Nancy Price Mendenhall and Zuofeng Li
The burgeoning interest in proton therapy is related to its potential to improve the therapeutic ratio for many malignancies treated with radiation. This chapter will touch on the rationale, evolving technology, potential applications, efficacy, toxicity, comparative effectiveness, and economic challenges associated with proton therapy.
RATIONALE
Therapeutic Ratio
With any medical intervention there is an optimum balance between the potential to benefit and the potential to harm the patient, called the therapeutic ratio. In radiation oncology, it is the balance between radiation effects on the tumor and radiation effects on normal, nontargeted tissues. This ratio informs every physician and patient decision, but different priorities may mandate different decisions in different clinical situations. For example, in a patient with an advanced paranasal sinus tumor adjacent to the optic nerves and chiasm, one physician might recommend a treatment that minimizes the risk of optic nerve damage at the price of a lower probability of tumor control, while another physician might recommend a treatment with a greater risk of optic nerve damage to provide a greater probability of tumor control. In another example, the potential risks of low-dose radiation exposure to a large volume of the brain tissue may be of greater concern to the parents of a young child than to an elderly patient with a short life expectancy. In evaluating new radiation modalities, the first question is whether there exists a potential to improve the therapeutic ratio.
The Impact of Dose Distribution on the Therapeutic Ratio
A number of factors influence radiation effects, and thus the therapeutic ratio, including total dose, dose intensity (the overall time or number of fractions in which a dose is given), relative biologic effectiveness (RBE), and modifying factors such as chemotherapy, patient age, and comorbidities. But none of these factors is as important as dose distribution—the relative dose to the target compared with the dose to nontargeted normal tissues—because radiation therapy is a nonspecific therapy that affects both normal and cancerous cells. A constant principle in radiation oncology is that the higher or more intense the radiation dose, the greater the probability of tumor control. The risk of treatment-related toxicity is similarly related to the dose and dose intensity to normal tissue, as well as to the volume of the normal tissue exposed to various radiation dose levels. This principle has been demonstrated in many malignancies (i.e., prostate, breast, head and neck) but may not always be apparent when tumoricidal doses cannot be delivered to the target because of excessive risks to normal tissues or when adequate tools for measuring toxicity are not available.
The primary barrier to maximizing local tumor control through dose escalation or intensification is the risk of damaging normal tissues either by delivering too high a dose or exposing too much of the normal tissue to radiation. Whereas tumor control is a dichotomized end point, toxicity occurs as a spectrum of end points over a long period of time in a variety of tissues and can be difficult to measure. When toxicities are easily defined, are functionally significant, and occur early, the risk of the radiation complication is often prioritized over the probability of cure, leading to a low therapeutic ratio because of low disease control rates. Conversely, when toxicities are ill-defined, are difficult to measure, or occur late, disease control will be prioritized. In most clinical settings, there is an opportunity for improvement of the therapeutic ratio by increasing disease control or by reducing toxicity. The most direct means of improving the therapeutic ratio is by reducing the radiation dose to nontargeted tissues, which both reduces toxicity and facilitates dose escalation for increased tumor control: herein lies the rationale for proton therapy.
FIGURE 19.1. The shape of the depth-dose curves for electrons (tan), photons/x-rays (orange), a pristine proton Bragg peak (purple), and a spread-out Bragg peak (blue), composed of multiple pristine Bragg peaks, differs significantly. Compared to photons or electrons, the entrance dose with protons is constant and reduced relative to the target dose and there is no exit dose. The dose fall-off at the end of the proton range is much sharper than for electrons. In summary, with photons, most of the radiation energy is deposited outside of the target, whereas with protons, most of the radiation energy is deposited inside the target.
The Problem with X-Rays and the Promise and Challenge of Protons
Most x-rays from an external beam pass through the patient without effect, but some are absorbed as they randomly interact with subatomic particles along their path, initiating a cascade of biochemical reactions that result in cell injury or cell death. X-rays thus leave a track of tissue damage from the skin-surface entrance to the skin-surface exit, much like the track of a bullet. With accumulating x-ray absorptions, the dose deposited along the beam path is attenuated as the number of x-rays available for interactions decreases. This pattern of radiation dose deposition with x-ray-based external-beam radiotherapy (EBRT; Fig. 19.1) is particularly problematic because the dose to tissues along the beam entrance path is always higher than the dose to the target. In addition, along the exit path of the beam, more normal tissues are exposed to doses of radiation only slightly less than the dose to the target. Thus, most of the radiation dose with x-ray-based therapy is deposited in the patient outside of the target.
Many elegant strategies have been developed for offsetting this basic problem (i.e., multiple fields, four-field box, stereotactic radiosurgery and stereotactic radiation therapy [SRT], three-dimensional conformal radiation therapy [3DCRT], intensity-modulated radiation therapy [IMRT], tomotherapy, arc therapy, robotic radiosurgery, etc.), but even with these technologies, radiation to nontargeted tissues is primarily redistributed, rather than reduced, and more radiation dose is deposited outside, versus inside, the target. Because the total dose to a given tissue is one of the key determinants of toxicity, certain toxicities are reduced when using these sophisticated techniques. Nevertheless, it is likely that other toxicities, influenced by the volume of tissue exposed to radiation, will increase by these integral dose redistribution methods as greater volumes of normal tissue are exposed to radiation. Some of the effects related to low-dose exposure to larger volumes of tissue may require more time for clinical manifestation (second malignancies), and some injuries may require more subtle measurement tools for detection (neurocognitive testing), leading to an early underestimation of toxicity with these sophisticated methods.
The pattern of dose deposition with protons differs significantly from that of x-rays. A proton is simply a hydrogen atom that has lost its electron. It has a mass of approximately 1,800 times that of an electron and carries a positive charge. As protons traverse matter, they lose energy primarily through interactions with atomic electrons. Due to the significantly larger mass of protons relative to electrons, protons lose only a small portion of their energy in each interaction (in contrast to x-rays) and experience only small directional changes.
As protons traverse matter, the rate of energy loss in electronic interactions is described by the linear stopping power, S(E), defined as dE/dx, where dE is the mean energy lost by a proton in electronic collisions over a distance of dx in media. A more commonly used function is the mass stopping power, s = S(E)/ρ, which denotes the energy loss dE when protons travel through a distance of dx in a media of density ρ.1
Linear energy transfer, or LET, is a closely related concept to stopping power. Whereas stopping power indicates energy lost by the protons, LET denotes the energy transferred to a medium as the protons traverse it. LET therefore represents the density of energy deposition in media and is directly related to the local RBE of the radiation, similar to its significance in other radiation modalities. At higher energies (e.g., along the entrance path), protons have a small stopping power and low LET values; their LET values increase sharply by up to two orders of magnitude as the proton’s energy decreases just before coming to a complete stop. This phenomenon creates the Bragg peak of radiation dose deposition characteristic of a proton beam, with a small, nearly constant, dose along the entrance path followed by a sharp peak in dose deposition immediately before the protons stop in media. In addition, the LET values rise sharply, corresponding to a sharp increase in RBE values near the end of a proton beam range. Much work has been done to determine in vitro and in vivo proton beam RBE values.2 For clinical applications of protons, it is currently recommended that an RBE multiplicative value of 1.10 should be applied to the physical dose.1 Currently, this RBE factor is applied uniformly to the physical dose distribution, without explicitly accounting for the increased RBE values near the end of a proton beam range. There is a rapid rise in RBE during the last several mm3 of the proton range, so that the actual RBE-corrected dose at the very end of the range may exceed the physical dose by up to 25%, thus producing an RBE value up to 1.3 at the very end of the proton range. RBE effects are much more significant in heavier ions like carbon due to their significant variations along beam paths; nascent carbon ion treatment-planning systems are attempting to account for these variations in RBE along the carbon beam path.4
In addition to energy loss through electronic collisions, high-energy protons also experience nuclear interactions in media. Such interactions, from the perspective of clinical proton therapy, remove protons from the initial proton fluence and produce secondary protons and heavier particles, such as deuterons, tritons, 3He, and α; these particles contribute only a small percentage of the dose along the beam path.5 Although negligible, most of the dose from the neutrons produced in these nuclear interactions is deposited downstream from the stopping point of the proton beam, beyond the Bragg peak. The neutrons produced by protons within the patient, as well as those produced in the beam delivery mechanical components, have been under intense investigation for their potential secondary cancer-inducing effects. Hall6 raised concern that such scattered neutrons may increase the incidence of secondary tumors in patients treated with historical proton therapy systems, similar to the transition from conventional 3DCRT to IMRT.7 Much research has been performed estimating secondary cancer risks from contemporary proton therapy systems using either scattering-beam techniques or scanning-beam techniques. A review of literature on this subject finds consistent advantages of proton therapy over photon therapy techniques in the form of reduced secondary cancer risks in the treatment of medulloblastoma, prostate cancer, and liver cancer.8
Accordingly, the rationale and promise of protons, compared with x-rays, lie in the significant reduction in dose to nontargeted tissues along the entrance path and the absence of exit dose. This reduction in dose (physical and RBE dose) to normal tissues should translate into a lower risk of complications for a given target dose and probability of tumor control (i.e., an improved therapeutic ratio compared with x-rays). A lower risk of complications could be leveraged for dose escalation or intensification to yield higher disease control in certain settings or for hypofractionation to reduce costs or increase tumor control in other settings.
TECHNOLOGY
Beam Production and Transport
Protons are produced either from hydrogen gas obtained from electrolysis of deionized water or from commercially available high-purity hydrogen gas. Application of a high-voltage electric current to the hydrogen gas strips the electrons off the hydrogen atoms, leaving positively charged protons. The protons are then accelerated to energy applicable for clinical proton therapy with either a cyclotron or a synchrotron (Fig. 19.2). Energies on the order of 250 MeV are required for penetrating approximately 32 cm into tissue. Cyclotrons produce a continuous beam of nearly monoenergetic and unidirectional protons. The beam must then be degraded to meet specific requirements for each treatment field in each patient. This process takes place shortly after the beam exits the cyclotron in a device composed of a material of variable thickness with a low atomic number known as the “energy degrader.” The proton beam exiting the energy degrader will have an energy spread centered around the desired final beam energy and direction variations that reduce the quality of the final treatment beam. An energy selection system (ESS), consisting of energy slits, bending magnets, and focusing magnets, is then used to eliminate protons with excessive energy or deviations in angular direction. Cyclotrons can produce a large proton beam current of up to 300 nA and thus deliver proton therapy at a high dose rate using the traditional double-scattering technique.
Synchrotrons, however, produce proton beams of selectable energy, thereby eliminating the need for the energy degrader and energy selection devices. A proton pulse exiting a preaccelerator, with energy typically up to 7 MeV, is injected into the ring-shaped accelerator. Each complete circuit of the proton pulse through the accelerator ring structure incrementally increases the proton pulse energy. When the desired beam energy is reached, the proton pulse is extracted from the accelerator. The time segment between pulses depends on the final energy required of the proton beam, ranging from subseconds to several seconds for the highest beam energy. The pulsed nature of the beam introduces additional complexity in certain treatment delivery scenarios, such as gated treatment of mobile targets and intensity-modulated proton therapy (IMPT).9 Beam currents from synchrotrons are typically much lower than with cyclotrons, thus limiting the maximum dose rates that can be used for patient treatment, especially for larger field sizes. The maximum dose rate available from a commercially available synchrotron-based proton therapy system for a 25 × 25-cm2 field has been specified at 0.8 Gy per minute.10 The elimination of the energy degrader and selection system removes a major source of radiation production in the accelerator vault. A synchrotron vault is therefore accessible immediately after the beam is stopped for maintenance, whereas for a cyclotron vault, approximately 30 minutes is necessary to allow the activated parts of the accelerator and ESS to “cool down” before maintenance can be performed, or longer if access to internal parts of the cyclotron is required. The shielding requirements for cyclotrons are higher than for synchrotrons as well due to the high radiation produced by the ESS. However, the overall footprint of a cyclotron vault, including the additional shielding required, is actually similar to or smaller than that of a synchrotron vault because of the smaller physical dimensions of cyclotrons.
The proton beam, whether exiting the ESS for a cyclotron-based system or exiting the accelerator for a synchrotron-based system, is transported to the treatment room(s) via the beam transport system. Maintenance of beam focusing, centering, spot size, and divergence throughout the beam transport system is critical to maintaining a high-quality proton beam for treatment delivery. Paganetti et al.11 performed Monte Carlo calculations of the dosimetric effect of proton beam energy spread, spot size, and angular energy spread on the shape of the Bragg peak. They found that small deviations in these parameters from their nominal values will result in a widening of the Bragg peak and increased entrance doses. Beam transport systems in clinical proton facilities therefore include bending and focusing magnets and beam profile monitors so that the proton beam quality may be monitored and adjusted (“tuned”) as it is transported through the beam transport system.
Recent progress in accelerating and control technologies has led to new designs of proton therapy systems in the form of single-treatment-room systems with small cyclotrons that may be mounted on a gantry, as well as dielectric wall accelerators (DWAs) and laser-accelerated accelerators.12 Single-room proton therapy systems are currently available from a number of commercial vendors, each with its own strengths and weaknesses, although none was in clinical operation at the time this chapter was written. DWA and laser-accelerated systems are currently under development. One of the primary drivers for development of these single-room proton therapy solutions is facility cost reduction.
FIGURE 19.2. Proton therapy system with cyclotron, energy selection system, beam line, gantry, and nozzle. Scattering is illustrated in the nozzle on the left and scanning in the nozzle on the right.
Beam Delivery
The proton beam exiting the transport system (Fig. 19.2) is a pencil-shaped beam with minimal energy and direction spread. The beam has a small spot size in its lateral direction and a narrow Bragg peak dose in its depth direction. Two basic techniques have been developed to convert this narrow pencil beam into a dose distribution suitable for treatment of a 3D target, broadly categorized into the “scattering technique” and the “scanning technique.” These techniques are implemented in the proton treatment “nozzle,” or treatment head.
The scattering technique aims to produce a dose distribution with a flat lateral profile, similar to what is achieved in linear accelerator–produced 3D conformal x-ray beams, but a depth dose with a low entrance dose and a plateau region, followed by the sharp fall-off of doses to zero. The depth-dose curve, with a plateau of adequate width to cover the full thickness of the target, is produced by summing a number of Bragg peaks, each with progressively reduced maximum energy and range in the patient. The weights of the individual Bragg peaks are optimized to achieve a flat top region of the depth-dose curve. A constant energy proton beam may be used for such treatments, with range-reducing materials inserted into the beam path to create subsequent Bragg peaks of reduced ranges. Range modulation wheels consisting of variable thicknesses of “acrylic glass” (polymerized methyl methacrylate) or graphite steps are traditionally used for this purpose.11 The proton beam travels through the variable-thickness steps, with each step creating a Bragg peak of a precalibrated range. The widths and thicknesses of the modulation wheels are calibrated to achieve a flat depth dose, or “spread-out Bragg peak” (SOBP). The width of the SOBP is controlled by turning the beam off when a prescribed width is reached. Alternatively, a “ridge filter” may be used to create an SOBP,13 with the advantage of eliminating sensitivity to organ motion in the formation of the SOBP but the added complexity that each ridge filter is designed to achieve a single SOBP width.
The scattering technique produces a large, flat lateral dose profile through physical scattering filters in the proton beam path.11 Either a single scattering filter may be used, creating a wide beam with a Gaussian dose profile (single scattering beam), or a second scattering filter can be added into the beam that flattens the beam to create a flat lateral dose profile (double scattering beam). Apertures fabricated out of brass or other metal are used to confine the treatment field to the target. A tissue-equivalent range compensator (or bolus) is designed to adjust the beam range throughout the field to conform to the distal profile of the target.14
In the scanning technique, as the pencil beam exits the transport system, it is magnetically steered in the lateral directions to deliver dose to a large treatment field.10,15–17 The proton beam intensity may be modulated as the beam is moved across the field, resulting in the modulated scanning beam technique, or IMPT, delivery. Current implementation of IMPT uses the so-called spot scanning technique, in which the beam spot is moved to a location within the target and the prescribed dose delivered to the spot, before it is moved to the next spot to deliver its prescribed dose. In the beam axis direction, IMPT treatments are delivered using the layer-stacking technique. A pencil beam with a pristine peak is scanned through the deepest layer of the target to deliver the intensity-modulated dose distribution for the layer before a range shifter—effectively an energy degrader of predetermined thickness—is inserted into the beam path to deliver dose to a depth immediately proximal to the deepest layer. Doses to each subsequent layer are delivered by inserting additional range shifters. The size of the spot, represented by the sigma of the Gaussian function describing the pencil-beam profile, is a critical parameter of the pencil beam. Smaller beam sigma values allow intricate sculpting of the intensity-modulated dose distribution, at the cost of increased control system complexity and delivery time; larger beam sigma values result in greater lateral penumbra of the delivered dose distribution and reduce the dose gradient achievable at interface regions of target and critical organs.
Similar to the interplay effect reported for IMRT,18–21 scanning techniques are sensitive to organ motion because moving the pencil beam across the treatment field for dose delivery to a given layer and inserting range shifters for dose delivery to subsequent layers are time consuming. Additional motion mitigation measures, such as breath holding, gated therapy, and abdominal compression, may be necessary to minimize the dosimetric effects of organ motion.
Treatment Planning
Treatment planning for proton therapy requires a volumetric patient computed tomography (CT) scan dataset. The CT Hounsfield Unit numbers are converted to proton stopping-power values for calculating the proton range required for the treatment field.22,23 Unlike the relatively reliable conversion of CT numbers to relative electron density for photon dose calculations, errors and uncertainties in the conversion of CT numbers to proton stopping power in proton dose calculations translate linearly into proton range calculation uncertainties and errors. In clinical practice, these uncertainties are handled during the treatment planning process by bracketing the intended SOBP with a distal margin beyond the target and a proximal margin before the target in the range calculation of each treatment field.24 Other considerations in determining the values for distal and proximal margins include target motion, daily setup errors, beam delivery uncertainties, and uncertainties in patient anatomy and physiology changes throughout treatment that could affect the water-equivalent depth of the target. In the lateral direction of the beam’s eye view (BEV) of a proton field, planning target volume (PTV) margins are necessary to accommodate setup errors and target movements, identical to their usage in x-ray treatments. It is therefore worth noting that the concept of PTV, as defined in the various International Commission on Radiation Units and Measurements (ICRU) reports,1,25,26 does not strictly apply to proton therapy. Generally, a PTV expansion, with either uniform or nonuniform margins, is used for x-ray planning, which accommodates maximum target motion and setup error in each of the patient axes (longitudinal, lateral, and anterior/posterior). In contrast to x-ray planning, the PTV for proton therapy is specific for each treatment field. Lateral margins are identical to traditional definitions, but the distal and proximal margins along the beam axis are calculated to account for proton-specific uncertainties. The lower entrance dose and absence of exit dose of a proton beam afford proton therapy practitioners the additional flexibility to select beam angles where lateral target motion and setup error are minimal; thus, the PTV expansion margin may be less for a given proton beam than would be required in IMRT to accommodate for uncertainties for all beam angles. State-of-the-art proton therapy dose calculations use pencil-beam algorithms,27–30 which model proton interaction and scattering in various heterogeneous media of the beam path, including the nozzle, range compensators, and the patient. Monte Carlo calculations have been used to study the accuracy of such dose calculation algorithms, with results indicating errors near tissue interfaces, particularly near interfaces of media differing significantly in density and composition, such as air cavity and bone in head and neck treatments.31–33 A number of fast Monte Carlo calculation algorithms have been proposed to improve proton therapy dose calculation accuracy,34–36 although none is currently available from commercial treatment-planning system vendors.
POTENTIAL PROTON THERAPY APPLICATIONS
Many publications have reported significant differences in dose distribution with proton treatment plans compared with x-ray-based treatment plans in a wide range of malignancies and benign lesions throughout the body, including the eye,37 brain,38,39 sinonasal structures,40,41 oropharynx, nasopharynx, skull base, lung,42,43 lymphoma,44,45 pancreas,46 esophagus,47 bladder, rectum, pediatric malignancies, prostate,48,49 cervix,50 breast,51,52 sarcomas, and standard target volumes such as pelvic lymph nodes53 and craniospinal irradiation.54 In all cases there is a striking advantage with protons for reduction in the volume of nontargeted normal tissue receiving low- to medium-range radiation doses. In some cases, there is also a reduction in the volume of nontargeted tissue receiving moderate- to high-dose irradiation. With double-scattered proton delivery modes currently in common usage, the target dose homogeneity and conformality index can sometimes, but not always, be inferior to that of IMRT. With scanned proton delivery modes, IMPT plans generally have not only significant reductions in low and moderate integral doses but also improved dose homogeneity and conformality indexes when compared with IMRT. Each case within each tumor type is different, and until comparative plans are performed, it may not be clear whether protons will be helpful or whether IMPT would be useful. At this stage in the development of proton therapy, there are no clear class solutions to treatment planning. In addition, the full potential for dose distribution improvements with protons has not been realized because of uncertainties in both treatment-planning algorithms and delivery modes. Strategies for motion management and quality assurance necessary to implement advanced proton delivery modes into routine practice are not fully developed. Finally, the clinical impact of some patterns of dose distribution improvements achievable with proton therapy may not be anticipated at this time and/or may require time, careful trial design, and special assessments to define.
Currently, proton therapy is a rare medical resource best used in situations where outcomes with commonly available radiation strategies present opportunities for improvement in the therapeutic ratio via improvements in dose distributions. Listed in Figures 19.3 through 19.10 are eight clinical settings that exemplify particular kinds of opportunities for proton therapy. The proton plans are all double-scattered mode, as few practices today are able to deliver IMPT. The first two examples (Figs. 19.3 and 19.4) represent opportunities where reductions in relatively high doses to critical structures make it possible to deliver a sufficiently high target dose with proton therapy to make disease control possible in cases not often cured with photon therapy. The next three examples (Figs. 19.5 through 19.7) are opportunities to reduce early functional and potentially fatal late toxicities with proton therapy by reducing the volume of tissue exposed to low and moderate radiation doses. Figures 19.8 and 19.9 represent opportunities to increase the effectiveness of combined modality therapy by increasing dose or dose intensity of either chemotherapy and/or radiation therapy made feasible through reducing moderate-dose irradiation to critical structures subject to acute toxicity. The last example (Fig. 19.10) represents an opportunity to improve the therapeutic ratio in a tumor generally considered to be well treated with sophisticated x-ray techniques; in this case, a small reduction in toxicity can be leveraged to facilitate hypofractionation, which may both enhance disease control and reduce health care costs without increasing toxicity.
FIGURE 19.3. Axial and sagittal planes from intensity-modulated radiation therapy (IMRT) plans are shown on the left and proton plans on the right for a skull-base sarcoma. In this particular case, the major advantage to the proton plan over the IMRT plan is the sharp gradient between the target and brainstem achievable with the proton plan. The maximum and mean relative doses to the brainstem are 71% and 42% with IMRT compared to 59% and 11% with protons, respectively. The lower maximum and mean relative doses to the brainstem permit delivery of a higher dose to the clivus sarcoma target. In addition, there is a substantial reduction in low-dose exposure to nontargeted tissue such as the posterior fossa, spinal cord, and nasal cavity, which might result in more acute tolerance of treatment or fewer late neurocognitive sequelae.
FIGURE 19.4. Axial and sagittal planes from intensity-modulated radiation therapy (IMRT) plans are shown on the left and proton plans on the right for a paranasal sinus tumor. In this particular case, the major advantages to the proton plan compared to the IMRT plan are reductions in mean dose to the chiasm (4,356 to 3,590 cGy [relative biologic effectiveness, RBE]), the right optic nerve (5,237 to 4,228 cGy [RBE]), and the brainstem (3,698 to 2,743 cGy [RBE]). Target coverage was more homogeneous with proton therapy as well. When 100% of the prescribed dose of 75 Gy/cGy (RBE) was delivered to at least 90% of the target volume with each plan, mean and maximum doses were, respectively, 7,860 cGy (RBE) and 8,680 cGy (RBE) with protons and 8,096 cGy and 9,463 cGy with IMRT.
FIGURE 19.5. Axial and coronal planes are shown from intensity-modulated radiation therapy (IMRT) plans (left), stereotactic radiation therapy (SRT) plans (center), and proton therapy plans (right) for a small craniopharyngioma. As apparent, the volume of tissue exposed to low- and intermediate-dose irradiation is reduced with the proton plan. Mean body and mean brain dose excluding the planning target volume (PTV) are 233 cGy and 888 cGy with IMRT, 215 cGy and 810 cGy with SRT, and 63 cGy (relative biologic effectiveness [RBE]) and 358 cGy (RBE) with protons, respectively. The mean right temporal lobe dose is 904 cGy with IMRT, 1,090 cGy with SRT, and 297 cGy (RBE) with protons. The mean left temporal dose is 951 cGy with IMRT, 1,200 cGy with SRT, and 370 cGy (RBE) with protons. The mean left hippocampal dose is 2,749 cGy with IMRT, 3,299 cGy with SRT, and 815 cGy (RBE) with protons. The mean right cochlear dose is 807 cGy with IMRT, 388 cGy with SRT, and 7 cGy (RBE) with protons. The mean left cochlear dose is 792 cGy with IMRT, 887 cGy with SRT, and 5 cGy (RBE) with protons.
FIGURE 19.6. Sagittal planes from three-dimensional conformal radiation therapy (3DCRT) plans are shown on the left and proton plans on the right for craniospinal axis irradiation necessary in a variety of brain tumors, most of which occur in young patients at risk for late effects. As apparent in the figure, the risks for functional and neoplastic effects in the thyroid, heart, lungs, breast, gut, and gonads from photon therapy can be avoided with proton therapy because of the lack of an exit dose. The total-body V10 and total body integral dose are 37.2% and 0.223 Gy-m3 with 3DCRT compared with 28.7% and 0.185 Gy-m3 with proton therapy, respectively.
FIGURE 19.7. Axial, coronal, and sagittal planes are shown for three-dimensional conformal radiation therapy (3DCRT; left), intensity-modulated radiation therapy (IMRT; center), and proton therapy (right) for a female patient with neck and mediastinal involvement by Hodgkin lymphoma. As apparent, the proton therapy plans expose a smaller volume of nontargeted tissue (particularly heart, lung, spinal cord, and breast) to radiation than either photon plan. Lung V4 and V20 are 59% and 25% with 3DCRT, 62% and 9% with IMRT, and 32% and 16% with proton therapy, respectively. Heart V4 and V20 are 79% and 54% with 3DCRT, 76% and 26% with IMRT, and 40% and 26% with proton therapy, respectively.
FIGURE 19.8. Axial, coronal, and sagittal planes are shown for three-dimensional conformal radiation therapy (3DCRT; left), intensity-modulated radiation therapy (IMRT; center) and proton therapy (right) for a patient with lung cancer. As apparent, both the IMRT and proton therapy plans reduce the volume of nontargeted tissue receiving high-dose irradiation. The proton therapy plan further reduces the volume of nontargeted tissue receiving low- and intermediate-dose irradiation compared to both the 3DCRT and IMRT plans. The 3DCRT, IMRT, and proton therapy plans respectively produced mean lung doses of 1,170 cGy, 1,424 cGy, and 819 cGy (relative biologic effectiveness [RBE]); lung V5 of 36%, 50%, and 21%; lung V10 of 27%, 37%, and 18%; lung V20 of 23%, 22%, and 13%; mean heart doses of 1,156 cGy, 714 cGy, and 548 cGy (RBE); and mean esophageal doses of 3,401 cGy, 2,617 cGy, and 2,171 cGy (RBE).
FIGURE 19.9. Axial, coronal, and sagittal planes of intensity-modulated radiation therapy (IMRT; left) and proton therapy (right) plans for a carcinoma of the pancreatic head. As apparent, there is much less low to intermediate dose to the bowel, kidney, and liver with the proton plan. Mean doses with IMRT and proton therapy are 1,174 cGy and 760 cGy (relative biologic effectiveness [RBE]) to the liver and 1,705 cGy and 443 cGy (RBE) to the small bowel, respectively. While mean right and left kidney doses are similar for the IMRT and proton therapy plans, over 40% of the left kidney tissue is unirradiated with the proton therapy plan.
FIGURE 19.10. Axial, coronal, and sagittal planes of intensity-modulated radiation therapy (IMRT; left) and proton therapy (right) plans for prostate cancer. The five-field IMRT plan appears more conformal than the two-field double-scatter proton therapy plan; however, the proton therapy plan exposes a much smaller volume of nontargeted tissue to radiation. The rectal wall V30, V40, and V50 are 29%, 23%, and 17% with the IMRT plan compared to 18%, 16%, and 14% with the proton therapy plan, respectively.
Skull-Base Sarcomas
Skull-base sarcomas frequently are not amenable to complete resection and require very high radiation doses for disease control. Their adjacency to the brainstem, optic chiasm, and optic nerves often precludes delivery of optimal tumor doses for fear of fatal or severe functional radiation injuries. Proton therapy in these cases can achieve dose distributions that often permit the delivery of potentially curative doses of radiation to the tumor with minimal risk of brainstem necrosis or blindness, and may offer the patient the only realistic chance of cure. Figure 19.3 demonstrates IMRT and proton plans in a patient with skull-base chondrosarcoma. As apparent in Figure 19.3, the IMRT plan delivers low-dose irradiation to a much larger volume of nontargeted tissue than the proton plan, in particular, to the nose, posterior fossa, brainstem, and spinal cord.
The relative minimum dose received by 98% of the volume (D98%), the relative maximum dose received by 2% of the volume (D2%), and the relative mean dose to the spinal cord are 62%, 56%, and 61% with IMRT compared with 16%, 15%, and 1% with proton therapy, respectively. For the brainstem, the D98%, D2%, and mean doses are 20%, 66%, and 42% with IMRT compared with 0%, 39%, and 10% with proton therapy, respectively. The significant reduction in relative dose to the brainstem and spinal cord permits the delivery of higher doses to the tumor with proton therapy. In other skull-base sarcomas, the benefits from proton therapy may differ or differ in degree. With the currently available double-scatter mode of proton delivery, “patch” fields are often required (as in this case) to avoid critical structures at the expense of a considerable increase in planning and delivery complexity to minimize dose inhomogeneities within the target. In essence, separate small fields (which cover parts of the target volume from optimal angles to avoid critical structures) are “patched” together to cover the entire target, somewhat similar to IMRT. In the aforementioned case, D98%, D2%, and mean target doses are 96%, 106%, and 104% with IMRT compared with 94%, 118%, and 110% with proton therapy, respectively. Care must be taken to avoid placing potential hot spots in critical structures. In the future, scanning-mode proton therapy will eliminate the need for patch fields and reduce the need for apertures and compensators, providing increased dose homogeneity within the target.
It is highly likely that the benefits of proton therapy in skull-base tumors will increase with the availability of IMPT. In addition, future improved treatment-planning systems may also more accurately estimate the actual dose inhomogeneity at tissue interfaces and other heterogeneities, including possible surgical hardware in the base of the skull. The dosimetric validation of treatment-planning systems in these complex anatomic regions is difficult and the use of Monte Carlo calculations can be helpful.32
Paranasal Sinus Tumors
Paranasal sinus tumors frequently extend into the orbit or anterior cranial fossa adjacent to critical optic structures, such as the chiasm, optic nerves, retinae, lacrimal glands, cornea, and lens. With photon-based therapy, it is often difficult to deliver adequate doses to the entire tumor target without injury to at least one of the critical optic structures. The physician must choose between prioritizing tumor control and preserving vision. Figure 19.4 shows a comparison of IMRT and proton plans in a patient with a paranasal sinus tumor. Both plans were specified to deliver the prescribed dose to at least 90% of the target volume. Compared to the proton plan, the IMRT plan exposes a much larger volume of nontargeted tissue to low-dose radiation, which includes the right temporal lobe, posterior fossa, oral cavity, and supratentorial brain. In addition, the IMRT plan produces a less homogeneous dose within the target with D98%, D2%, and mean doses of 82%, 118%, and 109% compared to 93%, 112%, and 106% with proton therapy, respectively. If 90% of the target receives the prescribed dose of 75 Gy/cobalt gray equivalent (CGE), likely necessary to control such tumors, the target dose ranges from approximately 61.5 to 88.5 Gy with a mean dose of 88.5 Gy with IMRT and from 69.8 (RBE) to 84 Gy (RBE) with a mean dose of 79.5 Gy (RBE) with protons. There is relative sparing of most of the optic structures with proton therapy with mean doses to the chiasm, right optic nerve, left optic nerve, and brainstem of 44 and 36 Gy (RBE), 53 and 43 Gy (RBE), 44 and 36 Gy (RBE), and 43 and 29 Gy (RBE) with the IMRT and proton plans, respectively. As with skull-base sarcomas, the benefits and degree of benefits with proton therapy will vary among different paranasal sinus tumors and are highly likely to increase with the availability of IMPT.
Craniopharyngioma
Craniopharyngioma is usually diagnosed in children and adolescents. Its suprasellar location places the temporal lobes, hippocampi, hypothalamus, optic chiasm, and nerves at risk for radiation injury. Figure 19.5 shows IMRT, SRT, and proton plans in an adolescent patient. All three plans achieve target coverage goals of 95% of the prescribed dose (54 Gy) to 100% of the target and 100% of the prescribed dose to at least 95% of the target. Mean body and mean brain doses excluding the PTV are, respectively, 4% (2.2 Gy) and 17% (9.2 Gy) with IMRT, 4% (2.2 Gy) and 15% (8.1 Gy) with SRT, and 1% (0.5 Gy [RBE]) and 6% (3.2 Gy [RBE]) with proton therapy. The relative mean doses with IMRT, SRT, and proton therapy for the following structures are right temporal lobe—17%, 20%, and 8%; left temporal lobe—18%, 22%, and 10%; left hippocampus—50%, 61%, and 16%; right cochlea—16%, 7%, and 0%; and left cochlea—14%, 16%, and 1%, respectively. These reductions in dose to nontargeted brain tissues with proton therapy are likely to result in reduced loss in neurocognitive and auditory function.
Craniospinal Axis Irradiation
Craniospinal axis irradiation is required in most medulloblastomas and occasionally in other brain tumors, such as advanced or metastatic germ cell tumors, primitive neuroectodermal tumors (PNETs), and ependymomas. Most patients with these tumors are young and at risk for late effects of radiation. As shown in Figure 19.6, the exit dose from photon therapy exposes the thyroid, heart, lung, gut, and gonads to functional and neoplastic risks that can be avoided with proton therapy. The total-body V10 and total-body integral dose are, respectively, 37.2% and 0.223 Gy-m3 with 3DCRT compared with 28.7% and 0.185 Gy-m3 with proton therapy, a reduction likely to result in a lower risk of second malignancy.
Lymphomas
Lymphomas frequently involve the mediastinum but typically require only a moderate dose of radiation therapy in conjunction with chemotherapy for disease control. Unfortunately, even low to moderate radiation doses place the patient at risk for late cardiac injury and second cancers, particularly breast cancers. Figure 19.7 shows a comparison of 3DCRT, IMRT, and proton therapy plans in a young woman with Hodgkin lymphoma. The proton plan shows a significant reduction in the volume of heart, lung, breast, spinal cord, and other soft tissues exposed to low-dose irradiation. Mean relative lung dose, lung V4, and lung V20 are 48%, 59%, and 25% with 3DCRT; 43%, 62%, and 10% with IMRT; and 27%, 31%, and 16% with proton therapy, respectively. Mean relative cardiac dose, cardiac V4, and cardiac V20 are 72%, 79%, and 54% with 3DCRT; 57%, 76%, and 26% with IMRT; and 37%, 40%, and 26% with proton therapy, respectively. These reductions are likely to result in lower risks of late cardiac injury and second malignancy.
Lung Cancers
Lung cancers typically are diagnosed at an advanced stage and occur in patients with underlying lung damage. Consequently, concern for protection of unaffected lung tissue often mandates compromise in the tumor dose. Figure 19.8 shows IMRT and proton plans in a patient with stage III lung cancer. As apparent, a smaller volume of nontargeted lung tissue, spinal cord, esophagus, and heart is exposed to radiation with proton therapy. The mean relative lung dose was 23% with 3DCRT, 19% with IMRT, and 11% with proton therapy. Lung V4 and V20 were 40% and 23% with 3DCRT, 54% and 22% with IMRT, and 22% and 13% with proton therapy, respectively. The mean relative heart dose was 15% with 3DCRT, 9% with IMRT, and 7% with proton therapy. The relative mean esophagus dose was 45% with 3DCRT, 35% with IMRT, and 28.7% with protons. In this case, the proton plan lowers the risk of acute (potentially fatal) pneumonitis and acute esophagitis, likely impacting the delivery of chemotherapy, as well as the cardiac exposure, likely correlating with greater chance of survival.
Pancreatic Cancers
Pancreatic cancers have an extremely low therapeutic ratio with radiation alone or combined with surgery and chemotherapy. The disease is frequently localized for a window of time before spreading, providing a potential opportunity to improve the overall outcome by intensifying local therapy. Figure 19.9 shows IMRT and proton therapy plans for a patient with cancer in the head of the pancreas. As apparent, there is much less low to intermediate dose to the bowel, kidney, and liver with the proton therapy plan. Mean relative doses with IMRT and proton therapy are 23% and 15% to the liver and 34% and 9% to the small bowel, respectively. While mean right and left kidney doses are similar for the IMRT and proton therapy plans, over 40% of the left kidney tissue is unirradiated with the proton therapy plan. This savings in normal-tissue exposure may be leveraged to permit either radiation or chemotherapy dose escalation or intensification, potentially increasing the opportunity for complete surgical resection, cure, or both.
Prostate Cancer
Prostate cancer results with IMRT are generally excellent, but dose-escalation trials from the M.D. Anderson Cancer Center (Houston, TX)55 show that the volumes of rectum and rectal wall receiving low- to moderate-dose radiation with x-ray-based therapy are significantly associated with the incidence of gastrointestinal toxicity. Dosimetry studies48 show that the low to moderate doses delivered to the rectum with proton therapy are less than with IMRT. Figure 19.10 shows a five-field IMRT plan and a two-field double-scattered proton plan for a patient with low-risk prostate cancer. Rectal wall V30, V40, and V50were 29%, 23%, and 17% with IMRT compared with 18%, 16%, and 14% with proton therapy, respectively, potentially providing a lower risk of rectal injury (see discussion regarding clinical significance later).
CLINICAL EVIDENCE FOR PROTON THERAPY
Efficacy and Toxicity
Despite the existence of only a few facilities with technology capable of proton delivery to most cancers, the efficacy of proton therapy has already been demonstrated in a variety of malignancies and benign lesions including representative series in eye cancers56–60; base of skull sarcomas61,62; brain and spinal cord tumors63,64; paranasal sinus tumors65,66; oropharyngeal carcinoma67; esophageal cancer68; early- and advanced-stage lung cancer69,70; low-, intermediate-, and high-risk prostate cancer71–73; Hodgkin and other lymphomas74,75; sarcomas76; a variety of pediatric malignancies77–81; hepatocellular carcinoma82; pancreatic cancer83; cervical cancer84; and benign lesions like acoustic neuroma, vestibular schwannoma, arteriovenous malformations (AVMs), age-related macular degeneration, pituitary adenomas, craniopharyngiomas, and meningioma.85,86 In general, toxicity rates reported with proton therapy appear to be low; however, comparisons with contemporary x-ray-based therapies are often difficult because of the absence of controlled studies, small patient numbers, patient selection, lack of appropriate comparative groups, and variable criteria for toxicity assessment.
Comparative Effectiveness
In most clinical situations, level 1 evidence of comparative effectiveness is desirable; however, it has been difficult to conduct randomized controlled trials in proton therapy. While there may be small differences in relative biologic effectiveness with proton therapy compared to photon therapy, laboratory and clinical data suggest that these differences are quite small. Therefore, the basic difference between protons and photons is simply the difference in entrance dose and exit dose to nontargeted tissues. The essence of a randomized controlled clinical trial of proton therapy and x-ray-based therapy would be the question of whether low to intermediate doses to nontargeted tissue would result in measureable good or harm to a patient. The nature of this question raises some ethical concerns. In addition, there are practical issues: patients may not choose to enroll in such a trial, particularly if they must travel to one of the few currently available proton facilities to face randomization between proton and photon therapy. Finally, the numbers of patients and resources required for randomized controlled trials raise concerns regarding the best use of currently available proton therapy facilities, specifically whether more questions could be answered using the same time frame, number of patients, and level of resources with a series of well-designed sequential studies aimed at optimizing proton therapy through dose escalation or hypofractionation. For a variety of ethical and practical reasons, there are currently no completed prospective randomized comparative-effectiveness trials between proton therapy and other radiation modalities, although two are under way. One randomized trial between proton therapy and IMRT in stage III lung cancer is ongoing as collaboration between the M.D. Anderson Cancer Center and Massachusetts General Hospital (Boston, MA). A second trial at the University of Heidelberg (Germany) in skull-base sarcomas randomizes patients to either proton therapy or carbon ion therapy.87
When there is a compelling dosimetric advantage to proton therapy, or strongly suggestive level 2 and 3 clinical evidence, level 1 evidence of an increased therapeutic ratio may not be necessary. Examples of clinical situations in which level 1 evidence may not be necessary include base-of-skull and paranasal sinus tumors, pediatric and young adult cancers, and eye malignancies.
The therapeutic ratio in paranasal sinus tumors with conventional x-ray-based radiation therapy has been low, with either low cure rates88 or high visual toxicity rates,89 which suggests physician prioritization of toxicity avoidance in the first scenario and tumor control in the second. Dosimetry studies comparing proton therapy and IMRT in patients with paranasal sinus tumors have suggested a substantial benefit in dose distribution from proton therapy40,90that could lead to clinical benefits. Reports of the clinical experience with proton therapy for paranasal tumors from Massachusetts General Hospital suggest a notably higher therapeutic ratio, with both high disease control rates and low toxicity rates,65,66 than previously reported.88,89 Such remarkable improvements in the therapeutic ratio for proton therapy compared with contemporary x-ray-based experiences may obviate the need for a controlled trial. Similarly, excellent disease control rates have been achieved in skull-base sarcomas at Massachusetts General Hospital with proton therapy dose regimens that are not feasible with photon-based therapy, likely obviating the desire for level 1 clinical evidence for proton therapy.
In pediatric cancer survivors, an increased risk of second malignancies has been associated with radiation doses as low as 4 to 15 Gy.91,92 Similar doses are believed to be associated with increased late cardiac injury.93 In addition, elegant dose-modeling studies at St. Jude Children’s Hospital (Memphis, TN) have suggested no threshold dose for radiation injury in the childhood brain.94 Thousands of children and many years of follow-up would be required to assess the results of a hypothetical controlled trial of proton therapy and conventional radiation to determine whether the reduction in low-dose exposure to the brain and total body with proton therapy would indeed result in a lower incidence of second malignancies or late cardiac injury. In the process, many children would intentionally be exposed to low-dose radiation in nontargeted tissues in the brain, likely resulting in permanent decreased neurocognitive function. Therefore, most clinicians consider the dosimetric evidence of significant reduction in the volume of tissue exposed to low- and moderate-dose irradiation sufficiently compelling to obviate the need for a controlled trial of proton therapy in children. Similar arguments follow for Hodgkin lymphoma, another malignancy occurring in relatively young patients. Hodgkin lymphoma has a high cure rate and long life expectancy but also high risks for second malignancy and late cardiac injury.44,45
There is substantial experience with proton therapy for choroidal and uveal melanomas. Several large retrospective comparative studies of long-term outcomes using various strategies for choroidal or uveal melanoma have provided solid level 2 evidence of the benefits from proton therapy57 with respect to disease control and preservation of vision, likely making a controlled trial unnecessary.
The area of greatest controversy is prostate cancer. More than 10,000 men with prostate cancer have been treated with proton therapy or a combination of proton therapy and photon therapy with excellent results. Disease control rates appear comparable to outcomes with photon therapy given with the same dose-fractionation schedules72; however, toxicity, particularly rectal toxicity, appears to be lower with proton therapy than with photon therapy at the same dose-fractionation schedules.71–73 While dose-escalation trials in prostate cancer have typically demonstrated a significant rise in toxicity with higher doses,55,95,96 when protons were used for approximately one-third of the dose in one dose-escalation trial,73 there was no significant increase in grade 3 toxicity with the higher doses, and toxicity rates for both dose levels appeared to be lower than reported with contemporary photon dose-escalation trials. A Radiation Therapy Oncology Group (RTOG) study of 3DCRT in localized prostate cancer97 found that increasing both the daily dose to 2 Gy and the total dose to 78 Gy resulted in a higher risk of grade 2 or higher gastrointestinal toxicity than observed with regimens of 79.2 Gy or less delivered in 1.8-Gy once-daily fractions or 74 Gy delivered in once-daily fractions of 2 Gy. The rate of grade 3 or higher gastrointestinal toxicity with 78 Gy delivered in 2-Gy fractions was 4% when the prostate only was treated and 7% when the prostate and part of the seminal vesicles were treated. In contrast, with proton therapy on three prospective trials at a single institution, a grade 3 toxicity rate of <1% was observed with doses of 78 Gy (RBE) delivered in once-daily 2-Gy (RBE) fractions,71 suggesting that even small reductions in the volume of rectum exposed to moderate radiation doses may impact the rate of radiation toxicity and that, in the case of prostate cancer, these relatively small toxicity reductions may facilitate the delivery of hypofractionated regimens, which could reduce health care costs and potentially increase disease control.98Because increased toxicity observed with dose escalation and hypofractionation (or dose intensification) with photons55,95–97 has not been observed thus far in the proton experiences,71,73 there is speculation that the potential benefits from proton therapy in prostate cancer may be increasingly apparent with either dose-escalated or hypofractionated regimens. One driver of interest in comparative effectiveness has been economic concern—proton therapy appears more expensive than photon therapy. Differential potential for safe hypofractionation of the treatment regimen might reverse the economic concerns regarding the cost-effectiveness of proton therapy and photon therapy.
ECONOMICS
Many factors must be accounted for in comparing the costliness and cost-effectiveness of different radiation technologies. Some of these include facility and equipment costs, operating and maintenance costs, patient throughput, and the global and individual financial impact of clinical outcomes. The perspectives of the patient, the providing physician and institution, the payor, and the nation on these factors may differ.
Current proton therapy equipment and facilities are up to 10-fold more expensive than current conventional radiation therapy facilities capable of treating a similar patient volume. In addition, operating costs are higher by one- to threefold. However, there are substantial differences between patient throughput in conventional radiation and proton facilities that must be accounted for in any financial comparison.
Conventional radiation therapy business models are based on treating a single—or slightly extended—shift each day, 5 days a week. The equipment generally depreciates over 7 to 10 years and replacement occurs when equipment begins to fail, becomes technically obsolescent, or is less promising than an alternative technology. When capacity with a single shift is exceeded, most facilities choose to add another treatment machine rather than add another shift of personnel. It is possible that many conventional radiation facilities are underutilized, replaced prematurely, or both; however, either economic considerations or equipment durability issues continue to support the single-shift model of conventional radiation therapy delivery and the 7- to 10-year replacement schedules.
In contrast, most proton therapy facilities operate two shifts a day and equipment is sufficiently robust to last 20 or more years. The Harvard cyclotron was in use for over 40 years and closed primarily because of the availability of a new gantry-based clinical facility. The Loma Linda University (Loma Linda, CA) synchrotron and most of its gantries have been in continuous operation for over 20 years. Proton therapy equipment appears to be substantially more durable than linear accelerators. Patient mix in conventional radiation therapy and proton therapy facilities may differ, but there does not appear to be a significant difference in time required for treatment of a given clinical condition. Thus, the actual patient throughput (for a given patient mix) in a proton facility may be substantially higher than typical for a conventional radiation therapy facility, perhaps double the patients per day (because of double shifts) for two to six times as many years of operation, leading to a four- to 12-fold increased throughput with proton facilities compared to photon facilities. Proton therapy throughput may be further enhanced if hypofractionation is more feasible with proton therapy than with x-rays. (As discussed earlier, proton therapy has been delivered to early-stage prostate cancer patients in 8 weeks with the same or less toxicity as a 9-week course of IMRT.71,97,99)
While initial capital costs and operational costs of proton facilities are substantially higher than conventional radiation facilities, long-term patient throughput with a proton facility may more than offset the initial capital costs and increased operational costs. If patient outcomes are also improved with proton therapy, as the rationale and early data in some tumor sites suggest, there will be additional savings related to reduced costs associated with treatment toxicity and disease recurrences. Accordingly, a blanket statement comparing cost-effectiveness of proton therapy and photon therapy based solely on initial capital costs of equipment is naïve and misleading.
Nevertheless, a shift to proton therapy requires a paradigm shift in philosophy about profit: it will not be a rapid “return on investment” but rather a benefit realized over time by way of increased throughput and improved outcomes. Daily operations must be focused not on work completion within a shift, but on maximal utilization of beam time, which implies focused attention on preventative maintenance, facility staffing, patient expectations, and program culture.
CONCLUSIONS
Proton therapy offers the promise of reduced toxicity to patients compared with photon therapy by reducing the radiation dose to nontargeted tissues. Reduced toxicity may be leveraged to increase disease control through dose escalation or intensification (hypofractionation). Hypofractionation may result in lower health care costs, as well as increased disease control and reduced toxicity. Additional research is needed to optimize treatment planning and delivery of proton therapy, to document and maximize its potential clinical benefits, and to understand its full impact on health care economics.
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