Pre-report on the dissertation Development of a treatment verification system for continuous scanning in proton therapy

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1 Research Collection Report Pre-report on the dissertation Development of a treatment verification system for continuous scanning in proton therapy Author(s): Klimpki, Grischa Publication Date: 2015 Permanent Link: Rights / License: In Copyright - Non-Commercial Use Permitted This page was generated automatically upon download from the ETH Zurich Research Collection. For more information please consult the Terms of use. ETH Library

2 Swiss Federal Institute of Technology Department of Physics PRE-REPORT on the Dissertation carried out by Grischa M. Klimpki at the Paul Scherrer Institute (Center for Proton Therapy) under the supervision of Dr. David Meer and Prof. Dr. Antony Lomax February 2015

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4 Development of a treatment verification system for continuous scanning in proton therapy The treatment of malignant tumors with scanned proton beams has gained increasing interest over the past years. Patients with deep-seated, inoperable tumors benefit from high dose conformity and enhanced sparing of surrounding healthy tissue. Large dose gradients can be planned with millimeter precision because of the finite proton range. This enables treating tumors in close proximity to critical structures under the condition of a well-known, static patient anatomy. But imaging and positioning uncertainties as well as organ motion largely hamper irradiation precision. In fact, large dose gradients could lead to undesired underdosages in the tumor volume or harmful overdosages to organs at risk induced by cardiac or respiratory motion. A promising strategy to tackle this issue is rescanning (possibly combined with irradiation gating). However, a significant reduction of the interplay between target motion and dose delivery calls for fast scanning techniques in order to apply as many rescans as possible without increasing treatment time. PSI Gantry 2 was designed to meet the technical requirements for such an approach. The delivery time of a single field could drastically be reduced by introducing continuous scanning along lines or meanders, which avoids time-consuming spot breaks. This opens up the possibility for a clinical integration of rescanning. However, substantial enhancements to the spot scanning verification system are required to monitor continuously scanned proton beams. Their analysis, development and validation is the objective of this dissertation. Safety requirements will be analyzed in a first, conceptual phase of this study. The realization phase will focus on the practical implementation of dynamic dose and position verification. The final validation phase shall then test accuracy and reproducibility of the adapted verification system and summarize the results in an addendum to the Gantry 2 safety report.

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6 CONTENTS 1 INTRODUCTION 1 2 FUNDAMENTAL QUANTITIES Heavy Charged Particle Interactions in Matter Ionizing radiation Linear stopping power Strong nuclear interactions Particle Fluence Linear Energy Transfer Absorbed Dose Physical dose Biological dose Depth-dose distribution BEAM DELIVERY AND DIAGNOSTICS Beam Delivery Spot scanning Raster scanning Line scanning Beam Diagnostics Intensity measurements Position measurements PROSCAN FACILITY Beamline Beam generation Energy selection Beam transportation Gantry Layout Beamline Nozzle Therapy Control System

7 5 RESEARCH PLAN Project Motivation and Context Discussion of Safety Goals Dose deposition Scan trajectory Measurement Devices Integrating monitors Dynamic Monitors Research Phases Conceptual phase Realization phase Validation phase APPENDICES A Lists iii A.1 List of abbreviations iii A.2 List of figures iv A.3 List of tables iv B Acknowledgments v REFERENCES C Bibliography ix

8 1 INTRODUCTION In 1946, Wilson outlined the radiological use of fast protons and heavier ions based on their physical properties. He claimed: "Higher-energy machines are now under construction [... ] and the ions from them will in general be energetic enough to have a range in tissue comparable to body dimensions. It must have occurred to many people that the particles themselves now become of considerable therapeutic interest." Robert R. Wilson (1946) [66] Only eight years later, in 1954, the first proton treatment was performed at the Lawrence Berkeley Laboratory in California with high energy proton beams that penetrated entirely through the patient s head [64], not yet exploiting the favorable depth-dose profile outlined by Wilson. Patient numbers increased continuously with the evolution of proton therapy. By the end of 2013, over 100, 000 patients in total received proton therapy treatment around the world [44]. protons and their range in matter Protons traverse tissue on nearly straight lines and lose most of their energy in the last centimeters of their path before being brought to rest [21]. Their range depends on the initial energy and the material composition along their track of ionization. By imaging the patient anatomy and selecting the corresponding energy, it is, therefore, possible to point a proton beam directly towards target structures while sparing healthy tissue from undesired radiation exposure. For this reason, optimized proton therapy treatment plans show very low energy deposition (or dose) proximal to the tumor and almost no dose a few centimeters distal to the tumor. This yields potential dosimetric advantages such as reduced integral dose and pronounced sparing of organs at risk (OARs) located in tumor proximity over conventional radiotherapy with high-energetic photons [14]. However, full exploitation of these benefits requires precise knowledge on the patient anatomy at all times during treatment. In fact, an unexpected shift of the sharp distal dose falloff (or penumbra) induced by errors in tissue composition characterization along the beam path might cause significant underdosage of the tumor [40]. The well-defined proton range must, therefore, be regarded as both benefit and risk. 1

9 1 INTRODUCTION benefits and drawbacks of actively scanned pencil beams The shift from passively scattered [30] to actively scanned beams [51] yielded further dose conformity in the tumor volume, opened up the possibility of dose modulation within a single field [32] and removed the need for patient-specific hardware (e.g. collimators and compensators) in the beamline. The latter has a large impact on the neutron background dose, which can be reduced significantly outside of the primary radiation field using active pencil beam scanning (PBS) [24]. But PBS comes at the cost of two main drawbacks: (1) The pencil beam width is energy-dependent and increasing with decreasing penetration depth [49]. Thus, irradiation fields of shallow targets exhibit smeared out lateral penumbras [53]. Suit et al. (2010) even conclude that it is generally difficult to achieve clearly advantageous lateral dose profiles with PBS compared to intensity modulated photon therapy [63]. (2) Furthermore, PBS shows increased sensitivity to (periodic) intrafractional anatomical variations, commonly referred to as organ motion during treatment. Engelsman et al. (2013), therefore, conclude that PBS "is not yet a mature therapy when it comes to dealing with treatment-related uncertainties in patients" [14], referring to the following sources of error listed by increasing residual impact: beam energy uncertainties: Dedicated accelerators and beams optics minimize this source of error to a very large extent. Range fluctuations originating from the energy spread of the delivered pencil beam are generally much smaller than inevitable range straggling in the patient. positioning uncertainties: A false patient position would cause false irradiation of entire regions containing healthy tissue or OARs. Due to this high severity, verification images are acquired prior to each treatment and mapped onto the planning reference in order to correct for positioning offsets. Additional immobilization devices or anesthesia (mostly for pediatric treatment) ensure patient stability throughout the fraction. imaging uncertainties: Inherent CT uncertainties (e.g. Hounsfield unit conversion and beam hardening) and CT artifacts challenge both tumor identification and tissue characterization. Both errors translate into large range uncertainties, which remain significant even with most advanced imaging hardware. anatomy variations: Interfractional changes in the patient anatomy such as tumor shrinkage or weight loss could be tackled by daily adaptive proton therapy [33]. The concept behind this strategy is to image, plan and deliver treatment fields on a daily basis. However, computational and workflow-related challenges hamper this approach. Intrafractional anatomical variations are even more difficult to address. Most prominent effects of organ motion during treatment are smoothing of dose gradients (dose blurring) [8], fluctuations in proton 2

10 beam range (e.g. due to the diaphragm moving in and out of the beam path) [6] and interferences between the PBS time structure and the target motion (interplay effect) [50]. The list of suggestions to these issues is long and briefly summarized in the following section. organ motion mitigation techniques Cardiac and respiratory motion contribute most to intrafractional anatomical variations with amplitudes up to several cm. (Periodically) changing offsets from the static irradiation plan can cause significant underdosages in the tumor and overdosages in neighboring OARs. In order to maximize dose conformity and minimize undesired radiation exposure, the following two strategies are being pursued and applied clinically: motion minimization: The most universal approach to large-amplitude motion mitigation is to freeze the patient anatomy during irradiation. This goal can, to some extend, be achieved by controlling the respiratory motion. It is either suppressed for a short time of beam delivery (breath-hold technique) or constantly monitored with beam delivery triggers at certain phases of the breathing cycle (gating technique [39]). Irregularities call for gating systems that operate online and in real-time. Furthermore, the accelerator and energyselection system must permit multiple beam on/off cycles. A major disadvantage of motion minimization by respiration control is the increased treatment time due to frequent beam interruptions. But residual motion is small and mitigable by beam-specific margins. safety margins: Motion effects induced by smaller amplitudes do not necessarily require minimization strategies and can be mitigated merely by adding a (beam-specific) safety margin around the treatment volume [29]. The enlarged target assures tumor coverage even under considerable anatomical variations, but yields suboptimal sparing of surrounding healthy tissue or OARs. Safety margins increase robustness of single field uniform dose (SFUD) plans with low in-field gradients but provide only marginal improvement concerning homogeneity and coverage for highly intensity-modulated treatment plans [1]. Furthermore, margins do not address the interplay between PBS time structure and target motion. Even with motion minimization and safety margins, intrafractional anatomical variations remain as major obstacles in particle therapy. They hamper accurate delivery of conformal dose distributions. Additional motion mitigation techniques such as tracking and rescanning are, therefore, being investigated extensively. The former would nominally be most precise, since it theoretically mitigates the entire spectrum of motion-related uncertainties, but the latter is by far easier to realize technically. Hence, clinical integration of rescanning (possibly combined with motion minimization techniques) can be expected in the near future. The underlying principles of both mitigation approaches mentioned are: 3

11 1 INTRODUCTION tumor tracking: Continuous real-time imaging of the patient anatomy would in principle allow for an online treatment plan adaptation to the current geometry. Recalculating treatment plans according to the tracked tumor position would maximize dose delivery efficiency and minimize dose outside of the target volume. But since the technical requirements on the planning and control system to realize this nominally most precise motion mitigation method are substantial, clinical implementation is not yet feasible [6]. rescanning: A promising approach to mitigate dose deformation due to small intrafractional anatomical variations and interplay effects is to deliver the planned irradiation field multiple times with a proportionally reduced dose. This idea of rescanning or repainting was already proposed by Phillips et al. in 1992 by showing that dose uniformity roughly increases with the square root of the number of rescans [50]. Recent research results could confirm the effectiveness of rescanning in terms of motion mitigation [56], thus a more detailed description on possible practical implementations is provided in the following section. motion mitigation by rescanning In PBS, uniform tumor coverage is achieved by superimposing dose profiles of single pencil beams [17]. The treatment planning algorithm optimizes the respective beam spot weights to achieve a final dose distribution of high lateral and distal homogeneity. Intrafractional motion causes range shifts of single pencil beams according to the changes in tissue composition along their paths. The resulting superposition is no longer homogeneous but exhibits hot and cold spots of over- and underdosage [50]. And since hot spots in healthy tissue could increase the risk of secondary cancer while cold spots inside the target volume weaken local tumor control, this interplay effect between the time structure of PBS delivery and target motion is of major concern. The interplay pattern in the resulting dose distribution strongly depends on motion period, phase and amplitude as well as on scanning speed and direction [7]. Multiple scans will, therefore, lead to an averaging effect of the interference patterns if motion parameters are distributed statistically over the set of rescans. The two core techniques are called layered (rescanning energy slices) and volumetric repainting (rescanning the entire volume). The latter requires very fast energy switches as the whole sequence is repeated multiple times. A study by Seco et al. (2009) on layered spot repainting revealed that using 5 10 rescans per field for an assumed tumor volume of cm 3 is sufficient to keep dose delivery errors below 5% if the breathing amplitude is less than 10 mm [60]. The smaller the tumor volume, the more rescans will be needed to average over interplay effects. For larger breathing amplitudes, rescanning should be combined with gating or breath-hold to reduce necessary safety margins [58]. Here, it is beneficial to distribute the number of rescans evenly over the irradiation window for optimal field homogeneity [60]. 4

12 Further proposed enhancements of the rescanning workflow are broader Gaussian shaped pencil beams [22] that provide greater uniformity at the expense of a decreased lateral penumbra or scaled repaintings accounting for the fact that layers of high energy deliver the major part of the target dose while layers of low energy may not require any rescans at all [60]. But despite all potential benefits, repainting increases treatment time significantly in the case of discrete spot scanning because of duplicated dead times. Clinical integration of this approach, thus, requires very fast continuous scanning techniques such as raster [23] or line scanning [54] as well as rapid energy switches in order to ensure acceptable delivery times. This would not only open up the possibility to treat moving tumors with rescanning but, of course, also help to reduce the overall treatment time for static tumor irradiations. continuous scanning at PSI Gantry 2 The Gantry 2 scanning system of the Paul Scherrer Institute (PSI) was designed to meet the technical requirements of fast scanning. The pencil beam can be swept continuously over the scan field with velocities of up to 5 mm /ms in U and 2 cm /ms in T direction 1 [49]. A vertical deflector plate installed in the proton accelerator (COMET cylotron [59]) near the ion source adds a second degree of freedom in dose modulation besides changing the scan speed. Combination of scan speed and beam intensity modulation principally enables delivery of whole fluence-shaped energy layers in hundreds of ms [54]. Additional fast energy switches (80 ms for 5 mm distal range shifts in water [54]) triggered by multiple carbon wedges in the degraded unit decrease overall treatment time even further and potentially allow for both layered and volumetric rescanning. The PSI Gantry 2 scanning system, therefore, fulfills the technical requirements for a successful clinical integration of rescanning. Recent studies confirmed the feasibility to treat moderately moving tumors with repainting at Gantry 2 [67] and showed only a modest degradation of the lateral penumbra in scanning direction when replacing discrete spot by continuous line scanning [57]. Thus, delivered dose distributions are expected to be comparable. However, a clinical integration of continuous scanning still requires substantial enhancements of the therapy verification system (TVS) to cope with continuous beam dynamics (preferably in real-time). It is of fundamental importance concerning patient safety to immediately pause or abort treatment if the applied spot dose or position deviates from the planned reference. The current TVS exploits dead times between spots to perform verification calculations such as spot position reconstruction. But since the advantageous feature of continuous scanning is the removal of spot breaks, dynamic verification must be enabled and implemented. This is the objective of this dissertation. For this 1 T and U refer to the lateral coordinates in beam direction. 5

13 1 INTRODUCTION purpose, careful analysis of safety requirements will be followed by experiments on time-resolved monitoring of intensity and position dynamics (conceptual phase). It will be the primary aim to modify devices already in operation to meet the verification demands, but installation of system enhancements including firm- and software integration will also be considered (realization phase). Final reproducibility and accuracy tests will then be used to validate the verification system (validation phase). pre-report structure This document is intended to be a literature review on the physical and technological aspects of proton therapy in order to describe the field of research of this dissertation. Quantities and physical descriptions required for the fundamental understanding of proton therapy are outlined in chapter 2. An overview on beam scanning techniques and their online verification is given in chapter 3. The PROSCAN facility at PSI is mapped out in chapter 4. Chapter 5 provides a three-year timeline structuring the different research phases in the development of a treatment verification system for continuous line scanning. 6

14 2 FUNDAMENTAL QUANTITIES This chapter summarizes definitions and equations required for the fundamental understanding of proton therapy. The clinical rationale is based on the characteristic stopping of protons in matter. Thus, different forms of energy loss are being summarized in section 2.1. The particle fluence Φ is introduced in section 2.2. Additional knowledge on the linear energy transfer (LET) (section 2.3) allows for energy deposition calculation. The mean energy imparted by ionizing radiation to matter is generally quantified as absorbed dose. Its characteristic distribution in depth is outlined and explained in section Heavy Charged Particle Interactions in Matter Ionizing radiation The International Commission on Radiation Units and Measurements (ICRU) defines ionization as the liberation of one or more electrons in collision of particles with matter [26]. This process can be triggered by charged (e.g. light electrons or heavy ions) and uncharged particles (e.g. photons or neutrons). They either produce ionization in a medium themselves or initiate transformations that then cause ionization or produce ionizing radiation. Incident heavy charged particles (HCPs) transfer a small fraction of their kinetic energy in each electron collision. As they gradually slow down, ionization probability decreases. Excitation (transfer of electrons to higher energy levels) and elastic scattering become the primary process of energy dissipation. Originally ionizing HCPs can, thus, be considered non-ionizing near the end of their range Linear stopping power Therapeutic proton energies E are much larger than those of thermal agitation (E th < 1 ev). The stopping of protons in matter results from collisions with atomic electrons and target nuclei. Hence, it can be described based on the laws of electromagnetic and strong interaction. The quantity of interest is the linear stopping power S, which considers electromagnetic interactions only. It is defined as the mean energy lost by the charged particle de in traversing a distance dx. According to the ICRU [26], S can be written as a sum of three independent components: S = S el + S rad + S nuc = ( ) de + dx el ( ) de + dx rad ( ) de dx nuc (2.1) 7

15 2 FUNDAMENTAL QUANTITIES S el denotes the linear electronic stopping power due to inelastic interactions with atomic electrons. S rad, the linear radiative stopping power, results from emission of bremsstrahlung in the electric fields of atomic nuclei or electrons. The linear nuclear stopping power S nuc originates from elastic Coulomb interactions, in which recoil energy is imparted to atoms. It does not refer to strong nuclear interactions generally being discussed outside of the context of linear stopping power. The impact parameter b, defined as the perpendicular distance between the projectile s trajectory and a scattering center, is commonly used to differentiate between different forms of electronic and radiative stopping. Their major characteristics are outlined in the following. Additionally, a brief description of nuclear stopping is provided at the end of this section. inelastic collisions with electrons Inelastic collisions of incident HCPs with atomic electrons can be distinguished in two different types according to the impact parameter b: soft collisions: For impact parameters much greater than the atomic radius (b >> r atom ), the particle s Coulomb force field distorts the electron cloud of target atoms. As a result, electrons can be excited into higher states or (with less probability) ejected if they occupy outer shells. Furthermore, polarization can occur in liquids and solids. The net kinetic energy transferred is very small ( 1 ev). But because large impact parameters hold for most interactions, soft collisions account for approximately 50% [3] of the total energy loss and give rise to a continuous slowing down of traversing ions. hard collisions: For impact parameters in the order of the atomic dimension (b r atom ), HCPs will primarily interact with single bound atomic electrons. A substantial amount of kinetic energy can be transfered in knock-on collisions. The ejection of inner-shell electrons ionizes and excites target atoms simultaneously. Characteristic X-rays and/or Auger electrons are emitted. Although hard collisions have a much smaller probability compared to soft collisions, they account for a comparable total energy loss [3]. Both kinds of inelastic collisions with electrons can be described analytically by the Bethe-Bloch equation [25]. The linear electronic stopping power depends on the electron density n e and the mean excitation energy I of the target material as well as the projectile charge Z and its relativistic velocity β: ( ) de = K n ez 2 L(β, I) dx el β2 (2.2) with K = 51 MeV fm 2 8

16 2.1 Heavy Charged Particle Interactions in Matter The leading term L 0 (β, I) of the stopping number L(β, I) causes the relativistic rise of S el for high projectile velocities β, since ( ) L 0 (β, I) 2me c 2 β 2 = ln β 2 C I(1 β 2 ) Z δ 2. (2.3) The shell correction C/Z has a significant contribution to L 0 (β, I), when the particle velocity is of the order of the bound electron velocities. The assumption of atomic electrons at rest breaks down and capture processes have an increased probability. The density-effect correction δ/2 considers the density-dependent polarization of the target. At high velocities, polarization shields the electrical field far from the particle path cutting off its long-range contribution and, hence, decreasing S el. Further correction factors can be added to L 0 (β, I), e.g. Barkas ZL 1 (β) and Bloch correction Z 2 L 2 (β) [25]. The Bethe-Bloch equation (2.2) states that the linear electronic stopping power increases with decreasing velocity accelerating the stopping of ions in matter. HCPs, therefore, show a drastic decline in velocity as well as a drastic rise in S el at the very end of their track. This effect is compensated around β 0 by charge screening. Z will drop down to zero just before HCPs come to rest. Thus, the energy loss of a single HCP is sharply peaked near the stopping point. An ion beam consisting of many particles shows a slightly different behavior. Statistical fluctuations of the energy loss in the large number of collisions yields a continuous spectrum of stopping points, which are responsible for the broadening of the characteristic Bragg peak. elastic collisions with electrons In elastic collisions with atomic electrons, energy and momentum are conserved. Thus, the energy transfer is typically less than the lowest excitation potential of bound electrons. Such collisions only occur for low-energy incident electrons (E < 100 ev) with large impact parameters (b >> r atom ) [15]. Therefore, their contribution to the stopping of HCPs in matter can be neglected. inelastic collisions with nuclei For very small impact parameters (b << r atom ), incident particles can encounter electromagnetic interactions with a target nucleus. If they are inelastic, charged particles are most likely deflected in the Coulomb field of the nucleus under emittance of bremsstrahlung. The probability of nuclear excitation is generally much smaller. The corresponding linear stopping power S rad depends on the following projectile parameters: ( ) de Z2 E (2.4) dx rad m 2 9

17 2 FUNDAMENTAL QUANTITIES Since the radiative contribution to the total linear stopping power decreases quadratically with increasing mass m, it is negligible for protons and heavier ions at clinical energies. This kind of interaction is most important for comparably light incident electrons. elastic collisions with nuclei Elastic collisions with atomic nuclei, in which incident particles are deflected without emitting radiation or exciting the nucleus, have a minor contribution to S when considering HCPs. The recoil energy imparted to atoms conserves momentum and yields projectile deflection. Although the energy transfer S nuc is small, elastic collisions with nuclei are the main source for the lateral heavy ion beam spread [55] Strong nuclear interactions The previous section shows that stopping processes in thick absorbers are governed by inelastic collisions with atomic electrons for high-energy ions. In contrast to electromagnetic interactions, strong nuclear interactions can result in a complete disintegration of both projectile and target nucleus [55]. For impact parameters b in the order of the nuclear radius r nuc, peripheral collisions are most probable. They yield partial fragmentation commonly described by the abrasion-ablation model [20]. In proton therapy, strong nuclear interactions lead to a significant neutron background dose. 2.2 Particle Fluence A fundamental scalar quantity of a radiation field is its particle fluence Φ. ICRU advises the following definition: The Φ A.= dn da, (2.5) where dn denotes the total number of particles incident on a sphere of cross-sectional area da. But since every radiation field has a finite particle density, the infinitesimal definition suffers from statistical fluctuations. Thus, particle fluence must be considered a macroscopic concept with lower da limit defined by the graininess of the field itself [9]. Figure 2.1 illustrates the formal ICRU definition. Papiez and Battista (1994) proposed a generalized volumetric definition, in which the sum of track length segments dl i contained within any sampling volume dv can be used to calculate Φ [42]: Φ V.= N i=1 dl i dv, (2.6) 10

18 2.3 Linear Energy Transfer They proved equivalence when considering straight-line trajectories and, in this case, advise against the use of the ICRU definition [43]. sphere of crosssectional area da P particle track Figure 2.1: The ICRU formally defines the particle fluence Φ at a point P on an elementary sphere around P with cross-sectional area da intersected by dn particles. 2.3 Linear Energy Transfer The ionization density along a particle s trajectory influences its radiobiological impact in hadron therapy [9]. Large amounts of energetic secondary electrons that carry away kinetic energy from the track core diminish the local ionization density. The restricted linear energy transfer L δ a predictor of radiation quality is, therefore, defined as the difference of linear electronic stopping power S el and the mean sum of the kinetic energies Ekin δ in excess of δ of all the electrons released by the charged particle [26]: L δ.= S el deδ kin dx. (2.7) Thus, the restricted LET can be interpreted as that fraction of S el, which includes all soft collisions and those hard collisions resulting in secondary electrons with energies less than δ. The unrestricted LET, denoted as L or simply L, is equal to the linear electronic stopping power S el (cf. equation 2.2). 11

19 2 FUNDAMENTAL QUANTITIES 2.4 Absorbed Dose Physical dose The absorbed dose D phys is the most important physical quantity in radiotherapy [55]. It considers all energy deposits ε i resulting from N single interactions of ionizing radiation in a given volume V of homogeneous density ρ. In order to compensate for statistical fluctuations in energy deposition, the ICRU defined absorbed dose as the mean energy imparted per mass element dm = ρdv : D phys.= 1 N N dε i dm = 1 1 ρ N i=1 N i=1 de i dx dl i dv (2.8) For monoenergetic and unidirectional heavy charged particle fields incident on a thin slice of absorber material, this expression simplifies to D phys = 1 ρ S Φ = 1 ρ L Φ. (2.9) Both equations show a strong material dependency. In radiation therapy, water is used as a tissue-equivalent reference medium. Dose measurements performed with air-filled ionization chambers have to be converted using the corresponding stopping power ratios Biological dose In radiotherapy with X-rays, the absorbed dose D phys serves as an estimator of clinical outcome. This attribution is hampered in hadron therapy due to a higher ionization density in the track core. Irradiation plans show enhanced LET distributions in the tumor volume which can lead to an increased biological effect. In order to transfer knowledge gathered in conventional radiotherapy to hadron therapy, the biological dose D biol is introduced. It can be calculated by applying a quality factor, commonly referred to as relative biological effectiveness (RBE), to D phys : D biol = RBE D phys (2.10) The RBE is formally defined as the ratio of 60 Co dose D γ to ion dose D ion required to achieve the same biological effect ζ 0 (e.g. cell survival or tumor control). RBE.= D γ (2.11) ζ=ζ0 D ion In proton therapy, the RBE can generally be approximated by a constant factor of 1.15 in the central irradiated tumor region neglecting dependencies on LET, tissue composition and clinical endpoint [41]. 12

20 2.4 Absorbed Dose Depth-dose distribution single Bragg peak The combination of numerous soft and hard electron collisions (mainly responsible for the broadened stopping point spectrum), elastic collisions with target nuclei (mainly responsible for lateral scattering) and strong nuclear interactions between primary ions and traversed matter (responsible for fragmentation and production of secondary particles) determines the characteristic shape of the depth-dose distribution. Figure 2.2 displays such Bragg peaks for monoenergetic proton beams incident on a large water tank E = 79 MeV E = 117 MeV E = 147 MeV E = 173 MeV E = 197 MeV absorbed dose [mgy] depth in water [cm] Figure 2.2: Single Bragg peaks for 79, 117, 147, 173 and 197 MeV proton beams with a fluence of Φ = /cm 2. The maximum absorbed dose decreases with increasing beam energy and the peak broadens as range straggling increases. The overall shape of the curves is governed by the increase of linear stopping power S as protons slow down continuously. Nuclear reactions along the beam path reduce the peak height and contribute dose to the plateau region. Due to fluctuations in the number of electron collisions, not all protons of the quasi-monoenergetic beam will stop at the same point. This range straggling broadens the peak. The effect of lateral scattering cannot be seen, since the Bragg peak is generally understood as the integral dose distribution projected onto the geometrical depth in the target [21]. range definition Not all protons of the same energy have the same range due to range straggling. The ideal range definition for a monoenergetic proton beam would be the position, 13

21 2 FUNDAMENTAL QUANTITIES where the integral depth-dose has decreased to 80% of the maximum dose. This point in the distal falloff coincides with the range at which half of the protons in the beam have stopped and is, thus, independent of the beam s energy spread [40]. spread-out Bragg peak The objective of particle therapy is to deliver the same amount of biological dose to every tissue element in the target volume [34]. This should trigger the same level of cell response and yield uniform local tumor control. In order to reach this aim with active beam scanning, single pencil beams with approximately Gaussian shaped, energy-dependent cross sections are pointed towards critical structures. Homogeneity in depth can be achieved by superimposing many single Bragg peaks with proper intensity and energy to form a spread out Bragg peak (SOBP). Optimization algorithms in the treatment planning software ensure high dose conformity across the target volume and a sharp distal falloff. Figure 2.3 displays a proton SOBP to cover a target volume ranging from 15 to 20 cm depth absorbed dose [Gy] depth in water [cm] Figure 2.3: Superposition of 24 single Bragg peaks (gray curves) with optimized beam weights to achieve a uniform dose distribution between 15 and 20 cm water depth (blue curve). The sharp distal falloff can be preserved by weighing the highest energy the strongest. 14

22 3 BEAM DELIVERY AND DIAGNOSTICS In order to cover the clinical target volume homogeneously with dose, the accelerated particle beam has to be spread out in three dimension. In active proton therapy, this includes energy and intensity modulation to achieve a uniform distal dose distribution as well as two-dimensional beam scanning to preserve lateral homogeneity. The clinical outcome depends on the accuracy of beam delivery. Thus, diagnostic devices such as ionization chambers and profile grids are used to constantly monitor the beam during treatment. Interlocks are being raised when certain beam parameters exceed or fall below their safety margins and the irradiation is paused or aborted. This chapter describes common active beam scanning techniques (section 3.1) and provides an overview on non-destructive beam diagnostics (section 3.2). 3.1 Beam Delivery Spot scanning Beam scanning can generally be understood as the process of moving the beam in both lateral dimensions across iso-energy slices [17]. In discrete spot scanning, introduced by Pedroni et al. (1995) [46], a rectangular grid of spot positions is defined. The number of applied protons per spot depends on the target geometry and will be optimized by the treatment planning system (TPS). There are principally two different approaches to deliver the planned proton distribution. Time-driven spot scanning requires a stable beam current, since the spot dose is only determined by the time τ i the beam spends at each pre-defined position (T i, U i ). Intensity-driven spot scanning allows for active beam current variations from spot to spot. In both cases, the beam is shut off (or rather deflected by a fast kicker magnet into a beam dump) when the total number of prescribed protons is reached. Currents in the sweeper magnets are changed according to the next spot position and beam is applied again after the magnetic field stabilized. The TPS optimization algorithm relies on accurate pencil beam models to minimize errors in the dose calculation. Characterizations for the PSI scanning systems Gantry 1 and Gantry 2 can be found under [48] and [49], respectively. Both publications show that the lateral beam shape can be described by a symmetric, 2D Gaussian distribution in first order approximation. To account for primary particles with large scattering angles and secondary particles resulting from nuclear interactions, higher order corrections need to be considered. Figure 3.1 displays 1D proton distributions along the central T -axis of 17 superimposed Gaussian beams. The 15

23 3 BEAM DELIVERY AND DIAGNOSTICS lateral penumbra, defined as the distance between the 80% and 20% dose falloff, can be decreased by increasing the outer beam weights and pulling in spots towards the center (edge enhancement). The resulting proton density is in both cases given by N p (T ) = n I i τ i B i (T, T i ), (3.1) i=1 where n denotes the total number of spots, I i the beam current (in protons per second), τ i the spot dwell time and B i (T, T i ) the actual beam shape in T. In this particular example, B i was approximated by a single Gaussian distribution: B i (T, T i ) = 1 σ 2π exp [ (T T i) 2 2σ 2 ]. (3.2) relative proton density [%] relative proton density [%] T position [cm] T position [cm] Figure 3.1: Lateral proton distributions for uniform (left) and modulated superposition (right) of 17 Gaussian-shaped beams with a width of σ = 5 mm and a spot spacing of λ = 7 mm. By emphasizing the outer beam weights and pulling low-weighted spots towards the center, the lateral penumbra can be reduced from 1.56 σ (left) to 1.21 σ (right). Dose homogeneity is slightly impaired. The discretization of irradiation in spot scanning results in an accumulation of dead times. Assuming a cubic target volume of cm 3, λ = 5 mm lateral spot spacing and 21 iso-energy slices, the optimized treatment plan would contain approximately 10,000 spots. Thus, a dead time of 3 ms in between spots delays the application of the entire field by roughly half a minute (not considering beam-off time during energy switches). On the other hand, dead times simplify steering and safety system analysis as well as dose calculation algorithms in the TPS. Spot breaks can, for example, be used to reconstruct and verify the actual beam position of the previous spot. 16

24 3.1 Beam Delivery Raster scanning The raster scanning technique was originally designed to move the particle beam continuously over each iso-energy slice without turning it off. The first technical realization consisted of a pair of dipole magnets, two power supplies, a set of ionization chambers and secondary emission monitors as well as a computer control system [51]. Initial tests with a 456 MeV /u neon ion beam resulted in reasonable dose uniformity when scanning the target area on a zig-zag path (cf. figure 3.2). Figure 3.2: Originally considered raster scan pattern for a rectangular area (dashed zig-zag line) and an irregular tumor slice (solid line). Additional shortcuts allow for beam on over the entire scan path while spearing surrounding healthy tissue. Reprinted form [23]. In 1993, the Gesellschaft für Schwerionenforschung (GSI) reported on therapeutic integration of the raster scanning method [23]. But in contrast to proposed continuous scan lines, their technical realization was based on a large amount of discrete scan steps (much smaller than those of the spot scanning grid). Treatment planning is nearly identical for this hybrid raster technique and spot scanning and results in a list of optimized spot positions and intensities. The desired dose distribution is delivered by varying the scan speed when the particle beam is swept across the raster. Overall treatment time of a single iso-energy layer is, therefore, limited by the beam current and the maximum scan speed and not by the time to switch the beam on and off and reach a stable magnetic field according to the next spot position. 17

25 3 BEAM DELIVERY AND DIAGNOSTICS Line scanning Continuous scanning is an active field of research at PSI [54, 57], since it would allow for fast delivery of single iso-energy layers with a fraction of the prescribed dose demanded by efficient repainting strategies. Line scanning is intended to be continuous in one lateral direction (here T ) and discrete in the other (here U). The PSI Gantry 2 was designed to fulfill the technical requirements on the delivery system. Scan velocities of up to 2 cm /ms [49] combined with reduced spot breaks enable very fast scanning of low-dose layers. Single lines L i (T, T i ) are delivered deterministically according to a list of spot breaks T i that define arrival times t i and beam currents I i. The line shape along the T -axis can be derived from the proton distribution along T for discrete spot scanning (cf. equation 3.1). The superposition of n Gaussian shaped spots yields: with the error function L i (T, T i ) = 1 [ ( ) T Ti erf 2 σ 2 erf(x) = 2 π x 0 ( )] Ti+1 T + erf σ, (3.3) 2 e α2 dα. (3.4) Modulation of arbitrary proton densities along T can be realized by superimposing (b 1) lines according to the list of b break points: b 1 N p (T ) = I i v 1 i L i (T, T i ). (3.5) i=1 The scan speed v i in the interval [T i, T i+1 ) is determined by the difference in arrival times t i+1 and t i : v i = T i+1 T i t i+1 t i (3.6) Equation 3.5 outlines the possibility to modulate the delivered dose distribution by varying two parameters. In speed modulation mode, the beam currents are fixed I 1 = I 2 =... = I b (preferably at the highest value for full beam extraction) and scan speed is variable. Very high doses can be achieved by decreasing the scan speed, but the lowest applicable dose is limited by the maximum scan speed. In intensity modulation mode, I is adjusted according to the prescribed line distribution while scan speed is constant (preferably at its maximum for short irradiation times). Very low doses can be handled in this mode, but the maximal available extracted beam 18

26 3.2 Beam Diagnostics current sets a high dose limit. Therapeutic practice should, therefore, combine both modes to enable a large dynamic dose range and optimal delivery flexibility [54]. Figure 3.3 depicts a relative proton distribution for a line scanned continuously along T using both speed (red) and intensity modulation (green). At the first break point (T = 10 cm), scan speed and beam current are initialized. At the second break point (T = 15 cm), the scan speed is doubled while the beam current remains constant. The resulting proton density drops down to 50%. This decrease is compensated at the third break point (T = 20 cm), where the beam current is doubled without changing the scan speed. The last break point at T = 25 cm defines the final arriving time of the line and sets the beam current back to zero relative proton density [%] v = 1 cm/ms I = 0.5 na v = 2 cm/ms I = 1.0 na T position [cm] Figure 3.3: Relative, lateral proton density (blue) of a continuously scanned line. Speed and intensity modulation are illustrated by the red and green lines, respectively. The proton density is decreased by half after doubling the scan speed at the second break point. This effect can be compensated by also doubling the beam current. The lateral penumbra of a scanned line measures 1.68 σ. It sets the upper limit for uniform spot superpositions with decreasing spot spacings λ 0 mm. 3.2 Beam Diagnostics Continuous beam monitoring and optimization is crucial for safe and successful particle therapy. Especially beam energy, intensity and position must be measured precisely with minimal beam disturbance. Thus, destructive devices such as Faraday cups or thermally isolated calorimeters can only be used for quality assurance (QA) or calibration purposes. Online treatment verification relies or quasi non-destructive devices presented in this section. 19

27 3 BEAM DELIVERY AND DIAGNOSTICS Intensity measurements Beam intensity measurements during patient irradiation monitor the number of delivered protons per spot position and trigger the subsequent action (e.g. beam off in case of discrete spot scanning) once the planned quantity is reached. Table 3.1 lists four different measurement devices and their applicable beam intensity range. beam intensity monitor beam intensity range beam current range beam current transformers > /s > 1.6 µa secondary electron monitors /s 16 pa 16 na ionization chambers /s 16 fa 16 na scintillation counters < /s < 0.16 pa Table 3.1: General comparison of four non-destructive beam intensity monitors and their measurement ranges. Stated values hold for protons with a kinetic energy of 1 GeV and a spot size of 1 cm 2. A beam intensity of protons per second corresponds to a beam current of 1.6 na. Reproduced according to [62]. beam current transformers Beam current transformers (BCTs) measure the magnetic field carried by a charged particle beam. They have the ability to monitor pulsed beams of high intensity nearly independent of their actual position and size [62]. An additional advantage is the direct proportionality between extracted signal and beam current. Thus, an absolute calibration can be performed based on well-known current pulses from an external beam source. Passive BCTs require installation of a torus inside the beam pipe. Traversing beam pulses create a periodically changing magnetic flux Φ B in the torus. These changes induce a current in the windings around the torus according to Faraday s law (third Maxwell equation): t Φ B = E d s (3.7) The output pulse amplitude is proportional to the bunch charge if beam bunches are sufficiently short (< 1 ns) and isolated (±50 ns) [65]. The necessary isolating gap in the beam pipe is commonly created using a ceramic insulator. The 0.8 ns pulse length of the cyclotron beam at PSI principally enables intensity measurements with BCTs. However, pulse spacing (14 ns) and extracted proton intensity (< /s) are both below detection capability. 20

28 3.2 Beam Diagnostics The most important features of a passive transformer are its sensitivity S and the droop time constant τ droop, which corresponds to a decrease in signal amplitude by 1/e. Deformation of the measured signal occurs if τ droop is smaller than the beam pulse length. The disadvantageous combination of S 1/N and τ droop N 2, N denoting the number of secondary windings, calls for technical enhancements in order to use BCTs in beam diagnostics. The current transformer of an active BCT is placed in the feedback loop of an operational amplifier. The droop time constant can be increased significantly by using a feedback resistor and amplification instead of a constant resistor to measure the output voltage [65]. This modification allows for a reduction of secondary windings and, therefore, yields a considerable gain in sensitivity. BCTs can also be used as integrating beam current monitors by including a capacitance and digitalizing the output signal via an analog-to-digital converter (ADC). Structure information on individual beam pulses are lost but continuous beam monitoring is enabled. secondary electron monitors Secondary electron monitors (SEMs) can be used to measure proton beam intensities between 10 8 and /s. They comprise an arrangement of thin metallic foils (typically some micrometers) positioned in the beam path. They are slightly curved in order to increase mechanical strength and reduce acoustic noise. The traversing beam interacts with the peripheral electrons of the foil surface and the free electrons contained within the metal. Due to the small energy transfer, only electrons down to several hundred Å can escape from the surface [62]. An appropriate electric field between the foils assures complete charge collection. Secondary emission efficiencies for 450 MeV proton beams are optimal for (gold-coated) aluminum foils [16]. Beam degradation due to large-angle Rutherford scattering or inelastic nuclear processes is typically small for therapeutic proton beams. ionization chambers Ionization chambers (ICs) cover the medium beam intensity range. They are compact, direct reading devices and, therefore, the dosimeters of choice in medical physics and radiotherapy. The term ionization chamber conventionally refers to gas-filled detection chambers with two electrodes (e.g. in the form of parallel plates or in cylindrical arrangement). Multiple ion pairs are created in the gas under incident ionizing radiation. The detection principle is based on the collection of these charges through the application of an electric field. In contrast to Geiger-Müller tubes, ICs are operated in proportional counting and not gas multiplication mode. The field strength must, nevertheless, be sufficiently high to minimize losses due to recombination. The polarity of the applied voltage determines whether collected ions or electrons define the output signal. 21

29 3 BEAM DELIVERY AND DIAGNOSTICS The plane parallel dose monitors at PSI are operated in ion-collection mode and, therefore, measure the cathode current [31]. The average drift velocity v drift of positive gas ions is proportional to the applied electric field strength E = U/d: v drift = µe. (3.8) The mobility µ remains fairly constant over wide ranges of electric field strength and gas pressure, but decreases with increasing mass of the ion. The ion collection time t coll scales with the square of the distance between cathode and anode plane d: t coll = d v drift = d2 µu. (3.9) A typical field strength of E = 4 kv /cm yields collection times of the order of 100 µs [31]. Dissociated electrons move approximately 1,000 times faster than positive gas ions due to their much lower mass. Thus, measuring the small ionization current at the anode is the common principle of a dc IC. Assuming negligible recombination and spatially homogeneous charge collection, the ionization current I IC is proportional to the total energy loss S el in the detection volume: I IC S el (E) k pt e N ion. (3.10) N ion denotes the number of incident ions, e the elementary charge and k pt the pressure and temperature correction factor: k pt = p p 0 T 0 T, with p 0 = hpa and T 0 = K. (3.11) The proportionality factor of equation 3.10 is generally referred to as chamber gain g IC and must be calculated from calibration measurements with e.g. Faraday cups. scintillation counters Scintillation counters are suitable to measure proton fluences precisely if the counting rate is below /s. The fluorescence process of plastic scintillators is triggered by fluors such as anthracene suspended in a solid polymer matrix (e.g. polyvinyltoluene). Traversing ions raise π-electrons into higher energy levels. Prompt fluorescence is emitted in transitions back form the excited to the ground state. Vibrational modes of each electronic state yield a broadened fluorescence spectrum and a Stokes shift of the peak value towards larger wavelengths. Scintillation light can be guided to charge-coupled device (CCD) cameras or photomultiplier tubes (PMTs). 22

30 3.2 Beam Diagnostics Scintillation counters are typically manufactured as thin foils in order to reduce beam disturbance. The following technical requirements must be fulfilled to enable accurate beam intensity measurements: The light output should be proportional to the number of traversing ions and show little quenching in the investigated intensity range. The decay time of π-electron excitation should be a short as possible to achieve high counting rates. Modern materials exhibit nanosecond response [27]. The scintillation material must be transparent to its fluorescence photons and the refractive index should be around 1.5 for optimal further transmission through a light guide [62]. The fluorescence wavelength λ fluor should be in the range of 350 nm < λ fluor < 500 nm to ensure high efficiency at the photo-cathode or CCD sensor Position measurements The density distributions of particles along the two transversal beam coordinates T and U are called beam profiles. They contain information on local beam width and position. Feedback loops that compare nominal and measured values can be used to adjust the phase space of the ion beam. In scanned particle therapy, profile measurements are an important verification step to monitor spot positions online and ensure accurate patient irradiation. Since measurement devices must be placed in the beamline, the following section outlines frequently used non-destructive beam profile monitors. multiwire proportional chambers Energies of therapeutic proton beams range between 50 and 250 MeV. Beam currents of the order of several na yield very small signals in conventional profile grids. As a consequence, multiwire proportional chambers (MWPCs), introduced by Breskin et al. [10], are often used as profile monitors in proton therapy. A unit module consists of a plane of thin wires ( µm diameter and mm spacing) positioned between two kv cathode planes [12]. Secondary electrons liberated by the traversing ion beam are multiplied in the high electric field and the resulting Townsend avalanche is detected. The current measured in a single wire will be highest in close proximity to the beam center and independent of the deposited dose because of charge amplification. The counting ability is merely limited by electrostatic field shielding due to slowly drifting positive gas ions in the chamber. Combination of horizontal and vertical modules allows for 2D lateral beam profile measurements. Position and width can be determined using fit routines or center-of-gravity algorithms. 23

31 3 BEAM DELIVERY AND DIAGNOSTICS strip chambers The lateral beam position can also be measured based on the distribution of positive gas ions in the detection volume. For this purpose, two planes of orthogonal strips (millimeter width and submillimeter spacing) are separated by a high voltage plane (anode). Signal planes are typically Kapton isolated and several 10 µm thick [62]. Thus, the penetration depth of the particles is large in comparison to the thickness of the planes and the resulting beam disturbance is minimal. The Gantry 2 strip chamber is operated in counting and not amplification mode at 1.8 kv. The distance between the anode planes and the strips measures 1 cm [31]. Because the drift velocity of positive ions is small compared to liberated electrons, strip chambers require longer integration times than MWPCs to accumulate sufficient signal for beam profile analysis. linear wire scanners Instead of many wires covering the beam cross section, only two of them are moved through the beam in linear wire scanners [4]. Modifications in the vertical and horizontal scan path allow for profile sampling at different locations and different times. The measurement principle of linear wire scanners appears well-suited regarding verification of continuously scanned particle beams (cf. section 3.1.3). Wire movement could be synchronized with the scan trajectory reducing dynamic line shapes to quasi static, spot-like profiles. However, this approach would require comparable wire and scan speeds of up to 2 cm /ms in the case of PSI Gantry 2. scintillation screens Combining a scintillation screen with a CCD camera can be a very simple and reliable profile monitor of high resolution (cf. section on scintillation counters). Information on beam position and width can easily be accessed by observing light spots on the scintillation foil. Hence, viewing screens can be very useful during commissioning or troubleshooting. Permanent installations in the beam line require coating of the screen with a thin conductive layer to avoid charging the non-conducting screen material. Furthermore, the scintillation material needs to be matched to particle type and energy as well as expected beam intensity. A high dynamic range and linearity between particle fluence and light output are essential requirements. Fast variations in the beam profile additionally call for fast fluorescence decay times. residual gas ionization monitors MWPCs, strip chambers, wire scanners and scintillation screens introduce additional material in the beamline to obtain profile information. Residual gas ionization monitors (RGIMs), on the other hand, surround the central beam axis and can, therefore, 24

32 3.2 Beam Diagnostics be considered as nearly non-destructive. They measure the ionization of the residual gas within the beamline by projecting positive gas ions onto a position-sensitive micro-channel plate [2]. The applied electric field must be large enough to enable sufficiently fast charge collection with negligible beam steering effect. Figure 3.4 depicts a sketch of an RGIM. The collector rods are aligned parallel to the beam axis in order to be independent of momentum transfer along this direction. Installation of both plates on a supporting frame rotating around the beam allows for profile measurements in T and U. Figure 3.4: Schematic drawing of an RGIM. Positive residual gas ions created within the RGIM are guided by a transverse electrostatic field onto a micro-channel plate amplifier. Its output is connected to the collector rods providing signals proportional to the spatial beam density. Rotation around the central beam axis enables 2D, nearly non-destructive profile measurements. Reprinted from [62]. The collection of positively charged ions is preferred for beam profile measurements with RGIMs, since the transversal momentum transfer of the space charge field to the liberated electrons is much larger and can cause significant geometrical distortions. Positive ions are much heavier and, therefore, less affected by space charge effects. However, their mobility is smaller than that of a free electron, which results in longer integration times. The use of RGIMs for low-energy particle beams is limited due to steering effects. The electric field between the two position-sensitive plates acts perpendicular to the beam direction. Thus, considerable deflection has to be taken into account at low beam energies (E < 5 MeV in case of protons). 25

33

34 4 PROSCAN FACILITY The Center for Proton Therapy at PSI launched the PROSCAN project in 2000 with the intention to extend the medical program and expand activities in dynamic proton beam scanning [47]. A dedicated superconducting cyclotron was installed in 2003 [59] and constructions on the new Gantry 2 were finalized in 2010 [49]. The first patient treatment on Gantry 2 was performed in November This chapter describes the technical aspects behind proton therapy at PSI. The layout of the PROSCAN facility will be of vast importance for the development of a line scanning verification system. Thus, central components of the hardware (sections 4.1 and 4.2) and software (section 4.3) will be reviewed in the following. 4.1 Beamline The beamline to the Gantry 2 can generally be divided into three compartments: (1) beam generation using the compact medical therapy (COMET) cyclotron, (2) energy selection using multiple low-mass carbon wedges and momentum slits and (3) beam transportation using laminated steering, bending and focusing magnets. Each compartment is being outlined based on a schematic plan in this section Beam generation The COMET cyclotron accelerates protons to a fixed energy of 250 MeV. Superconducting coils provide a magnetic field of 3.8 T [59]. The high extraction efficiency of up to 80% minimizes the amount of radioactivity in the machine and enables short waiting times for service. The extracted proton current used for patient irradiations on Gantry 2 is, on average, around 200 to 400 na and it shows fluctuations of the order of a few percent on a 200 µs time scale [49]. Beam stability is crucial for timedriven spot scanning in order to precisely deliver planned proton distributions. The proton beam intensity can be varied rapidly ( 100 µs) between 0 and 100% by adjusting the electric field on a vertical deflector plate [49]. Positioned in the central region of the cyclotron, it deflects slow protons in close proximity to the ion source. A set of collimators along the first orbits of acceleration reduces the beam transmission by blocking off the deflected part of the beam. This fast, dynamic modulation of the beam intensity will be exploited for shaping dose distributions when scanning the beam at maximum velocity along lines or contours. 27

35 4 PROSCAN FACILITY Figure 4.1 depicts the very first part of the PROSCAN beamline. Small dipole steering magnets (SM) center extracted protons in the transversal x, y-plane. Collimators such as KMA1 reduce the phase space of the proton beam to particles with little angular offset. The permanent diagnostic devices MMAC1 and MMAC2 monitor the extracted beam current. The first pair of quadrupole magnets (QMA1 and QMA2) focuses the beam onto the central z-axis. COMET SC cyclotron SMJ1xy VMAD1 SMA1xy KMA1 MMAC1 MMAC2 MMAP 1/2 QMA1 MMAP 3/4 QMA2 AMAKI QMA3 Steering magnets Magnets Beam blockers Diagnostics Vacuum devices Coll imator s/slits to the energy selection unit y z x Figure 4.1: The superconducting COMET cyclotron accelerates protons to 250 MeV. The extracted beam intensity can be adjusted by changing the voltage on a vertical deflector plate positioned in close proximity to the ion source. Small dipole and larger quadrupole magnets steer and focus the proton beam. The kicker magnet AMAKI1 acts as a fast beam on/off switch by deflecting protons into a beam dump. Diagnostic devices such as MMAC1 and MMAC2 monitor the extracted beam current. Image courtesy of PSI, Center for Proton Therapy. Spot scanning requires many on/off cycles of the accelerated ion beam when moving from one position to the next. The kicker magnet AMAKI1 acts as a fast switch with a 50 µs beam off reaction time and up to approximately 1 khz repetition rate capability [49]. The 250 MeV proton beam is steered out of the beamline and deflected into a dump with little leakage when AMAKI1 is activated. In the case of kicker magnet failure, the power of the high frequency generator can be reduced to 80% in less than 50 µs, which prevents protons from exiting the cyclotrons [11]. If diagnostic monitors still measure significant current along the beamline, the ion source is shut off with less than 20 ms reaction time [11]. 28

36 4.1 Beamline Energy selection Patient irradiations on Gantry 2 make use of the fast degrader system DMAD1 L/R. The final proton range is controlled by adjusting the beam energy dynamically with two opposed sets of carbon wedges, which are moved mechanically in and out of the beam (cf. figure 4.2). Energy changes that correspond to 5 mm proton range differences in water can be realized within 50 ms [59]. Setting the corresponding beam tune (e.g. changing the current in the laminated magnets along the beamline) requires additional 30 ms, which yields a total dead time for energy changes of 80 ms [54]. Combination of continuous scanning and rapid energy switches opens up the possibility for a clinical integration of volumetric rescanning without increasing overall treatment time beyond tolerable limits. to the cyclotron KMA2 BMA1 DMAD1 L/R QMA4 QMA5 VMAD2 MMAP 5/6 MMAC3 MMAP 7/8 MMAC4 DMAF1 KMA3 KMA4 SMA2xy MMAP 9/10 KMA5 MMAP 11/12 AMA1 15 to Gantry 2 1x Steering magnets Magnets Beam blockers Diagnostics Vacuum devices Coll imator s/slits to Gantry 1 MA Q Figure 4.2: Fast energy changes are realized by opposing carbon wedges in the degrader (DMAD1 L/R) and laminated magnets along the beamline. Two adjustable collimator stacks (KMA3 and KMA5) correct for phase space confusions such as beam broadening due to scattering. Nevertheless, energy degradation down to 70 MeV results in a significant decrease in transmission. The planar ionization chamber MMAC3 can register beam even if the kicker magnet is activated. Hence, its signal is well suited to control the voltage on the vertical deflector plate. Image courtesy of PSI, Center for Proton Therapy. 29

37 4 PROSCAN FACILITY Energy degradation down to 70 MeV reduces the transmission by more than two orders of magnitude, since phase space disturbances (beam broadening and lateral scattering) must be corrected by two sets of round collimators. Each one of them is comprised of five apertures. KMA3 defines the size of the proton beam and KMA5 the angular acceptance. After the degrader unit, the residual proton beam is again focused by a pair of quadrupoles and finally bent by the first, large dipole magnet AMA1. Figure 4.2 displays another final element: the beam blocker BMA1. Its response time is less than 1 s. The diagnostic devices MMAC3 (planar ionization chamber) and MMAC4 (secondary emission monitor) are installed such that they register beam even if it is deflected by the kicker magnet. It is, therefore, worth considering to use the signal of MMAC3 to control the voltage on the vertical deflector plate near the ion source Beam transportation The beamline between the degrader and the coupling point to Gantry 2 can be subdivided into two sections: The achromatic beamline section between AMA1 and AMA2 analyzes the energy content of the beam. Momentum band defining slits (FMA1 L/R) mark the symmetry point. The slit positions define the width of further transported momentum band. Currently, δp/p is set to approximately ±1% [49]. The following section diverts the beam into three different areas. In order to introduce artificial intensity losses for high energies and achieve a constant, energy-independent beam current of 0.5 na at iso-center, an intensity suppression collimator was installed upstream of BMB1. The quadrupole doublets QMA12/13 and QMA1/2 at the degrader entrance can be used to variably defocus the beam according to its energy in front of the collimators. After traversing the separation wall to Gantry 2 (Riegel), the proton beam is focused with a triplet at the coupling point to the gantry, where the vacuum is interrupted. The separation is depicted in the form of two vertical lines downstream of the air-pressured beam blocker BMB2 in figure 4.3 (< 60 ms reaction time [11]). A rotational symmetric double focus with achromatic phase space is established. The Frisch-grid ionization chamber MMBC 2 (or monitor 3) is one of the three beam current verification monitors. It is filled with dry nitrogen gas at 10 mbar overpressure [11]. The anode and cathode electrodes are made of 5 mm thick aluminum plates to reduce acoustic noise disturbances. The wire plane between them shields and collects positive ions slowly drifting to the cathode and minimizes charge collection time [31]. The measured anode current is directly proportional to the proton beam current at iso-center and serves as an independent check of the two primary dose delivery monitors. 30

38 A Riegel 4.2 Gantry 2 to the energy selection unit Steering magnets Magnets Beam blockers Diagnostics Vacuum devices Colimator l s/slits AMA1 QMA6 KMA6 MMAP 13/14 KMA7 QMA7 QMA8 VMAD3 MMAP 15/16 FMA1x SMA3x QMA9 MMAP 17/18 MMAP 19/20 SMA4y MMAP 21/22 KMA8 MMAP 23/24 MMAC5 MMAE1 MMAP 25/26 QMA13 intensity suppression collimator AMA3 QMB1 QMB2 QMD5 to PIF and Gantry 3 AMC1 KMBV1 MMBP 5/6 SMB2xy to OPTIS 2 QMB3 QMB4 QMB5 SMB3xy MMBP 7/8 BMB2 MMBC 2 MMBC 1 KMB2 MMBP 9/10 KMB1 to Gantry 2 AMA2 QMA11 QMA12 SMB1xy BMB1 MMBP 1/2 VMBD1 MMBP 3/4 to Gantry 1 QMA10 Figure 4.3: The first part of the beamline from the degrader unit to Gantry 2 acts as an energy analyzer. Momentum band defining slits (FMA1 L/R) select the energy content of the beam. Defocusing high-energy beams prior to the intensity suppression collimator accounts for comparable transmission losses low-energy beams experience in the degrader. Thus, a constant, energy-independent current of 0.5 na can be granted at isocenter between 100 and 200 MeV. The grid chamber MMBC 2 measures the residual current just in front of the coupling point and serves as an independent check of the two primary dose monitors. Image courtesy of PSI, Center for Proton Therapy. 4.2 Gantry Layout The second generation proton beam gantry built at PSI features advanced magnetic upstream scanning in both lateral directions T and U. The compact but eccentric design of Gantry 1 has been compromised to grant easy access to the patient table at any time on a permanent fixed floor. For this purpose, the isocentric rotation angle α has been limited to the interval [ 30, 180 ]. The patient table, set up opposite of the half-cylindric false floor, has full rotational flexibility in the horizontal plane. Thereby, any beam incidence within a 4π solid angle can be achieved. Figure 4.4 shows an image of the Gantry 2 treatment room [45] with α being set to approximately 80. The technical specifications of Gantry 2 meet the requirements of very fast beam scanning. This opens up the possibility to apply multiple target repaintings without increasing overall treatment time. Clinical integration of rescanning would reduce the sensitivity of PBS to organ motion and allow for treating moderately moving tumors. Two in-room imaging devices have been installed to support this objective: A sliding CT is used for daily positioning and includes the possibility to acquire timeresolved (4D) tomograms. An X-ray source in beams-eye-view can shine diagnostic radiation through a hole in the external return yoke of the last bending magnet on the gantry onto a panel that is extractable behind the patient. The system is used for redundant position checks and could potentially be synchronized with the proton beam to explore image-guided treatment of moving targets. 31

39 4 PROSCAN FACILITY Figure 4.4: Gantry 2 treatment room at PSI. The isocentric layout features onesided rotation between 30 and 180 and, thus, comfortable access to the patient table on a permanent fixed floor. The static, wooden ceiling height of about 2.3 m mimics typical rooms. Image courtesy of PSI, Center for Proton Therapy. Reprinted from [45] Beamline Figure 4.5 depicts the layout of the Gantry 2 beamline [49]. Downstream of the coupling point air gap, the beam is bent by two counteracting 58 dipole magnets (AMF1 and AMF2). The parallel displacement of the rotating beamline from the central axis of rotation measures 3.5 m. The sweeper magnets WMFT and WMFU deflect the beam laterally and realize 2D scanning. The pencil beam can be swept along T with a maximum speed of 2 cm /ms at iso-center. The maximum scan speed along U is four times slower (5 mm /ms). The thin quadrupole corrector QMFC placed between the first doublet keeps the beam focused at iso-center nearly independent of scan direction and amplitude. The last 90 bending magnet directs the beam towards the patient. A large 15 cm pole gap was necessary to transport the proton beam even at highest T deflection with minimal loss. The complex and challenging design of AMF3 is described in [19]. The final rectangular field size at iso-center measures T U = cm 2. Beam optic properties of Gantry 2 are based on a rotational symmetric achromatic phase space at the coupling point in order to achieve an invariant, rotationindependent beam transport through the gantry. The injected phase space is imaged 1:1 to the iso-center. Parallelism of the beam in both scanning directions was achieved by optimizing the entrance and exit angles of the last bending magnet to 12.2 and 24.4, respectively [49]. This simplifies treatment planning and dosimetry, since delivered dose distributions will be independent of the source-to-target distance. 32