Radiation Physics: Charged Particle Therapy

Charged Particle Acceleration/Steering

The source of protons for a treatment facility is a tank of hydrogen gas or a hydrogen gas generator that produces hydrogen from water. Electrons are removed from the atoms during ionization in an ion source, and the remaining protons are then accelerated by an electric field. Because the proton is a charged particle, it can be steered by a magnetic field perpendicular to its trajectory, with the resulting radius of curvature a function of the proton energy and the strength of the magnetic field. The proton beam is further focused with quadruple and higher-order magnetic fields. The beam pipe for particle transport is maintained at high vacuum, because collisions with air molecules would widen the beam and the energy distribution.

A clinical proton beam that will reach 32 cm of water equivalent depth needs to be accelerated to 59% the speed of light (kinetic energy of 228 MeV). This energy is difficult to reach with a single linear accelerator. Circular accelerating devices, cyclotrons and synchrotrons, use magnets to bend the proton paths so the same smaller accelerator elements can repeatedly push the proton to a higher energy each “lap.”

A cyclotron, as shown in Fig. 8.2 , had two Dees in its original design, with an alternating potential across the gap between them and a magnetic field perpendicular to the plane containing the particles. Modern designs have two to four “Dees” twisted into a spiral shape to improve beam focusing. Protons start in the cyclotron's center and, as the protons gain energy, they circulate at larger radii. A cyclotron produces a nearly continuous beam at a fixed energy.

A synchrotron circulates protons in a ring of fixed radius, as shown in Fig. 8.3 . It is fed by a linear accelerator that injects protons into the ring and there is at least one accelerating cavity in the synchrotron itself. As the protons are accelerated with each lap, the bending magnet fields ramp up to keep the protons in the beam pipe. The synchrotron can accelerate protons to a range of predefined energies.

Protons beams from a single cyclotron or synchrotron can be shared by multiple treatment rooms. Treatment rooms can have fixed-beam angles, a full gantry, or a partial gantry. Although gantry delivery permits greater beam angle choice, it is more technically challenging to maintain submillimeter precision while isocentrically rotating the bending and focusing magnets, and it requires two more building floors of vertical space for clearance. Compact cyclotron systems may be gantry-mounted.

Scattered Beam

With passive scattering, a spatially uniform dose distribution is achieved by scattering and degrading the primary proton beam in a set of distributed absorbers. Passive systems use either single- or double-scattering foils. The foil is made of high-Z materials, such as lead, and effectively scatters the beam while keeping energy loss to a minimum. Generally speaking, double-scatter systems are preferred because larger, more uniform beams can be generated ( Fig. 8.4 ).

Beam shaping is accomplished with the use of some devices that are patient- and field-specific and some that are not. Range modulation refers to the concept that pristine Bragg peaks must be spread out to be clinically useful. In the passive scatter system this is done with either a propeller-shaped modulator ( Fig. 8.5 ) or a ridge filter, which is not patient-specific but may be range-specific, with a small library of range modulators available to address the different range modulator options that might be clinically desired. The Bragg peaks are spread out by placing these devices, which have variable thicknesses, in the path of a given beam. The thicker the modulator, the more the beam is shifted in range.

Additional beam conformality is achieved with the use of patient-specific devices ( Fig. 8.6 ). Brass apertures are formulated to be used for a specific field for a given patient. Apertures are equivalent to the blocks used in conventional radiotherapy. A range compensator is another patient-specific beam modulator. This device is a block of plastic that is created by a computer-driven milling machine for a specific field and is responsible for conforming the distal edge of the beam, necessary because of the variability in individual proton range secondary to differences in radiologic density in the tissue along their particular trajectory. For historical reasons, a range compensator is sometimes referred to as a “bolus,” but this should not be confused with the bolus that is directly applied to the patient's skin in conventional radiation therapy. This device is located adjacent to the brass aperture in the snout of the accelerator, above the patient. Use of patient-specific devices that must be formulated for each beam angle is not a trivial matter in terms of labor and cost. Furthermore, insertion and removal of the heavy apertures for each field is time-consuming and potentially dangerous for therapists and technicians. This process adds a significant amount of time to each patient's treatment and limits the total number of patients treated at a facility.

Although proton therapy has the potential to spare more healthy tissue than x-ray therapy, by virtue of physical qualities of the Bragg peak, there is a concern for increased neutron production. Neutrons can be generated whenever high-energy protons are slowed down by nuclear interactions. Such events can take place inside or outside of the patient. A certain amount of neutron production takes place from proton interactions within the patient, and this is unavoidable and relatively modest . But with the passive scatter system, neutrons are primarily produced from proton interactions with patient-specific scattering material and collimators in the accelerator system.

Spot Scanning

Virtually all proton facilities under construction or recently constructed utilize active (spot or pencil beam) scanning. With this technique, the beam is not spread out or scattered; rather a narrow (2-mm to 10-mm wide) pencil beam is precisely steered through the treatment volume by magnets in the beam line nozzle. One of the benefits of this system is that there may not be any field-specific or patient-specific hardware; therefore, the time and cost of fabrication and daily insertion are avoided. Most importantly, neutron production and proton beam scatter are greatly diminished with the absence of this hardware. Apertures can still have a role in spot-scanned pencil beam treatments for targets that require a very sharp lateral penumbra. As the lowest energy protons have the largest spot sizes, more superficial targets, such as ocular melanomas, could benefit the most from an aperture system.

These pencil beams deposit the dose layer by layer, with the distal edge treated first and the more superficial layers treated thereafter ( Fig. 8.7 ). Pencil beam scanning also allows for good conformation of the proximal edge of the treatment volume, which is typically not possible with scattered beams. The ability to paint dose spot-by-spot makes pencil beam scanning the best technique for the treatment of irregularly shaped tumors and tumors tightly constrained by critical structures.

Because proton accelerators have a minimum as well as a maximum energy, the treatment of more superficial targets requires a range shifter . This device is a low-Z material of uniform thickness placed in the path of the beam, typically in the beam nozzle. The resulting Bragg peaks of the beam will be shifted by a fixed distance toward the surface of the patient. The range shifter will also scatter the beam. To minimize its effect on the spot size, it may be desirable to move the range shifter closer to the patient, either placing the device on the treatment couch or even incorporating it as a thicker couch.

Facility Design

The design of a proton facility must incorporate several elements in the treatment rooms, control areas, and building that are different than those in an x-ray facility. Neutron production requires thicker concrete shielding and a longer maze. A proton gantry is also substantially larger than a modern tomotherapy unit or x-ray linear accelerator, usually requiring at least two stories of shielded construction. The vendor will also require an on-site control room and storage for machine maintenance and quality assurance elements. To reserve the gantries for beam delivery, some facilities utilize a setup room for initial positioning and a gurney for transport of the couch and patient to the treatment room. The patient support assemblies are typically robotically controlled, able to adjust three translational and three rotational degrees-of-freedom in modern facilities. Standard or specialized couch tops and extensions can be mounted on the patient support assembly, but care must be taken to understand the couch top impact on proton treatment beam (heterogeneities, etc.). Robotic arms in setup and treatment rooms allow for transfer of the couch and patient between rooms, with the patient fully immobilized.

Image guidance is very similar to that in the photon clinic. In-room imaging can include 2D and 3D modalities for 2D/2D, 2D/3D, and 3D/3D alignment and optical surface imaging. Imaging equipment can move with the gantry or couch or have a fixed location in the room. A computed tomography (CT)-on-rails or a cone-beam CT can provide localization based on soft tissues and information about any changes in the beam path that could be significant enough to have an impact on dose accuracy, allowing the team to make an informed decision about patient treatment. It is envisioned that volumetric imaging will be used for rapid adaptive replanning while the patient is in the treatment room as image acquisition, target delineation, and computational planning tools become more automated and efficient.

Spread Out Bragg Peak and Water Equivalent Thickness

With an understanding of the fundamental interactions which form the basis of the Bragg peak, we can now turn our attention to the more macroscopic attributes of therapeutic charged particle beams. Fig. 8.1 presents the dose deposited in water as a function of depth by a series of monoenergetic, clinical proton beams. In addition to the depth of the Bragg peak increasing with beam energy, the width of the peak also increases owing to an effect known as “range straggling.” Even for a beam of monoenergetic particles, the individual particles do not all stop at exactly the same depth. Differences in the trajectory of each particle are caused by stochastic variations in energy loss and scattering dynamics. These path length differences become more pronounced as the particles penetrate deeper, which leads to a broadening of the Bragg peak as the beam goes deeper.

Because the width of a single Bragg peak is typically small compared with the size of the clinical target in the beam's eye view, the superposition of multiple Bragg peaks (i.e., multiple energies) is needed to create a uniform dose across the target. In a clinical beam, the “spread out Bragg peak,” or SOBP, is formed when appropriate relative weightings are applied to a collection of single-energy incident beams (see Fig. 8.1 ). The SOBP can be parameterized by well-known metrics. The “d80” refers to the depth at which the dose distal to the SOBP falls off to 80% of the maximum. It turns out that in a monoenergetic proton beam, which creates a single Bragg peak, this depth closely corresponds to the average range of the beam. Analogous to d80, “d20” is the depth at which the dose reaches 20% of the SOBP maximum, and the difference “d20-d80” characterizes how rapid the dose falls off distal to the SOBP. The modulation width refers generally to the length of the region of uniform dose in an SOBP. As the modulation width increases, the magnitude of the entrance dose increases relative to the SOBP.

Although proton depth dose curves are typically shown in water, it is critical to understand the differences associated with propagation through heterogeneous materials such as the human body. The penetration depth of a charged particle beam depends on the stopping powers of the various anatomic regions through which it passes. Higher stopping power implies more energy transfer from the charged particle beam to the material, resulting in a shorter particle range. Generally speaking, stopping power increases with material density, although there is also some dependence on the molecular composition of the material itself. Given two points separated by a distance d in a heterogeneous material , a “water equivalent thickness,” or “WET,” can be defined. This represents the equivalent thickness of water through which a charged particle beam would lose the same amount of energy as it would traversing the heterogeneous distance d .