BASIC UNDERSTANDING OF THERMODYNAMICS IN BIOLOGICAL
TISSUES
THERMOTHERAPY METHODS
CLINICAL RESULTS AND INDICATIONS
REFERENCES
SUMMARY
A broad range of temperatures is useful in oncology. Thermoablation (heat alone)requires temperatures of >45 to 50°C and is only clinically possible in circumscribedlesions. High-intensity focused ultrasound (HIFU) and nanotherapy are suitable meth-ods. The largest volume heated is the whole body, using whole-body hyperthermia(WBHT); 42°C is the highest temperature permitted. Clinical experience and somepositive studies suggest, however, that higher temperatures (e.g., 43°C) are required atleast in certain specific (e.g., hypoxic) parts of the tumors to increase local control in
conjunction with radiotherapy and/or chemotherapy and to be beneficial for patients.Dedicated multiantenna applicators operating in the radiofrequency range (60–200MHz) must be designed for each indication accounting for the anatomical region.Magnetic resonance monitoring is the first candidate for noninvasive control.
The technical problems have been solved to integrate such applicators into anMR-tomograph. Although commercially available systems (for regional hyperther-mia) are adequate for pelvic and extremity tumors, adaption/optimization is still desiredfor abdominally disseminated disease. Here, the term part-body hyperthermia has beencreated, for which a large number of clinical indications (gastrointestinal tumors) exist.
Heat has a reproducible and predictable effect on cells (1): the effect is quantifiedaccording to the thermal dose concept, taking the sum over a weighted product of tem-perature and exposure time. From the dependence of chemical reactions on temperature(the Arrhenius equation), we derive that above 42 to 43°C (the so-called breaking point)
a temperature increase of 1°C shortens the treatment time by a factor of 2 to achieve anisoeffective thermal dose, e.g., 60 min at 43°C is equivalent to 30 min at 44°C. However,below the breaking point, a further decrease of temperature by 1°C enforces a fourfoldtreatment time for the same effect, i.e., 4 h at 40°C instead of 1 h at 41°C. We conclude
that at temperatures < 43°C effective heat treatments are increasingly difficult.From typical survival curves it follows that 60 min at 43°C can reduce the cell number by a factor of 10 (1). However, to control a macroscopic tumor, we need to eradicate 109
cells or more, resulting theoretically in a required time of 10 h (at 43°C) or approximately10 min at 50°C. We conclude that for thermal ablation (destruction of tumors with heatalone), temperatures >45°C, or better, 50°C, are required (depending on the exposuretime). This is only realistic under clinical conditions for small volumes (<4–5 cm; seeSubheading 2.). Excessive treatment times at lower temperatures (e.g., for hours) or ahigher number of fractions are also problematic because of the phenomenon of thermotolerance.
For locally advanced tumors and large regions, we must employ the sensitizing
effect of heat that has been found in conjunction with radiation and/or certain cytotoxicdrugs. Here, the thermal enhancement ratio (TER) has been introduced. The TERdescribes the factor by which the dose might be reduced to get the same effect. Typicalvalues of TER 1.2 to 1.5 for 41 to 42°C (60 min) are known from laboratory studies.Therefore, lower temperatures in the range of 40 to 41°C might also be effective for acombination of heat with radiation (or chemotherapy). However, a precise estimationbecomes more difficult, because of the high number of variables including physiologi-cal parameters (tumor environment) and the timing between both modalities. Note thata true simultaneous application of radiotherapy and heat has not been realized, for technical reasons.
Other temperature-dependent effects are also known, such as various immunologicaltrigger points, intracellular processes, and physiological changes (especially perfusionchanges). The clinical impact of these mechanisms is not yet clear until now. In addition,the required temperature is not precisely fixed, even though for certain immunologicaleffects in particular (T-cell activation), a so-called danger signal (cell necrosis at tem-peratures 43°C) is postulated. In addition to the well-known enhancement of radiationand cytostatic drugs, other agents has been found that can be modulated by a temperatureincrease. Among them are thermosensitive liposomes (2) and heat control of gene expres-sion (3).
All in all, the actual required temperatures for a certain therapeutic objective are notwell defined. Therefore, a reasonable strategy is the attempt to get a temperature as highas possible. In fact, in a number of clinical studies, higher thermal parameters are statis-tically correlated with clinical end points describing local efficacy (see Subheading 3.).
From in vitro and in vivo studies there is a sound rationale to apply heat in cancer treatment. The problem in patients is technical realization of the heating itself and theappropriate temperatures under clinical conditions.
2. BASIC UNDERSTANDING OF THERMODYNAMICS INBIOLOGICAL TISSUES
Temperature increase in any material requires deposition of energy and is character-
ized by a material constant called heat capacity in kJ/kg (kilojoule per kilogram) required
to increase the temperature of 1 kg of this material by 1°C. For instance, for water, theheat capacity is 4.2 kJ/kg/°C (at 20°C), and this value is a good approximation for humantissue.
We note that 1000 Ws = 1 kJ, which implies that 1 W/kg over 4200 s (roughly 1 hr)
results in a temperature increase of 1°C in a closed system (i.e., no energy is transferredout of the specimen). Because the basal metabolic rate in humans is 1 W/kg, thermalisolation would lead to a core temperature of approx 42°C in 3 to 4 h (considering thatthe metabolic rate increases with temperature). This appears to be the easiest way toensure a temperature of 42°C in tumors, and for metastatic diseases it is the only optionto achieve a therapeutic temperature in all tumor lesions. The speed of heating (warm-upperiod) can be enhanced (down to 90–120 min) by external (infrared) radiation, whichdeposits several hundred watts in the body (4).
However, for whole-body hyperthermia (WBHT), 42°C is the absolute limit of sys-temic temperature (and consequently in the tumor also) and includes various risks for complications. The dose-effect estimations of Subheading 2. suggest that 42°C might be
the lower threshold of temperatures that must be exceeded for beneficial effects intumors. In fact, clinical studies evaluating WBHT in conjunction with chemotherapyhave revealed only slight additional efficacy together with some toxicity and nonneglect-able burden for most patients (5). On the other hand, a systemic temperature increaseabove 39 to 40°C is only tolerated by patients in deep sedation or even general anesthesia,contributing to further patient stress (6).
In consequence, other local or regional heating techniques are required that potentiallyachieve temperatures >42°C in tumors (for higher efficacy), at least in selected parts,without relevant systemic heating (for better tolerance). For that purpose, any methodshould be considered that can be used to apply a certain amount of power in a target
volume, given in W/kg. This power density is termed the specific absorption rate (SAR)and can be deduced from the gradient of temperature rise or fall-off (after switching onor switching off of power) by a simple formula:
SAR [W/kg] = 66.7 × T /time [°C/min] (1)
Therefore, a low temperature increase (T ) of 0.1°C per minute is caused by a power density of approx 7 W/kg, i.e., seven times the basal metabolic rate. For a higher tempera-ture increase of 1°C/min, a SAR 70 W/kg is needed, which is difficult to achieve withavailable systems (see Subheading 3.).
If the material is unperfused, e.g., in a static phantom or in necrotic areas of a tumor,and the exposed volume is high enough, the temperature increase becomes strictly linear.In biological tissues, perfusion w (in mL/100 g/min) and (thermal) conduction (inW/°C) will finally limit the temperature increase and result in a steady-state temperaturedistribution (plateau), mathematically described by the stationary bioheat-transfer equa-tion. This partial differential equation can be numerically solved (7,8), and the solutionwill depend on the perfusion, the SAR (as a source term), the thermal conductivity, andtheir spatial distribution, including the boundary conditions describing the heat transfer to the surroundings.
For sufficiently large volumes with constant perfusion w, SAR, and the temperatureelevation T in the steady state is given by another simple formula:
Therefore, in a tumor with a mean perfusion of 10 mL/100 g/min, we need a SAR of 30 to 40 W/kg to achieve ( 42°C (T = 4–6°C). The validity of this formula depends onthe size of the lesion and is illustrated in Fig. 1. We will discuss two cases:
1. If we lower the perfusion in a circumscribed volume of diameterd , the temperature increasesif a constant SAR everywhere is assumed. For d , the formula is exact. However, for smalld = 1.5 cm, only 20% of this temperature increase is achieved, and for d = 3.5 cm, 60% isachieved. Nevertheless, a necrotic region in a well-perfused tumor (w = 20 mL/100 g/min)exposed to 50 W/kg will rise 0.2°C for d = 5 mm (40.8°C vs 40.6°C), 0.7°C for d = 10 mm(41.3°C vs 40.6°C), 1.4°C ford = 15 mm (42.0°C vs 40.6°C), and 2.4°C ford = 20 mm (43.0°Cvs 40.6°C) in comparison with the well-perfused part of the tumor. We conclude that athermotherapy technique achieving only 40.6°C in well-perfused parts of a tumor (SAR =50 W/kg,w= 20 mL/100 g/min) will achieve 43°C in every necrosis of 2 cm extension or more.Therefore, this external technique (e.g., radiofrequency hyperthermia) might be more effec-tive than WBHT. Even a single necrosis of 5 mm extension achieves a few tenths of °C higher
temperature. Of course, an accumulation of necrotic areas will further enhance this effect,even if they are small and distributed.
2. If we deposit a constant SAR in a circumscribed lesion (with less SAR outside), again arelevant temperature increase is only achieved for diameters of a few centimeters, similar to case 1. The increase of T with d is steeper and shifted about 1 cm to smaller diameters(Fig. 1). For d = 2 cm, 50% of the theoretical T is achieved, and for d = 4 cm, 85%. Inconclusion, isolated SAR elevations of <5 up to 10 mm extension must be excessivelyhigh (hundreds of W/kg) to induce a relevant temperature elevation. If the diameter is<1 mm, we are unable to raise the temperature to some °C for realistic SAR. This meansin particular, that an isolated intracellular heating (cell diameter of about 10 µm) isimpossible for physical reasons.
Fig. 1. Illustration of Eq. 2. The accuracy depends on the diameter of a lesion if either the perfusion
w is lowered or the SAR is increased inside d . The background temperature is 40.6°C (50 W/kg
Achievable SAR and the temperature elevation induced strongly depend on the volume.We can roughly classify thermotherapy methods into thermoablative techniques (often inter-stitial) and locoregional techniques (mostly external). Very large SARs of 2500 W/kg areapplied with high-intensity focused ultrasound (HIFU), but only in a focus of1 to 3 mL (9),
which corresponds to temperature increases of 10°C per minute (see Eq. 1). For larger volumes (of 100 mL or more), we have to scan along the target. Considering the coolingtime between every sonification, the total treatment time can be considerable. HIFU ispresently under consideration for uterine myoma, breast cancer, prostate cancer, and liver malignoma. Similar power densities are deposited in limited volumes by laser-inducedthermotherapy (LITT) or radiofrequency ablation (RFA) (up to 4 cm extension).
Nanotherapy is an interstitial treatment with SAR of some 100 W/kg that is theoreti-cally and practically achieved (10) in a volume of up to 50 to 100 mL at maximum (seeSubheading 3. for further details).
Much less SAR is achieved by use of locoregional thermotherapy. For local applica-tions (11), we achieve a mean SAR of 100 W/kg depending on the tissue depth.
In regional hyperthermia for pelvic tumors we have measured SARs of 30 to 50 W/kg(12,13). This is a strong limitation to achieve >42°C reliably in human tumors, and further technical improvement is required (see Subheading 3.).
3. THERMOTHERAPY METHODS
3.1. Nanofluids and Other Interstitial Methods
Generally, interstitial methods are employed to heat circumscribed tumor lesions; theycan be used as a single treatment in localized tumors or as an enhancing procedure duringsystemic treatment in resistant areas. Thermoablation (e.g., achieving60°C in the target
volume for some minutes) is achieved by laser fibers (LITT), radiofrequency electrodesor antennas (RFA or high-frequency induced thermotherapy [HITT]), and currents(bipolar electrodes). These devices have a given power deposition pattern around their tip and are aiming to coagulate a certain range of the surrounding tissue with therapeuticranges up to 2 cm (typically less). Therefore, the greatest lesion diameter for a single-application use is about 4 to 5 cm.
Recently, an interstitial technique has been introduced based on a fluid withnanoparticles (14). This fluid can be instilled and distributed in a tissue by thin needles;it absorbs power from an external alternating magnetic field in a controlled manner, asdescribed elsewhere (15). Thus the quality of the pattern (and therefore the effectivity)depends on the intervention, which is an advantage (flexibility) and a disadvantage (user
dependency) at the same time.The physical characteristics (the power absorption in particular, and consequently the
heat generation) per nanoparticle in a magnetic alternating field are known for the mag-netic fluid MFL AS 082A (14). For example, a moderate power density of 50 W/kg isachieved if 1 mL of nanofluid is distributed in 10 mL of tumor tissue, and a 100-kHzmagnetic field of 5 kA/m field strength (continuous wave, peak value) is applied.
This calibration allows an estimation of the power density (SAR) for employing anyvolume v[mL] of nanofluid homogeneously distributed in a target volume V [mL] in amagnetic field H [kA/m] of frequency f [100 kHz]:
We can create from this formula a broad range of SARs covering moderate levels of 50 W/kg (typical for locoregional hyperthermia), up to several thousand W/kg withhigher concentrations (e.g., v = V ) and/or higher fields (e.g., H = 10–15 kA/m). Such highlevels are usable for thermoablation.
The distribution of the nanoparticles in the target volume can be quantitatively deter-mined via computed tomography (CT) utilizing the known relationship betweenHounsfield unit (HU)-elevation above the HU of the tissue of interest and the iron (Fe)concentration (according to Gneveckow et al. [14]).
In short, an undiluted instillation of v = 1 mL nanofluid in V = 1 mL target volume (i.e.,75 mg Fe in 1 mL) results in an HU elevation of 600, decreasing to 60 for a dilution of 1:10 (i.e., 7.5 mg Fe in 1 mL tissue). The lowest measurable concentration is 1:50,0.02 mL nanofluid in 1 mL tissue, resulting in an HU elevation of 20. Considering bothcharacteristics of nanofluids together (the potential for calculating the absolute SAR inevery point for given parameters v/V, H, f , and the ability to scan the amount of nanofluid),makes this approach very attractive for controlled interstitial treatment.
For clinical use we have exclusion criteria, in particular, metallic implants <30 cmdistant from the treatment area. For lesions in the head and neck region, amalgam fillingsor gold crowns must be replaced by ceramics. Metallic clips or seeds of some millimeter length and less than 1 mm diameter are not a contraindication, because we have found apower absorption of only a few mW per seed in a magnetic field of H = 10 kA/m.
The first clinical experiences are available with various nonresectable and recurrent(i.e., pretreated) solitary lesions (most of them preirradiated), which were treated in aphase I study (10). Two different methods of nanofluid application were evaluated.
3.1.1. PROSPECTIVE PLANNING
According to the first strategy, prospective planning of nanofluid distribution on athree-dimensional CT data set is performed by selecting a 2.5-mm slice distance in thetreatment position. In a planning system, a segmentation module is included to defineregions of different extension and shape with a given constant SAR (in W/kg) in andaround the target volume. This target volume is contoured by the oncologist beforehand.Every SAR region is targeted with a specific amount of nanofluid v in mL, which can becalculated from Eq. 3. For this preplan, an assumption about the achieved magnetic field
H [kA/m] is made. Then, for a given tumor perfusion [mL/100 g/min], the temperaturedistribution is calculated by solving the bioheat-transfer equation on a tetrahedral gridgenerated by utilizing the finite-element method.
For realization of this plan, the number of nanofluid depots is limited for practical
reasons. If the depots are injected under CT control, the direction and positions of thepuncture tracks are also preplanned to guide the radiologist. The total amount of nanofluid,the amount per depot, and their positions are optimized to achieve a given minimumtemperature (e.g., 42°C) enclosing the target volume and to give a certain homogeneity,i.e., to limit the maximum temperature. The predicted temperature value critically dependson the assumed perfusion, which is only approximately known.
The nanofluid depots are shaped like either spheres (i.e., injection in a defined point)or cylinders (i.e., gradual injection by retracting the canule along the puncture track).The tracks and positions of the depots are selected according to anatomical consider-ations (sparing of critical structures), if possible in standardized settings (e.g., trans-
verse planes). Then the number and distribution of the depots are manually modified
until the 42°C isotemperature line (or any other defined temperature) is covering thetarget volume.
The preplanned tracks and coordinates are targeted by the interventional radiologistunder CT fluoroscopy. The standardized technique is described by Ricke et al. (16). It
is advisable to leave a catheter in (<1 mm diameter) for a direct temperature measure-ment in the tumor after implantation. Permanent seed implantation with 125-I seeds isbased on a standardized procedure to place the seeds with millimeter precision under transrectal ultrasound and X-fluoroscopy (17). For the special indication of recurrentprostate cancer (after definitive external radiotherapy), the nanofluid is distributedalong the needle tracks on the same template as that used for seed implantation in thesame session.
3.1.2. INTRAOPERATIVE IMPLANTATION
The second strategy is an intraoperative implantation of the nanofluid intraoperativelyunder visual control. Here a prospective plan is not possible, because after tumor
debulking, an individual (nonresectable) tumor rest or risky area (R1/R2 situation) mustbe infiltrated. The quality and coverage of the nanofluid distribution depends on theability of the surgeon and his/her knowledge of the location and extension of the tumor-involved area. The surgeon is also advised to leave a catheter in for temperature measure-ment in situ. Cervical cancer recurrences at the pelvic wall (after primary treatment) aretypical clinical examples. The nanofluid concentration in the target area should be as highas possible. In these patients retrospective planning based on a postoperative CT data setis performed.
We treated 20 patients with recurrent lesions by the three methods described, i.e.,using either CT guidance, transrectal ultrasound (TRUS)-guidance for the prostate, or during operation under visual control. Different amounts of nanofluid (1.5–10 mL)were applied. The tolerated magnetic field strength ranged from 3 to 8 kA/m, with thetypical value in the pelvis being 4.5 kA/m. Discomfort in skin folds is a limiting factor.In the upper thorax and neck the tolerance is higher (6–8 kA/m). Postimplantationanalysis via CT revealed a high mean SAR of 190 W/kg after CT-guided implantationand 50 to 70 W/kg for the prostate and R1/R2 regions. These values are higher thanthose achieved with external locoregional systems. Therefore, the mean maximumtemperatures are estimated in the lesions as 44.5°C after CT-guided implantation and42 to 42.5°C for the other implantation methods, which is again superior to regionalhyperthermia (13). Dependency on the skill of the therapist is a shortcoming. There-fore, standardization and refinement of application modes are needed for further
improvement.A particularly effective method to increase the SAR further (and the temperature) is
the use of higher magnetic fields, because SAR is proportional to the square of H . Adjust-ing H to 10 kA/m, i.e., doubling of H , would increase the SAR up to 200 W/kg. Technicalsolutions are under way.
3.2. Locoregional Hyperthermia
Antennas radiating electromagnetic waves are suitable to generate SAR distributionsof a given shape and extension, if they are positioned outside the patient and near thetarget volume. Suitable frequencies lie in the radiofrequency (RF) range between 70 and
200 MHz. Theoretical considerations have shown that the best frequency depends on the
special problem (location of the target) (18). Because the penetration of electromagneticwaves with lower frequencies increases, the frequencies (e.g., 70 MHz) are better suitedfor deeply located and largely extended volumes.
The optimal arrangement of RF antenna arrays must be formulated as a general prob-lem, which is shown in Fig. 2. Antennas are radiating structures fed by an amplifier,which have a certain radiation profile at a given frequency. The matching to this fre-quency must be ensured by a transforming network between the amplifier and antennafeed point. The near-field range of such antennas is typically not suitable to generate aclinically useful power deposition pattern in patients.
Recently we have investigated attributes of antennas to design better dedicated appli-
cations. Among these attributes are form, thickness, coating, and special environment
Fig. 2. Schematic arrangement of small-dipole antennas around the body either to heat a pelvic
tumor or to expose the whole abdomen (called part-body hyperthermia). This applicator is called
(dielectric structures) of the antennas. We have found possibilities to reduce the near-field range. Outside the near-field range the field pattern is easier to control. A water load(presently used as a water bolus) is required to direct the radiation to the patient. However,this bolus probably does not need direct pressure/contact to the patients’ surface in an
advanced hyperthermia system. This would contribute to a much easier and more com-fortable application of RF hyperthermia.Another feature of antenna arrays are the coupling conditions, which depend on both
the network and the particular position of the antennas (field coupling). It has been foundthat flexibility (control) and efficiency of antenna arrays are inversely correlated. If theantennas of an array approach each other, they are automatically coupled for physicalreasons. Then, the technical possibilities for decreasing the coupling are limited. Theantenna arrays generate so-called resonance modes. In or near such a resonance mode thepower (efficiency), which is deposited in the patient, is quite high. Adjustments far awayfrom those resonance modes are difficult to realize and have certain disadvantages (seenext paragraph).
Planning systems are now available for calculating and optimizing the SAR distribu-tion (and consequently the temperature distribution) generated by these antenna arrays(12,20). The solution in the patient is quite reliable, if the antenna models are correct.However, the antenna models and the transforming networks are probably too simple.Recently, improved algorithms have been developed that take into account special anten-nas and networks. These algorithms are the basis for the improved design of applications(8). The most important parameters for controlling SAR patterns are the phases in the feedpoints of the antennas. Theoretically, using simplified antenna models we can calculatecomplicated SAR patterns with sometimes atypical and erratic phase values. In doing thiswe overestimate the flexibility of SAR control. However, if we want to adjust these
phases under practical conditions, we measure increasing phase errors from ±10° (near the resonance modes) up to ±20 to 30° (far away from the resonance modes). Further-more, high reactive power is then deposited for these adjustments for off-resonance in thewater bolus or network. Of course, this high reactive power results in ohmic losses inreality, which further decreases the efficiency of the power deposition. As a consequence,we have found technical limitations (amplifier power) to achieve the desired tempera-tures (e.g., >42°C). We conclude that these “optimized” patterns are not as useful for practical applications, and we have developed other strategies to improve the temperaturedistribution in tumors.
3.2.1. SWITCHING BETWEEN STANDARD PATTERNS CREATING HOT SPOTS AT DIFFERENT
LOCATIONS
In the pelvis, total power is mostly limited by a single hot spot, which might beeither at the symphysis or at the dorsal sacral bone (Fig. 3A). Switching between bothSAR patterns (by changing the phases) at time intervals of 1 to 8 min results intemperature-time curves (Fig. 3B) that still successfully heat the tumor up to 43°C(solid gray line), but distribute the thermal dose between both hot spot regions (fluc-tuating lines). The shorter the switching time, the smoother the curves, but there aretechnical/practical limitations. Increasing the number of patterns with different hotspots further improves the heating possibilities. Note that for this strategy we usestandard patterns (with only slight phase delays of maximally 30°), which minimizes
With a special applicator, a reasonably homogeneous SAR in the abdomen (with atypical volume of 15–20 L), including the liver, should be possible. The Sigma-Eyeapplicator (Fig. 2) positioned with the central plane at the umbilicus is a possible solution,if phase delay is adjusted at the outer rings.
However, better applicator concepts exist. For a SAR distribution to be as homoge-neous as possible, lower frequencies are better (e.g., 70 MHz instead of 100 MHz for theSigma-Eye applicator). Then, a large aperture angle, a small near-field range, and a goodmatch (with a minimum of transforming network components) are desirable. It is evidentthat small antennas (as in Fig. 2) are not optimal for this special task.
The objective of such an applicator is the generation of a constant SAR in the wholeabdomen according to the hatched region shown in Fig. 4. A SAR of 35 W/kg in 15 L(525 W) results in a temperature elevation of 1.4×35 W/kg/10 mL/100 g/min, i.e., approx5°C, corresponding to >42°C. This conforms to common heat treatments with regionalhyperthermia, but a well-adapted applicator would be a precondition. We note that 15 Lfits in a sphere of 30-cm diameter (which corresponds to the volume, in which a constant
SAR should be generated by the array to be designed). This concept is consistent with the
Fig. 4. Concept of part-body hyperthermia with a realistic (fairly constant) SAR of 35 W/kg in the
abdomen. With a mean perfusion of <10 mL/100 g/min, we achieve 42 to 43°C in the peritoneum,and we consider the preheating effect in the liver as well as effective temperatures in liver
observation that perfusion through the abdomen (and entering the liver via the portalvein) is not increased during part-body hyperthermia.
3.2.3. LOCAL APPLICATORS
In principle, every additional clinical application needs special arrangements of anten-nas or applicators. For malignant lymph nodes in the neck, we used single applicators(11,21). Dedicated equipment (dual-applicator operation) has been developed by theRotterdam group for recurrent breast cancer of the chest wall (22). The group at Duke hasdeveloped applicators for primary breast carcinoma (23).
Locoregional hyperthermia is feasible, but improved understanding and manufactureof applicators will be indispensable, if we want to optimize the heating conditions for different indications. Dedicated applicators are needed for the abdomen (part-body hy-perthermia), the neck region (including the primary tumor), the chest wall (includingadvanced and extended recurrences), and other areas.
For the pelvic region and lower extremities, the available applicators are already
appropriate (see the Sigma-Eye applicator in Fig. 2), but nevertheless they must becontrolled with sufficient certainty.
3.3. Thermotherapy and Monitoring
We clearly need a measuring parameter to classify our thermotherapy as effective(or not) and to detect potentially risky constellations, e.g., temperatures that are toohigh in normal tissues. When hyperthermia began, an intratumoral temperature mea-surement was demanded (24). This measurement has been justified by various corre-lations between the local efficacy of the treatment (clinically described by responseaccording to the World Health Organization [WHO] or a similar parameter) and ther-
mal parameters, e.g., the thermal dose achieved in 90% of the measurement points(25,26). To increase the acceptance, tolerance, and practicality of heat treatments(27,28), endoluminal reference points were introduced as substitutes in pelvic tumors,as indicated in Fig. 3 (29). These points can be arrived at with catheters by usingminimally invasive procedures.
Endoluminal SAR measurements (see Eq. 1) do not differ from intratumoral SARmeasurements, if interpreted properly (12,13). Therefore, endoluminal data are useful for a check of the SAR distribution. Not unexpectedly, endoluminal temperature measure-ments display characteristic differences in comparison with intratumoral measurements,as described in Wust and colleagues in 1998 (12). Nevertheless, the same correlationswere found as in intratumoral measurements, as documented for rectal carcinomas (30),
prostate carcinomas (31), and cervical carcinomas (32). For standard situations in thepelvis, including bladder cancer, an endoluminal temperature measurement for controlis equivalent.
For regional hyperthermia in the abdomen, especially part-body hyperthermia (Figs.3 and 4), reference points are not available under patient-friendly conditions. This is thefirst reason to integrate antenna arrays (applicators) into a magnetic-resonance (MR)tomograph in order to establish a noninvasive monitoring capability. Second, by thistechnique a full three-dimensional data set is acquired comprising much more informa-tion than any direct thermometry. Third, online monitoring of thermotherapy can accountfor online corrections or improvements in heating and provide the basis for optimization
Basically, the hybridization of both systems is electrotechnically realized by separat-ing the signal paths, filtering out the heating frequency (e.g., 100 MHz) in the receivingpath of the MR, and filtering out the resonance frequency (64 MHz at 1.5 Tesla) in thepower path to the antennas (33). The mechanical integration is shown in Fig. 5; the patientand the applicator are moved together from the back into the gantry (34). At the front, thediagnostic table is held ready, and a change between heating under MR monitoring or MR
examination for diagnostic purposes is accomplished in a few minutes.Magnetic resonance imaging (MRI) has been used for monitoring of thermoablativeinterventions for more than a decade. While researchers were evaluating MRI approachesfor this indication, it became clear that the MR signals are influenced by a mixture of tissue-effects (e.g., coagulation) in addition to temperature changes (35). As a result,more recent studies attempted to separate the temperature out as an observable factor inlaser-induced thermotherapy (36,37) and focused ultrasound (38,39) and to elucidate thereliability and limitations of MR-guided thermography. Basically, three methods are avail-able for MR thermography and are discussed in the literature. They are based on thetemperature dependency of the relaxation time T1 (40), the diffusion (apparent diffusioncoefficient [ADC-value]) (41,42), or the proton resonance frequency shift (PRFS) (43,44).
As T1-weighted sequences are particularly dependent on the type and status of the tissue(45,46), they are mainly recommended in low-field systems because of their higher contrast. On the other hand, diffusion methods still suffer from technical and methodicallimitations in clinical practice (47). Therefore, for magnetic fields 1 Tesla, PRFS-basedmethods are preferred by most investigators today (48–50).
However, temperature accuracies of MR-based thermography used for ablation tech-niques are still in the range ±3 to 4.5°C even under favorable in vitro conditions with liver specimens (36,37,51). This is because in vivo MR thermography is hampered not only bymotion but also by various (treatment-induced) tissue alterations, which are reversible or irreversible and depend on the temperature level, heating time, and tissue type. Therefore,
MR thermography of larger target volumes, e.g., during regional hyperthermia, appears
Fig. 5. In the hybrid system at the Berlin group, a multiantenna applicator (here the Sigma-Eye
applicator) is moved together with the patient into the gantry of a tunnel MR tomograph (here the
Magnetom Symphony, Siemens). On the right, we see more details of the components, whichpartially interact.
to be even more challenging, especially as the expected temperature differences in thecourse of a treatment (37–43°C) are near the uncertain values referenced for high-temperature applications.
In a small series, Carter et al. (52) employed a self-developed cylindrical phased-array
hyperthermia applicator (25-cm diameter) to treat sarcomas of the lower leg in a 1.5-TeslaMR scanner (Signa General Electric, Milwaukee, WI). The investigators acquired phaseimages during regional hyperthermia and demonstrated a satisfactory correlation withmeasured temperatures (standard error ±1°C) at moderate power levels of about 150 W.Peller et al. (53) described a hybrid system consisting of a Sigma-Eye applicator (BSDMedical, Salt Lake City, UT) integrated into a low-field system (0.2 Tesla, MagnetomOpen Viva, Siemens, Erlangen, Germany). With T1 imaging, they claimed to control theheat of treatments (hot spot detection) and presented selected cases (lower extremity)with good correlation of T1 and directly measured temperatures (54).
A hybrid system installed by Wust et al. 2004 (34) has been validated with the PRSFmethod in an anthropomorphic phantom (33). Special acquisition and postprocessing
methods have been developed. We note that for rectal recurrences after abdominoperinealresection, an endoluminal reference point is difficult to specify or does not exist, andtherefore the introduction of noninvasive MR thermography represents real progresswith respect to tolerance and safety. For rectal recurrences and soft tissue sarcomas, ananalogue correlation between response and mean MR temperature in the tumor has beenfound (55).
In the abdomen, the present standard method for noninvasive thermometry (PRFS)cannot be applied because of motion (breath-dependent) and strong susceptibility gradi-ents leading to severe artifacts (especially because of air conglomerations in the intes-tine). Navigated acquisition techniques might compensate for motion in the liver and
allow the PRFS method. The most stable monitoring technique is gained by perfusion(contrast media dynamics), which is not dependent on susceptibility discontinuities andis less vulnerable to motion. With the known SAR distribution (at least from a computer plan), the temperature distribution can be estimated.
4. CLINICAL RESULTS AND INDICATIONS
A series of randomized studies have demonstrated the enhancing effect of heat inaddition to radiotherapy, as summarized by Wust et al. (56) or Falk and Issels (57), andIssels et al. (58). The most convincing study employed the standard system for regionalhyperthermia (RHT) (BSD 2000) for pelvic tumors and found a survival benefit in the
study arm (radiotherapy plus hyperthermia) for cervical cancer IIB (59). Recently,radiochemotherapy has become the new standard for this patient group, and the role of regional hyperthermia will be defined in a running study comparing radiochemotherapywith hyperthermic radiochemotherapy (60). Other prospective randomized studies evalu-ating chemotherapy (±RHT) in soft tissue sarcomas (58) or radiochemotherapy (±RHT)in rectal cancer (61) are not yet completed.
Further randomized studies have succeeded in demonstrating the enhancing effect for superficial tumors of different histologies (62–64), with increased local control and/or response in the hyperthermia arm. Recently, considerably improved local control (recur-rence-free survival) was found in the hyperthermia (plus intravesical mitomycin C) armfor patients with bladder cancer by Colombo et al. (65).
There are various problems with these studies related to the clinical use of hyperthermia:
1. There are not many clinical data, and they are not sufficient to give unequivocal answers.2. For oncological reasons, a survival benefit was found only in the study of Van der Zee
(59). The improved local control found in other studies might be informative for the users,
but is not accepted as a clinical end point constituting an imperative indication.3. Further clinical studies would be strongly desirable. However, these studies are demand-
ing and difficult to perform, because only a few centers are involved, which prolongs therecruiting time.
4. Technological developments are still in progress (see Subheadings 2 and 3), which mightbe crucial for achievement of effective temperatures.
Existing commercial systems are suitable for heating pelvic tumors (prostate, cervix,rectum, bladder) and tumors of the lower extremity (soft tissue sarcoma).
Abdominal and liver metastases are frequently seen. Primary tumors are gastrointes-tinal (colorectal, pancreas, stomach) and ovarian cancers. Part-body hyperthermia might
be helpful to enhance the chemotherapy effect (± innovative substances in addition). Wefound that WBHT is not ideal for that indication, because of high burden and temperaturelimits (42°C). Better applications have been described (part-body hyperthermia), whichheat only the abdomen to (ideally) >42°C, and do not greatly affect the systemic tempera-ture. The adaption of present technology (Fig. 2) for this indication is under way, andclinical studies are in preparation. The objective of these studies must be survival improve-ment in patients receiving systemic standard chemotherapy.
Other oncological indications exist, when standard radiochemotherapy will not leadto a fully satisfactory outcome, e.g., inoperable head and neck carcinoma (with a presentlong-term survival of at best 30–40%, see, e.g., ref. 66 ) or malignant brain tumors (withpresent curative rates near zero) (67).
Dedicated applicators and treatment strategies are required for these indications, butthey are not yet available. Planning tools and technical know-how are evolving to makethe design of such applicators possible (e.g., antenna arrays) for every specific location— and to integrate these applicators into an MR gantry for noninvasive monitoring. Inconclusion, suitable and effective applicators must be designed for every specific onco-logical indication (when they are needed by clinical reasons), and the general rules for constructing them must be understood.
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