Controlling intensity of a particle beam

Abstract

In an example, a synchrocyclotron includes a particle source to provide pulses of ionized plasma to a cavity; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column outwardly; and an extraction channel to receive a beam of particles from the cavity for output from the particle accelerator. The particle source is configured to control pulse widths of the ionized plasma in order to control an intensity of the beam of particles. This example synchrocyclotron may include one or more of the following features, either alone or in combination.

Description

TECHNICAL FIELD

This disclosure relates generally to controlling the intensity of a particle beam, such as a proton or ion beam used in a particle therapy system.

BACKGROUND

Particle therapy systems use an accelerator to generate a particle beam for treating afflictions, such as tumors. In operation, the particle beam is accelerated inside a cavity of the particle accelerator, and removed from the cavity through an extraction channel. Various elements are used to focus the particle beam and apply it to appropriate areas of a patent.

Different patients may require different doses and dose rates of particles. The dose and dose rate applied to a patient is a function of the intensity of the particle beam. Controlling the intensity of the particle beam therefore enables control over the dose and dose rate.

SUMMARY

In an example, a synchrocyclotron includes a particle source to provide pulses of ionized plasma to a cavity; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column outwardly; and an extraction channel to receive a beam of particles from the cavity for output from the particle accelerator. The particle source is configured to control pulse widths of the ionized plasma in order to control an intensity of the beam of particles. This example synchrocyclotron may include one or more of the following features, either alone or in combination.

The particle source may be configured to activate for a period of time in response to control signal, where the particle source generates a pulse of ionized plasma when activated. The particle source may be configured to generate pulses of ionized plasma periodically. The particle beam may be output for a duration of about 0.1 μs to 100 μs (e.g., 1 μs to 10 μs). The particle beam may be output for a duration of about 0.1 μs to 100 μs (e.g., 1 μs to 10 μs) about every 2 ms. The particle source may include cathodes to provide voltage to ionize hydrogen to produce the ionized plasma. The cathodes may be unheated by an external source.

In an example, a proton therapy system includes the foregoing synchrocyclotron, and a gantry on which the synchrocyclotron is mounted, where the gantry is rotatable relative to a patient position. Protons are output essentially directly from the synchrocyclotron to the patient position.

In an example, a synchrocyclotron includes a particle source to provide pulses of ionized plasma to a cavity, where the particle source includes cathodes to provide voltage to ionize hydrogen to produce the ionized plasma; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column outwardly; and an extraction channel to receive a beam of particles from the cavity for output from the particle accelerator. A voltage associated with the cathodes is controllable in order to control an intensity of the beam of particles. This example synchrocyclotron may include one or more of the following features, either alone or in combination.

The cathodes may be unheated by an external source. The voltage may be controllable such that increasing the voltage increases an intensity of the beam of particles and such that decreasing the voltage decreases the intensity of the beam of particles.

In an example, a proton therapy system includes the foregoing synchrocyclotron, and a gantry on which the synchrocyclotron is mounted. The gantry is rotatable relative to a patient position. Protons are output essentially directly from the synchrocyclotron to the patient position.

In an example, a synchrocyclotron includes a particle source to provide pulses of ionized plasma to a cavity, where the particle source includes cathodes to provide voltage to ionize hydrogen to produce the ionized plasma; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column outwardly; and an extraction channel to receive a beam of particles from the cavity for output from the particle accelerator. The particle source is controllable to adjust an amount of the hydrogen between the cathodes in order to control an intensity of the beam of particles. This example synchrocyclotron may include one or more of the following features, either alone or in combination.

The cathodes may be unheated by an external source. The amount of hydrogen may be adjustable such that increasing the amount of hydrogen increases an intensity of the beam of particles and such that decreasing the amount of hydrogen decreases the intensity of the beam of particles.

In an example, a proton therapy system includes the foregoing synchrocyclotron, and a gantry on which the synchrocyclotron is mounted. The gantry is rotatable relative to a patient position. Protons are output essentially directly from the synchrocyclotron to the patient position.

In an example, a synchrocyclotron includes a particle source to provide pulses of ionized plasma to a cavity; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column outwardly; and an extraction channel to receive a beam of particles from the cavity for output from the particle accelerator. The voltage source is controllable to control the RF voltage rate in order to control an intensity of the beam of particles. This example synchrocyclotron may include one or more of the following features, either alone or in combination.

The particle source may include cathodes to provide voltage to ionize hydrogen to produce the ionized plasma, where the cathodes are unheated by an external source. A magnitude of the RF voltage may be adjustable such that increasing the magnitude increases an intensity of the beam of particles and such that decreasing the magnitude decreases the intensity of the beam of particles.

In an example, a proton therapy system includes the foregoing synchrocyclotron, and a gantry on which the synchrocyclotron is mounted. The gantry is rotatable relative to a patient position. Protons are output essentially directly from the synchrocyclotron to the patient position.

In an example, a synchrocyclotron includes a particle source to provide pulses of ionized plasma to a cavity; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column outwardly, where the RF voltage sweeps between a maximum frequency and a minimum frequency; and an extraction channel to receive a beam of particles from the cavity for output from the particle accelerator. The particle source is controllable to provide pulses of the ionized plasma at specific frequencies proximate to a decrease from the maximum frequency of the RF voltage to the minimum frequency of the RF voltage. This example synchrocyclotron may include one or more of the following features, either alone or in combination.

The particle accelerator may be controllable to provide pulses of the ionized plasma between 132 MHz of RF voltage and 131 MHz of RF voltage from a decrease from a maximum frequency of about 135 MHz of the RF voltage. The particle source may include cathodes to provide voltage to ionize hydrogen to produce the ionized plasma. The cathodes may be unheated by an external source.

In an example, a proton therapy system includes the foregoing synchrocyclotron, and a gantry on which the synchrocyclotron is mounted. The gantry is rotatable relative to a patient position. Protons are output essentially directly from the synchrocyclotron to the patient position.

In an example, a synchrocyclotron includes a particle source to provide pulses of ionized plasma to a cavity; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column outwardly; and an extraction channel to receive a beam of particles from the cavity for output from the particle accelerator. The particle source is configured to selectively output pulses of the ionized plasma in order to control an intensity of the beam of particles. This example synchrocyclotron may include one or more of the following features, either alone or in combination.

The RF voltage may sweep periodically from a maximum frequency to a minimum frequency. Selectively outputting the pulses may include outputting pulses in certain ones of the RF voltage sweeps and not in others of the RF voltage sweeps. Selectively outputting the pulses may include skipping pulse output in every Nth (N>1) sweep.

The synchrocyclotron may include a controller for performing operations that include: determining the intensity of the beam of particles; and selectively outputting the pulses based on the determined intensity.

In an example, a proton therapy system includes the foregoing synchrocyclotron, and a gantry on which the synchrocyclotron is mounted. The gantry is rotatable relative to a patient position. Protons are output essentially directly from the synchrocyclotron to the patient position.

In an example, a synchrocyclotron includes a particle source to provide pulses of ionized plasma to a cavity; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column outwardly; and an extraction channel to receive a beam of particles from the cavity for output from the particle accelerator. The voltage source is configurable to vary a slope of the RF voltage in order to control an intensity of the beam of particles.

In an example, a proton therapy system includes the foregoing synchrocyclotron, and a gantry on which the synchrocyclotron is mounted. The gantry is rotatable relative to a patient position. Protons are output essentially directly from the synchrocyclotron to the patient position.

In an example, a synchrocyclotron includes a particle source to provide pulses of ionized plasma to a cavity; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column outwardly, where the voltage source includes a first dee and a second dee, and where at least one of the first dee and the second dee has a bias voltage applied thereto; and an extraction channel to receive a beam of particles from the cavity for output from the particle accelerator. This example synchrocyclotron may include one or more of the following features, either alone or in combination.

The first dee may have a first bias voltage applied thereto and the second dee may have a second bias voltage applied thereto, where the first bias voltage is different from the second bias voltage. The first dee may have the bias voltage applied thereto and the second dee may be electrically grounded.

In an example, a proton therapy system includes the foregoing synchrocyclotron, and a gantry on which the synchrocyclotron is mounted. The gantry is rotatable relative to a patient position. Protons are output essentially directly from the synchrocyclotron to the patient position.

In an example, a particle therapy system may include: a synchrocyclotron to output a particle beam comprised of pulses; and a scanning system for the synchrocyclotron to scan the particle beam across at least part of an irradiation target. The scanning system may be configured to scan the particle beam in two dimensions that are angled relative to (e.g., perpendicular to) a longitudinal direction of the particle beam. The particle beam makes a spot at the irradiation target. The synchrocyclotron is controllable to vary a width of the pulses so as to vary an intensity of the particle beam between different spots on the irradiation target during scanning. Implementations of the particle therapy system may include one or more of the following features, either alone or in combination.

The synchrocyclotron may include a particle source, and the particle source may be controllable to activate for periods of time to generate pulses of the particle beam that vary in width. The synchrocyclotron may be configured to sweep between low and high voltages, and a rate (or speed) of the voltage sweep may be controllable to vary a width of the pulses. The particle source may include first and second cathodes to generate a plasma stream from gas. The pulses of particle beam are extractable from the plasma stream. The gas may be a combination of hydrogen and less than 25% of a noble gas or a combination of hydrogen and less than 10% of a noble gas. The gas may be a combination of hydrogen and helium. The helium may be less than 25% of a composition of the gas. In another example, the helium may be than 10% of a composition of the gas.

The scanning system may include: a magnet to affect a direction of the particle beam to scan the particle beam in the two dimensions across at least part of the irradiation target; and a degrader to change an energy of the beam prior to output of the particle beam to the irradiation target. The degrader may be down-beam of the magnet relative to the synchrocyclotron.

The synchrocyclotron may include a voltage source to provide a radio frequency (RF) voltage to a cavity to accelerate particles from a plasma column, where the cavity has a magnetic field causing particles accelerated from the plasma column to move orbitally within the cavity; an extraction channel to receive the particles accelerated from the plasma column and to output the received particles from the cavity; and a regenerator to provide a magnetic field bump within the cavity to thereby change successive orbits of the particles accelerated from the plasma column so that, eventually, particles output to the extraction channel. The magnetic field may be between 4 Tesla (T) and 20 T (or between 6 T and 20 T) and the magnetic field bump may be at most 2 Tesla.

Two or more of the features described in this disclosure (e.g., two or more methods of controlling the intensity of a particle beam), including those described in this summary section, may be combined to form implementations not specifically described herein.

Control of the various systems described herein, or portions thereof, may be implemented via a computer program product that includes instructions that are stored on one or more non-transitory machine-readable storage media, and that are executable on one or more processing devices. The systems described herein, or portions thereof, may be implemented as an apparatus, method, or electronic system that may include one or more processing devices and memory to store executable instructions to implement control of the stated functions.

The details of one or more implementations are set forth in the accompanying drawings and the description below. Other features, objects, and advantages will be apparent from the description and drawings, and from the claims.

DESCRIPTION OF DRAWINGS

FIG. 1 is a perspective view of an example therapy system.

FIG. 2 is an exploded perspective view of components of an example synchrocyclotron.

FIGS. 3, 4, and 5 are cross-sectional views of an example synchrocyclotron.

FIG. 6 is a perspective view of an example synchrocyclotron.

FIG. 7 is a cross-sectional view of a portion of an example reverse bobbin and windings.

FIG. 8 is a cross sectional view of an example cable-in-channel composite conductor.

FIG. 9 is a cross-sectional view of an example particle source.

FIG. 10 is a perspective view of an example dee plate and an example dummy dee.

FIG. 11 is a perspective view of an example vault.

FIG. 12 is a perspective view of an example treatment room with a vault.

FIG. 13 shows a patient positioned relative to an example particle accelerator.

FIG. 14 shows a patient positioned within an example inner gantry in a treatment room.

FIGS. 15 and 16 show an example particle source.

FIG. 17 is a graph showing an example voltage sweep, extraction window, and particle source pulse width.

FIG. 18 is a perspective view of example active and dummy dees that may be used with the particle therapy system described herein.

FIG. 19 is a side view of an example scanning system.

FIG. 20 is a perspective view of the example scanning system.

FIGS. 21 and 22 are front and perspective views, respectively, of an example scanning magnet that may be used in the example scanning system.

FIG. 23 is a perspective view of an example range modulator that may be used in the example scanning system.

FIG. 24 is a perspective view of motion of a plate from the range modulator into/out of the beam path.

Like reference symbols in the various drawings indicate like elements.

DETAILED DESCRIPTION

Overview

Described herein is an example of a particle accelerator for use in an example system, such as a proton or ion therapy system. The system includes a particle accelerator—in this example, a synchrocyclotron—mounted on a gantry. The gantry enables the accelerator to be rotated around a patient position, as explained in more detail below. In some implementations, the gantry is steel and has two legs mounted for rotation on two respective bearings that lie on opposite sides of a patient. The particle accelerator is supported by a steel truss that is long enough to span a treatment area in which the patient lies and that is attached stably at both ends to the rotating legs of the gantry. As a result of rotation of the gantry around the patient, the particle accelerator also rotates.

In an example implementation, the particle accelerator (e.g., the synchrocyclotron) includes a cryostat that holds a superconducting coil for conducting a current that generates a magnetic field (B). In this example, the cryostat uses liquid helium (He) to maintain the coil at superconducting temperatures, e.g., 4° Kelvin (K). Magnetic yokes are adjacent (e.g., around) the cryostat, and define a cavity in which particles are accelerated. The cryostat is attached to the magnetic yokes through straps or the like.

In an example implementation, the particle accelerator includes a particle source (e.g., a Penning Ion Gauge—PIG source) to provide a plasma column to the cavity. Hydrogen gas is ionized to produce the plasma column. A voltage source provides a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column. As noted, in this example, the particle accelerator is a synchrocyclotron. Accordingly, the RF voltage is swept across a range of frequencies to account for relativistic effects on the particles (e.g., increasing particle mass) as their velocity increases during acceleration and the decreasing magnetic field produced to maintain axial focusing of the particles. The magnetic field produced by the coil causes particles accelerated from the plasma column to accelerate orbitally within the cavity. A magnetic field regenerator is positioned near the outer edge of the cavity and may be used to adjust the existing magnetic field at this location to thereby change locations of successive orbits of the particles accelerated from the plasma column so that, eventually, the particles output to an extraction channel that passes through the yokes. The extraction channel receives particles accelerated from the plasma column and outputs the received particles from the cavity. Elements both inside and outside the extraction channel shape and focus the particle beam.

A control system may be used to select the intensity of the particle beam. For example, one or more parameters or features of the particle accelerator may be controlled or otherwise adjusted to output the particle beam with the selected intensity. The selected intensity may be constant or variable. The example systems described herein use techniques to control the intensity of a particle beam, e.g., to vary the dose and dose rate of a particle beam delivered to a patient. A description of these techniques is provided below, followed by a description of an example particle therapy system in which those techniques may be implemented.

In an example technique, the intensity of the particle beam can be controlled by varying the time duration of the pulse of particle pulses extracted from the plasma column. In more detail, the RF voltage sweeps from a starting (e.g., maximum) frequency (e.g., 135 MHz) to an ending (e.g., minimum) frequency (e.g., 90 MHz). The particle source is activated for a period of time during the RF sweep to produce a plasma column. For example, in some implementations, the particle source is activated at 132 MHz for a period of time. During that time, particles are extracted from the plasma column by the electric field produced by the RF voltage. The extracted particles accelerate outwardly in orbits as the RF voltage frequency drops, keeping pace with the decreasing magnetic field and increasing relativistic mass until the particles are swept out a time (e.g., about 600 microseconds) later. Changing the duration for which the particle source is activated changes the width of the pulse of particles that is extracted from the plasma column during a frequency sweep. Increasing the pulse width causes an increase in the amount of particles extracted and thus an increase in the intensity of the particle beam. Decreasing the pulse width causes a decrease in the amount of particles extracted and thus a decrease in the intensity of the particle beam.

In another example technique, the intensity of the particle beam can be controlled by changing a voltage applied to cathodes in the particle source. In this regard, the plasma column is generated by applying a voltage to two cathodes of the particle source, and by outputting a gas, such as hydrogen (H₂), in the vicinity of the cathodes. The voltage applied to the cathodes ionizes the hydrogen and the background magnetic field collimates the ionized hydrogen to thereby produce the plasma column. Increasing the cathode voltage causes an increase in the amount of ions in the plasma column, and decreasing the cathode voltage causes a decrease in the amount of ions in the plasma column. When more ions are present in the plasma column, more ions can be extracted during the RF voltage sweep, thereby increasing the intensity of the particle beam. When fewer ions are present in the plasma column, fewer ions can be extracted during the RF voltage sweep, thereby decreasing the intensity of the particle beam.

In another example technique, the intensity of the particle beam can be controlled by varying the amount of hydrogen supplied to the particle source. For example, increasing the amount of hydrogen supplied to the particle source results in more opportunity for ionization in the plasma column in response to the cathode voltage. Conversely, decreasing the amount of hydrogen supplied to the particle source results in less opportunity for ionization in the plasma column in response to the cathode voltage. As noted above, when more particles are present in the plasma column, more particles are extracted during the RF voltage sweep, thereby increasing the intensity of the particle beam. When fewer particles are present in the plasma column, fewer particles are extracted during the RF voltage sweep, thereby decreasing the intensity of the particle beam.

In another example technique, the intensity of the particle beam can be controlled by varying the magnitude of the RF voltage used to extract particles from the plasma column. For example, increasing the magnitude of the RF voltage causes more particles to be extracted from the plasma column. Conversely, decreasing the magnitude of the RF voltage causes fewer particles to be extracted from the plasma column. When more particles are extracted, the particle beam increases in intensity. Conversely, when fewer particles are extracted, the particle beam decreases in intensity.

In another example technique, the intensity of the particle beam can be controlled by varying the starting time during the frequency sweep at which the particle source is activated and, thus, during which particles are extracted. More specifically, there is a finite window during the frequency sweep during which particles can be extracted from the plasma column. In an example implementation, the frequency sweeps from about 135 MHz to about 90 MHz at a substantially constant rate. In this example, particles can be extracted at about the beginning of the downward slope between starting and ending frequencies, e.g., between 132 MHz and 131 MHz respectively, and the particle source can be activated for a period of time, e.g., for about 0.1 μs to 100 μs (or e.g., 1 μs to 10 μs up to about 40 μs). Changing the frequency at which the particle source is activated affects the amount of particles that are extracted from the particle beam and therefore the intensity of the particle beam.

In another example technique, pulse blanking may be used to control the intensity of the particle beam. In this regard, the frequency sweep is repeated a number of times per second (e.g., 500 times/second). The particle source could be activated for each frequency sweep (e.g., every 2 ms). Pulse blanking reduces the number of particles extracted from the particle beam by not activating the particle source during every frequency sweep. To achieve maximum beam intensity, the particle source may be activated every frequency sweep. To reduce beam intensity, the particle source may be activated less frequently, e.g., every second, third, hundredth, etc. sweep.

In another example technique, the intensity of the particle beam can be controlled by applying a DC bias voltage to one or more dees used to apply the RF voltage to the particle accelerator cavity. In this regard, the particle accelerator includes an active dee plate (or simply “dee”) that is a hollow metal structure having two semicircular surfaces that enclose a cavity in which the protons are accelerated during their rotation around the space enclosed by the magnet structure. The active dee is driven by a RF signal that is applied at the end of a radio-frequency transmission line to impart an electric field into the cavity. The RF field is made to vary in time as the accelerated particle beam increases in distance from the geometric center. A “dummy” dee comprises a rectangular metal wall with a slot for the beam that is spaced near to the exposed rim of the active dee. In some implementations, the dummy dee is grounded to the vacuum chamber and magnet yoke.

Applying RF voltage in the presence of a strong magnetic field can cause multi-pactoring, which can reduce the magnitude of the RF field and, in some cases, cause an electrical short. To reduce the amount of multi-pactoring, and thereby maintain the RF field, DC bias voltage may be applied to the active dee and, in some implementations, also to the dummy dee. In some implementations, the differential bias voltage between the active dee and dummy dee may be controlled to reduce multi-pactoring and thereby increase beam intensity. For example, in some implementations, there may be a 50% differential between the DC bias voltage on the active dee and dummy dee (e.g., a −1.9 KV DC bias voltage may be applied to the dummy dee and a −1.5 KV DC bias voltage may be applied to the active dee).

In another example technique, the intensity of the particle beam can be controlled by controlling the rate at which the RF voltage is swept (e.g., the slope of the decrease). By decreasing the slope, it is possible to increase the amount of time during which particles can be extracted from the plasma column. As a result, more particles can be extracted, thereby increasing the intensity of the particle beam. The converse is also true, e.g., by increasing the slope, the amount of time during which particles can be extracted from the plasma column can be decreased, which can result in a decrease in particle beam intensity.

The foregoing techniques for controlling the intensity of a particle beam in a particle accelerator may be used individually in a single particle accelerator, or any two or more of those techniques may be used in any appropriate combination in a single particle accelerator. The techniques are not limited to use with a particle therapy system, but rather may be used in any appropriate particle accelerator.

An example of a particle therapy system in which the foregoing techniques may be used is provided below.

Example Particle Therapy System

Referring to FIG. 1, a charged particle radiation therapy system 500 includes a beam-producing particle accelerator 502 having a weight and size small enough to permit it to be mounted on a rotating gantry 504 with its output directed straight (that is, essentially directly) from the accelerator housing toward a patient 506.

In some implementations, the steel gantry has two legs 508, 510 mounted for rotation on two respective bearings 512, 514 that lie on opposite sides of the patient. The accelerator is supported by a steel truss 516 that is long enough to span a treatment area 518 in which the patient lies (e.g., twice as long as a tall person, to permit the person to be rotated fully within the space with any desired target area of the patient remaining in the line of the beam) and is attached stably at both ends to the rotating legs of the gantry.

In some examples, the rotation of the gantry is limited to a range 520 of less than 360 degrees, e.g., about 180 degrees, to permit a floor 522 to extend from a wall of the vault 524 that houses the therapy system into the patient treatment area. The limited rotation range of the gantry also reduces the required thickness of some of the walls, which provide radiation shielding of people outside the treatment area. A range of 180 degrees of gantry rotation is enough to cover all treatment approach angles, but providing a larger range of travel can be useful. For example the range of rotation may be between 180 and 330 degrees and still provide clearance for the therapy floor space.

The horizontal rotational axis 532 of the gantry is located nominally one meter above the floor where the patient and therapist interact with the therapy system. This floor is positioned about 3 meters above the bottom floor of the therapy system shielded vault. The accelerator can swing under the raised floor for delivery of treatment beams from below the rotational axis. The patient couch moves and rotates in a substantially horizontal plane parallel to the rotational axis of the gantry. The couch can rotate through a range 534 of about 270 degrees in the horizontal plane with this configuration. This combination of gantry and patient rotational ranges and degrees of freedom allow the therapist to select virtually any approach angle for the beam. If needed, the patient can be placed on the couch in the opposite orientation and then all possible angles can be used.

In some implementations, the accelerator uses a synchrocyclotron configuration having a very high magnetic field superconducting electromagnetic structure. Because the bend radius of a charged particle of a given kinetic energy is reduced in direct proportion to an increase in the magnetic field applied to it, the very high magnetic field superconducting magnetic structure permits the accelerator to be made smaller and lighter. The synchrocyclotron uses a magnetic field that is uniform in rotation angle and falls off in strength with increasing radius. Such a field shape can be achieved regardless of the magnitude of the magnetic field, so in theory there is no upper limit to the magnetic field strength (and therefore the resulting particle energy at a fixed radius) that can be used in a synchrocyclotron.

Superconducting materials lose their superconducting properties in the presence of very high magnetic fields. High performance superconducting wire windings are used to allow very high magnetic fields to be achieved.

Superconducting materials typically need to be cooled to low temperatures for their superconducting properties to be realized. In some examples described here, cryo-coolers are used to bring the superconducting coil windings to temperatures near absolute zero. Using cryo-coolers can reduce complexity and cost.

The synchrocyclotron is supported on the gantry so that the beam is generated directly in line with the patient. The gantry permits rotation of the cyclotron about a horizontal rotational axis that contains a point (isocenter 540) within, or near, the patient. The split truss that is parallel to the rotational axis, supports the cyclotron on both sides.

Because the rotational range of the gantry is limited, a patient support area can be accommodated in a wide area around the isocenter. Because the floor can be extended broadly around the isocenter, a patient support table can be positioned to move relative to and to rotate about a vertical axis 542 through the isocenter so that, by a combination of gantry rotation and table motion and rotation, any angle of beam direction into any part of the patient can be achieved. The two gantry arms are separated by more than twice the height of a tall patient, allowing the couch with patient to rotate and translate in a horizontal plane above the raised floor.

Limiting the gantry rotation angle allows for a reduction in the thickness of at least one of the walls surrounding the treatment room. Thick walls, typically constructed of concrete, provide radiation protection to individuals outside the treatment room. A wall downstream of a stopping proton beam may be about twice as thick as a wall at the opposite end of the room to provide an equivalent level of protection. Limiting the range of gantry rotation enables the treatment room to be sited below earth grade on three sides, while allowing an occupied area adjacent to the thinnest wall reducing the cost of constructing the treatment room.

In the example implementation shown in FIG. 1, the superconducting synchrocyclotron 502 operates with a peak magnetic field in a pole gap of the synchrocyclotron of 8.8 Tesla. The synchrocyclotron produces a beam of protons having an energy of 250 MeV. In other implementations the field strength could be in the range of 6 to 20 Tesla or 4 to 20 Tesla and the proton energy could be in the range of 150 to 300 MeV

The radiation therapy system described in this example is used for proton radiation therapy, but the same principles and details can be applied in analogous systems for use in heavy ion (ion) treatment systems.

As shown in FIGS. 2, 3, 4, 5, and 6, an example synchrocyclotron 10 (e.g., 502 in FIG. 1) includes a magnet system 12 that contains an particle source 90, a radiofrequency drive system 91, and a beam extraction system 38. The magnetic field established by the magnet system has a shape appropriate to maintain focus of a contained proton beam using a combination of a split pair of annular superconducting coils 40, 42 and a pair of shaped ferromagnetic (e.g., low carbon steel) pole faces 44, 46.

The two superconducting magnet coils are centered on a common axis 47 and are spaced apart along the axis. As shown in FIGS. 7 and 8, the coils are formed by of Nb₃Sn-based superconducting 0.8 mm diameter strands 48 (that initially comprise a niobium-tin core surrounded by a copper sheath) deployed in a twisted cable-in-channel conductor geometry. After seven individual strands are cabled together, they are heated to cause a reaction that forms the final (brittle) superconducting material of the wire. After the material has been reacted, the wires are soldered into the copper channel (outer dimensions 3.18×2.54 mm and inner dimensions 2.08×2.08 mm) and covered with insulation 52 (in this example, a woven fiberglass material). The copper channel containing the wires 53 is then wound in a coil having a rectangular cross-section of 8.55 cm×19.02 cm, having 26 layers and 49 turns per layer. The wound coil is then vacuum impregnated with an epoxy compound. The finished coils are mounted on an annular stainless steel reverse bobbin 56. Heater blankets 55 are placed at intervals in the layers of the windings to protect the assembly in the event of a magnet quench.

The entire coil can then be covered with copper sheets to provide thermal conductivity and mechanical stability and then contained in an additional layer of epoxy. The precompression of the coil can be provided by heating the stainless steel reverse bobbin and fitting the coils within the reverse bobbin. The reverse bobbin inner diameter is chosen so that when the entire mass is cooled to 4 K, the reverse bobbin stays in contact with the coil and provides some compression. Heating the stainless steel reverse bobbin to approximately 50 degrees C. and fitting coils at a temperature of 100 degrees Kelvin can achieve this.

The geometry of the coil is maintained by mounting the coils in a reverse rectangular bobbin 56 to exert a restorative force 60 that works against the distorting force produced when the coils are energized. As shown in FIG. 5, the coil position is maintained relative to the magnet yoke and cryostat using a set of warm-to-cold support straps 402, 404, 406. Supporting the cold mass with thin straps reduces the heat leakage imparted to the cold mass by the rigid support system. The straps are arranged to withstand the varying gravitational force on the coil as the magnet rotates on board the gantry. They withstand the combined effects of gravity and the large de-centering force realized by the coil when it is perturbed from a perfectly symmetric position relative to the magnet yoke. Additionally the links act to reduce dynamic forces imparted on the coil as the gantry accelerates and decelerates when its position is changed. Each warm-to-cold support includes one S2 fiberglass link and one carbon fiber link. The carbon fiber link is supported across pins between the warm yoke and an intermediate temperature (50-70 K), and the S2 fiberglass link 408 is supported across the intermediate temperature pin and a pin attached to the cold mass. Each link is 5 cm long (pin center to pin center) and is 17 mm wide. The link thickness is 9 mm. Each pin is made of high strength stainless steel and is 40 mm in diameter.

Referring to FIG. 3, the field strength profile as a function of radius is determined largely by choice of coil geometry and pole face shape; the pole faces 44, 46 of the permeable yoke material can be contoured to fine tune the shape of the magnetic field to ensure that the particle beam remains focused during acceleration.

The superconducting coils are maintained at temperatures near absolute zero (e.g., about 4 degrees Kelvin) by enclosing the coil assembly (the coils and the bobbin) inside an evacuated annular aluminum or stainless steel cryostatic chamber 70 that provides a free space around the coil structure, except at a limited set of support points 71, 73. In an alternate version (FIG. 4) the outer wall of the cryostat may be made of low carbon steel to provide an additional return flux path for the magnetic field.

In some implementations, the temperature near absolute zero is achieved and maintained using one single-stage Gifford-McMahon cryo-cooler and three two-stage Gifford McMahon cryo-coolers. Each two stage cryo-cooler has a second stage cold end attached to a condenser that recondenses Helium vapor into liquid Helium. The cryo-cooler heads are supplied with compressed Helium from a compressor. The single-stage Gifford-McMahon cryo-cooler is arranged to cool high temperature (e.g., 50-70 degrees Kelvin) leads that supply current to the superconducting windings.

In some implementations, the temperature near absolute zero is achieved and maintained using two Gifford-McMahon cryo-coolers 72, 74 that are arranged at different positions on the coil assembly. Each cryo-cooler has a cold end 76 in contact with the coil assembly. The cryo-cooler heads 78 are supplied with compressed Helium from a compressor 80. Two other Gifford-McMahon cryo-coolers 77, 79 are arranged to cool high temperature (e.g., 60-80 degrees Kelvin) leads that supply current to the superconducting windings.

The coil assembly and cryostatic chambers are mounted within and fully enclosed by two halves 81, 83 of a pillbox-shaped magnet yoke 82. In this example, the inner diameter of the coil assembly is about 74.6 cm. The iron yoke 82 provides a path for the return magnetic field flux 84 and magnetically shields the volume 86 between the pole faces 44, 46 to prevent external magnetic influences from perturbing the shape of the magnetic field within that volume. The yoke also serves to decrease the stray magnetic field in the vicinity of the accelerator. In some implementations, the synchrocyclotron may have an active return system to reduce stray magnetic fields. An example of an active return system is described in U.S. patent application Ser. No. 13/907,601, which was filed on May 31, 2013, the contents of which are incorporated herein by reference. In the active return system, the relatively large magnetic yokes described herein are replaced by smaller magnetic structures, referred to as pole pieces. Superconducting coils run current opposite to the main coils described herein in order to provide magnetic return and thereby reduce stray magnetic fields

As shown in FIGS. 3 and 9, the synchrocyclotron includes a particle source 90 of a Penning ion gauge geometry located near the geometric center 92 of the magnet structure 82. The particle source may be as described below, or the particle source may be of the type described in U.S. patent application Ser. No. 11/948,662 incorporated herein by reference.

Particle source 90 is fed from a supply 99 of hydrogen through a gas line 101 and tube 194 that delivers gaseous hydrogen. Electric cables 94 carry an electric current from a current source 95 to stimulate electron discharge from cathodes 192, 190 that are aligned with the magnetic field, 200.

In this example, the discharged electrons ionize the gas exiting through a small hole from tube 194 to create a supply of positive ions (protons) for acceleration by one semicircular (dee-shaped) radio-frequency plate 100 that spans half of the space enclosed by the magnet structure and one dummy dee plate 102. In the case of an interrupted particle source (an example of which is described in U.S. patent application Ser. No. 11/948,662), all (or a substantial part) of the tube containing plasma is removed at the acceleration region, thereby allowing ions to be more rapidly accelerated in a relatively high magnetic field.

As shown in FIG. 10, the dee plate 100 is a hollow metal structure that has two semicircular surfaces 103, 105 that enclose a space 107 in which the protons are accelerated during half of their rotation around the space enclosed by the magnet structure. A duct 109 opening into the space 107 extends through the yoke to an external location from which a vacuum pump 111 can be attached to evacuate the space 107 and the rest of the space within a vacuum chamber 119 in which the acceleration takes place. The dummy dee 102 comprises a rectangular metal ring that is spaced near to the exposed rim of the dee plate. The dummy dee is grounded to the vacuum chamber and magnet yoke. The dee plate 100 is driven by a radio-frequency signal that is applied at the end of a radio-frequency transmission line to impart an electric field in the space 107. The radio frequency electric field is made to vary in time as the accelerated particle beam increases in distance from the geometric center. The radio frequency electric field may be controlled in the manner described in U.S. patent application Ser. No. 11/948,359, entitled “Matching A Resonant Frequency Of A Resonant Cavity To A Frequency Of An Input Voltage”, the contents of which are incorporated herein by reference.

For the beam emerging from the centrally located particle source to clear the particle source structure as it begins to spiral outward, a large voltage difference is required across the radio frequency plates. 20,000 Volts is applied across the radio frequency plates. In some versions from 8,000 to 20,000 Volts may be applied across the radio frequency plates. To reduce the power required to drive this large voltage, the magnet structure is arranged to reduce the capacitance between the radio frequency plates and ground. This is done by forming holes with sufficient clearance from the radio frequency structures through the outer yoke and the cryostat housing and making sufficient space between the magnet pole faces.

The high voltage alternating potential that drives the dee plate has a frequency that is swept downward during the accelerating cycle to account for the increasing relativistic mass of the protons and the decreasing magnetic field. The dummy dee does not require a hollow semi-cylindrical structure as it is at ground potential along with the vacuum chamber walls. Other plate arrangements could be used such as more than one pair of accelerating electrodes driven with different electrical phases or multiples of the fundamental frequency. The RF structure can be tuned to keep the Q high during the required frequency sweep by using, for example, a rotating capacitor having intermeshing rotating and stationary blades. During each meshing of the blades, the capacitance increases, thus lowering the resonant frequency of the RF structure. The blades can be shaped to create a precise frequency sweep required. A drive motor for the rotating condenser can be phase locked to the RF generator for precise control. One bunch of particles is accelerated during each meshing of the blades of the rotating condenser.

The vacuum chamber 119 in which the acceleration occurs is a generally cylindrical container that is thinner in the center and thicker at the rim. The vacuum chamber encloses the RF plates and the particle source and is evacuated by the vacuum pump 111. Maintaining a high vacuum insures that accelerating ions are not lost to collisions with gas molecules and enables the RF voltage to be kept at a higher level without arcing to ground.

Protons traverse a generally spiral orbital path beginning at the particle source. In half of each loop of the spiral path, the protons gain energy as they pass through the RF electric field in space 107. As the ions gain energy, the radius of the central orbit of each successive loop of their spiral path is larger than the prior loop until the loop radius reaches the maximum radius of the pole face. At that location a magnetic and electric field perturbation directs ions into an area where the magnetic field rapidly decreases, and the ions depart the area of the high magnetic field and are directed through an evacuated tube 38, referred to herein as the extraction channel, to exit the yoke of the cyclotron. A magnetic regenerator may be used to change the magnetic field perturbation to direct the ions. The ions exiting the cyclotron will tend to disperse as they enter the area of markedly decreased magnetic field that exists in the room around the cyclotron. Beam shaping elements 107, 109 in the extraction channel 38 redirect the ions so that they stay in a straight beam of limited spatial extent.

The magnetic field within the pole gap needs to have certain properties to maintain the beam within the evacuated chamber as it accelerates. The magnetic field index n, which is shown below,n=−(r/B)dB/dr, should be kept positive to maintain this “weak” focusing. Here r is the radius of the beam and B is the magnetic field. Additionally, in some implementations, the field index needs to be maintained below 0.2, because at this value the periodicity of radial oscillations and vertical oscillations of the beam coincide in a vr=2 v_z resonance. The betatron frequencies are defined by v_r=(1−n)^1/2 and v_z=n^1/2. The ferromagnetic pole face is designed to shape the magnetic field generated by the coils so that the field index n is maintained positive and less than 0.2 in the smallest diameter consistent with a 250 MeV beam in the given magnetic field.

As the beam exits the extraction channel it is passed through a beam formation system 125 (FIG. 5) that can be programmably controlled to create a desired combination of scanning, scattering, and/or range modulation for the beam. Beam formation system 125 may be used in conjunction with an inner gantry 601 (FIG. 14) to direct a beam to the patient.

During operation, the plates absorb energy from the applied radio frequency field as a result of conductive resistance along the surfaces of the plates. This energy appears as heat and is removed from the plates using water cooling lines 108 that release the heat in a heat exchanger 113 (FIG. 3).

Stray magnetic fields exiting from the cyclotron are limited by both the pillbox magnet yoke (which also serves as a shield) and a separate magnetic shield 114. The separate magnetic shield includes of a layer 117 of ferromagnetic material (e.g., steel or iron) that encloses the pillbox yoke, separated by a space 116. This configuration that includes a sandwich of a yoke, a space, and a shield achieves adequate shielding for a given leakage magnetic field at lower weight.

As mentioned, the gantry allows the synchrocyclotron to be rotated about the horizontal rotational axis 532. The truss structure 516 has two generally parallel spans 580, 582. The synchrocyclotron is cradled between the spans about midway between the legs. The gantry is balanced for rotation about the bearings using counterweights 122, 124 mounted on ends of the legs opposite the truss.

The gantry is driven to rotate by an electric motor mounted to one or both of the gantry legs and connected to the bearing housings by drive gears. The rotational position of the gantry is derived from signals provided by shaft angle encoders incorporated into the gantry drive motors and the drive gears.

At the location at which the ion beam exits the cyclotron, the beam formation system 125 acts on the ion beam to give it properties suitable for patient treatment. For example, the beam may be spread and its depth of penetration varied to provide uniform radiation across a given target volume. The beam formation system can include passive scattering elements as well as active scanning elements.

All of the active systems of the synchrocyclotron (the current driven superconducting coils, the RF-driven plates, the vacuum pumps for the vacuum acceleration chamber and for the superconducting coil cooling chamber, the current driven particle source, the hydrogen gas source, and the RF plate coolers, for example), may be controlled by appropriate synchrocyclotron control electronics (not shown), which may include, e.g., one or more computers programmed with appropriate programs to effect control.

The control of the gantry, the patient support, the active beam shaping elements, and the synchrocyclotron to perform a therapy session is achieved by appropriate therapy control electronics (not shown).

As shown in FIGS. 1, 11, and 12, the gantry bearings are supported by the walls of a cyclotron vault 524. The gantry enables the cyclotron to be swung through a range 520 of 180 degrees (or more) including positions above, to the side of, and below the patient. The vault is tall enough to clear the gantry at the top and bottom extremes of its motion. A maze 146 sided by walls 148, 150 provides an entry and exit route for therapists and patients. Because at least one wall 152 is not in line with the proton beam directly from the cyclotron, it can be made relatively thin and still perform its shielding function. The other three side walls 154, 156, 150/148 of the room, which may need to be more heavily shielded, can be buried within an earthen hill (not shown). The required thickness of walls 154, 156, and 158 can be reduced, because the earth can itself provide some of the needed shielding.

Referring to FIGS. 12 and 13, for safety and aesthetic reasons, a therapy room 160 may be constructed within the vault. The therapy room is cantilevered from walls 154, 156, 150 and the base 162 of the containing room into the space between the gantry legs in a manner that clears the swinging gantry and also maximizes the extent of the floor space 164 of the therapy room. Periodic servicing of the accelerator can be accomplished in the space below the raised floor. When the accelerator is rotated to the down position on the gantry, full access to the accelerator is possible in a space separate from the treatment area. Power supplies, cooling equipment, vacuum pumps and other support equipment can be located under the raised floor in this separate space. Within the treatment room, the patient support 170 can be mounted in a variety of ways that permit the support to be raised and lowered and the patient to be rotated and moved to a variety of positions and orientations.

In system 602 of FIG. 14, a beam-producing particle accelerator of the type described herein, in this case synchrocyclotron 604, is mounted on rotating gantry 605. Rotating gantry 605 is of the type described herein, and can angularly rotate around patient support 606. This feature enables synchrocyclotron 604 to provide a particle beam directly to the patient from various angles. For example, as in FIG. 14, if synchrocyclotron 604 is above patient support 606, the particle beam may be directed downwards toward the patient. Alternatively, if synchrocyclotron 604 is below patient support 606, the particle beam may be directed upwards toward the patient. The particle beam is applied directly to the patient in the sense that an intermediary beam routing mechanism is not required. A routing mechanism, in this context, is different from a shaping or sizing mechanism in that a shaping or sizing mechanism does not re-route the beam, but rather sizes and/or shapes the beam while maintaining the same general trajectory of the beam.

Further details regarding an example implementation of the foregoing system may be found in U.S. Pat. No. 7,728,311, filed on Nov. 16, 2006 and entitled “Charged Particle Radiation Therapy”, and in U.S. patent application Ser. No. 12/275,103, filed on Nov. 20, 2008 and entitled “Inner Gantry”. The contents of U.S. Pat. No. 7,728,311 and in U.S. patent application Ser. No. 12/275,103 are hereby incorporated by reference into this disclosure.

In some implementations, the synchrocyclotron may be a variable-energy device, such as that described in U.S. patent application Ser. No. 13/916,401, filed on Jun. 12, 2013, the contents of which are incorporated by reference into this disclosure.

Example Implementations

Referring to FIG. 3, particle source 90 is deployed near to the magnetic center of synchrocyclotron 10 so that particles are present at the synchrocyclotron mid-plane, where they can be acted upon by the RF voltage field. As noted above, the particle source may have a Penning ion gauge (PIG) geometry. In the PIG geometry, two high voltage cathodes are placed about opposite each other so that they are aligned linearly. For example, one cathode may be on one side of the acceleration region and one cathode may be on the other side of the acceleration region and in line with the magnetic field lines. A gas tube 101 extends toward the acceleration region proximate to the particle source. When a relatively small amount of a gas (e.g., hydrogen/H₂) occupies a region in the tube between the cathodes, a plasma column may be formed from the gas by applying a voltage to the cathodes. The applied voltage causes electrons to stream along the magnetic field lines, essentially parallel to the tube walls, and to ionize gas molecules that are concentrated inside the tube. The background magnetic field prevents scattering of the ionized gas particles and creates a plasma column between the cathodes.

In some implementations, the gas in gas tube 101 may include a mixture of hydrogen and one or more other gases. For example, the mixture may contain hydrogen and one or more of the noble gases, e.g., helium, neon, argon, krypton, xenon and/or radon (although the mixture is not limited to use with the noble gases). In some implementations, the mixture may be a mixture of hydrogen and helium. For example, the mixture may contain about 75% or more of hydrogen and about 25% or less of helium (with possible trace gases included). In another example, the mixture may contain about 90% or more of hydrogen and about 10% or less of helium (with possible trace gases included). In examples, the hydrogen/helium mixture may be any of the following: >95%/<5%, >90%/<10%, >85%/<15%, >80%/<20%, >75%/<20%, and so forth.

Possible advantages of using a noble (or other) gas in combination with hydrogen in the particle source may include: increased beam intensity, increased cathode longevity, and increased consistency of beam output.

An example of a PIG geometry particle source 700 that may be used in synchrocyclotron 10 is shown in FIGS. 15 and 16. Referring to FIG. 15, particle source 700 includes an emitter side 701 containing a gas feed 702 for receiving gas (e.g., hydrogen (H₂), and a reflector side 704. A housing, or tube, 706 holds the gas. FIG. 16 shows particle source 700 passing through dummy dee 710 and adjacent to active (RF) dee 711. In operation, the magnetic field between active dee 711 and dummy dee 710 causes particles (e.g., protons) to accelerate outwardly. The acceleration is spiral to create orbits about the plasma column, with the particle-to-plasma-column radius progressively increasing. The radii of curvature of the spirals depend on a particle's mass, energy imparted to the particle by the RF field, and a strength of the magnetic field.

When the magnetic field is high, it can become difficult to impart enough energy to a particle so that it has a large enough radius of curvature to clear the physical housing of the particle source on its initial turn(s) during acceleration. The magnetic field is relatively high in the region of the particle source, e.g., on the order of 2 Tesla (T) or more (e.g., 4 T, 5 T, 6 T, 8 T, 8.8 T, 8.9 T, 9 T, 10.5 T, or more). As a result of this relatively high magnetic field, the initial particle-to-ion-source radius is relatively small for low energy particles, where low energy particles include particles that are first drawn from the plasma column. For example, such a radius may be on the order of 1 mm. Because the radii are so small, at least initially, some particles may come into contact with the particle source's housing area, thereby preventing further outward acceleration of such particles. Accordingly, the housing of particle source 700 is interrupted, or separated to form two parts, as shown in FIG. 16. That is, a portion of the particle source's housing may be entirely removed at the acceleration region 714, e.g., at about the point where the particles are to be drawn from the particle source. This interruption is labeled 715 in FIG. 16. The housing may also be removed for distances above, and below, the acceleration region. In an alternative implementation, a substantial portion (e.g., 30%, 40%, 50% or more), but not all, of the PIG source housing is removed, leaving the plasma beam partly exposed. Thus, portions of the PIG housing are separated from their counterpart portions, but there is not complete separation as was the case above.

In the synchrocyclotron described herein, a particle beam is extracted using a resonant extraction system. That is, radial oscillation of the beam is created by a magnetic perturbation inside the accelerator, which establishes a resonance of these oscillations. When a resonant extraction system is used, extraction efficiency is improved by limiting the phase space extent of the internal beam. With attention to the design of the magnetic and RF field generating structures, the phase space extent of the beam at extraction is determined by the phase space extent at the beginning of acceleration (e.g., at emergence from the particle source). As a result, relatively little beam may be lost at the entrance to the extraction channel and background radiation from the accelerator can be reduced.

Cathodes 717 may be “cold” cathodes. A cold cathode may be a cathode that is not heated by an external heat source. Also, the cathodes may be pulsed, meaning that they output plasma burst(s) periodically rather than continuously. When the cathodes are cold, and are pulsed, the cathodes are less subject to wear and can therefore last relatively long. Furthermore, pulsing the cathodes can eliminate the need to water-cool the cathodes. In one implementation, cathodes 717 pulse at a relatively high voltage, e.g., about 1 kV to about 4 kV, and moderate peak cathode discharge currents of about 50 mA to about 200 mA at a duty cycle between about 0.1% and about 1% or 2% at repetition rates between about 200 Hz to about 1 KHz. However, the particle source is not limited to these values.

Various aspects of example particle therapy system described herein may be computer-controlled. Computer controls may be effected through one or more signals output from the computer to various electronics on the particle therapy system. For example, the intensity of the particle beam generated by the particle therapy system may be measured, and the particle therapy system may be adjusted to control the intensity of the particle beam. The measurement and adjustment may occur once, at each use of the particle therapy system, in real-time (e.g., during treatment), or at other frequencies. In some implementations, the various parameters or other features of the particle accelerator described below may be varied, and the intensity of the resulting particle beam measured in order to determine if the appropriate result was achieved. If the appropriate result was not achieved, then the parameters or other features may be varied again and the results measured until the appropriate results are achieved.

In an example implementation, the time-width of pulses output by the particle source may be varied to control the intensity of the particle beam. In other words, the amount of time that the particle source is intermittently (e.g., periodically) activated is varied, thereby providing the plasma column for different periods of time and enabling extraction of different numbers of particles. For example, if the pulse width is increased, the number of particles extracted increases and, if the pulse width decreases, the number of particles extracted decreases. In some implementations, there is a linear relationship between the time that the particle source is on and the intensity of the particle beam. For example, the relationship may be one-to-one plus an offset. In an example implementation, the particle source may be pulsed within a frequency window that occurs during a frequency sweep between a maximum frequency of about 135 MHz and a minimum frequency of about 95 MHz or 90 MHz. For example, the particle source may be pulsed between 132 MHz and 131 MHz for a period of time. In an implementation, this period of time is about 40 μs; however, these values may vary or be different in other implementations. Failing to pulse the particle source outside of the frequency window can inhibit extraction of particles from the plasma column.

FIG. 17 is a graph showing the voltage sweep in the resonant cavity over time from a maximum frequency (e.g., 135 MHz) to a minimum frequency (e.g., 90 MHz or 95 MHz). The extraction window 720 occurs, in this example, between 132 MHz and 131 MHz. The width of pulse 721 may be varied to control the intensity of the particle beam output by the particle accelerator.

In other implementations, the voltage of cathodes 717 may be adjusted in order to control the amount of ionization in the plasma column, and thereby control the intensity of the particle beam output from the accelerator. Varying the voltage of cold cathodes can produce particularly sharp pulse edges.

In other implementations, the gas flow in tube 101 may be adjusted to increase or decrease the amount of hydrogen in the plasma column. As explained above, this increase or decrease in hydrogen can cause an increase or decrease in the amount of particles in the plasma column that are available for extraction. Thus, by varying the amount/flow of hydrogen provided by the particle source, it is possible to control the amount of particles available for extraction and thus the intensity of the resulting particle beam. For example, as explained above, when more particles are present in the plasma column, more particles are extracted during the RF voltage sweep, thereby increasing the intensity of the particle beam. When fewer particles are present in the plasma column, fewer particles are extracted during the RF voltage sweep, thereby decreasing the intensity of the particle beam.

In other implementations, increasing the magnitude of the RF voltage during the extraction period can increase the amount of particles extracted and thereby increase the intensity of the particle beam. In this regard, the magnitude of the RF voltage may be varied during the entire period of the RF sweep or it may be varied only during times that particles are extractable from the plasma column. For example, in some implementations, particles are extracted from the plasma column during the sweep period from 132 MHz to 131 MHz. The magnitude of the RF voltage may be increased during that period only or, in some cases, during periods that precede and follow the extraction period. In some example implementations, the time period during which the magnitude may be increased is 20-40 μs. Notably, these values are specific to one example particle accelerator, and the values, including frequency window and time period, may be different for different systems.

In other implementations, particle source 700 is controllable to provide pulses of the ionized plasma at specific frequencies proximate to a decrease from the maximum RF frequency to the minimum RF frequency during the voltage sweep. For example, referring to FIG. 17, pulse width 721 may be controlled to occur at any point between a starting (e.g., maximum) frequency 722 and an ending (e.g., minimum) frequency 723. The amount of particles extracted at various frequencies may be measured in order to determine the best location.

In some implementations, pulse widths may be varied by controlling the rate of the RF sweep. For example, slower RF sweeps may result on longer pulses and, thus, more particles (intensity) per pulse.

In other implementations, the voltage source is configurable to vary the RF voltage in order to control an intensity of the beam of particles. For example, the RF voltage may be swept over a time-scale from a high value to a low value. Initially, the RF voltage may be at a high value (e.g., constant for a period of time). There, the voltage is applied for an initial period of time (e.g., 20-40 μs). Then, the voltage is reduced, e.g., every 20 μs during the sweep so that its amplitude is adjusted to control the intensity of the particle beam. The slope of the decreasing magnetic field (versus time) may be increased or decreased in order to change the amount of particle extracted. In some implementations, the voltage may be applied in steps to control the particle beam output.

In other implementations, pulse-blanking may be used to control the intensity of the particle beam. For example, particle source 700 may be controlled to selectively output pulses of the ionized plasma. For example, the pulses may be output for a period during every voltage sweep; however, pulse output may be skipped in every N^th (N>1) sweep. So, for example, the control system may detect that there is 1% too much beam, in which case every 100^th pulse may be skipped. In other implementations, pulses may be skipped more frequently, e.g., every second, third, tenth or any other appropriate numbered pulse may be skipped.

In other implementations, a bias voltage may be applied to the active dee and/or to the dummy dee to reduce the effects of multi-pactoring and thereby increase the intensity of the particle beam. In this regard, multi-pactoring occurs when electrons bounce between dee plates, causing additional electrons to be cast-off of the dee plates upon impact. The result can adversely affect the operation of the dee plates to the point of electrically shorting the dee plates.

To reduce the effects of multi-pactoring, a DC bias voltage may be applied to the active dee and/or to the dummy dee. This causes the background RF voltage to oscillate further away from ground than would otherwise be the case, thereby reducing electron transfer between dees. In an implementation, a DC bias voltage is applied to the dummy dee only, and the dummy dee is isolated from ground. In other implementations, differential DC bias voltages are applied to the dummy dee and the active dee. For example, a larger DC bias voltage may be applied to the dummy dee and a smaller DC bias voltage may be applied to the active dee. In some implementations, DC the bias plates 800, 801 of FIG. 18 may be added to the dummy dee 102. In this figure, the active dee is labeled 100.

In some implementations, the DC bias voltage differential (that is, the difference between the bias voltages applied to the active and dummy dees) may be within the +/−50% range. The specific amounts of the DC bias voltage may vary based on the level of the RF voltage. For example, 2.1 KV DC voltage may be applied to the dummy dee and 1.7 KV DC voltage may be applied to the active dee. In another example, 1.5 KV DC voltage may be applied to the dummy dee and 1.0 KV DC voltage may be applied to the active dee. In another example, 1.9 KV DC voltage may be applied to the dummy dee and 1.5 KV DC voltage may be applied to the active dee. In other implementations, different DC bias voltage may be used.

Referring to FIG. 19, at the output of extraction channel 802 of a particle accelerator (which may have the configuration shown in FIGS. 1 and 2) is a beam formation system, such as beam formation system 125. The beam formation system may be a scanning system. An example scanning system 806 is shown in FIG. 19, which may be used to scan the particle beam across at least part of an irradiation target. FIG. 20 also shows examples of the components of the scanning system include a scanning magnet 808, an ion chamber 809, and an energy degrader 810. Other components of the scanning system are not shown in FIG. 20.

In an example operation, scanning magnet 808 is controllable in two dimensions (e.g., Cartesian XY dimensions) to direct the particle beam across a part (e.g., a cross-section) of an irradiation target. Ion chamber 809 detects the dosage of the beam and feeds-back that information to a control system. Energy degrader 810 is controllable to move material into, and out of, the path of the particle beam to change the energy of the particle beam and therefore the depth to which the particle beam will penetrate the irradiation target.

FIGS. 21 and 22 shows views of an example scanning magnet 808. Scanning magnet 808 includes two coils 811, which control particle beam movement in the X direction, and two coils 812, which control particle beam movement in the Y direction. Control is achieved, in some implementations, by varying current through one or both sets of coils to thereby vary the magnetic field(s) produced thereby. By varying the magnetic field(s) appropriately, the particle beam can be moved in the X and/or Y direction across the irradiation target. In some implementations, the scanning magnet is not movable physically relative to the particle accelerator. In other implementations, the scanning magnet may be movable relative to the accelerator (e.g., in addition to the movement provided by the gantry).

In this example, ion chamber 809 detects dosage applied by the particle beam by detecting the numbers of ion pairs created within a gas caused by incident radiation. The numbers of ion pairs correspond to the dosage provided by the particle beam. That information is fed-back to a computer system that controls operation of the particle therapy system. The computer system (not shown), which may include memory and one or more processing devices, determines if the dosage detected by ion chamber is the intended dose. If the dosage is not as intended, the computer system may control the accelerator to interrupt production and/or output of the particle beam, and/or control the scanning magnet to prevent output of the particle beam to the irradiation target.

FIG. 23 shows a range modulator 815, which is an example implementation of energy degrader 810. In some implementations, such as that shown in FIG. 23, range modulator includes a series of plates 816. The plates may be made of one or more energy absorbing materials.

One or more of the plates is movable into, or out of, the beam path to thereby affect the energy of the particle beam and, thus, the depth of penetration of the particle beam within the irradiation target. For example, the more plates that are moved into the path of the particle beam, the more energy that will be absorbed by the plates, and the less energy the particle beam will have. Conversely, the fewer plates that are moved into the path of the particle beam, the less energy that will be absorbed by the plates, and the more energy the particle beam will have. Higher energy particle beams penetrate deeper into the irradiation target than do lower energy particle beams. In this context, “higher” and “lower” are meant as relative terms, and do not have any specific numeric connotations.

Plates are moved physically into, and out of, the path of the particle beam. For example, as shown in FIG. 24, a plate 816a moves along the direction of arrow 817 between positions in the path of the particle beam and outside the path of the particle beam. The plates are computer-controlled. Generally, the number of plates that are moved into the path of the particle beam corresponds to the depth at which scanning of an irradiation target is to take place. For example, the irradiation target can be divided into cross-sections, each of which corresponds to an irradiation depth. One or more plates of the range modulator can be moved into, or out of, the beam path to the irradiation target in order to achieve the appropriate energy to irradiate each of these cross-sections of the irradiation target.

In some implementations, a treatment plan is established prior to treating the irradiation target using scanning. The treatment plan may specify how scanning is to be performed for a particular irradiation target. In some implementations, the treatment plan specifies the following information: a type of scanning (e.g., spot scanning or raster scanning); scan locations (e.g., locations of spots to be scanned); magnet current per scan location; dosage-per-spot; locations (e.g., depths) of irradiation target cross-sections; particle beam energy per cross-section; plates or other types of pieces to move into the beam path for each particle beam energy; and so forth. Generally, spot scanning involves applying irradiation at discrete spots on an irradiation target and raster scanning involves moving a radiation spot across the radiation target. The concept of spot size therefore applies for both raster and spot scanning.

In some implementations, the intensity of spots in the scanning system may vary from spot-to-spot. Any of the techniques described herein may be used to vary the intensity of the particle beam from spot-to-spot. For example, the intensity of the particle beam may be varied from individual spot to individual spot, or from one group of spots to another group of spots, and so forth.

The pulse-width modulation techniques (PWM) described herein to vary the pulse width of pulses of the particle beam (and thereby vary the number of particles per pulse, i.e., pulse intensity) may be particularly useful for varying the intensity from spot-to-spot in the scanning system. PWM techniques may be particularly useful in a scanning context because they enable variation in spot intensity rather quickly, e.g., in a sub-second time-frame, and have a relatively wide dynamic range (although non-PWM techniques are still usable).

Any two more of the foregoing implementations may be used in an appropriate combination to affect the energy of a particle beam in the extraction channel. Likewise, individual features of any two more of the foregoing implementations may be used in an appropriate combination for the same purpose.

Elements of different implementations described herein may be combined to form other implementations not specifically set forth above. Elements may be left out of the processes, systems, apparatus, etc., described herein without adversely affecting their operation. Various separate elements may be combined into one or more individual elements to perform the functions described herein.

The example implementations described herein are not limited to use with a particle therapy system or to use with the example particle therapy systems described herein. Rather, the example implementations can be used in any appropriate system that directs accelerated particles to an output.

Additional information concerning the design of an example implementation of a particle accelerator that may be used in a system as described herein can be found in U.S. Provisional Application No. 60/760,788, entitled “High-Field Superconducting Synchrocyclotron” and filed Jan. 20, 2006; U.S. patent application Ser. No. 11/463,402, entitled “Magnet Structure For Particle Acceleration” and filed Aug. 9, 2006; and U.S. Provisional Application No. 60/850,565, entitled “Cryogenic Vacuum Break Pneumatic Thermal Coupler” and filed Oct. 10, 2006, all of which are incorporated herein by reference.

The following applications, all of which are filed on the same date as the subject application (entitled “CONTROLLING INTENSITY OF A PARTICLE BEAM” (Application No. 61/707,466)), are incorporated by reference into the subject application: the U.S. Provisional Application entitled “ADJUSTING ENERGY OF A PARTICLE BEAM” (Application No. 61/707,515), the U.S. Provisional Application entitled “ADJUSTING COIL POSITION” (Application No. 61/707,548), the U.S. Provisional Application entitled “FOCUSING A PARTICLE BEAM USING MAGNETIC FIELD FLUTTER” (Application No. 61/707,572), the U.S. Provisional Application entitled “MAGNETIC FIELD REGENERATOR” (Application No. 61/707,590), the U.S. Provisional Application entitled “FOCUSING A PARTICLE BEAM” (Application No. 61/707,704), the U.S. Provisional Application entitled “CONTROLLING PARTICLE THERAPY (Application No. 61/707,624), and the U.S. Provisional Application entitled “CONTROL SYSTEM FOR A PARTICLE ACCELERATOR” (Application No. 61/707,645).

The following are also incorporated by reference into the subject application: U.S. Pat. No. 7,728,311 which issued on Jun. 1, 2010, U.S. patent application Ser. No. 11/948,359 which was filed on Nov. 30, 2007, U.S. patent application Ser. No. 12/275,103 which was filed on Nov. 20, 2008, U.S. patent application Ser. No. 11/948,662 which was filed on Nov. 30, 2007, U.S. Provisional Application No. 60/991,454 which was filed on Nov. 30, 2007, U.S. Pat. No. 8,003,964 which issued on Aug. 23, 2011, U.S. Pat. No. 7,208,748 which issued on Apr. 24, 2007, U.S. Pat. No. 7,402,963 which issued on Jul. 22, 2008, U.S. patent application Ser. No. 13/148,000 filed Feb. 9, 2010, U.S. patent application Ser. No. 11/937,573 filed on Nov. 9, 2007, U.S. patent application Ser. No. 11/187,633, titled “A Programmable Radio Frequency Waveform Generator for a Synchrocyclotron,” filed Jul. 21, 2005, U.S. Provisional Application No. 60/590,089, filed on Jul. 21, 2004, U.S. patent application Ser. No. 10/949,734, titled “A Programmable Particle Scatterer for Radiation Therapy Beam Formation”, filed Sep. 24, 2004, and U.S. Provisional Application No. 60/590,088, filed Jul. 21, 2005.

Any features of the subject application may be combined with one or more appropriate features of the following: the U.S. Provisional Application entitled “ADJUSTING ENERGY OF A PARTICLE BEAM” (Application No. 61/707,515), the U.S. Provisional Application entitled “ADJUSTING COIL POSITION” (Application No. 61/707,548), the U.S. Provisional Application entitled “FOCUSING A PARTICLE BEAM USING MAGNETIC FIELD FLUTTER” (Application No. 61/707,572), the U.S. Provisional Application entitled “MAGNETIC FIELD REGENERATOR” (Application No. 61/707,590), the U.S. Provisional Application entitled “FOCUSING A PARTICLE BEAM” (Application No. 61/707,704), the U.S. Provisional Application entitled “CONTROLLING PARTICLE THERAPY (Application No. 61/707,624), and the U.S. Provisional Application entitled “CONTROL SYSTEM FOR A PARTICLE ACCELERATOR” (Application No. 61/707,645), U.S. Pat. No. 7,728,311 which issued on Jun. 1, 2010, U.S. patent application Ser. No. 11/948,359 which was filed on Nov. 30, 2007, U.S. patent application Ser. No. 12/275,103 which was filed on Nov. 20, 2008, U.S. patent application Ser. No. 11/948,662 which was filed on Nov. 30, 2007, U.S. Provisional Application No. 60/991,454 which was filed on Nov. 30, 2007, U.S. Pat. No. 8,003,964 which issued on Aug. 23, 2011, U.S. Pat. No. 7,208,748 which issued on Apr. 24, 2007, U.S. Pat. No. 7,402,963 which issued on Jul. 22, 2008, U.S. patent application Ser. No. 13/148,000 filed Feb. 9, 2010, U.S. patent application Ser. No. 11/937,573 filed on Nov. 9, 2007, U.S. patent application Ser. No. 11/187,633, titled “A Programmable Radio Frequency Waveform Generator for a Synchrocyclotron,” filed Jul. 21, 2005, U.S. Provisional Application No. 60/590,089, filed on Jul. 21, 2004, U.S. patent application Ser. No. 10/949,734, titled “A Programmable Particle Scatterer for Radiation Therapy Beam Formation”, filed Sep. 24, 2004, and U.S. Provisional Application No. 60/590,088, filed Jul. 21, 2005.

Except for the provisional application from which this patent application claims priority and the documents incorporated by reference above, no other documents are incorporated by reference into this patent application.

Other implementations not specifically described herein are also within the scope of the following claims.

Claims

1. A synchrocyclotron comprising: a particle source to provide pulses of ionized plasma to a cavity;a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the ionized plasma outwardly;an extraction channel to receive a beam of particles from the cavity for output from the synchrocyclotron; anda control system comprising one or more processing devices to control the particle source to provide pulse widths of the ionized plasma in order to output the beam of particles at an intensity selected from different intensities;wherein the particle source comprises cathodes to provide voltage to ionize hydrogen to produce the ionized plasma, the cathodes being unheated by an external source.

2. The synchrocyclotron of claim 1, wherein the particle source is configured to activate for a period of time in response to a control signal from the control system, where the particle source generates a pulse of ionized plasma when activated.

3. The synchrocyclotron of claim 1, wherein the particle source is configured to generate pulses of ionized plasma periodically.

4. The synchrocyclotron of claim 3, wherein the particle beam is output for a duration of about 0.1 μs to 100 μs.

5. The synchrocyclotron of claim 3, wherein the particle beam is output for a duration of about 0.1 μs to 100 μs about every 2 ms.

6. A proton therapy system comprising: the synchrocyclotron of claim 1; anda gantry on which the synchrocyclotron is mounted, the gantry for rotating the synchrocyclotron around a patient position.

7. The synchrocyclotron of claim 1, wherein the voltage source comprises a first dee and a second dee, and wherein at least one of the first dee or the second dee has a bias voltage applied thereto.

8. The synchrocyclotron of claim 1, wherein the voltage source is configured to sweep the RF voltage between a maximum frequency and a minimum frequency, and the particle source is controllable to affect the intensity by providing pulses of the ionized plasma at specific frequencies during a decrease from the maximum frequency of the RF voltage to the minimum frequency of the RF voltage.

9. A synchrocyclotron comprising: a particle source to provide pulses of ionized plasma to a cavity, the particle source comprising cathodes to provide voltage to ionize hydrogen to produce the ionized plasma;a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the ionized plasma outwardly;an extraction channel to receive a beam of particles from the cavity for output from the synchrocyclotron; anda control system comprising one or more processing devices to control a voltage provided by the cathodes in order to output the beam of particles at an intensity selected from among different intensities;wherein the cathodes are unheated by an external source.

10. The synchrocyclotron of claim 9, wherein the voltage is controllable such that increasing the voltage increases an intensity of the beam of particles and such that decreasing the voltage decreases the intensity of the beam of particles.

11. A proton therapy system comprising: the synchrocyclotron of claim 9; anda gantry on which the synchrocyclotron is mounted, the gantry for rotating the synchrocyclotron around a patient position.

12. The synchrocyclotron of claim 9, wherein the voltage source comprises a first dee and a second dee, and wherein at least one of the first dee or the second dee has a bias voltage applied thereto.

13. The synchrocyclotron of claim 9, wherein the voltage source is configured to sweep the RF voltage between a maximum frequency and a minimum frequency, and the particle source is controllable to affect the intensity by providing pulses of the ionized plasma at specific frequencies during a decrease from the maximum frequency of the RF voltage to the minimum frequency of the RF voltage.

14. A synchrocyclotron comprising: a particle source to provide pulses of ionized plasma to a cavity, the particle source comprising cathodes to provide voltage to ionize hydrogen to produce the ionized plasma;a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the ionized plasma outwardly;an extraction channel to receive a beam of particles from the cavity for output from the synchrocyclotron; anda control system comprising one or more processing devices to control the particle source to adjust an amount of the hydrogen between the cathodes in order to output the beam of particles at an intensity selected from among different intensities;wherein the cathodes are unheated by an external source.

15. The synchrocyclotron of claim 14, wherein the amount of hydrogen is adjustable such that increasing the amount of hydrogen increases an intensity of the beam of particles and such that decreasing the amount of hydrogen decreases the intensity of the beam of particles.

16. A proton therapy system comprising: the synchrocyclotron of claim 14; anda gantry on which the synchrocyclotron is mounted, the gantry for rotating the synchrocyclotron around a patient position.

17. The synchrocyclotron of claim 14, wherein the voltage source comprises a first dee and a second dee, and wherein at least one of the first dee or the second dee has a bias voltage applied thereto.

18. The synchrocyclotron of claim 14, wherein the voltage source is configured to sweep the RF voltage between a maximum frequency and a minimum frequency, and the particle source is controllable to affect the intensity by providing pulses of the ionized plasma at specific frequencies during a decrease from the maximum frequency of the RF voltage to the minimum frequency of the RF voltage.

19. A synchrocyclotron comprising: a particle source to provide pulses of ionized plasma to a cavity;a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the ionized plasma outwardly;an extraction channel to receive a beam of particles from the cavity for output from the synchrocyclotron; anda control system comprising one or more processing devices to control the voltage source to provide a magnitude of the RF voltage in order to output the beam of particles at an intensity selected from among different intensities;wherein the particle source comprises cathodes to provide voltage to ionize hydrogen to produce the ionized plasma, the cathodes being unheated by an external source.

20. The synchrocyclotron of claim 19, wherein the magnitude of the RF voltage is adjustable such that increasing the magnitude increases an intensity of the beam of particles and such that decreasing the magnitude decreases the intensity of the beam of particles.

21. A proton therapy system comprising: the synchrocyclotron of claim 19; anda gantry on which the synchrocyclotron is mounted, the gantry for rotating the synchrocyclotron around a patient position.

22. The synchrocyclotron of claim 19, wherein the voltage source comprises a first dee and a second dee, and wherein at least one of the first dee or the second dee has a bias voltage applied thereto.

23. The synchrocyclotron of claim 19, wherein the voltage source is configured to sweep the RF voltage between a maximum frequency and a minimum frequency, and the particle source is controllable to affect the intensity by providing pulses of the ionized plasma at specific frequencies during a decrease from the maximum frequency of the RF voltage to the minimum frequency of the RF voltage.

24. A synchrocyclotron comprising: a particle source to provide pulses of ionized plasma to a cavity;a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the ionized plasma outwardly, the RF voltage sweeping between a maximum frequency and a minimum frequency;an extraction channel to receive a beam of particles from the cavity for output from the synchrocyclotron; anda control system comprising one or more processing devices to control the particle source to output the beam of particles at an intensity selected from among different intensities, the particle source being controllable by the control system to provide pulses of the ionized plasma at specific frequencies during a decrease from the maximum frequency of the RF voltage to the minimum frequency of the RF voltage;wherein the particle source comprises cathodes to provide voltage to ionize hydrogen to produce the ionized plasma, the cathodes being unheated by an external source.

25. The synchrocyclotron of claim 24, wherein the particle source is controllable to provide pulses of the ionized plasma between 132 MHz of RF voltage and 131 MHz of RF voltage from a decrease from a maximum frequency of about 135 MHz of the RF voltage.

26. A proton therapy system comprising: the synchrocyclotron of claim 24; anda gantry on which the synchrocyclotron is mounted, the gantry for rotating the synchrocyclotron around a patient position.

27. The synchrocyclotron of claim 24, wherein the voltage source comprises a first dee and a second dee, and wherein at least one of the first dee or the second dee has a bias voltage applied thereto.

28. A synchrocyclotron comprising: a particle source to provide pulses of ionized plasma to a cavity;a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the ionized plasma outwardly;an extraction channel to receive a beam of particles from the cavity for output from the synchrocyclotron; anda control system comprising one or more processing devices to control the particle source to selectively output pulses of the ionized plasma in order to output the beam of particles at an intensity selected from among different intensities;wherein the particle source comprises cathodes to provide voltage to ionize hydrogen to produce the ionized plasma, the cathodes being unheated by an external source.

29. The synchrocyclotron of claim 28, wherein the voltage source is configured to sweep the RF voltage periodically from a maximum frequency to a minimum frequency; and wherein selectively outputting the pulses comprises outputting pulses in certain ones of the RF voltage sweeps and not in others of the RF voltage sweeps.

30. The synchrocyclotron of claim 28, wherein the voltage source is configured to sweep the RF voltage periodically from a maximum frequency to a minimum frequency; and wherein selectively outputting the pulses comprises skipping pulse output in every Nth (N>1) sweep.

31. The synchrocyclotron of claim 28, wherein the control system is configured to perform operations comprising: selecting the intensity of the beam of particles.

32. A proton therapy system comprising: the synchrocyclotron of claim 28; anda gantry on which the synchrocyclotron is mounted, the gantry for rotating the synchrocyclotron around a patient position.

33. The synchrocyclotron of claim 28, wherein the voltage source comprises a first dee and a second dee, and wherein at least one of the first dee or the second dee has a bias voltage applied thereto.

34. The synchrocyclotron of claim 28, wherein the voltage source is configured to sweep the RF voltage between a maximum frequency and a minimum frequency, and the particle source is controllable to affect the intensity by providing pulses of the ionized plasma at specific frequencies during a decrease from the maximum frequency of the RF voltage to the minimum frequency of the RF voltage.

35. A synchrocyclotron comprising: a particle source to provide pulses of ionized plasma to a cavity;a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the ionized plasma outwardly;an extraction channel to receive a beam of particles from the cavity for output from the synchrocyclotron; anda control system comprising one or more processing devices to control the voltage source to vary a rate of change of the RF voltage in order to output the beam of particles at an intensity selected from among different intensities;wherein the particle source comprises cathodes to provide voltage to ionize hydrogen to produce the ionized plasma, the cathodes being unheated by an external source.

36. A proton therapy system comprising: the synchrocyclotron of claim 35; anda gantry on which the synchrocyclotron is mounted, the gantry for rotating the synchrocyclotron around a patient position.

37. The synchrocyclotron of claim 35, wherein the voltage source comprises a first dee and a second dee, and wherein at least one of the first dee or the second dee has a bias voltage applied thereto.

38. The synchrocyclotron of claim 35, wherein the voltage source is configured to sweep the RF voltage between a maximum frequency and a minimum frequency, and the particle source is controllable to affect the intensity by providing pulses of the ionized plasma at specific frequencies during a decrease from the maximum frequency of the RF voltage to the minimum frequency of the RF voltage.

39. A synchrocyclotron comprising: a particle source to provide pulses of ionized plasma to a cavity;a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the ionized plasma outwardly, the voltage source comprising a first dee and a second dee, wherein at least one of the first dee or the second dee has a bias voltage applied thereto; andan extraction channel to receive a beam of particles from the cavity for output from the synchrocyclotron;wherein the particle source comprises cathodes to provide voltage to ionize hydrogen to produce the ionized plasma, the cathodes being unheated by an external source.

40. The synchrocyclotron of claim 39, wherein the first dee has a first bias voltage applied thereto and the second dee has a second bias voltage applied thereto, the first bias voltage being different from the second bias voltage.

41. The synchrocyclotron of claim 39, wherein the first dee has the bias voltage applied thereto and the second dee is electrically grounded.

42. A proton therapy system comprising: the synchrocyclotron of claim 39; anda gantry on which the synchrocyclotron is mounted, the gantry for rotating the synchrocyclotron around a patient position.

43. The synchrocyclotron of claim 39, wherein the voltage source is configured to sweep the RF voltage between a maximum frequency and a minimum frequency, and the particle source is controllable to affect the intensity by providing pulses of the ionized plasma at specific frequencies during a decrease from the maximum frequency of the RF voltage to the minimum frequency of the RF voltage.

44. The synchrocyclotron of claim 39, wherein the first dee comprises a DC bias plate, and the bias voltage is a DC bias voltage.