 4.2 mrad and 0.39 mm when incident on the bremsstrahlungtarget of the MM50 racetrack at Karolinska (Svensson and Brahme, 1996)and were used as initial values. The treatment head geometry used in thesimulations is shown in Figure 2 in paper III. Figure 5.5a shows Monte Carlosimulation of the dose distribution at isocenter, below 95 mm of water, froma scanned 50 MV bremsstrahlung kernel, incident on the water phantom,produced in a target of 6 mm beryllium. Also in Fig 5.5a is the total dosedistribution at the same depth. In Fig 5.5b the total dose distribution has beenseparated into dose distribution from the transmitted electrons, dose fromparticles that have interacted with the block or MLC and dose distributionfrom the photon kernel. The dose contribution from the transmitted electronshave a maximum of 3.5%, normalized to the total dose distribution. Thiscan be compared to the 2.0% from 3 mm beryllium as shown in paper III.Although the mean energy of the transmitted electrons for a 6 mm berylliumtarget is lower then from a 3 mm target, multiple scatter in the target resultsin an increased electron contamination of the photon beam.
16
5.5 Monte Carlo simulations
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Figure 5.5: a) Monte Carlo simulation of the the total dose distribution and of the dosedistribution from particles that were photons incident on the water phantom and whichdid not interact with the block or MLC. b) The lower region complemented with twoMonte Carlo simulations: Dose distribution, normalized to the total dose distribution,from particles that have interacted with the block or MLC. Dose distribution, normal-ized to the total dose distribution from particles that were electrons incident on the topof the water phantom and had not interacted with the block or MLC.
17
Chapter 6 Positron emission tomography measurements
6. Positron emission tomographymeasurements
6.1 IntroductionIn positron emission tomography (PET) two gamma rays created from theannihilation of a positron and an electron are detected, in coincidence, bydetectors that surround the volume containing the positron emitter. The gen-erated 511 keV gamma rays are almost anti-parallel (emitted close to 180 degrelative each other) which gives the PET system information on the line, be-tween the two detector elements, where the annihilation most likely took place(line of response). By registering a large number of annihilations it is possibleto reconstruct an image of the annihilation distribution and thus the effectivesource of the positron emitters inside the volume. During the last decade thecombination of PET and X-ray computer tomography (CT) has offered a newlevel of accuracy in the delineation of the tumor on the background of the nor-mal tissue anatomy (Brahme, 1987). A number of new radiopharmaceuticals(Apisarnthanarax and Chao, 2005), such as FDG and FLT, which allow thestudy of tumor metabolism for proliferation, and Fmiso for hypoxic regions,are under investigation making PET-CT one of the leading research areas toimprove radiotherapy.
There are a number of error sources, affecting the resolution of the pro-duced image, which need to be addressed. In general, it is the position of thepositron emitting radionuclide and not the annihilation position which is of in-terest. The positron has an average range of up to a few mm depending on theradionuclide and the material. The spatial resolution of the detectors naturallyaffects the resolution of the image but also a small uncertainty in the directionof the two gamma rays (4mrad) plays a role. For a detector ring of 80 cm di-ameter and a detector coincidence pair resolution of 0 mm, the loss of spatialresolution in the image when detecting 11C produced positron annihilations is2.3 mm. Increasing the detector pair resolution to 2 mm or 6 mm gives a imageresolution loss of 3.2 mm and 6.6 mm respectively (Wernick and Aarsvold,2004). The detector coincidence pair resolution is thus critical for the imageresolution and for the value of the functional information it contains. The im-age resolution is also affected by the parallax error, where one or both of thegamma rays interacts in one of the detector crystal elements after having tra-versed an adjacent crystal, creating a false line of response. The reconstructedimage is fraught with noise from scattered and random coincidence. If two an-
18
6.1 Introduction
nihilations are close enough in time the produced gamma rays can be detectedsimultaneously, even though they are not correlated, generating a false line ofresponse and noise in the image. Similarly one of the gamma rays might bescattered in the material before reaching the detector, creating a false line ofresponse.
One of the most useful approaches to increasing the detected coincidenceevents and reducing the noise is to increase the solid angle of the PET detector.This can be done be reducing the diameter of the detector ring or removing theaxial collimation and increasing the axial dimension (3D volume imaging).The last approach increases the random coincidence but increases the truecoincidence to a higher degree.
The positron emitter is a radionuclide, normally produced in a cyclotron,attached to suitable marker and injected into the patient. The most commonradiopharmaceutical is F-18 FDG, which is used as a marker of glucosemetabolism, has a half-life of 118 min. Since tumor cells have a highmetabolism, the FDG uptake imaged by PET can be used to study the densityof clonogenic cells in the primary tumor and in possible metastasis locations.There are a number of different radiopharmaceuticals under investigation thatmight prove superior when trying to examine, for example, hypoxia.
By combining the PET system with a CT system the accuracy in locatingregions with high positron emission is considerable increased, compared tohaving a separate PET and CT system. The added information in the CT imageallows more accurate correction for attenuation in the PET image and helps inisolating the true tumor uptake of the radiopharmaceutical from non specific-uptake.
In radiotherapy with ions or high energy photons, positron emitters are cre-ated by the radiation beam which offers a possibility to do in vivo treatmentverification both during and between fractions. Comparing the PET image tothe treatment plan, deviations in the beam delivery can be detected and cor-rected for in the next fraction or possibly in the future, corrections could bemade on line. The BioArt (Brahme, 2003) concept combines all relevant as-pects, such as PET imaging through FDG and nuclei reactions with the treat-ment beam, into a clinically applicable technique for treatment optimization.
During ion therapy direct nuclear reactions create positron emitters and in-beam PET for clinical applications have been installed at GSI in Germany andHIMAC (Iseki et al, 2003) in Japan. For ion therapy PET imaging can be usedfor determining the position of the bragg peak at the end of the particle range(Enghardt et al., 2004).
With high energy photon beams, positron emitters are created through pho-tonuclear (γ ,n) reactions in the irradiated volume. For body tissue the pre-dominant target nuclei are 12C, 16O and 14N with the respective productionthresholds at 18.7 MeV, 15.7 MeV and 10.6 MeV (Chadwick et al, 2000). Thecomposition of these nuclei, in clinically relevant material, vary considerablywith oxygen dominating in bone and muscle while carbon dominates in adi-
19
Chapter 6 Positron emission tomography measurements
pose tissue (ICRU, 1992). The relative short half-life time of 15O, 2.04 min,compared to 11C at 20.4 min means that to take full advantage of the infor-mation gained by monitoring the positron annihilation from 15O, the PET ex-amination should be started shortly after the irradiation. Recently, an off-linestudy of the PET activation in PMMA and a frozen pig leg, by 50 MV pho-tons, compared the measured annihilation distribution to the treatment planfor the dose distribution (Janek et al., 2006). This study shows the potentialof PET measurements for treatment verification with high energy photons butalso emphasizes the need to start the measurements as soon as possible af-ter the irradiation. For a treatment unit that combines radiotherapy with PET(RT-PET-CT), there is also a possibility to monitor the positron annihilationsgenerated by pair productions during the irradiation. Monte Carlo simulations(Müller and Enghardt, 2006) and an experimental study (Kluge et al., 2007)have verified the possibility of having an on-line PET camera when irradiatingwith high energy photon beams. Given the differences in nuclei compositionin human tissue, a RT-PET-CT could potentially be used to monitor the patientmovement during irradiation (Müller and Enghardt, 2006).
6.2 Positron activation kernels of high energy photonsbeams in graphiteThe bremsstrahlung from a thin target has a higher mean energy and thus willgive a deeper dose maximum and depth dose distribution, than a full range tar-get. This is not only of interest for treatment planning but also for the level ofinduced positron activity in the body, which can be used for in vivo treatmentverification using PET-CT imaging (Janek et al., 2006; Brahme, 2003, 2009).Also when the narrow photon pencil beam is scanned the average photon en-ergy fluence is much more homogeneous than when using a broad beam gener-ated by a thick target followed by a flattening filter(Sätherberg and Karlsson,1998). In fact, a narrow scanned high energy photon beam is the ideal acti-vation source since the activation efficiency and the effective photon energywill be more uniform all over the target volume due to the absence of flat-tening filters and low multiple electron scattering with a thin Beryllium target(Karlssson et al., 1993). This makes narrow high energy beam scanning veryinteresting - not only from a physical, dosimetric and technical point of viewbut also from a diagnostic and biologically optimized adaptive therapy pointof view.
6.3 Treatment unit and phantomThe treatment unit was a racetrack microtron MM50 (Brahme, 1987) that candeliver photon beams of up to 50MV with a pulse repetition frequency of
20
6.4 PET-CT camera
200 Hz. The intrinsic electron beam of 50 MeV is scanned stigmatically, intwo perpendicular directions, so that the focus of the resulting bremsstrahlungbeam is located in the top part of the bremsstrahlung target. Below the target,a strong purging magnet is placed to remove the leptons produced by photoninteractions in the downstream end of the bremsstrahlung target. To producea narrow photon beam the full range target can be replaced with a thin trans-mission target while the transmitted electron beam is deflected away from thetherapeutic photon beam onto a dedicated electron stopper (Brahme, 1987;Svensson et al., 1998a). For this study we used a thin target of 3 mm beryl-lium and the transmitted high energy electron beam was deflected onto one ofthe fully extended block collimators acting as a electron stopper. The photonbeam was scanned 60 mm in the opposite direction of the deflection of thetransmitted electron beam to minimize the impact of the tail of the transmittedelectron beam (Andreassen et al., 2010).
A cylindrical graphite phantom was used to investigate the activity profiles,The phantom had a diameter of 19.5 cm, length 11.4 cm, density 1.787 g/cm2
with a detector channel for a cylindrical ion chamber placed 0.4 cm from thetop of the phantom.
6.4 PET-CT cameraThe distribution of the positron annihilations was measured in a PET/CT Bio-graph 64 TrueV (CTI/Siemens Knoxville, TN), located a few hundred metersaway from the treatment unit. Emission data were reconstructed by an iter-ative reconstruction algorithm using one iteration and eight subsets. Trans-mission scans for attenuation correction were performed with the CT and re-constructed by filtered back-projection. All data were corrected for randomcoincidences, dead time, scatter and attenuation. A 3 mm Gaussian filter wasapplied for the 3D reconstruction.
6.5 Monte CarloTo compare with the experimental results, Monte Carlo simulations withGeant4 (details in chapter 5.4) were run taking the magnetic field intoaccount. The treatment head geometry used in the simulations is given in Fig1a in paper III. The graphite phantom was simulated in detail. The stationarymagnetic field from the purging magnet, used during the experiments, wassimulated using Opera3D/Tosca and was incorporated into the Monte Carlosimulations.
The deposited dose, photon fluence differential in energy and photon energyfluence differential in energy where scored in 2x2x2 mm3 voxels. The threequantities were only scored in the graphite phantom, not in voxels outside
21
Chapter 6 Positron emission tomography measurements
60 40 20 0 20 40/ mm
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60
Figure 6.1: Monte Carlo simulation of the photon fluence in a graphite phantom, Thephoton fluence was scored in 2x2x2 mm3 voxels and only in graphite material, so thevoxels in the air filled detector hole or outside the phantom show zero fluence. a) agaussian filter with FWHM 5 mm is applied to have a resolution similar to the PETmeasurement b) No filter applied
the phantom or in the detector hole. A total of 5x108 histories were simu-lated resulting in an statistical uncertainty of less then 1.25 % at maximumdose. The voxels where convolved with a gaussian filter with FWHM 5 mmto have similar resolution in the simulation as in the measurement. In Fig 6.1aand b the photon fluence in the graphite phantom is shown with and withoutthe gaussian filter. When the filter is applied the fluence in the voxels outsidethe phantom are not zero and the maximum fluence is moved deeper into thephantom. To get the 11C activity distribution the crossection for (γ ,n) reactionsin graphite (IAEA Photonuclear Data Library) was convolved with the photonfluence differential in energy for each voxel. Geant4 can be used to both pro-duce and track the 11C produced positrons, but for these simulations this moresimple approach was deemed adequate.
6.6 Experimental results and simulationsThe irradiations were done during 300 s, with the PET measurement starting10 min after the end of the irradiation and lasting 30 min. The 11C generatedpositron annihilation distribution in a plane parallel to the incoming 50 MVphoton kernel, using a 3 mm beryllium target, is shown in Fig 6.2a. The pixeldata is shown in gray and isodose lines have been calculated to help interpretthe data. Since the resolution of the image from the PET camera, with ourchosen reconstruction parameters, is > 5 mm and positrons, originating from11C reactions, can annihilate in the air filled hole, the presence of the hole canonly be seen as a disturbance in the positron annihilation distribution. Thisclearly shows the limitation of PET imaging to distinguish material borders
22
6.6 Experimental results and simulations
due to the range of the generated positrons, especially in materials where therange of positrons is high. When materials of higher density are used, this ef-fect is less pronounced as can be seen in a previous publication(Möckel et al.,2007).
23
Chapter 6 Positron emission tomography measurements
120100806040200/ mm
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Figure 6.2: PET measurement and Geant4 simulations of scanned photon kernels in agraphite phantom. The electron energy is 50 MeV, incident on a bremsstrahlung tar-get of 3 mm beryllium. The column to the left shows raw data from the measurementand simulations, while the column to the right show processed values. a) PET mea-surement of the positron annihilation distribution. b) Geant 4 simulation of the 11Cdistribution c) Geant4 simulation of the dose distribution d) PET measurement wherethe 90 to 60 isodose lines have been manually smoothed. e) Geant 4 simulation of the11C distribution where the voxels have been convolved with a gaussian filter with aFWHM of 5 mm. f) Geant4 simulation of the dose distribution where the voxels havebeen convolved with a gaussian filter with a FWHM of 5 mm.
24
6.6 Experimental results and simulations
The 11C distribution simulated with Geant4, in the same plane as the pre-sented PET measurement, is shown in Fig 6.2b and the simulated dose distri-bution is shown in Fig 6.2c. Figure 6.2d is the same as Fig 6.2a except thatthe 90 to 60 isodose lines have been manually smoothed. In Fig 6.2e and f thesimulated 11C distribution and dose distribution have been convolved with agaussian filter with a FWHM 5 mm to have a resolution similar to the PETmeasurement.
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Figure 6.3: Line profiles across the Geant4 simulated 11C distribution (solid line) com-pared to PET measurement (dots) from a 50 MV photon kernel, using a 3 mm beryl-lium bremsstrahlung target, in a graphite phantom. a) 6 mm into the graphite phantom.The position of the detector hole, where the photon fluence was not scored, is markedwith a dashed oval. b) 50 mm into the phantom and c) 100 mm into the phantom.
Line profiles of both the MC simulated 11C distribution and the PET mea-sured positron annihilation distribution are presented in Fig 6.3 for 6, 50 and100 mm depth in the graphite phantom. For the MC simulations the voxelshave been convolved with a gaussian filter with 5 mm FWHM. The measuredand simulated profiles are in good agreement, although the MC simulationsunderestimate the width of the profile at 50 and 100 mm depth. This canpartly be attributed to the fact that the positron range is not considered inthe simulation. Line profiles of the simulated dose distribution are comparedto the PET measurement in Fig 6.4, the dose profile for 100 mm is somewhatlarger but still close to the measured positron annihilation profiles. Since onlyphotons above 18 MV contribute to (γ ,n) reactions the outer regions of thephoton kernel, with a higher fraction of low energy photons, are somewhat
25
Chapter 6 Positron emission tomography measurements
underestimated when comparing the PET measurement to the simulated dosedistribution.
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Figure 6.4: Line profiles across the Geant4 simulated dose deposition (solid line) com-pared to PET measured activity (dots) from a 50 MV photon kernel, using a 3 mmberyllium bremsstrahlung target, in a graphite phantom. a) 3 mm into the graphitephantom. The position of the detector hole, where the dose was not scored, is markedwith a dashed oval. b) 50 mm into the phantom and c) 100 mm into the phantom.
Depth profiles along the central axis of the photon kernel are presented inFig 6.5 for PET measurement, MC 11C distribution and MC dose depositionin the graphite phantom. The PET measurement and MC 11C distribution arein good agreement. The resolution of the PET camera and the range of thepositrons gives positron annihilations outside of the graphite phantom, visiblein the PET measurement. The Monte Carlo simulations did not score dose orphoton fluence outside the graphite phantom but since a gaussian filter with5 mm FWHM was applied the line profiles continue outside the edge of thephantom. The measured positron annihilate density is higher then the sim-ulated 11C activation near the surface of the phantom. The dose depositionshows the characteristic dose build up region and then a quasi linear profile.The depth profile for the dose deposition show as expected a large discrep-ancy in the dose build up region, where the positron annihilation distributionis the strongest and the dose deposition low. Beyond the dose build up regionthe profiles show similar shape as has been previously reported (Möckel et al.,2007). The angle of this later part of the dose profile is comparable to the PETmeasurement profile which could make a PET measurement to deposited doseconversion more feasible in this region.
26
6.6 Experimental results and simulations
0 50 100 150
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Figure 6.5: Line profiles along the central axis of a incoming 50 MV photon kernel,using a 3 mm beryllium bremsstrahlung target, in a graphite phantom. Geant4 simu-lation of the deposited dose (solid line), Geant4 simulation of the 11C activity in thephantom (dashed line) and a PET measurement of the 11C activity (dots).
27
Chapter 7 Conclusions
7. Conclusions
The main conclusions of this thesis are:
• An effective purging magnet design for a SID of 75 cm, suitable forremoving the primary electrons from fast scanned narrow high energy (50- 75 MV) photon beams has been developed. Analytical expressions forthe angular and lateral spread of the transmitted electron beams was usedto estimate the level of electron contamination in a photon field built upby scanned narrow photon beams. For clinically relevant field sizes theelectron contamination will be negligible, assuming that the transmittedelectron beams that are deflected to the electron collector are effectivelyabsorbed.
• A scaning system for high energy photons with a SID of 75 cm wasexamined. The available scan area for pencil beam scanning can be aslarge as 43 x 40 cm2 at isocenter for an incident electron energy of 50MeV and 28 x 40 cm2 at 75 MeV.
• Using the existing components in the racetrack microtron MM50 it waspossible to experimentally produce both homogeneous and complex, withsharp gradients, dose distributions, by scaning a 50 MV narrow photonbeam from a 3 mm beryllium target. The results indicate the potentialto deliver fast IMRT with scanned narrow photon beams alone or incombination with a penumbra trimming MLC. The treatment deliverytime depends on the intensity of the intrinsic electron beam and the size ofthe treated area.
• With scanned photon beams it is possible to sharpen the delivered dosedistributions, to better conform to the target volume and thus avoid organsat risk. This can be done in a fast and efficient way by placing photonkernel spots on the MLC edges.
• Verification of the delivered photon fluence is possible by PET imaging ofthe distribution of positron emitters from photonuclear reactions with thetherapeutic high energy photons. Narrow scanned high energy beams areideal for positron activation since the mean energy will be more uniform
28
over the target volume.
• With scanned narrow photon beams it is possible to fully intensitymodulate the delivered field within a few seconds, making scanned beamsan interesting possibility to improve intensity modulated Arc therapy.
• In a uniform phantom a uniform scanned photon beam will also generate aquasi uniform positron activation.
29
Acknowledgements
I would like to express my deepest gratitude to the following people:
Professor Anders Brahme, without your guidance, creative mind and positiveattitude this thesis would never have been written.
Docent Håkan Danared, your thoughtful comments, discussions and calm sci-entific advice helped more then you know. For making me feel welcome at theManne Siegbahn Laboratory those many years ago.
PhD Roger Svensson, for your guidance during the first years of my PhD andthe frustrating task of correcting my early attempts at scientific writing.
Peder Näfstadius, for sharing your physics and engineering expertise duringthe Sisyphus task of keeping the racetrack MM50 operational. Your optimisticapproach to our many problems gave us the strength to keep trying and, insome cases, succeeding.
PhD student Sara Strååt, we held out Sara. Through the good as well as thebad periods, working with you has been a pleasure.
Rickard Holmberg, without your expertise and unpaid/unselfish work on yourspare time, there wouldn’t be much to write about. I am in your debt.
Patrik Vreede and Marilyn Noz for proofreading the thesis at the last moment.Your positive, productive comments improved it a lot.
The Staff at the Manne Siegbahn Institute, for generous use of both the facilityand computers but most importantly inviting me into your group and makingme feel welcome.
The Staff at the Karolinska Hospital pysics department. For everything frommissordering a whole package of EDR2 film to generous use of both the fa-cility and measuring equipment,
Employees at Scanditronix AB, for all the professional assistance with theracetrack MM50 and patience with us in explaining electronic/mechanic is-sues.
30
Lil Engström, while your administrative skills are par to none, the kindnessyou have shown to me and all who passed through the department is amazing.
Magda and Bartosz Górka, for your scientific advice but most importantlyfriendship during these years has been invaluable.
Aleksandra Kazi, still to this day you bring a smile to my face.
My previous colleagues at MSF: Brigida, Johan, Johanna, Susanne, Malin S,Malin H, Robert, Sharif, Kristin, Takis and Andreas you are all part of whatmade it fun to go to work.
Shangin, Mariann and Ann-Charlotte, for your great administrative work andpositive attitude.
The Senior staff at MSF, your professional skill and assistance is greatly ap-preciated.
Docent Bo Nilsson, your care for everyone around you is most impressive.
The PhD students at MSF, I will miss all of you.
To my family, for your patience and unwavering support.
31
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