The first step in generating a proton beam is to obtain a source of protons which can then be accelerated to energies sufficient for treatment. This can be done using hydrogen as the starting product and separating the hydrogen's electron from its proton by using an electrical field. Once protons have been generated, they must be accelerated such that the proton energy is sufficient to reach the distal edge of a tumor. Particle accelerators use an electrical field to accelerate protons and a magnet field to steer the charged particles. Linear accelerators are commonly used in photon therapy. In photon therapy, electrons are accelerated with a single pass through a series of electrical fields, and at the end of the beam line, the electrons are steered toward a target to generate photons, or alternatively to a scatter foil for electron therapy. Proton therapy requires cyclical particle accelerators which cause the particle to pass through the electrical field repeated times until they reach an energy sufficient for clinical use.
Presently, the two most commonly used devices for proton acceleration are cyclotrons and synchrotrons. Cyclotrons (figure 1. IBA) are composed of two large semi-circles with a space between them. These two semi-circles are known as 'D's' or 'Dees' (figure 2. IBA). There is a magnetic field which is perpendicular to the plane of the dees that is kept constant. Protons are injected at the center on the two dees, and by alternating the voltage supplied to the dees, the protons are gradually accelerated. The magnetic field helps steer the protons such that they move in a spiral pattern. The magnetic field and the voltage changes between the dees are kept constant, and as the protons circle, they continue to gain energy and gradually move outward until they can be extracted for use. Cyclotrons produce a high, continuous current of protons (i.e. they can produce a large amount of protons continuously); however, they are only able to produce protons of a fixed energy.
With synchrotrons, as the protons are accelerated, the magnetic field as well as the rate of voltage oscillation are both continuously modulated to keep the protons traveling in a fixed loop, rather than having them move gradually outward. Hence, in cyclotrons the protons' path changes as the energy increases, whereas in synchrotrons the protons are held in a constant path via changes in the strength of the magnetic field and alteration in the rate of voltage oscillation. Thus, synchrotrons can produce protons of various energies by varying the magnetic and electrical fields. However, once a given quantity of protons starts to be accelerated, the synchrotron stays in step with this set of protons, and a new set of protons can not be accelerated until the first set has exited the synchrotron. In other words, synchrotrons do not have a continuous output.
Once the protons have been accelerated, they must be guided to the gantry for delivery to the patient. When protons first exit the accelerator, they come out as a thin beam of protons of a single energy. Protons must be guided in a mono-energetic form because the bending magnets, which are part of the beam line that transports the proton beam, can only bend mono-energetic beams (figure 3. IBA). The proton beam travels in a vacuum within the beam line and is guided by a variety of magnets which can both deflect (dipole magnets) and focus (quadruple magnets) the beam. These magnets can be precisely controlled to direct beams of different energies. Additionally, many facilities use a single accelerator to supply multiple gantries. These beam lines have multiple branches, and the precise control of the bending magnets can be used to supply each gantry individually. Since only one gantry can be supplied at a time, algorithms to optimize beam delivery to each gantry have are being developed to maximize beam usage between gantries. Currently, in most clinical centers, the beam is manually directed by controllers to the next treatment room that has a patient ready for treatment.
The energy of the proton beam itself can be modulated in a number of ways. In accelerators which are able to produce protons of variable energies, such as synchrotrons, protons can simply be extracted at the appropriate energy. However, in accelerators which produce mono-energetic beams, such as fixed energy cyclotrons, a beam degrader can be used to change the energy of the proton beam (figure 4. IBA). This energy selection system (ESS) degrades the initial beam produced by the cyclotron to produce several different lower energies. This allows the beam energy to be modulated such that a variety of depths within the tissue can be treated. At present, with 250 MeV proton beams, depths of approximately 40 cm can be treated with a degrader, allowing shallower depths to be treated. Generally, a low Z material, such as graphite, is used to degrade the proton beam as it attenuates the beam while minimizing scatter. Nonetheless, despite the use of low Z materials, there is still significant beam attenuation and scatter resulting from degradation of the beam and shielding is used to minimize scatter.
Once the desired proton beam energy has been produced, it still needs to be 'spread out' such that it can cover the entire tumor, as a mono-energetic proton beam would only cover a small portion of the tumor with its Bragg peak. To create a beam with multiple energies that can spread its Bragg peak over multiple depths (Spread out Bragg peak), a modulator wheel can be used (figure 5. IBA) ('range shifter wheel' in figure 6. PSI). The modulator wheel spins in the path of the proton beam and is comprised of areas of different thicknesses which attenuate the proton beam to different degrees. By doing this, it creates protons of various energies, resulting in a spread out Bragg peak, such that larger tumors can be treated. However, the modulator wheel has several limitations. The modulator wheel has a limited number of proton energies for which it is effective, and can only create a single given width of spread out Bragg peak. Hence, a large number of modulator wheels would be needed. However, this number can be reduced by using current modulation.
The proton beam can be turned off at specific points during the revolution of the modulator wheel, which changes the width of the treatment field. This does have the disadvantage of increasing treatment times (as the beam is off for a portion of treatment) as well as making the system more complex. The beam current can also be varied precisely with the rotation of the modulator wheel, which can further reduce the number of modulator wheels needed; however, the complexity of the system is increased even further. Using this technique also requires a very sensitive current detector and changes in current which are very stable and linear.
After the proton beam has been created and directed to the treatment room using the beam line, there are a variety of ways that the protons can be precisely directed to treat the tumor. One way is to use a gantry which can rotate in 360 degrees about the patient, allowing delivery of radiation from any angle within a single plane (Figure 7. IBA) In conjunction with an adjustable treatment table, nearly any treatment angle desired can be achieved. There is great interest in the integration of robotic tables with proton therapy because they have more flexibility then traditional treatment tables (discussed in greater detail in module 6). However, the gantries need to be quite large (three stories or approximately 10 meters in height) to appropriately guide protons to the patient, and the space at the center of the gantry must be large enough to accommodate the patient as well as imaging equipment, which is crucial for the precise delivery of protons.
The incline beam system uses two beams, a horizontal beam and a second beam which is angled to 30 degrees off of the vertical. These beams use a common isocenter and can be used together, in conjunction with a robotic patient positioner, to achieve a wide array of angles to treat the patient (Figure 8. Procure). There are also fixed beams which can only deliver protons in a single direction. These beams rely on the movement of the treatment table or chair around the beam to allow multiple angles to be treated.
Nozzles are used to deliver the protons to the patient and are comprised of multiple components (Figure 9. IBA). There are two main types of proton delivery systems used at present, passive scatter and scanning beams. Passive scattering will be discussed first.
In a passive scatter system, the nozzle contains the above mentioned components including the scatter foils, ridge filter or modulator wheel, the aperture and the range compensator. In addition to spreading the proton field out to cover the depth of the tumor, the lateral aspects of the field also need to be expanded from the original thin pencil beam. A double scatter foil system can be used to broaden the proton beam laterally (green discs in figure 6). The first scatter foil used in this system is uniform and creates a Gaussian, or bell shaped, distribution of protons. However, this must then be made flat prior to reaching the patient. Hence, a second, non-uniform scatter foil is needed. The central protons are scattered to a greater degree compared with the protons in the lateral aspect of the beam, which acts to flatten the proton beam. Much like a flattening filter in photon therapy, any misalignment of the non-uniform, second scatter foil can create a skewed beam. In a single scatter foil system the single scatter foil causes a Gaussian distribution of proton energies, however, only the center portion, where the proton energies are within about 5% of the highest energy protons produced, can be used. This greatly limits the beam size. Additionally, the double scatter foil system allows less energy loss than the single scatter foil system. An aperture can then be used to further shape the lateral borders of the beam ('collimator' in figure 6). These are generally made of brass and can be quite heavy and difficult to manipulate. Both apertures and compensators are inserted into the 'snout' of the nozzle. There is currently a MLC system in development for protons at the University of Pennsylvania which would prevent the need for custom apertures. The MLC development process is described in greater detail here: Designing a Multileaf Collimator for Proton Therapy.
As described above, the depth to which a proton can treat can be modulated by the energy of the proton beam. Furthermore, by using protons of multiple energies, the entire tumor can be treated. However, the distal edge of the proton field is still not conformal to the distal border of the tumor. Without further shaping of the beam, you are left with a block-shaped field. A compensator (figure 6), usually made of wax or acrylic, is used to control where the dose goes along the distal edge of the tumor. These can be milled onsite (figure 10) However, by shaping the dose distally, some of the dose is 'pushed' proximally. This results in some areas proximal to the tumor, which we do not wish to be treated, receiving 100% of the dose (note the purple areas proximal to the tumor in figure 6). Compensators and apertures are only needed in a passive scattering beam and are not required for distal shaping with a scanning beam system.
There are several disadvantages to the passive scatter system. Due to the need for an aperture and compensator, the proton beam must travel through several layers, creating more neutrons. Neutrons contamination delivers unwanted dose to the patient. As described above, the use of a compensator can shift the full dose into areas proximal to the tumor, which receive unwanted high doses. Passive scatter treatments also require many custom components to be made for each patient. Custom compensators and apertures must be made for each field used to treat a patient. The apertures become radioactive after treatment and need to be stored until they can be safely disposed. All of these components also need to have quality and assurance checks performed which can be time consuming.
The second type of proton therapy which is just starting to be used in the United States is scanning beam proton therapy. Scanning beams use magnets to move the proton beam precisely, such that it can 'paint' the area that is to be treated. In a spot scanning system the nozzle contains the magnets needed to steer the proton beam. Once an area at a given depth has been treated, the energy of the proton can be changed and the next 'layer' can be painted. By repeating this, it is possible to treat the entire tumor. This technique allows greater conformality with shaping of the distal and proximal ends of the proton field. Scanning proton beams also allow the use of Intensity Modulated Proton Therapy (IMPT). With IMPT, multiple beams are used and a computer algorithm calculates the optimal arrangement of individual Bragg peaks needed to cover the volume. The sum of these Bragg peaks can provide a precise distribution of dose throughout the three- dimensional volume. Fewer neutrons are produced with scanning beams, as a compensator, scatter foil and aperture are not needed. The major disadvantage to the scanning beam is the greater complexity and longer treatment times due to the multiple 'layers' which must be 'painted'. There are also significant challenges to using scanning beams in areas of organ motion, as this technology is more susceptible to problems with motion.
Similar to photon-based therapy, the dose of radiation delivered to a patient with protons is measured in monitor units. These monitor units correspond to the known amount of charge collected in an ionization chamber present within the beam. The ion chamber plays the critical role of being the absolute reference monitor for the delivered dose. This also means that the dose prescription in gray must be converted to monitor units. The conversion factor needed to convert dose to monitor units is known as an output factor, and several models for output factor have been developed for photons. For protons, an actual measurement of the output for individual fields is required to accurately determine output factors. Numerous models, predominantly algorithms which use Monte Carlo calculations, to determine the output factor for a given treatment field are under development. Determination of the dose delivered by a proton beam is generally performed with an ionization chamber which has been specially designed for use with protons. Usually, ither a cylindrical ionization chamber or a plane-parallel chamber (for large depth dose gradients or narrow spread out Bragg peaks) is used. Scanning beams present additional challenges. The spot scanning beam requires close monitoring to ensure that each spot is receiving the correct dose. For scanning beams, generally two monitor chambers, which encompass the entire range of the scanning beam, are used to provide two independent measures of beam flux. The ionization chambers also monitor the dose delivered to a given spot.
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