Active return system

Abstract

An example particle accelerator includes a magnet to generate a magnetic field, where the magnet includes first superconducting coils to pass current in a first direction to thereby generate the first magnetic field, and where the first magnetic field is at least 4 Tesla (T). The example particle accelerator also includes an active return system including second superconducting coils. Each of the second superconducting coils surrounds, and is concentric with, a corresponding first superconducting coil. The second superconducting coils are for passing current in a second direction that is opposite to the first direction to thereby generate a second magnetic field having a magnetic field of at least 2.5 T. The second magnetic field has a polarity that is opposite to a polarity of the first magnetic field.

Description

TECHNICAL FIELD

This disclosure relates generally to an active return system for a superconducting magnet.

BACKGROUND

Particle therapy systems use an accelerator to generate a particle beam for treating afflictions, such as tumors. In operation, particles are accelerated in orbits inside a cavity in the presence of a magnetic field, and removed from the cavity through an extraction channel. The particles are part of a beam, which is applied to the patient for treatment. The magnetic field is generated by a magnet, which produces magnetic flux. Too much stray magnetic flux can adversely affect the operation of the accelerator and of other components of the particle therapy system. A return may therefore be used to route the stray magnetic flux. Ferromagnetic returns can be heavy, and add considerable weight to the accelerator. This can be problematic in some cases.

SUMMARY

An example particle accelerator comprises a magnet to generate a magnetic field, where the magnet comprises first superconducting coils to pass current in a first direction to thereby generate the first magnetic field, and where the first magnetic field is at least 4 Tesla (T). The example particle accelerator also comprises an active return system including second superconducting coils. Each of the second superconducting coils surrounds, and is concentric with, a corresponding first superconducting coil. The second superconducting coils are for passing current in a second direction that is opposite to the first direction to thereby generate a second magnetic field having a magnetic field of at least 2.5 T. The second magnetic field has a polarity that is opposite to a polarity of the first magnetic field. The example particle accelerator may include one or more of the following features, either alone or in combination.

A power supply may provide current to both the first superconducting coils and the second superconducting coils. The first superconducting coils and the second superconducting coils may be mounted on a structure. The structure may comprise at least one of stainless steel and carbon fiber.

The first superconducting coils may be mounted on an interior of the structure and the second superconducting coils may be mounted on an exterior of the structure such that the second superconducting coils are separated from the first superconducting coils by at least part of the structure. A banding ring may be around the second superconducting coils.

Magnetic pole pieces may define the cavity, and the structure may be around at least part of the magnetic pole pieces. A cryostat cover may be around at least part of the structure and at least part of the magnetic pole pieces. The cryostat cover may comprise a non-ferromagnetic material.

The particle accelerator may weigh less than 15 tons, less than 10 tons, less than 9 tons, less than 8 tons, less than 7 tons, and so forth.

A proton therapy system may comprise the foregoing particle accelerator (and variations thereof), along with a gantry on which the particle accelerator is mounted. The gantry is rotatable relative to a patient position. Protons are output essentially directly from the particle accelerator to the patient position. The particle accelerator may be a synchrocyclotron. The proton therapy system may also comprise a particle source to provide ionized plasma to a cavity containing the first magnetic field and a voltage source to provide voltage to accelerate a beam comprised of pulses of ionized plasma towards an exit.

An example particle accelerator may comprise a voltage source to provide a radio frequency (RF) voltage to a cavity to accelerate particles to produce a particle beam, where the cavity has a first magnetic field for causing particles accelerated from the plasma column to move orbitally within the cavity, and where the RF voltage is controllable to vary in time as the particle beam increases in distance from the plasma column. The example particle accelerator may also comprise a magnet to generate the first magnetic field in the cavity, where the magnet comprises first superconducting coils to pass current in a first direction to thereby generate the first magnetic field. The example particle accelerator may also comprise an active return system comprising second superconducting coils, where each of the second superconducting coils surrounds, and is concentric with, a corresponding first superconducting coil. The second superconducting coils are for passing current in a second direction that is opposite to the first direction to thereby generate a second magnetic field having a magnetic field of at least 2.5 Tesla (T). The second magnetic field has a polarity that is opposite to a polarity of the first magnetic field. The example particle accelerator may include one or more of the following features, either alone or in combination.

The first magnetic field may be least 4 T. The second magnetic field may be at between 2.5 T and 12 T. The first magnetic field may be between 4 T and 20 T and the second magnetic field may be between 2.5 T and 12 T.

A single power supply may be used to provide current to both the first superconducting coils and to the second superconducting coils. The first superconducting coils and the second superconducting coils may be mounted on a structure. The structure may comprise at least one of stainless steel and carbon fiber. The first superconducting coils may be mounted on an interior of the structure and the second superconducting coils may be mounted on an exterior of the structure such that the second superconducting coils are separated from the first superconducting coils by at least part of the structure. A banding ring may be around the second superconducting coils.

Magnetic pole pieces may define the cavity, and the structure may be around at least part of the magnetic pole pieces. A cryostat cover may be around at least part of the structure and at least part of the magnetic pole pieces. The cryostat cover may comprise a non-ferromagnetic material.

The particle accelerator may weigh less than 15 tons, less than 10 tons, less than 9 tons, less than 8 tons, less than 7 tons, and so forth.

A proton therapy system may comprise the foregoing particle accelerator (and variations thereof), along with a gantry on which the particle accelerator is mounted. The gantry is rotatable relative to a patient position. Protons are output essentially directly from the particle accelerator to the patient position. The particle accelerator may be a synchrocyclotron. The proton therapy system may also comprise a particle source to provide ionized plasma to a cavity containing the first magnetic field and a voltage source to provide voltage to accelerate a beam comprised of pulses of ionized plasma towards an exit.

Two or more of the features described in this disclosure, 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 THE DRAWINGS

FIG. 1 is a side cut-away view of a superconducting magnet.

FIG. 2 is top view of example main and active return coils.

FIG. 3 is a front view of an example particle therapy system.

FIG. 4 is a perspective, cut-away view of example components of a superconducting magnet with active return coils.

FIG. 5 is a front, cut-away view of example components of a superconducting magnet with active return coils.

FIG. 6 is a cross-sectional view of part of an example support structure and example superconducting coil windings.

FIG. 7 is a cross-sectional view of an example cable-in-channel composite conductor.

FIG. 8 is a cross-sectional view of an example ion source.

FIG. 9 is a perspective view of an example dee plate and dummy dee.

FIG. 10 is a perspective view of an example vault containing an example gantry and particle accelerator.

Like reference symbols in the various drawings indicate like elements.

DETAILED DESCRIPTION

Described herein is an example of a particle accelerator for use in a system, such as a proton or ion therapy system. The example particle therapy 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 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 pole pieces are located inside the cryostat, and define a cavity in which particles are accelerated.

In this 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) when accelerating particles from the column. The magnetic field produced by running current through the superconducting coil causes particles accelerated from the plasma column to accelerate orbitally within the cavity.

A magnetic field regenerator (“regenerator”) is positioned near the outside of the cavity (e.g., at an interior edge thereof) to adjust the existing magnetic field inside the cavity to thereby change locations (e.g., the pitch and angle) 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 cryostat. The regenerator may increase the magnetic field at a point in the cavity (e.g., it may produce a magnetic field “bump” at an area of the cavity), thereby causing each successive orbit of particles at that point to precess outwardly toward the entry point of the extraction channel until it reaches the extraction channel. The extraction channel receives particles accelerated from the plasma column and outputs the received particles from the cavity as a particle beam.

The superconducting coil can produce relatively high magnetic fields. Traditionally, large ferromagnetic magnetic yokes acted as a return for stray magnetic field produced by the superconducting coil. For example, in some implementations, the superconducting magnet can generate a relatively high magnetic field of, e.g., 4 Tesla (T) or more, resulting in considerable stray magnetic fields. In some systems, such as that shown in FIG. 1, relatively large ferromagnetic return yokes 100 were used as a return for the magnetic field generated by superconducting coils 102. A magnetic shield 104 surrounded the pole pieces. The return yokes and the shield together dissipated stray magnetic field, thereby reducing the possibility that stray magnetic fields would adversely affect the operation of the accelerator. Drawbacks of this configuration may include size and weight. For example, in some such systems, the accelerator could have a weight on the order of 25 tons or more with correspondingly large dimensions.

In some implementations, therefore, the relatively large yokes and shield used because of the relatively high magnetic field may be replaced by an active return system. An example active return system includes one or more active return coils that conduct current in a direction opposite to current through the main superconducting coils. In some example implementations, there is an active return coil for each superconducting coil, e.g., two active return coils—one for each superconducting coil (referred to as a “main” coil). Each active return coil may also be a superconducting coil that surrounds the outside of a corresponding main superconducting coil. For example, a main coil 200 and an active return coil 201 may be arranged concentrically, as shown in FIG. 2.

Current passes through the active return coils in a direction that is opposite to the direction of current passing through the main coils. The current passing through the active return coils thus generates a magnetic field that is opposite in polarity to the magnetic field generated by the main coils. As a result, the magnetic field generated by an active return coil is able to dissipate the relatively strong stray magnetic field resulting from the corresponding main coil. In some implementations, each active return may be used to generate a magnetic field of between 2.5 T and 12 T or more. For example, an active return coil may be used to generate magnetic fields at, or that exceed, one or more of the following magnitudes: 2.5 T, 2.6 T, 2.7 T, 2.8 T, 2.9 T, 3.0 T, 3.1 T, 3.2 T, 3.3 T, 3.4 T, 3.5 T, 3.6 T, 3.7 T, 3.8 T, 3.9 T, 4.0 T, 4.1 T, 4.2 T, 4.3 T, 4.4 T, 4.5 T, 4.6 T, 4.7 T, 4.8 T, 4.9 T, 5.0 T, 5.1 T, 5.2 T, 5.3 T, 5.4 T, 5.5 T, 5.6 T, 5.7 T, 5.8 T, 5.9 T, 6.0 T, 6.1 T, 6.2 T, 6.3 T, 6.4 T, 6.5 T, 6.6 T, 6.7 T, 6.8 T, 6.9 T, 7.0 T, 7.1 T, 7.2 T, 7.3 T, 7.4 T, 7.5, 7.6 T, 7.7 T, 7.8 T, 7.9 T, 8.0 T, 8.1 T, 8.2 T, 8.3 T, 8.4 T, 8.5, 8.6 T, 8.7 T, 8.8 T, 8.9 T, 9.0 T, 9.1 T, 9.2 T, 9.3 T, 9.4 T, 9.5, 9.6 T, 9.7 T, 9.8 T, 9.9 T, 10.0 T, 10.1 T, 10.2 T, 10.3 T, 10.4 T, 10.5, 10.6 T, 10.7 T, 10.8 T, 10.9 T, 11.0 T, 11.1 T, 11.2 T, 11.3 T, 11.4 T, 11.5, 11.6 T, 11.7 T, 11.8 T, 11.9 T, 12.0 T, 12.1 T, 12.2 T, 12.3 T, 12.4 T, 12.5, or more. Furthermore, an active return coil may be used to generate magnetic fields that are within the range of 2.5 T to 12 T (or more) that are not specifically listed above.

The magnetic field generated by a main coil that may be within a range of 4 T to 20 T or more. For example, a main coil may be used to generate magnetic fields at, or that exceed, one or more of the following magnitudes: 4.0 T, 4.1 T, 4.2 T, 4.3 T, 4.4 T, 4.5 T, 4.6 T, 4.7 T, 4.8 T, 4.9 T, 5.0 T, 5.1 T, 5.2 T, 5.3 T, 5.4 T, 5.5 T, 5.6 T, 5.7 T, 5.8 T, 5.9 T, 6.0 T, 6.1 T, 6.2 T, 6.3 T, 6.4 T, 6.5 T, 6.6 T, 6.7 T, 6.8 T, 6.9 T, 7.0 T, 7.1 T, 7.2 T, 7.3 T, 7.4 T, 7.5 T, 7.6 T, 7.7 T, 7.8 T, 7.9 T, 8.0 T, 8.1 T, 8.2 T, 8.3 T, 8.4 T, 8.5 T, 8.6 T, 8.7 T, 8.8 T, 8.9 T, 9.0 T, 9.1 T, 9.2 T, 9.3 T, 9.4 T, 9.5 T, 9.6 T, 9.7 T, 9.8 T, 9.9 T, 10.0 T, 10.1 T, 10.2 T, 10.3 T, 10.4 T, 10.5 T, 10.6 T, 10.7 T, 10.8 T, 10.9 T, 11.0 T, 11.1 T, 11.2 T, 11.3 T, 11.4 T, 11.5 T, 11.6 T, 11.7 T, 11.8 T, 11.9 T, 12.0 T, 12.1 T, 12.2 T, 12.3 T, 12.4 T, 12.5 T, 12.6 T, 12.7 T, 12.8 T, 12.9 T, 13.0 T, 13.1 T, 13.2 T, 13.3 T, 13.4 T, 13.5 T, 13.6 T, 13.7 T, 13.8 T, 13.9 T, 14.0 T, 14.1 T, 14.2 T, 14.3 T, 14.4 T, 14.5 T, 14.6 T, 14.7 T, 14.8 T, 14.9 T, 15.0 T, 15.1 T, 15.2 T, 15.3 T, 15.4 T, 15.5 T, 15.6 T, 15.7 T, 15.8 T, 15.9 T, 16.0 T, 16.1 T, 16.2 T, 16.3 T, 16.4 T, 16.5 T, 16.6 T, 16.7 T, 16.8 T, 16.9 T, 17.0 T, 17.1 T, 17.2 T, 17.3 T, 17.4 T, 17.5 T, 17.6 T, 17.7 T, 17.8 T, 17.9 T, 18.0 T, 18.1 T, 18.2 T, 18.3 T, 18.4 T, 18.5 T, 18.6 T, 18.7 T, 18.8 T, 18.9 T, 19.0 T, 19.1 T, 19.2 T, 19.3 T, 19.4 T, 19.5 T, 19.6 T, 19.7 T, 19.8 T, 19.9 T, 20.0 T, 20.1 T, 20.2 T, 20.3 T, 20.4 T, 20.5 T, 20.6 T, 20.7 T, 20.8 T, 20.9 T, or more. Furthermore, a main coil may be used to generate magnetic fields that are within the range of 4 T to 20 T (or more) that are not specifically listed above. In some implementations, the currents through the active return coils and the main coils have the same (or about the same (e.g., within 10% difference)) magnitude. In some implementations, the currents through the active return coils and the main coils have different magnitudes.

In some implementations, each main coil is superconducting and made of niobium-3 tin (Nb₃Sn) and each active return coil is superconducting and made of niobium-titanium. However, in other implementations, each main coil and each return coil may be made of the same, different, and/or other materials than those noted above.

In some implementations, the same (e.g., a single) power supply may be used to generate current for both the main coil(s) in the magnet and the active return coil(s). This enables the current through all coils to ramp appropriately, and may be useful in example particle therapy systems.

The active return system described herein may be used in a single particle accelerator, and any two or more of the features thereof described herein may be combined in a single particle accelerator. The particle accelerator may be used in any type of medical or non-medical application. An example of a particle therapy system in which a superconducting magnet having the active return system described herein may be used is provided below.

Referring to FIG. 3, a charged particle radiation therapy system 300 includes a beam-producing particle accelerator 302 having a weight and size small enough to permit it to be mounted on a rotating gantry 304 with its output directed straight (that is, essentially directly) from the accelerator housing toward a patient 306. In some implementations, the weight of the particle accelerator may be less than, or about equal to, one of the following weights: 20 tons, 19 tons, 18 tons, 17 tons, 16 tons, 15 tons, 14 tons, 14 tons, 13 tons, 12 tons, 11 tons, 10 tons, 9 tons, 8 tons, 7 tons, 6 tons, 5 tons, or 4 tons. However, the particle accelerator may have any appropriate weight.

In some implementations, the steel gantry has two legs 308, 310 mounted for rotation on two respective bearings 312, 314 that lie on opposite sides of the patient. The accelerator is supported by a steel truss 316 that is long enough to span a treatment area 318 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 320 of less than 360 degrees, e.g., about 180 degrees, to permit a floor 322 to extend from a wall of the vault 324 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 are not directly aligned with the beam, e.g., wall 330), 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. Angles of rotation other than these may be used.

The horizontal rotational axis 332 of the gantry may be located nominally one meter above the floor where the patient and therapist interact with the therapy system. This floor may be positioned about three 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 334 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.

In the example implementation shown in FIG. 3, the superconducting synchrocyclotron 302 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 some implementations, the magnetic field strength may be in the range of 4 T to 20 T and the proton energy may be in the range of 150 to 300 MeV. In some implementations, the magnetic field strength of the active return coils may be in the range of 2.5 T to 12 T.

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.

An example synchrocyclotron includes a magnet system that contains a particle source, a radio frequency (RF) drive system, and a beam extraction system. In some implementations, types of particle accelerators may be used in which one or more of these elements is external to the accelerator.

Referring to FIGS. 4 and 5, 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 400, 401 and a pair of shaped ferromagnetic (e.g., low carbon steel) pole faces 403, 404.

The two superconducting magnet coils are centered on a common axis 405 and are spaced apart along the axis. Referring to FIGS. 6 and 7, the coils may be formed by of Nb₃Sn-based superconducting 0.8 mm diameter strands 701 (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 702 (in this example, a woven fiberglass material). The copper channel containing the wires 703 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 400, 401 are mounted on an annular stainless steel reverse support structure 601. Heater blankets 602 are placed at intervals in the layers of the windings to protect the assembly in the event of a magnet quench.

The geometry of the main coils is maintained by support structure 601, which exerts a restorative force 605 that works against the distorting (e.g., expansion) force produced when the coils are energized. The coil positions may be maintained relative to the magnet pole piece and cryostat using a set of tension links (not shown) that connect the support structure to a cryostat cover (described below) that defines the perimeter of the cryostat.

The main superconducting coils are maintained at temperatures near absolute zero (e.g., about 4 degrees Kelvin) by enclosing the coil assembly (the coils and the support structure) inside an evacuated annular aluminum or stainless steel cryostatic chamber that provides at least some free space around the coil structure. In some implementations, the temperature near absolute zero is achieved and maintained using a cooling channel (not shown) containing liquid helium, which is formed inside the support structure, and which contains a thermal connection between the liquid helium in the channel and the corresponding superconducting coil. An example of a liquid helium cooling system of the type described above, and that may be used is described in U.S. patent application Ser. No. 13/148,000 (Begg et al.).

In FIGS. 4 and 5, the superconducting coils 400, 401 are mounted on the interior of support structure 601. In some implementations, support structure 601 may be made of structural steel, such as stainless steel, or carbon fiber. Active return coils 409, 410 are mounted on the exterior of support structure 601, as shown in FIGS. 4 and 5. A banding ring 411, which may be made, e.g., of carbon fiber or other appropriate material, is mounted around active return coils 409, 410 to hold them in place during magnet operation and thereby maintain their shape (e.g., in response to expansive force resulting from operation). Each active return coil 409, 410 is concentric with respect to its corresponding main coil 400, 401.

The active return coils may be made of superconducting material, such as niobium-titanium or other appropriate materials. The active return coils may be constructed in the same manner as the main coils. In some implementations, the active return coils may be maintained at superconducting temperatures in the same manner as the main superconducting coils, e.g., by conducting heat to a liquid helium cooling channel (not shown in FIGS. 4 and 5). In some implementations, the active return coils may be cooled using other techniques.

Support structure 601, including the main and active return coils, surrounds ferromagnetic (e.g., iron) pole pieces 403, 404, which together define a cavity 412. An ion source is at about the center of cavity 412 to provide the particles for acceleration. In other examples, the ion source may be external to the accelerator. Particles are accelerated in cavity 412 and output as a beam to an extraction channel (not shown) inside the magnet assembly. From the extraction channel, the beam is output essentially directly to the patient.

The support structure, the pole pieces, the main coils and the active return coils (along with other structure, not described herein) are housed in a cryostat cover 415 which, among other things, maintains the temperature of the magnet assembly. Cryostat cover 415 may be may be made of stainless steel, carbon, or other appropriate, relatively lightweight material. Accordingly, as indicated above, in some implementations, a particle accelerator containing the example magnet assembly may have a weight that is less than, or about equal to, one of the following weights: 20 tons, 19 tons, 18 tons, 17 tons, 16 tons, 15 tons, 14 tons, 14 tons, 13 tons, 12 tons, 11 tons, 10 tons, 9 tons, 8 tons, 7 tons, 6 tons, 5 tons, or 4 tons. The actual weight of the particle accelerator and of the magnet assembly may depend on a variety of factors, and is not limited to the example weights provided here.

Examples of particle sources that may be included in cavity 412 are as follows. Referring to FIG. 8, in some implementations, a particle source 800 has a Penning ion gauge geometry. 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. U.S. patent application Ser. No. 11/948,662 describes a particle source in which a tube containing plasma is interrupted at at least a portion of its mid-plane. The remaining features of the particle source are similar to those described with respect to FIG. 8.

Particle source 800 is fed from a supply of hydrogen through a gas line and a tube that delivers gaseous hydrogen. Electric cables carry an electric current from a current source to stimulate electron discharge from cathodes 804, 805 that are aligned with the magnetic field, 810.

In this example, the discharged electrons ionize the gas exiting through a small hole from tube 811 to create a supply of positive ions (protons) for acceleration by one semicircular (dee-shaped) radio-frequency plate 900 that spans half of the space enclosed by the magnet structure and one dummy dee plate 902. 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. 9, the dee plate 900 is a hollow metal structure that has two semicircular surfaces 903, 905 that enclose a space 907 in which the protons are accelerated during half of their rotation around the space enclosed by the magnet structure. A duct 909 opening into the space 907 extends through the pole piece to an external location from which a vacuum pump can be attached to evacuate the space 907 and the rest of the space within a vacuum chamber in which the acceleration takes place. The dummy dee 902 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 pole piece. The dee plate 900 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 907. The radio frequency electric field is made to vary in time as the accelerated particle beam increases in distance from the geometric center. Examples of radio frequency waveform generators that are useful for this purpose are described in U.S. patent application Ser. No. 11/187,633, titled “A Programmable Radio Frequency Waveform Generator for a Synchrocyclotron,” filed Jul. 21, 2005, and in U.S. Provisional Application No. 60/590,089, same title, filed on Jul. 21, 2004, both of which are incorporated herein by reference. 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 applied across the radio frequency plates. 20,000 Volts may be 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 may be arranged to reduce the capacitance between the radio frequency plates and ground. This may be done by forming holes with sufficient clearance from the radio frequency structures through the outer pole piece 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 its Q high during the radio 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 (e.g., cavity 412) 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. Maintaining a high vacuum reduces the chances that accelerating ions will be 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 907. 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 (which is part of the accelerator), referred to herein as the extraction channel, to exit the pole piece 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 in the extraction channel redirect the ions so that they stay in a straight beam of limited spatial extent.

As the beam exits the extraction channel it may be passed through a beam formation system that can be programmably controlled to create a desired combination of scattering angle and range modulation for the beam. Examples of beam forming systems useful for that purpose are described in 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, both of which are incorporated herein by reference. The beam formation system may be used in conjunction with an inner gantry to direct a beam to the patient.

During operation, 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 may be removed from the plates using water cooling lines that release the heat in a heat exchanger.

Stray magnetic fields exiting from the cyclotron are limited by active return coils 409, 410. Accordingly, separate magnetic shielding is typically not required. However, in some implementations, a separate magnetic shield may be used. The separate magnetic shield may include a layer ferromagnetic material (e.g., steel or iron) that encloses the cryostat and is separated by a space.

As mentioned, the gantry allows the synchrocyclotron to be rotated about the horizontal rotational axis 332. 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.

Referring to FIG. 10, at the location at which the ion beam exits synchrocyclotron 302, a beam formation system 1001 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 may can include passive scattering elements as well as active scanning elements.

All of the active systems of the synchrocyclotron (current driven superconducting coils, RF-driven plates, vacuum pumps for the vacuum acceleration chamber and for a superconducting coil cooling chamber, current driven particle source, hydrogen gas source, and 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 (e.g., executable instructions) to effect control.

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

Further details regarding 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.

Any two more of the foregoing implementations may be used in an appropriate combination in an appropriate particle accelerator (e.g., a synchrocyclotron). Likewise, individual features of any two more of the foregoing implementations may be used in an appropriate combination.

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 the particle accelerator 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 as if set forth in full.

The following applications, which were filed on Sep. 28, 2012, are incorporated by reference into the subject application as if set forth herein in full: the U.S. Provisional Application entitled “CONTROLLING INTENSITY OF A PARTICLE BEAM” (Application No. 61/707,466), 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 as if set forth herein in full: 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, and U.S. patent application Ser. No. 11/937,573 filed on Nov. 9, 2007.

Any features of the subject application may be combined with one or more appropriate features of the following: the U.S. Provisional Application entitled “CONTROLLING INTENSITY OF A PARTICLE BEAM” (Application No. 61/707,466), 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, and U.S. patent application Ser. No. 11/937,573 filed on Nov. 9, 2007.

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

Claims

1. A particle accelerator comprising: a magnet to generate a magnetic field, the magnet comprising first superconducting coils to pass current in a first direction to thereby generate the first magnetic field, the first magnetic field being at least 4 Tesla (T);an active return system comprising second superconducting coils, each of the second superconducting coils surrounding, and being concentric with, a corresponding first superconducting coil, the second superconducting coils for passing current in a second direction that is opposite to the first direction to thereby generate a second magnetic field having a magnetic field of at least 2.5 T, the second magnetic field having a polarity that is opposite to a polarity of the first magnetic field; anda single structure on which at least one first superconducting coil and corresponding second superconducting coil are mounted.

2. The particle accelerator of claim 1, further comprising: a power supply to provide current to both the first superconducting coils and to the second superconducting coils.

3. The particle accelerator of claim 1, wherein the first superconducting coils and the second superconducting coils are all mounted on the single structure.

4. The particle accelerator of claim 3, wherein the first superconducting coils are mounted on an interior of the single structure and the second superconducting coils are mounted on an exterior of the single structure such that the second superconducting coils are separated from the first superconducting coils by at least part of the single structure.

5. The particle accelerator of claim 3, further comprising: a banding ring around at least one of the second superconducting coils.

6. The particle accelerator of claim 3, wherein the single structure comprises at least one of stainless steel and carbon fiber.

7. The particle accelerator of claim 1, further comprising: magnetic pole pieces defining the cavity, the single structure being around at least part of the magnetic pole pieces.

8. The particle accelerator of claim 7, further comprising: a cryostat cover around at least part of the single structure and at least part of the magnetic pole pieces, the cryostat cover comprising a non-ferromagnetic material.

9. The particle accelerator of claim 1, which weighs less than 15 tons.

10. The particle accelerator of claim 1, which weighs less than 10 tons.

11. A proton therapy system comprising: the particle accelerator of claim 1; anda gantry on which the particle accelerator is mounted, the gantry being rotatable relative to a patient position;wherein the proton therapy system is configured to output protons essentially directly from the particle accelerator to the patient position.

12. The proton therapy system of claim 11, wherein the particle accelerator comprises a synchrocyclotron.

13. The proton therapy system of claim 11, wherein the particle accelerator comprises: a particle source to provide ionized plasma to a cavity containing the first magnetic field; anda voltage source to provide voltage to accelerate a beam comprised of pulses of ionized plasma towards an exit.

14. A particle accelerator comprising: a voltage source to provide a radio frequency (RF) voltage to a cavity to accelerate particles to produce a particle beam, the cavity having a first magnetic field for causing particles accelerated from the plasma column to move orbitally within the cavity, the RF voltage being controllable to vary in time as the particle beam increases in distance from the plasma column;a magnet to generate the first magnetic field in the cavity, the magnet comprising first superconducting coils to pass current in a first direction to thereby generate the first magnetic field;an active return system comprising second superconducting coils, each of the second superconducting coils surrounding, and being concentric with, a corresponding first superconducting coil, the second superconducting coils for passing current in a second direction that is opposite to the first direction to thereby generate a second magnetic field having a magnetic field of at least 2.5 Tesla (T), the second magnetic field having a polarity that is opposite to a polarity of the first magnetic field; anda single structure on which at least one first superconducting coil and corresponding second superconducting coil are mounted.

15. The particle accelerator of claim 14, wherein the first magnetic field is least 4 T.

16. The particle accelerator of claim 15, wherein the second magnetic field is at between 2.5 T and 12 T.

17. The particle accelerator of claim 14, wherein the first magnetic field is between 4 T and 20 T and the second magnetic field is between 2.5 T and 12 T.

18. The particle accelerator of claim 14, further comprising: a single power supply to provide current to both the first superconducting coils and to the second superconducting coils.

19. The particle accelerator of claim 14, wherein the first superconducting coils and the second superconducting coils are all mounted on the single structure.

20. The particle accelerator of claim 19, wherein the first superconducting coils are mounted on an interior of the single structure and the second superconducting coils are mounted on an exterior of the single structure such that the second superconducting coils are separated from the first superconducting coils by at least part of the single structure.

21. The particle accelerator of claim 19, further comprising: a banding ring around at least one of the second superconducting coils.

22. The particle accelerator of claim 19, wherein the single structure comprises at least one of stainless steel and carbon fiber.

23. The particle accelerator of claim 14, further comprising: magnetic pole pieces defining the cavity, the single structure being around at least part of the magnetic pole pieces.

24. The particle accelerator of claim 23, further comprising: a cryostat cover around at least part of the single structure and at least part of the magnetic pole pieces, the cryostat cover comprising a non-ferromagnetic material.

25. The particle accelerator of claim 14, which weighs less than 15 tons.

26. The particle accelerator of claim 14, which weighs less than 10 tons.

27. A proton therapy system comprising: the particle accelerator of claim 14; anda gantry on which the particle accelerator is mounted, the gantry being rotatable relative to a patient position;wherein the proton therapy system is configured to output protons essentially directly from the particle accelerator to the patient position.