Intensity modulated proton therapy.
Journal: 2015/October - British Journal of Radiology
ISSN: 1748-880X
Abstract:
Intensity modulated proton therapy (IMPT) implies the electromagnetic spatial control of well-circumscribed "pencil beams" of protons of variable energy and intensity. Proton pencil beams take advantage of the charged-particle Bragg peak-the characteristic peak of dose at the end of range-combined with the modulation of pencil beam variables to create target-local modulations in dose that achieves the dose objectives. IMPT improves on X-ray intensity modulated beams (intensity modulated radiotherapy or volumetric modulated arc therapy) with dose modulation along the beam axis as well as lateral, in-field, dose modulation. The clinical practice of IMPT further improves the healthy tissue vs target dose differential in comparison with X-rays and thus allows increased target dose with dose reduction elsewhere. In addition, heavy-charged-particle beams allow for the modulation of biological effects, which is of active interest in combination with dose "painting" within a target. The clinical utilization of IMPT is actively pursued but technical, physical and clinical questions remain. Technical questions pertain to control processes for manipulating pencil beams from the creation of the proton beam to delivery within the patient within the accuracy requirement. Physical questions pertain to the interplay between the proton penetration and variations between planned and actual patient anatomical representation and the intrinsic uncertainty in tissue stopping powers (the measure of energy loss per unit distance). Clinical questions remain concerning the impact and management of the technical and physical questions within the context of the daily treatment delivery, the clinical benefit of IMPT and the biological response differential compared with X-rays against which clinical benefit will be judged. It is expected that IMPT will replace other modes of proton field delivery. Proton radiotherapy, since its first practice 50 years ago, always required the highest level of accuracy and pioneered volumetric treatment planning and imaging at a level of quality now standard in X-ray therapy. IMPT requires not only the highest precision tools but also the highest level of system integration of the services required to deliver high-precision radiotherapy.
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Br J Radiol 88(1051): 20150195

Intensity modulated proton therapy

Optimization

The physician's prescription is a set of quantified statements (i.e. maximum dose to target), suitable for numeric manipulation in an optimization algorithm, and generalized intent or objectives (i.e. minimize dose to the brain stem). The quality of a particular objective is limited by the prescription that fixes the boundaries of what remains possible in a trade-off consideration of competing objectives. IMPT has a large number of variables: the value for each of the 1000–10,000 pencil beam spots. Thus, the physician will have considerable choice to consider one or more trade-offs to meet the overall intent or to further “improve” the treatment plan given a prescription. The prescription typically is the pragmatic copy of an X-ray prescription and hence reflects the constraints of X-ray radiotherapy, which presumably can be improved by the improved IMPT dose distribution.

Multicriteria decision analysis [or multicriteria optimization (MCO)] is a subfield of operations research that formalizes the inherent conflict in multiple criteria in the decision-making process. Pareto optimality (named for its inventor, V. Pareto) is one formal technique that computes the set of optimal solutions in terms of values for each criterion such that improving one such criterion necessarily worsens all other criteria. Thus, any Pareto optimal treatment plan is the best achievable given set of trade-off values (i.e. reducing brain stem dose necessarily results in worsening target coverage). MCO treatment planning is emerging as a necessary improvement on current “single” plan optimization methods. For these “single” plan methods, the user has no choice but to achieve a “better” plan through trial and error while not, in fact, knowing the optimality of any achieved plan.

Common MCO techniques use multilevel optimization33 or Pareto optimization.34 For multilevel optimization, the optimization proceeds stage-wise whereby each stage takes the optimized plan from the previous stage and attempts to optimize the current stage given additional criteria (i.e. now minimize the brain stem dose for the plan that has maximized the minimum target dose). Each stage produces a single plan and requires re-execution of the pipeline beyond a current stage if their optimization criterion is changed. Pareto optimization computes the multispace of all Pareto optimal plans. Of course, this space is very large and in practice approximated by a sufficient set of plans that represent the whole space and where other plans are obtained through interpolation. The user can traverse this space by selecting the value for a particular constraint (i.e. minimized brain stem dose is 45 vs 50 Gy) and inspect the consequence of this choice in the presence of the value change on all other intent values (i.e. the target minimum dose drops to 65 vs 70 Gy as a consequence). Our clinical practice uses Pareto optimization to quantitatively assess clinical trade-off.35Figure 8 shows the use of MCO in the breast treatment case (Figure 1) to allow the physician to consider the effect of lowering the thyroid dose, albeit with an unavoidable lowering of target coverage.

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Demonstration of the use of multicriteria optimization to achieve improved thyroid minimum dose, albeit with worsened target coverage for the breast chestwall irradiation shown in Figure 1. The use of multicriteria optimization allows the physician to navigate the space of trade-offs while maintaining the absolute constraints of a treatment prescription. The sliders on the left show the value range for minimizing the mean thyroid dose. The slider shows the minimum and maximum achievable values of 9.5 Gy and 21.2 Gy (radiobiological effect) and the current values of 15.7 (top) and 9.5 (bottom). The white rectangle is the user control to change the trade-off value. The dose display shows how the dose “pulls” away from the oesophagus in the bottom, which results in reduced target coverage (dose inside the purple contour).

Dose accuracy

Accurate modelling of dose in the patient is limited by the presence of physical uncertainties such as stopping power accuracy, dose to tissue36 calculation, dose calibration13 and by the presence of geometric and anatomical uncertainties. Standard practice for clinical dose computations uses pencil beam algorithms (PBAs)37 whose spatial precision is limited (theoretically and at best) to the multiple Coulomb scatter (σ > 0.02 × R) resolution and whose dosimetric accuracy is limited by the approximation of water equivalence of any structure.22 An empirical pencil beam calculation can therefore not model deep heterogeneities better than the Coulomb scatter resolution. This may lead to underprediction or overprediction of dose around deep (relative to the penetration of the proton beam) heterogeneities. Only MC can resolve details of such effects, if any.

Figure 9 (see also Grassberger et al38) shows fields in three example patients (liver, base of skull and lung) and demonstrates how the discrepancy between MC simulation and PBA increases with increasing geometrical complexity.

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Dose distributions calculated using Monte Carlo (MC, left column), pencil beam algorithm (PBA, middle column) and their difference (right column) in three example patients with varying complexity (liver top row, base of skull middle row, lung bottom row). The PBA modelling of the increasingly complex topology of heterogeneities causes an increase differential with the more able MC. Adapted from Grassberger et al.38

Schuemann et al28 discuss the need for site-specific range margins (albeit applied to SOBP fields) based on a comparison of the distal dose surfaces achieved by a PBA (in-house implementation based on Hong et al37) and MC (TOPAS39) where each calculation produced the identical SOBP in water. They observe deviations (MC-PBA), for head and neck fields, for example, in the order of −2 mm (or −1.5% relative to the nominal range) for the points along the same ray where the 90% distal dose is achieved. Other treatment sites showed similar but significantly different systematic deviation. This result indicates that the PBA overestimates the proton penetration. Conversely, it implies, depending on prescription practices, that the distal target is underdosed by individual fields. The overall impact of the single-field deviation is mitigated, in our practice, through the use of multiple fields. It should also be emphasized that the transition to MC-based clinical computations is inevitable, if only because computing hardware technologies place such calculations within reasonable times,40 but even an MC calculation needs to be benchmarked and calibrated against measurements and the community must establish standard and traceable benchmarks for its clinical deployment.41

Any dose calculation is a systematic deviation from the “true” dose distribution yet compatible with clinical practice that empirically and formally equates dose computed within practice-guided accuracy to expected and observed outcome in the patient. For proton radiotherapy, this means that prescription and calculation practices must be consistent. It is precisely within the domain of IMPT with its various uncertainties that much effort is expended to ensure this consistency. Improvements such as those achievable through MC dose calculations28 must be translated into clinical practice while being cognizant of the empirical expectation.

Motion management

Dose prediction in the patient must consider the effect of temporal changes—set-up uncertainties, anatomical changes and organ motion—in the patient to which proton dose distributions are particularly sensitive. In X-ray therapy, such changes are well respected by (artificially) folding the uncertainties into a planning margin around a structure such that if the dose to the margined structure is achieved, dose to the structure itself certainly is achieved. The margin expansion technique is possible for X-ray therapy because the X-ray dose envelope is spatially invariant within the clinical uncertainties. Changes in radiological density change the X-ray attenuation by the order of 4% cm. Thus, geometrically placing the target within (or structure without) this envelope ensures compliance with the dose intent. In proton therapy, the concept of a planning (target) expansion volume is invalid as the dose envelope is sensitive to the uncertainties. Changes in radiological density change the position of the Bragg peak 100% cm. Changes must be considered explicitly. For SOBP treatments, these changes are considered by increasing the penetration range and modulation of the SOBP field, by increasing the aperture and by “smearing” the range compensator.42 None of these options apply to IMPT.

Current TPSs, typically, allow the user to specify the expected variances in position of the isocentre and of the range. The nominal computed plan is tested against its representation at combinations of variant values, and the user must consider whether the variant plans remain acceptable.

We wish, however, to compute a treatment plan in which the dose constraints and objectives are met for every possible set of uncertainty values. That is, the treatment plan must be “robust” (or insensitive) in the presence of presumed uncertainties. This computation requires the computation of many (in the order of 10 or more) combinations where the treatment plan is optimized simultaneously over all combinations.43,44

Similarly, in the consideration of motion, the treatment plan is optimized simultaneously over all volumetric representations of the moving anatomy such as are available through a four dimensional CT. This, of course, increases the computational combinations by a factor of 10 again!

Finally, both IMRT and IMPT have a time structure: the movement of the MLC in IMRT and the temporal sequence of the pencil beam spots in IMPT. The MLC temporal structure is largely decoupled from the motion problem, and, again, uncertainties can be folded into an appropriate margin.45 The spot sequence is fast (200–1000 cm s) but is constrained by the energy switching time, currently in the order of 1 s. Thus, the temporal evolution of the spot sequence must be synchronized with the temporal evolution of the motion46 to assess the correlation between the operational envelope of the TDS—magnet speed, energy switching time, spot size, etc.—and the patient dynamics. If these factors are not considered, the interference between the dynamic pencil beam delivery and the tumour motion, typically referred to as interplay effect, can lead to dose degradation within the target, as demonstrated in Figure 10.47 Panels Figure 10a–d show the dose distributions after a single fraction for delivery starting at varying points in the breathing cycle (T0, peak inhale; T50, end exhale), and Figure 10e shows how delivery over many fractions can mitigate the interplay effect.

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Sagittal view of the dose distributions for a patient with lung cancer with clinical target volume in end-exhale phase (red, small contour) and internal clinical target volume (pink, large contour). The tumour motion amplitude is 30 mm. Panels (a) through (d) denote the one-fraction cases for delivery starting in four breathing phases, panel (e) the dose distribution after a standard course of fractionation, and panel (f) the planned dose distribution on the static CT. Adapted from Grassberger et al.47

In short, treatment planning must move to a complete simulation system of the treatment and patient dynamics. Such simulations, however, must correlate with the actual state of the patient at any given point in the treatment course. Unlike the “population average” method by van Herk et al,45 IMPT must incorporate sufficient daily imaging and treatment plan dosimetry verification, and adaptation if needed, to ensure that the treatment plan remains within specification. This, in turn, will relax the computational burden of the pre-treatment treatment plan but increase this burden at the time of treatment, as uncertainties can be assessed directly and accounted for with increased precision.

Robust treatment planning

Robust treatment planning aims to achieve treatment plans whose delivery parameters create a dose distribution that satisfies the clinical prescription if the patient's treatment representation remains within an uncertainty envelope considered in the treatment plan. The robust plan is obtained by simultaneously optimizing the treatment plan parameters over all (or adequately represented) uncertainty scenarios such that the treatment plan satisfies the prescription. Robust treatment planning, in essence, perpetuates the concept that a single pre-treatment treatment plan can or must be achieved. In clinical practice, however, the uncertainty envelope can be managed through daily volumetric imaging to reduce the uncertainty to its minimum and to allow for per-treatment reoptimization of the treatment plan. Thus, site-specific analysis should identify the inherent robustness of the nominal treatment approach as quantified by dosimetric quality indicators and as quantified by the biological response variation. In general, any treatment plan should be computed with its uncertainty intervals, a necessity identified long ago48 yet still not available in clinical practice. Finally, robust treatment planning or its (proposed above) site-specific substitute is the correct method for managing the “PTV” concept in proton planning.49

Robust optimization considers primary physical and mathematical statements of the problem. The latter considerations include, amongst others, worst-case50 and likelihood analysis,51 which consider the details of the time evolution of the treatment course treatment delivery. Of specific significance in the latter is the actual simulation or estimation of the treatment delivery sequence in the presence of organ motion.46,52,53 In any case, the only definitive marker will be equivalence of biological response over the course of treatment that current treatment planning does not quantify well.

The level of robustness also depends on the quality of the optimization process itself. Indeed, a simplistic optimization algorithm54 might overemphasize the use of the distal, and sharpest, edge of the proton beam. Such an overemphasis might result in a plan that uses that distal edge to achieve the sharpest penumbral fall-off between the target and an OAR. This, however, ignores the currently largest uncertainty, i.e. the actual position of that edge. Thus, robustness invariably depends on how the dose contributions from multiple beams interdigitate, which favours smooth individual beam dose distributions (see, for example, Figure 4).

Radiobiological considerations

The clinical radiobiological response in tissue from proton interactions events is assumed equivalent to those from photon (X-ray) interactions except for a uniform scaling RBE factor of 1.10. The unit of proton dose is therefore given as Gy (RBE) and implies that the stated value [say 50 Gy (RBE)] is equal to that of C (as a traceable dosimetry standard) delivered dose.

The interaction profile of proton events is more differentiated as this single factor implies. Even in clinical practice, the biological dose penetration of a Bragg peak (say to 100 mm at 90% peak dose) is projected deeper in the order of 1–2 mm (i.e. 101–102 mm)55 and is one reason for not placing the distal penumbra between a target volume and a critical structure (the other reason, again, is range uncertainty). Beyond this obvious effect, the “real” RBE for proton is not known and is confounded with clinical experience based largely on SOBP treatments that tend to use a large number of fields (approximately 10 in our practice) for complex cases (such as may be exemplified in the treatment of chordoma) and where any local RBE enhancement is minimized because of the field arrangement. Paganetti et al55 do consider the clinical consequence of a 5–10% underestimation in the RBE value, which equates to an overdose in the patient (i.e. >1.1 C dose). An overdose effect on a complication response depends on the shape of the dose–response curve. The authors conclude that a 10% underestimation would result in an unacceptable rise in complication. Given the absence of such an increase in clinical practice, the clinical use of RBE = 1.1 appears justified.

The physical effect responsible for RBE variation is the variation in linear energy transfer (LET), which quantifies how much energy the proton particle transfers per unit traversed distance. The dose-averaged LET distribution56 from primary protons is in the order of 0.5 keV μm and rises steeply across the Bragg peak to 10 keV μm, while secondary protons have an LET of about 12 keV μm. Thus, secondary protons whose contribution to the dose is at most 10% (see also Clasie et al13) have a much larger effect on LET, and thus RBE, heterogeneity. The LET for SOBP fields increases at the distal edge of the SOBP field, as the most distal peak in the SOBP field carries a very significant weight (approximately 75%) compared with the other pristine peaks in the SOBP field. Figure 11 exemplifies this for a SOBP field in a patient, showing the marked increase in LET towards the end of range.

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Dose (left) and linear energy transfer (LET, in units of keV/Mm) (right) distributions for a scattered proton field in a base of skull patient. As expected for a scattered proton field, the LET increases toward where the most distal peak deposits most of the dose and hence its LET increase dominates the overall LET differential over the field. An intensity modulated (IMPT) field, in combination with other fields in the IMPT set, will smooth the LET differential over the target. The IMPT field set optimization, however, could include specific increase of LET differentials within the target such as may be of advantage in dose-painting scenarios. Adapted from Grassberger and Paganetti.56

For IMPT fields, the individual peaks have a very different contribution profile to the dose in patient. The individual peak weights are selected to achieve the dose objectives but without, in general, a consideration of the underlying LET heterogeneity. The large number of IMPT spots and the use of MC allow for optimization of the dose distribution, the desired LET profile and biological response.5759

Department of Radiation Oncology, Massachusetts General Hospital and Harvard Medical School, Boston, MA, USA
Corresponding author.
H M Kooy: moc.liamg@yook.ennah; C Grassberger: ude.dravrah.hgm@snemelC.regrebssarG
Address correspondence to: Dr Hanne M Kooy. E-mail: ude.dravrah.hgm@yookh
Department of Radiation Oncology, Massachusetts General Hospital and Harvard Medical School, Boston, MA, USA
H M Kooy: moc.liamg@yook.ennah; C Grassberger: ude.dravrah.hgm@snemelC.regrebssarG
Received Received on March 5, 2015; Revised Revised on May 22, 2015; Accepted Accepted on May 26, 2015.

Abstract

Intensity modulated proton therapy (IMPT) implies the electromagnetic spatial control of well-circumscribed “pencil beams” of protons of variable energy and intensity. Proton pencil beams take advantage of the charged-particle Bragg peak—the characteristic peak of dose at the end of range—combined with the modulation of pencil beam variables to create target-local modulations in dose that achieves the dose objectives. IMPT improves on X-ray intensity modulated beams (intensity modulated radiotherapy or volumetric modulated arc therapy) with dose modulation along the beam axis as well as lateral, in-field, dose modulation. The clinical practice of IMPT further improves the healthy tissue vs target dose differential in comparison with X-rays and thus allows increased target dose with dose reduction elsewhere. In addition, heavy-charged-particle beams allow for the modulation of biological effects, which is of active interest in combination with dose “painting” within a target. The clinical utilization of IMPT is actively pursued but technical, physical and clinical questions remain. Technical questions pertain to control processes for manipulating pencil beams from the creation of the proton beam to delivery within the patient within the accuracy requirement. Physical questions pertain to the interplay between the proton penetration and variations between planned and actual patient anatomical representation and the intrinsic uncertainty in tissue stopping powers (the measure of energy loss per unit distance). Clinical questions remain concerning the impact and management of the technical and physical questions within the context of the daily treatment delivery, the clinical benefit of IMPT and the biological response differential compared with X-rays against which clinical benefit will be judged. It is expected that IMPT will replace other modes of proton field delivery. Proton radiotherapy, since its first practice 50 years ago, always required the highest level of accuracy and pioneered volumetric treatment planning and imaging at a level of quality now standard in X-ray therapy. IMPT requires not only the highest precision tools but also the highest level of system integration of the services required to deliver high-precision radiotherapy.

Abstract

The practice of proton radiotherapy covers 50 years since the first proton patient at the Berkeley Lawrence Livermore Laboratory (Berkeley, CA). In that period, a few post-research proton accelerators have been transformed into semi-clinical facilities and commenced treatments. One such facility at the Harvard Cyclotron Laboratory (Cambridge, MA) had a 160 MeV accelerator well suited for the treatment of cranial neoplasms1 in parallel with similar practice in Sweden,2 eyes3 and large field treatments.4 These sites were managed in three semi-independent clinical programmes that persist today at the F H Burr Proton Therapy Center at the Massachusetts General Hospital in Boston.

The large field programme required the development of proton field scattering and energy modulation techniques to achieve uniform fields and spread-out Bragg peak modulated (SOBP) fields of constant penetration range and modulation. The large field programme was only possible after the introduction of CT to model these fields, with apertures and range compensators to control the lateral extent and penetration around the three dimensional (3D) target volume extent as identified on CT.5,6 The fields were created by mechanical means, which allowed their early clinical use in the absence of electronic controls.

The practice of SOBP proton radiotherapy required all the quality management features of modern radiotherapy: volumetric treatment planning, accurate immobilization and verification and on-treatment imaging. The practice of SOBP proton radiotherapy established the axiom of radiotherapy: accuracy improves healthy tissue dose avoidance and target coverage and higher target dose achieves cure. The promise and realization of cure was demonstrated in patients with otherwise incurable chordoma.7,8 The practice of SOBP proton radiotherapy persists today, and most patients are still treated with SOBP fields.

The primary proton beam out of an accelerator is, in the absence of scattering materials, a collimated well-circumscribed “pencil” beam and easily manipulated by electromagnetic means. The proton pencil beam allows dose modulation in the patient with four degrees of freedom: number of protons (NP) to control the local dose deposition, energy to control the local penetration and magnetic deflection to control the off-axis position. The size of the pencil beam is a fifth degree of freedom although not readily available. Spot size control would positively impact delivery efficiency, as “larger” spots can deliver more protons in vivo given safety constraints (see section on back-of-the-envelope calculations), albeit possibly with an increase of integral dose. The spot size is typically characterized by the gaussian width σ of the pencil beam lateral intensity distribution and quantified in air at the isocentre.

Proton pencil beams thus have one (or two) more degrees of freedom, penetration dose modulation, compared with intensity modulated radiotherapy [IMRT or volumetric arc therapy (VMAT)] fields. Proton fields (at dose equilibrium) exhibit the charged-particle Bragg peak depth dose characterized by a sharp dose increase, the “spot” at the energy characteristic penetration range and absence of dose beyond this distal range. The full electromagnetic control of the heavy-charged-particle pencil beam combined with the Bragg peak and absence of distal dose makes pencil beam scanning (PBS) an easier and more powerful delivery system for modulated therapy compared with the mechanical multileaf collimators (MLCs) required in X-ray IMRT (or VMATs), as well as the creation of SOBP fields.

We use the label “pencil beam (spot) scanning” (where “spot” refers to the location of the Bragg peak in the patient) for the technical mode of delivery and the label “intensity modulated proton therapy (IMPT)” for the clinical mode of PBS where each individual field is allowed to assume an arbitrary dose distribution, and only the full set of fields in the treatment fraction, as in IMRT, assumes the desired dose fraction distribution. Other clinical modes exist, but IMPT is simply the desired, although presumably the most challenging, goal of PBS and our focus here.

Clinical PBS was systematically developed and applied at the Paul Scherrer Institute in Villigen Switzerland.9,10 Their original clinical system consisted of a very compact isocentric gantry combined with a couch and a scanning system that scanned a single line of pencil beams (i.e. irradiating planes in the patient) and thus required patient movement to accommodate multiple planes. The gantry transported protons at a fixed set of constant energies, whose energy at the patient was modulated by a set of mechanical degraders. The system implemented full modulation of all pencil beam parameters, albeit by considerable mechanical means. This unique design demonstrates, amongst other things, the possible variability of delivery systems, although all modern systems employ near-complete electromagnetic modes to implement scanning. Nevertheless, modern system designs and choice of components will influence the technical and clinical quality of scanning.

As stated, technical, physical and clinical challenges remain for the effective clinical deployment of IMPT. A pre-IMPT point–counterpoint argued that while IMPT may in-silico outperform IMRT, its expense and complexity exceeds that of IMRT and that of SOBP treatments.11 A rebuttal12 argued that IMPT will become generally available and its use necessary to fully exploit the dosimetric advantages of proton radiotherapy. Indeed, IMPT (or more precisely PBS) delivery technology is now standard and is, in fact, more cost-effective and simpler in terms of commissioning13 and operation compared with other delivery modes of proton radiotherapy. Overall costs, depending on accounting, are generally assumed to be twice that of IMRT and remain an issue.

The sections below elaborate on these individual issues. We argue that clinical IMPT requires a system approach whereby the current (i.e. in X-ray radiotherapy) individuality of treatment management components must be integrated to achieve optimal performance. Optimal performance combined with exploitation of dosimetric advantages, in turn, can lead to an improved cost profile. The hypothesis is if IMRT is cost-effective in some end point (see, for example, Kohler et al14), then IMPT can exceed this cost-effectiveness criterion through additional dose advantages or through increased performance such as may be achieved through hypofractionation.

CTV, clinical target volume.

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