New Developments in Proton Therapy Systems

C-M Charlie Ma, Ph.D.

Radiation Oncology Department, Fox Chase Cancer Center, Philadelphia, USA

Abstract. Proton beams can provide better dose conformity to the treatment target compared to commonly used photon and electron beams allowing for dose escalation and/or hypofractionation to increase local tumor control, reduce normal tissue complications and/or treatment time/cost. This paper reviews three novel proton accelerator designs that aim at cost-effective solutions for widespread applications of advanced particle therapy. The basic concepts, the system designs and the potential clinical applications are discussed in detail for superconductor accelerators, dielectric wall accelerators and laser-particle accelerators.

Keywords: Particle Accelerators, Superconducting Techniques, Dielectric Wall Accelerator (DWA), Laser-Particle Acceleration, intensity-Modulated Particle Therapy (IMPT), Relative Biological Effectiveness (RBE), Oxygen Enhance Ratio (OER). PACS: 87.56,bd.

INTRODUCTION

Proton therapy can provide better dose conformity to the treatment target compared to commonly used photon and electron therapy because of the low entrance dose, sharp penumbra, and rapid falloff at the distal edge of the proton dose distribution, and the maximum rate of energy loss near the end of proton range, i.e., the Bragg Peak effect [1]. Proton, helium, and heavier ion beams were used for biomedical studies in the early 1950s. The first human patient was treated for a pituitary tumor in 1954. Since then, about fifty five thousand patients have been treated with proton beams worldwide [2].

In the last two decades proton therapy has lagged behind photon and electron beam therapy because of the prohibitively high cost of a conventional particle therapy system based on the cyclotron or synchrotron technology. An accelerator that is big enough to accelerate protons or ions to the required therapeutic energies can cost tens of millions of dollars. Protons and ions are difficult to handle, requiring large magnets and a vacuum delivery path. As a result, the gantry that directs a proton/ion beam into the patient is massive. Currently available proton/ion gantries are roughly three times the radius and three times the length of typical clinical Linac gantries. The amount of concrete and steel used to shield a proton/ion treatment room will thus be approximately nine times or greater than for a Linac room.

The use of an expensive accelerator is maximized by sharing the accelerator in multiple treatment rooms. Such large proton therapy facilities will require additional space for a switchyard, which houses the vacuum beam lines, dipole bending magnets, steering magnets, focusing quadrapole magnets, power supplies, and cooling equipment. The cost of the gantries and the building increases the total capital cost to

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about $100 million for such a large proton facility with multiple treatment rooms. Even if a proton facility can be amortized for 30 years or longer, its maintenance, upgrade, and operational cost will be significantly higher than that for a linac-based facility of similar treatment capacity. Alternative solutions are therefore needed to address both the size and cost issues in order to make proton or ion therapy a widely available treatment modality [3].

ADVANCES IN PARTICLE ACCELERATION TECHNIQUES

Particle accelerators are designed to control electric and magnetic fields in such a way as to efficiently accelerate charged particles. A linear accelerator (linac) accelerates particles in a linear path, and therefore its length is proportional to the electric field and the gain in particle energy. Conventional Linacs cannot provide sufficient electric field strengths to accelerate heavy charged particles such as protons or light ions with therapeutic energies for radiation oncology applications. For a given electric field strength, alternative solutions are cyclotrons and synchrotrons that repeatedly steer charged particles across the same electric field to achieve desired energies (see figure 1). A cyclotron provides a constant magnetic field to steer charged particles (different energies corresponding to different radii), and its size and weight are proportional to the strength of the magnetic field, and therefore to the size and weight of the magnets used. A synchrotron provides a variable magnetic field to steering charged particles of different energies through the same trajectory, and it is generally larger and heavier than cyclotrons.

For therapeutic applications, protons must be accelerated to 230 – 270 MeV

(corresponding to 33 – 43 cm range in water). The Francis H. Burr Proton Therapy Facility at the Massachusetts General Hospital (MGH) uses a room-temperature cyclotron containing 3 Tesla magnets, which is 4 m in diameter and its iron core and

a.

b.

c.

FIGURE 1. Schematic demonstration of different ways to accelerate charged particles using electric and magnetic fields: (a) linear acceleration; (b) cyclic acceleration; and (c) synchronized acceleration.

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copper coils weigh more than 220 tons [4]. Conventional cyclotrons are still too large and heavy for gantry mounted proton therapy systems but they have been used for large proton therapy facilities to provide beams for multiple treatment rooms. Synchrotrons have been used for proton therapy applications and they are the only sources used for ion therapy facilities [5]. A compromised accelerator design has combined the best aspects of both cyclotrons and synchrotrons; a fixed focusing, alternating gradient (FFAG) accelerator [6] consists of a ring of magnets (like a synchrotron) that provide a large, constant magnetic field to allow for different trajectory radii (like a cyclotron). The actual advantages of this accelerator design remain to be demonstrated. Conventional accelerators have been optimized in terms of performance, stability and cost, and are being improved with new technical capabilities for beam scanning, gating and intensity-modulated dose delivery.

Superconductor Accelerators

New methods have been developed to minimize the size and weight of particle accelerators so as to allow for one accelerator per treatment room as is in the case for conventional radiotherapy using electrons and photons. A significant improvement in reducing the magnet size and weight is achieved by using superconducting techniques. The current-carrying coils are contained in a superconducting cryostat so that higher currents can be achieved and consequently provide higher magnetic field strength. The Comet cyclotron built by ACCEL Instruments (Bergisch Gladbach, Germany) for proton therapy uses a superconducting accelerator that weighs only 80 tons [7]. The compact proton therapy system commercialized by Still River Systems (Littleton, MA) uses a superconducting cyclotron that provides a 11 Tesla magnetic field strength and is about 1 m in diameter, weighing about 17 tons. “Table-top” synchrotrons weighing < 2 tons have been proposed and the concept of superconducting cyclotrons for ion acceleration weighing ~600 tons has been proposed [8].

DWA Techniques

The dielectric wall accelerator (DWA) is a new type of induction linear accelerator, which is based on the development of a new class of insulators called high gradient insulators (HGI). HGI has significantly improved voltage-holding ability over conventional insulators. A conventional induction linac has an accelerating field only in the gap of the accelerating cells, which represent only a small fraction of the length. By replacing the conducting beam pipe by an insulating wall, accelerating fields can be applied uniformly over the entire length of the accelerator yielding a much higher gradient (around 100 MeV/m) allowing the design of much more compact linear accelerators. The pulsed acceleration field is developed by a series of so-called asymmetric Blumlein structures [9] incorporated into the insulator with fast switches (see figure 2), allowing for a single pulse traveling wave mode. Oil switch/Polypropylene Blumlein has achieved 100 MV/m stress in transmission lines for 5 ns pulses. At the Lawrence Livermore National Laboratory (LLNL), a prototype of a proton DWA is currently being built and design studies for a 250MeV proton accelerator with an overall length of less than 3 m have been carried out [10].

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FIGURE 2. Schematic diagram showing proton acceleration based on the DWA technique. A proton source will be needed to inject protons into the acceleration structure.

Laser-Particle Acceleration Techniques

Laser particle acceleration was first suggested for electrons [11]. Laser-electron acceleration flourished following the development of chirped pulse amplification and high fluence solid-state laser materials such as Ti:sapphire. In the last decade, charged particle acceleration using laser-induced plasmas has shown its potential for a compact, cost-effective proton accelerator. Proton acceleration is achieved by focusing a high-power laser pulse onto a thin target. The short femtosecond pulse width of the laser produces a high peak intensity that causes massive ionization in the target and expels a great number of relativistic electrons. This sudden loss of electrons leaves the target with a highly positive charge, yielding an effective electric field of 1012 V cm−1. This transient field then accelerates protons and other light positive ions, if present, to high energies. Laser-proton acceleration has been investigated by major laboratories worldwide, and energetic protons up to 58 MeV have been generated using high-intensity, short-pulse lasers [12]. Theoretical studies show that at a laser intensity of 1021–1022 W cm-1, protons may be accelerated up to 300 MeV with a broad spectrum and angular distribution, which are not suitable for direct radiotherapy applications. To solve this problem, compact particle selection and collimation devices [13] and compact proton therapy systems were proposed [13,14].

Complex target designs have been investigated to improve the ion acceleration efficiency and the beam characteristics. A double-layer target design has been shown to produce quasi-monoenergetic protons with narrow angular distributions [15,16]. In this design the target is composed of two layers: a high-Z base layer consisting of heavy ions and a large number of electrons and a second layer consisting of protons and electrons. The second layer is sufficiently thin in the laser propagation direction and sufficiently small in the transverse direction. In the field of the incident laser pulse, the target is ionized and many electrons escape from it under the action of the ponderomotive pressure of the laser radiation. Escaping electrons give rise to the quasi-static electric field of an unneutralized electric charge. The quasi-static electric field is localized in a finite region with dimensions comparable to the transverse dimension of the laser pulse. Since the total number of protons is small in comparison

protonHGI

Switch

Blumlein

HGI Transmission line

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with the number of electrons that have escaped from the target the effect of the electric field of protons on their dynamics is negligible and the proton acceleration can be described in the approximation of test particles moving in a prescribed electric field. If the thickness and width of the proton layer are negligible in comparison with a large electric field all protons will be accelerated in the same direction with the same energy. For a finite size proton layer, the relative energy spread of the particles is proportional to the ratio of the target thickness to the inhomogeneity scale of the accelerating electric field. The energy distribution of a quasi-monoenergetic laser-proton beam simulated using the PIC method for a double-layer target design shows [17] that the peak energy is 173 MeV with a 1% energy spread (FWHM). The total number of protons accelerated by the same laser pulse is 109. The proton layer in this case is 2.5 μm in diameter and 0.1 μm thick. The laser pulse contains 150 Joules with a 150 fs length and a 10 μm spot size (FWHM). Extensive PIC simulations were performed to investigate the combination of laser parameters (laser intensity, pulse length and spot size) and target dimensions [18]. Protons of therapeutic energies 150 MeV - 250 MeV can be accelerated using a PW laser with optimized laser intensity, pulse duration and spot size.

POTENTIAL APPLICATIONS OF COMPACT ACCELERATORS FOR RADIATION ONCOLOGY

Attempts have been made to utilize these novel particle acceleration techniques for medical applications especially in the field of radiation oncology. The goal is to develop compact particle therapy systems that can be used for single-room treatment facilities or that can be installed in existing treatment rooms designed for conventional photon and electron therapy. A brief description will be given below of two proton systems that are designed as single-room systems and a detailed discussion will follow on a more compact system based on laser-proton acceleration that can be placed in an existing radiotherapy treatment room.

Superconductor Accelerator Based Therapy Systems

Particle accelerators employing room-temperature magnets are generally too large and too heavy to be placed on a treatment gantry. They are therefore kept in separate locations and used to feed multiple treatment rooms to maximize the use of such an expensive accelerator. Novel designs of superconductor accelerators are being developed that are compact enough to be used for single-room proton therapy systems. The ACCEL Comet cyclotron weighs only 80 tons, which have been used in compact proton therapy designs for both single-room and multiple room facilities. The Still River superconductor cyclotron weighs about 17 tons, which is placed directly on the proton treatment gantry (see figure 3). The gantry only rotates 180 degrees while the 360 degree beam incidence is facilitated by a robotic couch. A portable c-arm cone-beam CT is used for accurate target localization prior to a treatment. The proton system only provides broad, passive scattered beams and a secondary collimator with a separate, more precise rotation system is used to deliver shaped proton fields at accurately defined incident angles. Such compact designs allow for single-room

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proton therapy systems with much reduced shielding and construction cost. The Still River proton therapy system requires a 50 ft x 50 ft x 40 ft room and the whole facility including the proton therapy system and the room will cost $21 million.

FIGURE 3. The Still River single-room proton therapy system design [31].

DWA Based Therapy Systems

The DWA technique is being developed for compact proton therapy systems. The LLNL prototype DWA is designed for a maximum of 250MeV proton energy with an overall length of less than 3 m [9,10]. This DWA will produce a 100mA beam current in a short pulse of 1ns with a 50 Hz pulse repetition rate and variable energy and focus. LLNL is collaborating with Tomotherapy, Inc. (Madison, WI) to develop a single-room proton therapy system based on the DWA technique [19]. The LLNL compact accelerator will be mounted directly on a gantry (see figure 4). The gantry will rotate 200 degrees while the 360 degree beam incidence will be provided by proper couch rotation. The vault dimensions are 20 ft x 20 ft x 14 ft. This DWA based proton therapy system is estimated to cost about $20 million.

FIGURE 4. The Tomotherapy single-room proton therapy system design [32].

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Laser-Proton Accelerator Based Therapy Systems

Starting from 1998, researchers at LLNL and Stanford University investigated the feasibility of using laser-accelerated protons for radiation therapy. As a result, two research initiatives were put forward focusing on target configuration to achieve therapeutic proton energies [20] and on particle selection and beam collimation to solve the problem of broad energy and angular distributions of laser-accelerated protons [21], respectively. Research continued at both Japanese Atomic Energy Agency (JAEA) and Fox Chase Cancer Center (FCCC) along these directions as the principal investigators moved to new institutions. With the support of the US Department of Health and Human Services, an experimental facility dedicated to laser-proton acceleration for cancer treatment has recently been established at FCCC and the research continued under this initiative [22-29]. The system being developed is based on the design of Ma [13] with the following three key elements: (1) a compact laser-proton source to produce high-energy protons, (2) a compact particle selection and beam collimating device for accurate beam delivery, and (3) a treatment optimization algorithm to achieve conformal dose distributions using laser-accelerated proton beams. The system was reviewed in detail by Ma et al [26].

Target Assembly

In this compact laser proton system, the target assembly consists of the final focusing system and the target holder. The target assembly is contained in a vacuum chamber (see figure 5) inside the treatment gantry together with the particle selection beam collimation system (see figure 6). Previous PIC simulations of optimal laser parameters and target geometry have been used to guide experimental studies. The results were also used to derive particle phase space data for further dose calculation and treatment optimization studies. Bi-layer targets are primary options where the front, thick layer provides a large number of heavy ions to form an intense electric field after relativistic electrons are expelled and the back, thin layer provides light ions to be accelerated by the electric field. The materials and thicknesses of the bi-layer target must be optimized based on the laser parameters and acceleration requirements.

FIGURE 5. The FCCC target chamber for proton acceleration.

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Particle Selection and Beam Collimation

A compact device for particle selection and beam modulation has been investigated, which utilizes a magnetic field to spread the laser-accelerated protons spatially, based on their energies and emitting angles, and apertures of different shapes will be used to select protons within a therapeutic window of energy and angle [13,26]. Such a compact device will eliminate the massive beam transportation and collimating equipment in a conventional proton therapy system. The laser-proton target assembly and the particle selection and collimating device can be installed on the treatment gantry to form a compact treatment unit, which may be installed in an existing radiotherapy treatment room.

FIGURE 6. A compact design for proton energy selection and beam collimation.

The particle selection and collimation system is schematically shown in figure 6,

where theoretical proton tracks in high magnetic fields (moving from left to right) are displayed. Only protons of energies within an energy range will be allowed to pass through the beam stoppers and refocused through an exit collimator. Collimators of different shapes, sizes and locations can be used to select particles of desired energies. Other protons will be stopped or scattered by the energy selection collimator so that they will not be able to reach the exit collimator. This collimator and particle selection design also serves as an independent quality assurance device to ensure that only the protons with the desired energies will be allowed to go through the system even if the protons have a quasi-monoenergetic distribution. Superconducting magnets are being investigated to reduce the size of the device [25]. The shielding for the whole system has been designed [27] to reduce the radiation leakage (from protons, electrons and other radiation particles) to the level required by state regulations (=0.1% of the treatment beam).

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Beam Monitoring System

Quasi-monoenergetic protons are advantageous for therapeutic applications in terms of dose rate and shielding design. On the other hand, protons with a broad energy spectrum provide opportunities for selecting protons of proper energies to deliver dose distributions with desired spread out Bragg Peaks (SOBP) that are essential to treating bulky tumors. Using the particle selection device described above we can provide proton beams of different energy spectra to realize “energy modulation”. By mapping the tumor volume with an array of laser proton beams of weights we can achieve “intensity modulation”. By combining energy modulation and intensity modulation, we can deliver more conformal dose distribution for radiation therapy using laser proton beams [26]. Energy and intensity modulated proton therapy (EIMPT) has been studied using monoenergetic proton beams [24]. The energy and direction of the proton beam are varied sequentially so that the Bragg peak will cover the whole tumor volume. With a known energy the dose rate of a conventional proton beam can be determined by a fluence-monitoring chamber. For laser-accelerated protons with an energy spectrum both the spectral shape and the fluence must be known in order to predict the dose rate and dose distribution. Our solution to this problem is a differential chamber in the particle selection system so that the fluence for individual energies (different spatial locations in the energy space) can be measured. A differential chamber consists of multiple electrodes to collect ionization from different parts of the cavity volume. An integral chamber will be used to monitor output of the combined beam based on the information from the differential chamber. This differential-integral chamber configuration will ensure accurate dose delivery with proper dose conformity for EIMPT using laser-accelerated proton beams [26].

System Design

The massive cost of a particle therapy facility is not only because of the expensive particle accelerator but also the associated costs for the gantries, the beam lines, the switchyard and the shielding required, which represent 50 –70% of the overall cost. Therefore, a cheaper particle source will not solve the problem if one still has to transport the particle beams through long beam lines to different treatment rooms/gantries. Also, laser-accelerated protons have a broad energy spectrum, which produce a clinically unfavorable dose distribution and cannot be transported using conventional beam lines that are designed for monoenergetic protons [22]. Our solution to this problem is to transport the laser beam to each of the treatment rooms and to design a compact gantry to include the target assembly, the particle/energy selection, the beam collimation and monitoring system so that we can retrofit it in a conventional linac room [13,26]. This will reduce the cost by at least an order of magnitude compared to that for a conventional proton therapy facility.

We show in figure 7 a schematic diagram of the laser-proton therapy unit [13,26]. The laser is transported directly to the gantry. The target assembly and the beam selection device will be placed on the rotating gantry and the laser beam will reach the final focusing mirror through a series of mirrors. The distances between some mirrors can be adjusted to scan the proton beam along x- and y-axis, respectively, which will

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generate a parallel scanned beam. An alternative method is to swing the target and beam selection device to achieve a scan pattern. This will generate a divergent scan beam. The treatment couch can be adjusted to allow for multiple beam arrangement with coplanar and non-coplanar, and isocentric and SSD (source-to-surface distance) treatments. To maximize the use of a PW laser, the laser beam can be shared by several treatment gantries for a multiple treatment room laser-proton therapy facility.

FIGURE 7. A compact design for a laser-proton radiotherapy system [13,29].

CONCLUSIONS

This paper has reviewed three novel proton accelerator designs as cost-effective solutions for widespread clinical applications of advanced particle therapy. Each technique has some potential advantages over the others and some technical challenges that may limit its clinical application and cost-effectiveness [2,3,5]. More compact particle therapy systems may be built for single-room particle therapy applications using superconductor-based accelerators. The size and weight of such systems are still too large to fit in an existing conventional radiotherapy vault. The cost of a single-room particle therapy system is 5-10 times of that of a conventional linac, in addition to the large space requirement. The DWA technique may be readily developed for a prototype system but many technical details must be first worked out for practical therapy applications. For example, one must first demonstrate a reliable system to produce >200 MeV protons with a reasonable dose rate for beam scanning [9,10,19]. The length of a DWA-based proton accelerator will be 2-3m, which will require a much larger gantry than that for conventional linacs. The cost of a DWA proton therapy system is estimated to be similar to that of a superconductor-based system. The laser-particle system has a great potential for a compact design to replace

Target and beam selection system

mirror

Adjustable distance for scanning along x

Adjustable distance for scanning along y

Gantry rotation

Main laser beam line

x

y

Couch

Proton beam

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conventional radiotherapy linacs and more cost-effective than the other particle systems. Some technical and engineering issues must be solved before a clinical prototype can be built including optimal target designs for therapeutic proton energies, high dose rate for beam scanning and high laser quality such as laser beam contrast and stability to meet the clinical stability and reliability requirements [3,26,30].

ACKNOWLEDGMENTS

The author would like to acknowledge the generous support from the US Department of Health and Human Services, Varian Medical Systems, the Strawbridge Family Foundation and the Kim Family Foundation.

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