Gantries for particle therapy systems

Abstract

An exemplary particle therapy system includes a gantry having a beamline structure configured to direct a monoenergetic particle beam from an output of a particle accelerator toward an irradiation target, the beamline structure including a magnetic bending element for bending the particle beam along a length of the beamline structure; and an energy degrader downstream of the beamline structure relative to the particle accelerator, the energy degrader configured and controllable to change the energy of the particle beam before at least a portion of the particle beam reaches the irradiation target.

Description

[Technical field]

[0001] This specification describes example particle therapy systems and gantries for use therewith. [Background technology]

[0002] Particle therapy systems use particle accelerators to generate particle beams for treating afflictions (such as tumors). Particle therapy systems can use a gantry to direct the particle beam toward the patient from multiple angles. In some examples, the gantry includes a device that supports the radiation delivery apparatus during treatment. [Prior art documents] [Patent documents]

[0003] [Patent Document 1] U.S. Patent No. 10,675,487 [Patent Document 2] US Patent Application Publication No. 2017 / 0128746 [Patent Document 3] U.S. Patent No. 8,791,656 [Patent Document 4] U.S. Patent No. 8,975,836 [Patent Document 5] U.S. Patent No. 9,730,308 [Patent Document 6] US Patent Application Publication No. 2020 / 0298025 Summary of the Invention [Means for solving the problem]

[0004] An exemplary system includes a gantry having a beamline structure configured to direct a monoenergetic particle beam from an output of a particle accelerator toward an irradiation target, the beamline structure including a magnetic bending element for bending the particle beam along a length of the beamline structure; and an energy degrader downstream of the beamline structure relative to the particle accelerator, the energy degrader configured and controllable to change the energy of the particle beam before at least a portion of the particle beam reaches the irradiation target. The exemplary system can include one or more of the following features, either alone or in combination.

[0005] The energy degrader may be the only mechanism for actively controlling the change in energy of the particle beam after it is output by the particle accelerator and before it reaches the irradiation target. The beamline structure may be configured not to actively control the energy of the particle beam after it is output by the particle accelerator and before it reaches the energy degrader. The magnetic bending element may include a magnet having a magnetic field of 2.5 Tesla (T) or greater. The magnetic bending element may include a magnet having a magnetic field of 3 Tesla (T) or greater.

[0006] The exemplary system can further include a collimator downstream of the energy degrader relative to the particle accelerator for intercepting at least a portion of the particle beam before the at least a portion of the particle beam reaches the patient. The gantry can include a support structure configured to move a portion of the beamline structure in a circular path around the irradiation target.

[0007] The support structure may have a dimension of 6 meters or less. The dimension may be a diameter of the support structure. The length of the beamline structure may be 6 meters (m) or less. The length of the beamline structure may be 5 meters (m) or less. The energy of the particle beam may not vary by more than 1% in the beamline structure. The distance between the output of the beamline structure and an isocenter containing the irradiation target may be 1.5 meters (m) or less.

[0008] An exemplary system may include a scanning system having one or more scanning magnets for moving the particle beam in at least two dimensions across at least a portion of a beam field covering at least a portion of an irradiation target. At least one of the scanning magnets may be or include a superconducting magnet. At least one of the scanning magnets may be positioned downstream of the beamline structure with respect to the particle accelerator. At least one of the scanning magnets may be positioned in the beamline structure. The one or more scanning magnets may include a first scanning magnet positioned in the beamline structure and a second scanning magnet positioned downstream of the first scanning magnet with respect to the particle accelerator.

[0009] The first scanning magnet can be separate from the second scanning magnet. The first scanning magnet can be configured to move the particle beam across at least a portion of the beam field in at least two dimensions. The second scanning magnet can be configured to move the particle beam across at least a portion of the beam field in at least two dimensions. The first scanning magnet can be configured to move the particle beam across at least a portion of the beam field in only a first dimension. The second scanning magnet can be configured to move the particle beam across at least a portion of the beam field in only a second dimension. The second dimension can be different from the first dimension. The first scanning magnet can be positioned between magnetic components included in the beamline structure, and the second scanning magnet can be positioned in a particle beam output device downstream of the beamline structure. The particle beam output device can be or include a nozzle. The first scanning magnet can be positioned in the beamline structure, and the second scanning magnet can be positioned in the beamline structure.

[0010] The beamline structure may include an output channel including magnetic dipoles arranged in series to bend the particle beam by at least 90°. The magnetic dipoles may include at least a first magnetic dipole and a second magnetic dipole. The first magnetic dipole and the second magnetic dipole may be between a first scanning magnet and a second scanning magnet. The beamline structure may include an output channel configured to bend the particle beam by at least 90° toward an irradiation target in the presence of a magnetic field of at least 3 Tesla (T). The beamline structure may include a magnetic dipole for bending the particle beam and two or more magnetic quadrupoles or magnetic hexapoles arranged between the magnetic dipoles along a length of the beamline structure to focus the particle beam.

[0011] The exemplary system may include a particle accelerator, and the particle accelerator may be a compact particle accelerator. The beamline structure may have an efficiency of 10% or more and a length of 6 meters or less. The efficiency of 10% or more may include 10% or more of the particles output from the particle accelerator being output from the beamline structure. The exemplary particle therapy system includes a particle accelerator configured to output particles as a monoenergetic particle beam, and an exemplary gantry. The exemplary gantry includes a beamline structure configured to direct the particle beam, the beamline structure having an efficiency of 10% or more and a length of 6 meters or less; and a support structure, a portion of the beamline structure mounted on the support structure, and a portion of the beamline structure configured to move over the support structure. The energy degrader may be downstream of the beamline structure relative to the particle accelerator. The energy degrader may be configured to vary the energy of the particle beam and may be controllable. The efficiency of 10% or more may include 10% or more of the particles output from the particle accelerator being output from the beamline structure. An exemplary particle therapy system can include one or more of the following features, either alone or in combination.

[0012] The energy degrader can be the only mechanism for actively controlling a change in the energy of the particle beam after it is output by the particle accelerator and before it reaches the irradiation target. The gantry can be configured to not actively control the energy of the particle beam after it is output by the particle accelerator and before it reaches the energy degrader. The particle accelerator can have a volume of 3 cubic meters or less.

[0013] An exemplary particle therapy system may include a collimator downstream of the energy degrader relative to the particle accelerator. The collimator may be for blocking at least a portion of the particle beam before the at least a portion of the particle beam reaches the patient. An exemplary particle therapy system may include a configurable collimator downstream of the energy degrader relative to the particle accelerator. The configurable collimator may include a plurality of leaves, the plurality of leaves being dynamically reconfigurable during movement of the particle beam to change a shape of an edge defined by the plurality of leaves. The edge may be movable between at least a portion of the particle beam and a target of the particle beam such that a first part of the particle beam on a first side of the edge is at least partially blocked by the plurality of leaves and a second part of the particle beam on a second side of the edge is allowed to reach the target.

[0014] The support structure may have a diameter of 6 meters or less. The beamline structure may include an output channel for directing the particle beam to an isocenter of the particle therapy system. A distance between an output of the output channel and the isocenter may be 2 meters or less. A distance between an output of the output channel and the isocenter may be 1 meter or less. The beamline structure may have a length of 5 meters or less. The beamline structure may include an output channel for directing the particle beam to an isocenter of the particle therapy system. The output channel may be configured to bend the particle beam by 90 degrees or more in the presence of a magnetic field of 2.5 Tesla or more. The beamline structure may include an output channel for directing the particle beam to an isocenter of the particle therapy system. The isocenter may be 6 meters or less from the particle accelerator. The particle therapy system may have a footprint of 93 square meters or less. At least one of the particle accelerator or the gantry may generate 10 millisieverts or less of neutrons per Gray of dose delivered by the particle beam.

[0015] The beamline structure can include an output channel for bending the particle beam by 90° or more. The particle therapy system can include a scanning system for moving the particle beam in at least two dimensions across at least a portion of the beam field. The scanning system can include a first scanning magnet and a second scanning magnet. The first scanning magnet can be in the path of the particle beam and can be in the output channel or upstream of the output channel with respect to the particle accelerator.

[0016] The second scanning magnet can be in the path of the particle beam and can be downstream of the output channel relative to the particle accelerator. The first scanning magnet can be in the output channel. The first scanning magnet can be in the beamline structure but not in the output channel. At least one of the first scanning magnet or the second scanning magnet can include a superconducting magnet. The beamline structure can include an output channel for bending the particle beam by 90° or more in combination with a scanning magnet downstream of the output channel relative to the particle accelerator, where the scanning magnet includes one or more superconducting magnets.

[0017] An exemplary particle therapy system includes a particle accelerator configured to output particles as a particle beam that is monoenergetic, and a gantry. The exemplary gantry includes an output channel including magnetic components configured to bend the particle beam by at least 90° in the presence of a magnetic field of at least 2.5 Tesla (T); the output channel includes a support structure mounted for movement at least in part around an irradiation target; and a conduit for directing the particle beam to the output channel. The exemplary particle therapy system includes an energy degrader downstream of the output channel relative to the particle accelerator, and a scanning system including two or more scanning magnets. At least one of the scanning magnets is upstream of at least some of the magnetic components in the output channel relative to the particle accelerator. The two or more scanning magnets can be separated by air or other magnetic components and can be configured to move the particle beam across at least a portion of the beam field in at least two dimensions. The exemplary particle therapy system can include one or more of the following features, either alone or in combination.

[0018] The energy degrader may be the only mechanism for actively controlling the change in energy of the particle beam after it is output by the particle accelerator and before it reaches the irradiation target. The gantry may be configured not to actively control the energy of the particle beam after it is output by the particle accelerator and before it reaches the energy degrader. The conduit and the output channel together may have an efficiency of 10% or more. An efficiency of 10% or more may include that 10% or more of the particles output from the particle accelerator are output from the output channel. The two or more scanning magnets may be or include one or more superconducting magnets. The two or more scanning magnets may include a first scanning magnet, the first scanning magnet being in the output channel or being upstream of the output channel relative to the particle accelerator. The scanning system may include a second scanning magnet downstream of the output channel. The first scanning magnet can be configured to move the particle beam across at least a portion of the beam field in at least two dimensions, and the second scanning magnet can be configured to move the particle beam across at least a portion of the beam field in at least two dimensions. The first scanning magnet can be configured to move the particle beam across at least a portion of the beam field in only a first dimension, and the second scanning magnet can be configured to move the particle beam across at least a portion of the beam field in only a second dimension. The second dimension can be different from the first dimension.

[0019] The gantry can be 5 meters or less in length as measured from the particle accelerator. The magnetic components can include a first magnetic dipole for bending the particle beam toward the irradiation target; and a second magnetic dipole upstream of the first magnetic dipole relative to the particle accelerator. The second magnetic dipole can be configured to bend the particle beam toward the first magnetic dipole. The gantry and the particle accelerator can be in the same space and may not be separated by a shielding external to the particle accelerator that is more than 30 centimeters in thickness.

[0020] An exemplary particle therapy system can have a footprint of 75 square meters or less. The particle accelerator can include a synchrocyclotron, and the particle therapy system can be sized to fit into a vault designed for a linear accelerator (LINAC). The energy degrader can be mounted on a gantry between the output channel and an isocenter of the particle therapy system. The energy degrader can include a structure movable into and out of the path of the particle beam. The structure can include a plate. The structure can include a wedge.

[0021] A distance between an output of the output channel and an isocenter of the particle therapy system can be 2 meters or less. A distance between an output of the output channel and an isocenter of the particle therapy system can be between 0.8 meters and 1.4 meters. The particle accelerator can include a synchrocyclotron. The particle accelerator can include a synchrotron.

[0022] The magnetic components in the output channel may include a first magnetic dipole; a second magnetic dipole upstream of the first magnetic dipole relative to the particle accelerator; and one or more focusing magnetic quadrupoles and one or more defocusing magnetic quadrupoles between the first and second magnetic dipoles. The particle accelerator may include an active return system. The active return system may include a superconducting coil that conducts current in a reverse direction relative to a main superconducting coil in the particle accelerator.

[0023] An exemplary particle therapy system may include one or more imaging devices mounted to a support structure; and a control system configured to control rotation of the one or more imaging devices along the support structure around the irradiation target. The one or more imaging devices may include one or more of a computed tomography (CT) scanner, a two-dimensional (2D) X-ray device, a magnetic resonance imaging (MRI) device, a fan-beam CT scanner, a 2D camera, a three-dimensional (3D) camera, a surface imaging device, or a cone-beam CT scanner.

[0024] An exemplary particle therapy system includes a particle accelerator configured to output a particle beam; and a gantry configured to connect to the particle accelerator. The gantry can include a beamline structure configured to receive the particle beam from the particle accelerator, and the gantry can include magnetic components for directing the particle beam to a particle beam output device downstream of the beamline structure relative to the particle accelerator. The scanning system is configured to move the particle beam in at least two dimensions across at least a portion of the beam field. The scanning system can include a first scanning magnet positioned in the beamline structure and a second scanning magnet positioned downstream of the first scanning magnet relative to the particle accelerator. The first scanning magnet can be separate from the second scanning magnet. At least one of the first scanning magnet or the second scanning magnet can include a superconducting magnet.

[0025] The first scanning magnet can be configured to move the particle beam across at least a portion of the beam field in at least two dimensions. The second scanning magnet can be configured to move the particle beam across at least a portion of the beam field in at least two dimensions. The first scanning magnet can be configured to move the particle beam across at least a portion of the beam field in only a first dimension. The second scanning magnet can be configured to move the particle beam across at least a portion of the beam field in only a second dimension. The second dimension can be different from the first dimension. The first scanning magnet can be positioned between magnetic components included in the beamline structure, and the second scanning magnet can be positioned in a particle beam output device. The particle beam output device can be or include a nozzle.

[0026] The first scanning magnet can be positioned in the beamline structure, and the second scanning magnet can be positioned downstream of the beamline structure with respect to the particle accelerator. The beamline structure can include an output channel configured to bend the particle beam by at least 90° toward the irradiation target in the presence of a magnetic field of at least 2.5 Tesla (T). The beamline structure can include a magnetic dipole for bending the particle beam and two or more magnetic quadrupoles or hexapoles disposed between the magnetic dipoles along the length of the beamline structure. The beamline structure can include an output channel having magnetic dipoles disposed in series to bend the particle beam by at least 90°. The magnetic dipoles can include at least a first magnetic dipole and a second magnetic dipole. The first magnetic dipole and the second magnetic dipole can be between the first scanning magnet and the second scanning magnet.

[0027] The particle accelerator and the gantry can be located in the same treatment space. The treatment space can include a first compartment and a second compartment, with no shielding between the first and second compartments or with a shielding between the first and second compartments that is less than 30 centimeters thick. The particle accelerator can be configured to output a particle beam at one of two energy levels. One of the energy levels can be greater than another of the energy levels.

[0028] The particle accelerator can include a main superconducting coil for generating a magnetic field for accelerating particles to create a particle beam. The particle accelerator can include an active return coil for conducting a current in an opposite direction to the main superconducting coil. The particle beam can be output from the particle beam output device at a FLASH dose. The particle beam can be output from the particle beam output device at a dose greater than 20 Gray per second for a duration of less than 5 seconds.

[0029] The gantry may include a support structure to which at least a portion of the beamline structure is mounted. The particle therapy system may include an imaging system, the imaging system coupled to the support structure and configured for rotation about the support structure. The imaging system may be configured to capture an image of the irradiation target, the imaging system may include one or more of a computed tomography (CT) scanner, a two-dimensional (2D) X-ray device, a magnetic resonance imaging (MRI) device, a fan-beam CT scanner, a 2D camera, a three-dimensional (3D) camera, a surface imaging device, or a cone-beam CT scanner.

[0030] The particle therapy system can include a control system programmed to control output of the particle beam to the irradiation target based on the image. The particle accelerator can be configured to output the particle beam at multiple energy levels.

[0031] The particle therapy system can include a control system for selecting between multiple energy levels.

[0032] An exemplary particle therapy system includes a particle accelerator having a volume of 2 cubic meters or less. The particle accelerator is configured to output particles as a particle beam that is monoenergetic. The particles are output to the irradiation target with an efficiency of 10% or more. An efficiency of 10% or more includes 10% or more of the particles output from the particle accelerator reaching the irradiation target. The particle accelerator can be stationary and the particle beam can be fixed to be output from the particle accelerator in only a single direction. The treatment couch can be configured to move with three or more degrees of freedom relative to the particle beam. An exemplary particle therapy system can include one or more of the following features, either alone or in combination.

[0033] The particle therapy system can have a footprint of 75 square meters or less. The treatment couch can be configured to move with at least four degrees of freedom relative to the particle beam. The particle therapy system can include an energy degrader, which is the only mechanism for actively controlling the change in energy of the particle beam after it is output by the particle accelerator and before it reaches the irradiation target.

[0034] An exemplary scanning magnet is for use in a particle therapy system to treat an irradiation target. The scanning magnet can be configured to move a particle beam in at least two dimensions relative to the irradiation target. The scanning magnet includes a first set of superconducting coils for moving the particle beam in a first dimension relative to the irradiation target; and a second set of superconducting coils for moving the particle beam in a second dimension relative to the irradiation target, the second dimension being different from the first dimension. The scanning magnet can include one or more of the following features, either alone or in combination:

[0035] The second dimension may be orthogonal to the first dimension. The scanning magnet may include an electrically non-superconducting material separating the first set of superconducting coils from the second set of superconducting coils. The scanning magnet may include an electrically non-conducting material separating the first set of superconducting coils from the second set of superconducting coils. The scanning magnet may include a cryocooler for maintaining the first set of superconducting coils and the second set of superconducting coils at a high temperature superconducting temperature.

[0036] Each coil in at least one of the first set of superconducting coils or the second set of superconducting coils may include a rare earth barium copper oxide (ReBCO) superconducting layer. Each coil in at least one of the first set of superconducting coils or the second set of superconducting coils may include a silver (Ag) cap layer adjacent to the rare earth barium copper oxide (ReBCO) superconducting layer. Each coil in at least one of the first set of superconducting coils or the second set of superconducting coils may include a buffer layer stack adjacent to the rare earth barium copper oxide (ReBCO) superconducting layer. Each coil in at least one of the first set of superconducting coils or the second set of superconducting coils may include a metal substrate layer adjacent to the buffer layer stack. Each coil in at least one of the first set of superconducting coils or the second set of superconducting coils may include a copper layer encasing the coil.

[0037] The scanning magnet can include a cryostat containing a first set of superconducting coils and a second set of superconducting coils. The cryostat can be for maintaining the first set of superconducting coils and the second set of superconducting coils at a high temperature superconducting temperature. Each coil in at least one of the first set of superconducting coils or the second set of superconducting coils can include a superconducting layer constructed from YBCO or BSCCO.

[0038] Any two or more of the features described in this specification (including this Summary section) may be combined to form an implementation not specifically described in this specification.

[0039] Control of the various systems, or portions thereof, described herein may be implemented via a computer program product including instructions stored on one or more non-transitory machine-readable storage media and executable on one or more processing devices (e.g., programmed logic such as a microprocessor, an application specific integrated circuit, or a field programmable gate array, etc.). The systems, or portions thereof, described herein may be implemented as an apparatus, method, or medical system that may include one or more processing devices and a computer memory for storing executable instructions for implementing control of the described functions. The devices, systems, and / or components described herein may be configured, for example, through design, construction, arrangement, installation, programming, operation, activation, deactivation, and / or control.

[0040] The details of one or more implementations are set forth in the accompanying drawings and the description below. Other features and advantages will be apparent from the description and drawings, and from the claims. [Brief description of the drawings]

[0041] [Figure 1] 1 is a diagram illustrating a partially transparent perspective view of an exemplary particle therapy system having an exemplary gantry of the type described herein. [Diagram 2] FIG. 2 is a cutaway side view of components of the particle therapy system shown in FIG. 1, including an exemplary gantry. [Diagram 3]FIG. 2 is a cutaway enlarged side view of a nozzle that may be part of the gantry shown in FIG. 1 and components included in an example beamline structure. [Figure 4] FIG. 2 is a cutaway enlarged side view of a nozzle that may be part of the gantry shown in FIG. 1 and components included in an example beamline structure. [Diagram 5] FIG. 2 is a front view of an exemplary scanner magnet configured to scan a particle beam in two orthogonal dimensions. [Figure 6] FIG. 2 is a cutaway enlarged side view of a nozzle that may be part of the gantry shown in FIG. 1 and components included in an example beamline structure. [Figure 7] FIG. 2 is a front view of an exemplary scanner magnet configured to scan a particle beam in a single dimension. [Figure 8] FIG. 2 is a front view of an exemplary scanner magnet configured to scan a particle beam in a single dimension. [Figure 9] FIG. 2 is a cutaway enlarged side view of a nozzle that may be part of the gantry shown in FIG. 1 and components included in an example beamline structure. [Figure 10] FIG. 2 is a cutaway enlarged side view of a nozzle that may be part of the gantry shown in FIG. 1 and components included in an example beamline structure. [Figure 11] FIG. 1 is a front view of an exemplary superconducting scanning magnet configured to scan a particle beam in two orthogonal dimensions. [Figure 12a] FIG. 2 is a front view of an exemplary superconducting scanning magnet configured to scan a particle beam in a single dimension. [Figure 12b] FIG. 12b is a front view of an exemplary superconducting scanning magnet configured to scan a particle beam in a single dimension orthogonal to the dimension of FIG. 12a. [Figure 13] 1 is a drawing showing a perspective view of an exemplary configurable collimator that may be part of the particle therapy system of claim 1. [Figure 14] FIG. 14 illustrates a partially transparent front view of the configurable collimator of FIG. 13. [Figure 15] FIG. 15 shows a partially transparent perspective view of the configurable collimator of FIGS. 13 and 14. [Figure 16] FIG. 2 is a block diagram of an example treatment space configured to accommodate all or a portion of the particle therapy system of FIG. [Figure 17] 4 is a graph illustrating an example horizontal (x) and vertical (y) particle beam envelope produced in an example gantry described herein. [Figure 18] 1 is a graph illustrating an example achromatic grating design for a beamline of an example gantry described herein. [Figure 19] 4 is a graph illustrating the results produced by scanning a particle beam in the horizontal (x) and vertical (y) planes using an example gantry described herein. [Figure 20] FIG. 2 is a cutaway side view of components within an example particle accelerator that may be used with the particle therapy systems described herein. [Figure 21] FIG. 1 is a side view of an exemplary fixed beam particle therapy system. [Figure 22] FIG. 1 is a perspective view of an exemplary energy degrader. [Diagram 23] 2 is a front cutaway view of an exemplary superconducting magnet that may be used as a scanning magnet in the particle therapy system of claim 1. [Figure 24] FIG. 24 is a cutaway view of an exemplary superconducting coil that may be used in the superconducting magnet of FIG. 23. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0042] Like reference numbers in the different drawings indicate like elements.

[0043] Described herein is an exemplary particle therapy system that can accommodate a patient and an accelerator in the same space. The exemplary system includes a particle accelerator, which can be, but is not limited to, a synchrocyclotron, which has low radiation leakage and is small enough to fit in a standard linear accelerator (LINAC) vault. The system also includes a medical gantry configured to deliver a charged particle beam (e.g., protons or ions, etc.) output from the accelerator to treat a tumor or other condition in a patient. The gantry includes beamline structures for directing the particle beam from the accelerator to a treatment location and for delivering the particle beam to the treatment location. The beamline structures include magnetic components (e.g., one or more magnetic dipoles and one or more magnetic quadrupoles, etc.) for directing the particle beam toward the treatment location. To enable delivery of the particle beam in the same space used for the procedure, particularly in a relatively small space such as a standard LINAC vault, at least some of the magnetic components in the beamline structure are configured to bend the particle beam at a right angle or at an obtuse angle. In one example, the magnetic components are constructed and arranged to bend the particle beam by 90° or more.

[0044] Additionally, the implementations of the particle therapy system described herein combine the functionality of large diameter superconducting magnets with the use of upstream scanning magnets, making the particle therapy system relatively compact. While compact in construction, the exemplary particle therapy system is configured to allow for beam focusing, beam scanning, beam bending, and beam rotation, as described below.

[0045] Figure 1 illustrates an exemplary implementation of a particle therapy system 10 of the type described in the preceding paragraph. As shown in Figure 1, the particle therapy system 10 includes a particle accelerator 12, an example of which is described herein.

[0046] In this example, the particle accelerator 12 is a synchrocyclotron having a superconducting electromagnetic structure that generates a maximum magnetic field strength of 2.5 Tesla (T) or more, or 3 T or more. In this regard, a superconductor is an element or metal alloy, such as niobium-tin (NbsSn), that loses most, if not all, electrical resistance when cooled below a threshold temperature. As a result, electrical current flows through the superconductor substantially unimpeded. Thus, in their superconducting state, superconducting coils can conduct larger currents than regular electrical wires of the same size. Due to the high amount of current they can conduct, superconducting coils are particularly useful in particle therapy applications.

[0047] An exemplary synchrocyclotron is configured to output protons or ions as a monoenergetic particle beam having an energy level of 150 megaelectron volts (MeV) or greater. An exemplary synchrocyclotron has a volume of 4.5 cubic meters (m 3 ) or less and a weight of 30 tons (T) or less. Due to its size, this type of particle accelerator is referred to as "compact." However, as described herein, synchrocyclotrons or other types of particle accelerators having other weights, dimensions, magnetic fields, and / or energy levels may also be used in particle therapy system 10.

[0048] The particle therapy system 10 also includes a gantry 14. The gantry 14 includes a ring-shaped or circular support structure 15 and a beamline structure 16. The combination of the support structure 15 and the beamline structure 16 can be referred to as a "compact gantry" due to its relatively small size. The beamline structure 16 includes an output channel 17 that mounts to the support structure 15 and a conduit 18 that directs the particle beam to the output channel. The gantry 14 also includes one or more motors (not shown) for moving the output channel 17 around the support structure 15 to a treatment location 19. The treatment location can include a system isocenter where a patient can be positioned for treatment. In one example, the motor can move the output channel 17 along a track on the structure 15, resulting in a rotation of the output channel 17 relative to the treatment location 19. In one example, the structure to which the output channel 17 is attached can rotate relative to the treatment location 19, resulting in a rotation of the output channel 17 relative to the treatment location. In some implementations, the rotation enabled by the gantry 14 allows the output channel 17 to be positioned at any angle relative to the treatment position. For example, the output channel 17 can rotate through 360°, and so the output channel 17 can be positioned at 0°, 90°, 270°, and back to 0° / 360°, or any angle between these rotational positions.

[0049] As previously mentioned, the beamline structure 16 is configured to direct the particle beam from the accelerator 12 to a treatment location 19. To this end, the output channel 17 includes magnetic components for bending the particle beam toward the treatment location. In addition, the beamline structure 16 includes a conduit 18 containing magnetic components along the beamline that directs the particle beam from the particle accelerator 12 to the output channel 17.

[0050] 2 and 3, the conduit 18 of the exemplary beamline structure 16 includes non-superconducting magnetic quadrupoles 21 and 22 and a superconducting magnetic dipole 23. The electromagnetically shielded outer shell of the beamline structure 16 is partially omitted from the illustrations of FIGS. 2 and 3. The magnetic quadrupoles 21 and 22 are configured to keep the particle beam focused and traveling straight or substantially straight (e.g., no more than 5% deviation from straight) through the beamline structure 16. The magnetic quadrupoles 21 and 22 are configured to focus the particle beam to maintain a substantially consistent cross-sectional area of ​​the particle beam (e.g., within a tolerance of ±5%). The magnetic dipole 23 is configured to bend the particle beam toward the output channel 17 as shown. The magnetic dipole 23 can be configured to bend the particle beam anywhere within a range of 20° to 80° with respect to a horizontal direction 24. In general, a larger bending angle can reduce the distance between the particle accelerator 12 and the treatment location 19 or system isocenter, thereby reducing the space required to accommodate the gantry and thus the size of the particle therapy system. For example, replacing the magnetic dipoles 23 with one or more superconducting magnetic dipoles that bend the particle beam by 80° or more (e.g., 90° or more) can further reduce the distance from the particle accelerator 12 to the support structure 15 and thus to the treatment location 19 and isocenter.

[0051] In some implementations, higher order magnetic components can be used instead of or in addition to any of the magnetic quadrupoles described herein. For example, the beamline structure can include one or more magnetic sextupoles instead of or in addition to the magnetic quadrupoles. The magnetic sextupoles are configured to keep the particle beam focused and traveling straight or substantially straight (e.g., no more than 5% deviation from straight) through the beamline structure 16. The magnetic sextupoles can also be configured to maintain a consistent cross-sectional area of ​​the particle beam (e.g., within a tolerance of ±5%). The sextupoles can also compensate for chromatic effects of the quadrupole magnets. Compared to magnetic quadrupoles, magnetic sextupoles have a greater focusing effect on particles that are displaced farther from the axis that defines the ideal location of the beamline.

[0052] Referring back to FIG. 3, in this example, the conduit 18 of the beamline structure 16 also includes two non-superconducting magnetic quadrupoles 26 and 27. The magnetic quadrupoles 26 and 27 are configured to maintain the particle beam focused and traveling straight or substantially straight (e.g., no more than 5% deviation from straight) through the beamline structure 16. The magnetic quadrupoles 26 and 27 are configured to maintain a consistent cross-sectional area of ​​the particle beam (e.g., within a tolerance of ±5%). As previously described, higher order magnetic components may be substituted for one or more of the magnetic quadrupoles to improve focusing.

[0053] The particle therapy system 10 also includes one or more scanning magnets 30 in the path of the particle beam, which are configured to move the particle beam across at least a portion of a beam field that covers all or a portion (i.e., at least a portion) of the irradiation target. In some examples, the beam field includes a maximum range in which the particle beam can be moved across a plane parallel to the treatment area on the patient for a given position of the compact gantry. Movement of the particle beam across the beam field results in movement across at least a portion of the irradiation target at the treatment location 19. The scanning magnet can be sized and configured to move the particle beam across a beam field having an area of ​​20 centimeters (cm) by 20 cm or more, although the system 10 is not limited to any particular beam field size. For example, the scanning magnet can have an aperture of 20 cm by 20 cm or more, although the scanning magnet is not limited to any particular aperture size.

[0054] The scanning magnets can be positioned at different locations in the particle therapy system. For example, in the beamline structure 16a shown in FIG. 4 (which is a variation of the beamline structure 16), all of the scanning magnets 30a can be positioned in the nozzle 40a along with the energy degrader 41a and the collimator 44a (both described below) on the path of the particle beam between the output channel 17a and the treatment location. With reference to FIG. 5, the exemplary scanning magnet 43 is controllable in two dimensions (e.g., Cartesian XY dimensions) to position the particle beam in those two dimensions and move the particle beam across at least a portion of the irradiation target. In this example, the scanning magnet 43 includes two coils of a first set 45 and two coils of a second set 46, where the two coils of the first set 45 control the particle beam movement in the Cartesian X dimension of a defined coordinate system, and the two coils of the second set 46 are orthogonal to the two coils of the first set and control the particle beam movement in the Cartesian Y dimension. Control over the movement of the particle beam can be achieved by varying the current through one or both sets of coils, thereby varying the magnetic field produced thereby. By appropriately varying the magnetic field, the magnetic field acts on the particle beam, causing it to move in the X and / or Y dimensions across the beam field, and thus across the irradiation target.

[0055] In some implementations, there may be more than one scanning magnet. An implementation including multiple scanning magnets at different points along the path of the particle beam and separated by air or a structure (e.g., magnets or beam absorbing plates, etc.) may be referred to as a split scanning system. For example, in the beamline structure 16b shown in FIG. 6 (which is a variation of the beamline structure 16), there may be multiple (e.g., two) scanning magnets 30b1 and 30b2 between the output channel 17b and the treatment location. The scanning magnets may be positioned in the nozzle 40b along with the energy degrader 41b and collimator 44b on the path of the particle beam between the output channel 17b and the treatment location. The scanning magnets may be in separate locations and separated by air or an energy degrading structure. For example, in this implementation, a first scanning magnet 30b1 is capable of moving the particle beam in two dimensions (e.g., Cartesian X and Y dimensions) and a second scanning magnet 30b2 is capable of moving the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, the scanning magnets 30b1 and 30b2 can have the same construction and operation as the scanning magnets shown in Figure 5. Each magnet 30b1 and 30b2 is capable of partially moving the beam, and the combined movement produced by the two magnets creates the desired movement specified in the treatment plan.

[0056] In a variation of the implementation of FIG. 6, the scanning magnet 30b1 can move the particle beam in only one dimension (e.g., the Cartesian X dimension) and the scanning magnet 30b2 can move the particle beam in only one dimension (e.g., the Cartesian Y dimension). One magnet 30b1 can be upstream of the other magnet 30b2 with respect to the particle accelerator as shown. They can be separated by air or an energy degrading structure as described above. FIGS. 7 and 8 show exemplary magnets 90 and 91, respectively, with orthogonal coils (coil 90a is orthogonal to coil 91a) for moving the particle beam in different dimensions. In this example, the scanning magnet 30b1 can be of the type shown in FIG. 7 and can include a first set of coils 90a, and the scanning magnet 30b2 can be of the type shown in FIG. 8 and can include a second set of coils 91a that are orthogonal to coils 90a. Each magnet 30b1, 30b2 is capable of partially moving the beam, and the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0057] In some implementations, one or more (e.g., all or less than all) of the scanning magnets can be positioned in the beamline structure. For example, in the beamline structure 16c shown in FIG. 9 (which is a variation of the beamline structure 16), there can be multiple (e.g., two) scanning magnets including a first scanning magnet 30c1 and a second scanning magnet 30c2, where the first scanning magnet 30c1 is positioned in the beamline structure 16c, and the second scanning magnet 30c2 is positioned in the nozzle 40c outside the beamline structure, together with the energy degrader 41c and the collimator 44c, between the output channel 17 and the treatment position. The first scanning magnet 30c1 can be positioned between the magnetic components included in the beamline structure 16c. For example, the first scanning magnet 30c1 can be positioned in the output channel 17c upstream of the magnetic dipole 32c relative to the particle accelerator, or as shown in FIG. 9, the first scanning magnet 30c1 can be positioned upstream of the output channel 17c relative to the particle accelerator. In one example, the first scanning magnet 30c1 can be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions), and the second scanning magnet 30c2 can be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, the scanning magnets 30c1 and 30c2 can have the same construction and operation as the scanning magnets shown in FIG. 5. Each magnet 30c1 and 30c2 can move the beam partially, and the combined movement produced by the two magnets creates the desired movement specified in the treatment plan.

[0058] In a variation of the implementation of FIG. 9, the first scanning magnet 30c1 can be configured to move the particle beam in only one dimension (e.g., the Cartesian X dimension) and the second scanning magnet 30c2 can be configured to move the particle beam in only one dimension (e.g., the Cartesian Y dimension). In this example, the scanning magnet 30c1 can include a first set of coils and the scanning magnet 30c2 can include a second set of coils that are orthogonal to the first set of coils. The magnets 30c1 and 30c2 can have a configuration similar to the magnets shown in FIG. 7 and FIG. 8 in this example. Each magnet 30c1 and 30c2 can be configured to move the beam partially, and the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0059] In some implementations, all of the scanning magnets can be positioned in the beamline structure. As shown in the split scanning system of FIG. 10, both the first scanning magnet 30d1 and the second scanning magnet 30d2 can be positioned in the beamline structure 16d (which is a variation of the beamline structure 16). The scanning magnets may not be positioned in the nozzle 40d, which in this example includes an energy degrader 41d and a collimator 44d. In other examples, one or more scanning magnets can also be present in the nozzle. The first scanning magnet 30d1 and the second scanning magnet 30d2 can be positioned between the magnetic components included in the beamline structure 16d. For example, as shown in FIG. 10, the first scanning magnet 30d1 can be positioned in the output channel 17d upstream of the magnetic dipole 32d relative to the particle accelerator, or the first scanning magnet can be positioned upstream of the output channel 17d relative to the particle accelerator. The second scanning magnet 30d2 can be positioned upstream of the first scanning magnet 30d1 with respect to the particle accelerator. In the example shown in FIG. 10, the second scanning magnet 30d2 precedes the output channel 17d in the beamline. The scanning magnets can be at separate locations in the beamline structure and separated by magnetic components (e.g., dipoles or quadrupoles, etc.) in the beamline structure and / or air. The separate locations can include different points or locations in series along the path of the particle beam or the length of the beamline structure. For example, as shown in FIG. 10, the magnetic dipole 31d is between the first scanning magnet 30d1 and the second scanning magnet 30d2. In another example, the scanning magnet 30d1 can be moved after the magnetic dipole 32d, such that both the magnetic dipoles 31d and 32d are between the scanning magnets 30d1 and 30d2. In another example, both scanning magnets 30d1 and 30d2 can be in output channel 17d, and magnetic dipoles 31d and 32d can be between scanning magnets 30d1 and 30d2.In one example, the first scanning magnet 30d1 can be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions) and the second scanning magnet 30d2 can be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, the scanning magnets 30d1 and 30d2 can have the same construction and operation as the scanning magnets shown in Figure 5. Each magnet 30d1 and 30d2 can partially move the particle beam, and the combined movement produced by the two scanning magnets creates the desired particle beam movement specified in the treatment plan.

[0060] In a variation of the implementation of FIG. 10, the first scanning magnet 30d1 can be configured to move the particle beam in only one dimension (e.g., the Cartesian X dimension) and the second scanning magnet 30d2 can be configured to move the particle beam in only one dimension (e.g., the Cartesian Y dimension). In this example, the scanning magnet 30d1 can include a first set of coils and the scanning magnet 30d2 can include a second set of coils that are orthogonal to the first set of coils. The magnets 30d1 and 30d2 can have a configuration similar to the magnets shown in FIGS. 7 and 8 in this example. Each magnet 30d1 and 30d2 can be configured to move the beam partially, and the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0061] In some implementations, there may be more than two scanning magnets positioned in the beamline structure and / or positioned between the output of the output channel and the treatment location. For example, there may be more than two scanning magnets positioned at various separate locations in the beamline structure. For example, there may be more than two scanning magnets positioned at various separate locations between the output of the output channel and the treatment location. In each case, the scanning magnets may be arranged in series.

[0062] In some implementations, there may be a single scanning magnet positioned upstream of the output of the output channel or elsewhere in the beamline structure. For example, as shown in Figures 2 and 3, a scanning magnet 30 may be positioned upstream of the output channel 17 and at the input of the output channel 17 relative to the particle accelerator. The scanning magnet 30 may be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, the scanning magnet 30 may have the same construction and operation as the scanning magnet shown in Figure 5. In this example, all particle beam movement is implemented by controlling current through one or more coils of the single scanning magnet.

[0063] In this regard, by positioning all or some of the scanning magnets within the beamline structure, it may be possible to reduce the size of the particle therapy system relative to systems that implement scanning outside the gantry.

[0064] In some implementations, one or more of the scanning magnets described herein can be superconducting. For example, one or more (including all) of the scanning magnets downstream of the output channel can be superconducting. For example, one or more (including all) of the scanning magnets in the beamline structure can be superconducting. In this regard, it can be difficult to accurately move the particle beam in the presence of a high magnetic field (such as that found in the beamline structure). The use of superconducting magnets for scanning allows the generation of a magnetic field of 2.5T or more or 3T or more to move the particle beam, which can overcome the effect of the high magnetic field (e.g., 2.5T or more or 3T or more) created by the beamline structure on the particle beam.

[0065] FIG. 11 illustrates an exemplary implementation of a superconducting scanning magnet 92 configured to move a particle beam in two dimensions, which can be used in the scanning implementations described herein. In this example, the scanning magnet 92 can have the same construction and operation as the scanning magnet 43 illustrated in FIG. 5. The superconducting magnet 92 includes a set of high-temperature superconducting coils 92a and 92b, which are similar in construction to the coils 46 and 45, respectively, of FIG. 5. Examples of high-temperature superconductors include, but are not limited to, YBCO (yttrium barium copper oxide) and BSCCO (bismuth strontium calcium copper oxide). The scanning magnet 92 is contained within a cryostat 94, which maintains the superconducting magnet at a superconducting temperature (e.g., a temperature above 77° Kelvin (K) or a temperature above 90° K). The cryostat can include a device configured to maintain the superconducting coils at a cryogenic temperature. The cryostat can maintain the temperature by thermally isolating the superconducting coils from room temperature. This is typically accomplished using vacuum insulation, thermal radiation shields, and / or superinsulation to reduce radiative heat transfer, as well as low thermal conductivity connections between room temperature and cryogenic temperatures. In some instances, liquid helium can be used to cool the coils to superconducting temperatures in a cryostat, for example, using conduction or immersion cooling. In conduction cooling, heat is transferred away from the superconducting coils using a thermal conductor. In immersion cooling, the superconducting coils can be in direct contact with the cryogen (such as, for example, liquid helium). In operation, a current is applied to coils 92a and 92b to generate the magnetic field used for scanning.

[0066] FIG. 12a shows an example of a superconducting magnet 95 configured to move a particle beam in only one dimension, which can be used in the scanning implementations described herein. The superconducting magnet includes a high-temperature superconducting coil set 95a configured to move a particle beam in only one dimension (e.g., Cartesian X or Y dimension). Examples of high-temperature superconductors include, but are not limited to, YBCO and BSCCO. The superconducting magnet 95 is contained within a cryostat 96 that maintains the superconducting magnet at a superconducting temperature (e.g., a temperature above 77° Kelvin (K)). For example, liquid helium can be used to cool the coils to the superconducting temperature. A current is applied to the coils 95a to generate a magnetic field used for scanning. FIG. 12b shows an example of a superconducting scanning magnet 97 configured to move a particle beam in only one dimension, which is different from (e.g., perpendicular to) the dimension in which the magnet 95 of FIG. 12a moves the particle beam. The superconducting magnet 97 includes a high temperature superconducting coil set 97a that is configured to move the particle beam in only one dimension (e.g., the Cartesian X or Y dimension). Examples of high temperature superconductors include, but are not limited to, YBCO and BSCCO. The superconducting magnet 95 is contained within a cryostat 98 that maintains the superconducting magnet at a superconducting temperature (e.g., above 77° Kelvin (K)). For example, liquid helium can be used to cool the coils to the superconducting temperature. A current is applied to the coils 97a to generate the magnetic field used for scanning.

[0067] FIG. 23 shows a front cutaway view of another exemplary implementation of a superconducting scanning magnet 150 configured to move a particle beam in two dimensions, which can be used in the scanning implementations described herein. In this example, the scanning magnet 150 can be contained in a cryostat (not shown) (e.g., such as those described above) to maintain the superconducting magnet at a superconducting temperature (e.g., between 30° K and 40° K in this example), although the cryostat is not limited to these temperatures. A cryocooler can be used to maintain the temperature of the cryostat at the superconducting temperature. The cryocooler includes a device for providing active cooling of the superconducting coils down to cryogenic temperatures. The cryocooler can be controlled by a control system described herein.

[0068] In FIG. 23, grid 151 shows the scanning beam aperture in both Cartesian X and Y dimensions (153 and 154, respectively). For example, grid 151 shows that scanning magnet 150 can move the particle beam ±5 cm in the X dimension and ±5 cm in the Y dimension relative to a reference 0,0 point 155. In other implementations, the scanning magnet can be configured to move the particle beam over a length that is greater than or less than ±5 cm in the X dimension and ±5 cm in the Y dimension. In FIG. 23, a set of superconducting coils 158 and 159 are wound around an electrically non-conducting or electrically non-superconducting material 160 to create an aperture 161 that contains grid 151. Inner superconducting coil 158 can be separated from outer superconducting coil 159 by an electrically non-conducting or electrically non-superconducting material 162. Superconducting coil 158 can be configured such that the magnetic field generated thereby is orthogonal to the magnetic field generated by superconducting coil 159. And superconducting coil 159 can be configured such that the magnetic field generated thereby is orthogonal to the magnetic field generated by superconducting coil 158. For example, the windings of superconducting coils 158 and 159 can be orthogonal to one another. In some implementations, the magnetic fields generated by superconducting coils 158 and 159 need not be orthogonal, but rather can be different (e.g., at an angle less than 90° to one another) and still allow for scanning in a grid, such as grid 151.

[0069] In this example, superconducting coils 158 control the movement of the particle beam in the X dimension. For example, a current runs through them to create a magnetic field. The strength of the magnetic field is proportional to the amount of current running through them. And the strength of the magnetic field is proportional to the amount the particle beam moves in the X dimension during a scan. In this example, superconducting coils 159 control the movement of the particle beam in the Y dimension. For example, a current runs through them to create a magnetic field. The strength of the magnetic field is proportional to the amount of current running through them. And the strength of the magnetic field is proportional to the amount the particle beam moves in the X dimension during a scan. Currents can run through superconducting coils 158 and 159 at the same time to create a cumulative magnetic field that moves the particle beam in both the X and Y dimensions. Currents can run through superconducting coils 158 and 159 at different times to cause the particle beam to move in the X or Y dimensions at separate times but still reach a target location.

[0070] An example of an electrically non-superconducting material that may be included in scanner magnet 150 is copper. However, scanner magnet 150 is not limited to the use of copper. The electrically non-superconducting material promotes heat dissipation, for example, during quenching of superconducting coils 158 and 159.

[0071] 24 illustrates a cross section of an exemplary superconducting coil 165 that may be used to implement each of the superconducting coils 158 and 159. The superconducting coil 165 includes a copper (Cu) stabilization layer 166 that wraps or surrounds the other layers of the superconducting coil 165. The superconducting coil 165 also includes a silver (Ag) cap layer 167, a rare earth barium copper oxide (ReBCO) superconducting layer 168 (or a layer of other high temperature superconducting material) adjacent to and in contact with the silver cap layer, a buffer layer stack 169 adjacent to and in contact with the ReBCO superconducting layer to prevent interdiffusion between the oxide and the metal substrate, and a substrate layer 170 adjacent to and in contact with the buffer layer stack. Examples of materials that may be included in the substrate layer include, but are not limited to, conductive metals (such as, for example, copper, nickel, or aluminum). Examples of materials that may be included in the buffer layer stack include, but are not limited to, SrRuO3 (strontium ruthenate - SRO) and LaNiO3 (LNO). The superconducting coil 165 may have a different configuration than that shown or may include different materials than that shown. For example, the copper stabilization layer may be omitted or materials other than copper may be used. Other types of superconducting materials may be used, such as YBCO and / or BSCCO.

[0072] Referring back to FIG. 3, the output channel 17 portion of the beamline structure 16 includes a large aperture superconducting magnetic dipole 31 arranged in series with a large aperture superconducting magnetic dipole 32. Examples of large apertures include, but are not limited to, 20 cm by 20 cm. Positioned between the magnetic dipole 31 and the magnetic dipole 32 are a plurality of large aperture superconducting magnetic quadrupoles 33, 34, and 35. In this example, the magnetic quadrupoles 33, 34, and 35 alternatively include one or more focusing magnets and one or more defocusing magnets for focusing and defocusing the particle beam, respectively, to maintain a substantially consistent cross-sectional area of ​​the particle beam. In this regard, the net effect on a particle beam passing through the alternating magnetic field gradient of the magnetic quadrupoles is to cause the beam to converge (i.e., to focus). In some implementations, magnetic quadrupole 33 includes a defocusing magnet, magnetic quadrupole 34 includes a focusing magnet, and magnetic quadrupole 35 includes a defocusing magnet. In some implementations, magnetic quadrupole 33 includes a focusing magnet, magnetic quadrupole 34 includes a defocusing magnet, and magnetic quadrupole 35 includes a focusing magnet. In some implementations, output channel 17 can include a different number of magnetic quadrupoles in different configurations and / or a different number of magnetic dipoles in different configurations. In some implementations, output channel 17 can include higher order magnetic components (e.g., hexapoles, etc.) instead of or in addition to the magnetic quadrupoles shown.

[0073] In some implementations, output channel 17 is configured to bend the particle beam in the presence of a magnetic field of 2.5 T, 3 T, or more in the beamline structure. For example, the magnetic field can be generated by running a current through one or more coils in a magnet in the beamline structure, and can be on the order of 2.5 T or more, 3 T or more, 4 T or more, 5 T or more, 6 T or more, 7 T or more, 8 T or more, 9 T or more, 10 T or more, 11 T or more, 12 T or more, 13 T or more, 14 T or more, or 15 T or more. In the presence of such magnetic fields, the magnetic components in the output channel 17 are configured to create a combined total bending angle of the particle beam anywhere in the range of 90° to 170° (e.g., 90°, 95°, 100°, 105°, 110°, 115°, 120°, 125°, 130°, 135°, 140°, 145°, 150°, 155°, 160°, 165°, or 170°). Alternatively, in some implementations, the output channel 17 is configured to bend the particle beam at a combined total bending angle less than 90° or greater than 170° (e.g., 180° or greater). In FIGS. 1-3, the output channel 17 is configured to bend the particle beam at a combined total bending angle of about 150° relative to the line 38. To achieve a bending magnitude having a value between 110° and 170°, magnetic dipole 31 can be configured to bend the particle beam within a range of 20° to 85° relative to line 38, and magnetic dipole 32 can be configured to bend the particle beam within a range of 20° to 85° relative to horizontal line 38.

[0074] In some implementations, the output channel 17 can include different numbers of magnetic structures in different configurations. For example, the output channel 17 can include a magnetic dipole of the type described herein, followed by three alternating magnetic quadrupoles of the type described herein, followed by a magnetic dipole, followed by three alternating magnetic quadrupoles of the type described herein, followed by a magnetic dipole of the type described herein. For example, additional magnetic components can be used to change where and how much the particle beam bends. Additional magnetic structures can also be used to focus the particle beam over a longer distance. Conversely, a fewer number of magnetic structures can be used to focus the particle beam over a shorter distance, for example, as shown in FIG. 1.

[0075] The nozzle 40 (FIG. 1) is positioned at the output or exit of the output channel 17. In the example of FIG. 1, the nozzle 40 is connected to the output channel 17 and moves with the output channel, if applicable. The nozzle 40 may or may not be considered part of the compact gantry. The nozzle 40 is an example of a particle beam output device. In this example, the nozzle 40 receives the particle beam from the output channel 17 and, in some implementations, conditions the particle beam for output to an irradiation target (such as, for example, a tumor in a patient) at a treatment location or isocenter. In this regard, as stated, the output channel 17 bends the particle beam by at least 90°. Thus, the particle beam is directed toward the treatment location or isocenter as it exits the output channel 17. In addition, as described herein, the scanning magnet 30 can move the particle beam in a plane and across the irradiation target.

[0076] In this respect, as previously described, the nozzle may contain one or more scanning magnets. The energy degrader is downstream of the scanning magnet, and the collimator is downstream of the scanning magnet. In Figures 2 and 3, the energy degrader 41 receives the scanning or moving particle beam from the scanning magnet. In this example, the energy degrader 41 is mounted on the gantry 14 (through the nozzle 40) between the output channel 17 and the irradiation target at the treatment position 19.

[0077] The energy degrader 41 is configured and controllable to change the energy of the particle beam before it reaches the irradiation target. In some implementations, the energy degrader is the only mechanism for actively controlling the change in the energy of the particle beam before it reaches the irradiation target. In some implementations, the energy of the particle beam is not actively controllable after it is output by the particle accelerator and before it reaches the energy degrader. For example, in such implementations, the components of the gantry between the particle accelerator and the energy degrader do not actively control and are not configured to control the beam energy. Stated further, the gantry or its beamline conduits are not configured to actively control the particle beam after it is output by the particle accelerator and before it reaches the energy degrader. In some cases, there may be some incidental changes in energy caused by travel through the beamline structures, but those changes are not actively controlled.

[0078] As previously mentioned, the particle beam output by the accelerator can be monoenergetic, and the energy degrader is the sole / only vehicle or primary vehicle for changing the beam energy during treatment of the irradiation target. A monoenergetic particle beam includes a particle beam having a single fixed energy level (e.g., 100 MeV, 150 Mev, 200 Mev, 250 Mev, etc.). A monoenergetic particle beam can deviate from the fixed energy level by a predetermined amount (e.g., ±10%, ±5%, ±2%, or ±1%, etc.) and still be considered monoenergetic. Switching the operation of the accelerator during a treatment, as required to switch particle beam energy during a treatment, can create excess stray neutrons, resulting in the need for increased shielding and reducing beamline efficiency. Neutrons can be generated by the particle accelerator and / or by magnetic components along the beamline structure. By using a particle beam that is monoenergetic during the procedure and by relying on energy degraders to vary the beam energy, the production of stray neutrons can be reduced or minimized and the efficiency of the beamline structure can be increased.

[0079] In one example, the energy degrader can include a plate that can be moved into or out of the path of the particle beam. In another example, the energy degrader can include a wedge that is at least partially overlapping and that can be moved in the path of the particle beam. An exemplary wedge is a polyhedron defined by two triangular and three trapezoidal faces. In either configuration, a variable amount of material can be moved into the path of the particle beam. The material absorbs energy from the particle beam, resulting in a reduced energy beam output. The more material there is in the path of the particle beam, the less energy the particle beam will have. In some implementations, the energy absorbing structure can be moved across all of the beam field or across only a portion of the beam field. As stated, in some examples, the beam field includes a maximum range that the particle beam can be moved across a plane parallel to the treatment area on the patient for a given position of the compact gantry.

[0080] Referring to FIG. 22, in one example, the energy degrader 48 is a range modulator that can be controlled to move the structure 42 into (and out of) the path of the particle beam to change the energy of the particle beam and thus the depth to which the dose of the particle beam will be deposited in the irradiation target. Examples of such energy absorbing structures include, but are not limited to, plates; polyhedrons, such as wedges, tetrahedrons, or toroidal polyhedrons; and curved three-dimensional shapes, such as cylinders, spheres, or cones. In this way, the energy degrader can cause the particle beam to deposit a dose of radiation into the interior of the irradiation target to treat a layer or column of the target. In this regard, as protons at a particular energy move through the tissue, the protons ionize the atoms of the tissue and deposit the dose primarily at a predefined tissue depth corresponding to that energy. Thus, the energy degrader is configured to move the particle beam in a Cartesian Z dimension through the target, thereby enabling the scanning magnet to perform scanning in a third dimension (Cartesian Z) in addition to the Cartesian X and Y dimensions. In some implementations, the energy absorbing structure of the energy degrader (e.g., a plate or wedge) can be configured to move during the movement (scanning) of the particle beam and to track or follow the particle beam during the movement. An exemplary energy degrader that tracks or follows particle beam movement is described in U.S. Pat. No. 10,675,487 (Zwart), entitled "High-Speed ​​Energy Switching." The contents of U.S. Pat. No. 10,675,487, particularly those related to energy degraders that track or follow particle beam movement (e.g., Figures 36 to 46 of U.S. Pat. No. 10,675,487 and the accompanying description), are incorporated herein by reference.

[0081] The Bragg peak is a prominent peak on the Bragg curve, which plots the energy loss of ionizing radiation during its travel through tissue. The Bragg peak represents the depth at which most of the radiation is deposited in the tissue. For protons, the Bragg peak occurs just before the particle comes to rest. Thus, the energy of the particle beam can be varied to change the location of its Bragg peak and thus the location at which the majority of the proton dose will be deposited at a depth in the tissue. In this regard, the particle accelerator can be a fixed energy particle accelerator. In a fixed energy particle accelerator, the particle beam always exits the particle accelerator at the same (or approximately the same) energy (e.g., within no more than 10%, 5%, or 1% deviation from the expected or target energy). In a fixed energy particle accelerator, the energy degrader is the primary or only vehicle for varying the energy of the beam applied to the irradiation target in the patient. In some implementations, the particle accelerators described herein are configured to output particle beams at a single energy or at two or more energies in a range between about 100 MeV and about 300 MeV (e.g., between 115 MeV and 250 MeV). The fixed energy output can be within that range (e.g., 250 MeV) or, in some examples, can be above or below that range.

[0082] In some implementations, the particle accelerator is a dual energy accelerator. In a dual energy particle accelerator, the particle beam exits the particle accelerator at one of two different energy levels (high energy level or low energy level). The terms "high" and "low" do not have a specific numerical connotation, but rather are intended to convey a relative magnitude. In some implementations, the particle accelerator described herein is configured to output a particle beam at two energies in a range between about 100 MeV and about 300 MeV. The high energy output and the low energy output can be values ​​within the range, or in some examples, can be values ​​above or below the range. The energy degraders described herein can be used with the dual energy particle accelerator to reduce the energy of the particle beam below one of the two energy levels and / or to fine tune between the two energy levels.

[0083] In the figures (FIGS. 2, 3), the nozzle 40 also includes a collimator 44 downstream of the energy degrader 41 relative to the particle accelerator (i.e., closer to the irradiation target). In one example, the collimator is a structure that can be controlled to allow some radiation to reach the target and block some radiation from reaching the patient. Typically, the radiation that passes is directed to the irradiation target to be treated, and the blocked radiation would otherwise strike and potentially damage healthy patient tissue. In operation, the collimator is placed in the radiation path between the output channel 17 and the irradiation target, and is controlled to create an opening of appropriate size and shape that allows some radiation to pass through the opening and reach the irradiation target, while the rest of the structure blocks some radiation from reaching adjacent tissue.

[0084] The collimator can be configurable, e.g., its aperture can be controlled and changed during a procedure. The collimator can be fixed or non-variable, e.g., the collimator can have a fixed shape that cannot be changed.

[0085] In some implementations, an exemplary configurable collimator component includes multiple leaves that are dynamically reconfigurable during movement of the particle beam to change a shape of an edge defined by the multiple leaves, the edge being movable between at least a portion of the particle beam and a target of the particle beam such that a first part of the particle beam on a first side of the edge is at least partially blocked by the multiple leaves and a second part of the particle beam on a second side of the edge is permitted to reach the target.

[0086] 13, 14, and 15 show an exemplary implementation of a configurable collimator 44a that can be used with the particle therapy system described herein. The collimator 44a includes carriages 113, 114, and 115 that are configured to hold (and move) the leaves described above in both vertical and horizontal directions relative to the irradiation target. As shown, vertical movement includes movement in a Cartesian Z dimension 117, and horizontal movement includes movement in a Cartesian X dimension 118 (the Cartesian Y dimension is the direction into (or out of) the page in FIGS. 13 and 14). FIGS. 14 and 15 show a portion of the carriage housing as transparent to show the components inside the housing, but the housing is not actually transparent.

[0087] Carriage 113 is referred to herein as the primary carriage, and carriages 114 and 115 are referred to herein as the secondary carriages. Secondary carriages 114, 115 are coupled to primary carriage 113 as shown in Figures 13-15. In this example, secondary carriages 114, 115 each include a housing that is secured to primary carriage 115 via corresponding members 118, 119. In this example, primary carriage 113 is movable vertically (Z dimension) along track 120 relative to the irradiation target and relative to the particle accelerator. Vertical movement of primary carriage 113 also causes secondary carriages to move vertically. In some implementations, secondary carriages move vertically in unison.

[0088] As shown in Figures 13-15, each secondary carriage 114, 115 is connected to a corresponding rod or rail 122, 123 along which the secondary carriages move. More specifically, in this example, motor 125 drives secondary carriage 114 to move along rod 122 towards or away from secondary carriage 115. Similarly, in this example, motor 126 drives secondary carriage 115 to move along rod 123 towards or away from secondary carriage 114. Control over the movement of the primary and secondary carriages is implemented to position the leaves relative to the irradiation target, as described herein. In addition, the leaves themselves are configured to move in and out of the carriage, as also described herein.

[0089] As shown in Fig. 15, the motor 130 drives the vertical movement of the primary carriage 113. For example, as shown in Fig. 15, the lead screw 131 is coupled to the housing 132, which holds the motors 125, 126 that drive the corresponding secondary carriages 114, 115, and the housing 132 is mounted on the track 120. The lead screw 131 is coupled to the motor 130 and is driven vertically by the motor 130. That is, the motor 130 drives the lead screw 131 in the vertical direction (Cartesian Z dimension). Because the lead screw 131 is fixed to the housing 132, this movement also causes the housing 132 (and thus the secondary carriages 114, 115) to move along the track 120, either towards or away from the irradiation target.

[0090] In this exemplary implementation, seven leaves 135, 136 are mounted on each secondary carriage 114, 115. Each secondary carriage can be configured to move its leaf horizontally into (or out of) the treatment area. Using linear motors, individual leaves on each secondary carriage can be independently and linearly movable in the X dimension relative to other leaves on the same secondary carriage. In some implementations, the leaves can be configured to move in the Y dimension. Additionally, the leaves on the secondary carriage 114 can be independently movable from the leaves on the other secondary carriages 115. These independent movements of the leaves on the secondary carriages, along with the vertical movements allowed by the primary carriage, allow the leaves to be moved into various configurations. As a result, the leaves can conform both horizontally and vertically to treatment areas that are randomly shaped with respect to both horizontal and vertical dimensions. The size and shape of the leaves can be varied to generate different conformances. For example, size and shape can be varied to address a single beam spot (and thus a single column). In some implementations, individual leaves on each secondary carriage may be independently and linearly movable using electric motors driving lead screws in the X dimension relative to other leaves on the same secondary carriage.

[0091] The leaves can be made of any suitable material that prevents or inhibits the penetration of radiation. The type of radiation used may dictate what material is used in the leaves. For example, if the radiation is an x-ray, the leaves can be made of lead. In the example described herein, the radiation is a proton beam or an ion beam. Thus, different types of metals or other materials can be used for the leaves. For example, the leaves can be made of nickel, tungsten, lead, brass, steel, iron, or any suitable combination thereof. The height of each leaf can determine how much it inhibits the penetration of radiation.

[0092] The implementation of the configurable collimator described with respect to Figures 13-15 is described in U.S. Patent Application Publication No. 2017 / 0128746 (Zwart), entitled "Adaptive Aperture." The contents of U.S. Patent Application Publication No. 2017 / 0128746, particularly with respect to the description of adaptive apertures (e.g., Figures 1-7 of U.S. Patent Application Publication No. 2017 / 0128746 and the accompanying description), are incorporated herein by reference.

[0093] Referring back to FIG. 1, as stated, the exemplary particle therapy system includes an isocentric gantry that is compact in size, which reduces the overall system size. In the implementation of the compact gantry 14, the diameter of the support structure 15 can be less than 6 meters (m), less than 5 m, or less than 4 m. In one example, the diameter of the support structure 15 is 4.8 m. The length of the beamline structure can be measured from the accelerator output and the system isocenter and equal to the distance between the accelerator output and the system isocenter. In the implementation of the compact gantry 14, the length of the beamline structure 16 can be less than 6 meters (m), less than 5 m, less than 4.5 m, or less than 4 m. In one example, the length of the beamline structure 16 is 4.2 m (FIG. 2). In this regard, the distance between the particle accelerator and the system isocenter or treatment position can be less than 6 m, less than 5 m, less than 4.5 m, or less than 4 m. In compact gantry 14 implementations, the distance between the output of output channel 17 and the system isocenter or treatment location is 2 m or less, 1.5 m or less, or 1 m or less. In compact gantry 14 implementations, the distance between the output of output channel 17 and the system isocenter or treatment location is between 0.8 m and 14 m. In one example, the distance between the output of output channel 17 and the system isocenter or treatment location is 1.01 m. Other implementations can have different dimensions than those listed here.

[0094] In some implementations, the particle therapy system has an area of ​​93 square meters (m 2 ) or less than 75m 2In some implementations, the particle therapy system is configured to fit into a vault designed for a LINAC. For example, the components of Figures 1-3 can be small enough to fit into a vault with the following dimensions: length 25 feet (7.62 m) or less, width 20 feet (6.09 m) or less, and height 11 feet (3.35 m) or less. For example, the components of Figures 1-3 can be small enough to fit into a vault with the following dimensions: length 25 feet (7.62 m) or less, width 26 feet (7.92 m) or less, and height 10 feet (3.05 m) or less. For example, the components of Figures 1-3 may be small enough to fit into, and have dimensions to fit into, a LINAC vault having a footprint of 26.09 feet (11 m) or less by 29.62 feet (9 m) or less with a height of 16.40 feet (5 m) or less. However, as noted, some implementations of particle therapy systems may have different dimensions, including but not limited to diameter, height, width, and length. In some implementations, the ceiling of an existing LINAC vault may not be high enough to support a full 360° rotation of or around the gantry. In such implementations, a pit 90 (Figure 1) may be dug under the floor of the LINAC vault to allow for rotation.

[0095] 1 and 16 show examples of treatment spaces 49 and 50 in which the particle therapy system 10 and its variations can be housed. The treatment spaces are implemented in LINAC vaults in these examples, which can be shielded using lead or other suitable materials (e.g., concrete, borated polyethylene, and / or steel, etc.). In this regard, particles (e.g., protons, etc.) generated by the particle accelerator but that do not reach the irradiation target generate secondary radiation through the production of high-energy neutrons. In one example, the particle accelerator 12 and / or gantry generate 10 millisieverts or less of such neutrons per Gray dose delivered by the particle beam.

[0096] The use of monoenergetic particle beams and reliance on energy degraders outside the beamline structures allows magnetic components in the beamline to efficiently guide the beam. More specifically, changes in beam energy in the beamline increase the production of stray neutrons, which in turn increases losses of the particle beam in the beamline, thereby worsening its efficiency. The monoenergetic particle beams used in the implementations of the systems described herein, combined with magnetic structures in the beamline, can lead to increased efficiency. In some cases, reducing the length of the beamline structures can also increase efficiency. In some implementations, the beamline structure variations described herein have efficiencies of 10% or more, 20% or more, 30% or more, 40% or more, 50% or more, 60% or more, 70% or more, 80% or more, or 90% or more. In some examples, efficiency is a measure of the percentage of particles output from the beamline structures that are output from the particle accelerator. Thus, an efficiency of 10% or more includes that 10% or more of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 20% or more includes that 20% or more of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 30% or more includes that 30% or more of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 40% or more includes that 40% or more of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 50% or more includes that 50% or more of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 60% or more includes that 60% or more of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 70% or more includes that 70% or more of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 80% or more includes that 80% or more of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 90% or more includes that 90% or more of the particles output from the particle accelerator are output from the beamline structure.In one example, the particle accelerator and gantry described herein delivers more than 70% of the proton beam to the patient, even at energies in the lower range of the accelerator.

[0097] The type of beamline efficiency described herein allows for a "single room" solution, where the particle accelerator, gantry, and patient are all in a single vault, as described above. In this vault, the particle accelerator itself may include shielding, but the separate compartments 60 and 61 (see FIG. 16) in the vault containing the patient and particle accelerator, respectively, do not need to be shielded from each other. In other words, in some implementations, there is no electromagnetic shielding external to the particle accelerator and gantry that separates the particle accelerator from the patient. Due to the low level of neutrons emitted by the system, shielding may not be required. In some implementations, there may be minimal shielding between the separate compartments 60 and 61. For example, the shielding may be 30 cm or less thick, 20 cm or less thick, or 10 cm or less thick.

[0098] 1, the particle therapy system 10 also includes a treatment couch 51. The treatment couch 51 is configured to move relative to a hole 53 in or through the gantry 14 to position a patient at a system isocenter or treatment position. In this example, the treatment couch 51 is mounted to a robotic arm 54. The arm 54 includes a first segment 55, a second segment 56, and a third segment 57. The first segment 55 is rotatably coupled to the second segment 56, which is rotatably coupled to the third segment 57. The treatment couch 51 is coupled to the third segment 57 as shown. The arm 54 is controllable to move the treatment couch 51 into and through the hole 53 to position a patient lying on the couch for treatment (i.e., to move the patient to the treatment position). In some implementations, the arm 54 can position the patient with two, three, four, five, or six degrees of freedom. Examples of two degrees of freedom are front-back and left-right movement; examples of three degrees of freedom are front-back, left-right, and up-down movement; examples of four degrees of freedom are front-back, left-right, up-down, and one of pitch, yaw, or roll movement; examples of five degrees of freedom are front-back, left-right, up-down, and two of pitch, yaw, or roll movement; examples of six degrees of freedom are front-back, left-right, up-down, pitch, yaw, and roll movement. In some implementations, the treatment couch can be replaced by or include a couch that can be at least partially tilted or converted into a chair and still be controllable with two, three, four, five, or six degrees of freedom to position the patient for treatment. In some implementations, the arm 54 can have a different configuration than that shown in FIG. 1. For example, the arm 54 can have two segments or four or more segments. Hydraulics, robotics, or both can control or implement non-planar movements of the treatment couch.

[0099] In some implementations, the output channel 17 may rotate at least partially (including completely) around the support structure 15, or the output channel may remain fixed on the support structure 15 and all or a portion of the support structure 15 may rotate around the treatment location. In some implementations, the output channel 17 may not rotate around the support structure 15, and the support structure may not rotate around the patient.

[0100] Instead, the output channel can remain stationary, thereby providing a particle beam that is fixed in one direction. In implementations such as these, the treatment couch or other seat moves relative to the fixed beam during the treatment. In some systems described herein, the location of the particle beam can be set through a rotation of the gantry, after which the beam remains fixed except for a scanning movement across the irradiation target, and the treatment couch or other seat moves during the treatment. In some implementations, the treatment can be implemented using a combination of gantry movement and treatment couch (or other seat) movement. For example, the output channel can be positioned, the beam can be temporarily fixed, during which time the treatment couch moves to implement the treatment. The output channel can then be repositioned to temporarily fix the beam at the new location. The treatment can be implemented at the new location through couch movement. These operations can be repeated as defined by a treatment plan designed for use with the particle therapy system.

[0101] The particle therapy system 10 can be an intensity modulated proton therapy (IMPT) system. IMPT systems allow spatial control of a localized beam of protons that can have variable energy and / or intensity. IMPT utilizes the charged particle Bragg peak (as stated, a characteristic peak of dose at the edge of the particle's delivery range) combined with modulation of particle beam variables to generate a target-local modulation of dose that achieves the objectives described in the treatment plan. IMPT can include directing a particle beam toward an irradiation target at different angles and different intensities to treat the target. In some implementations, the particle beam can be scanned (e.g., moved) across layers of the irradiation target, with each layer being treated one or more times from the same or different angles. Moving across the irradiation target to implement scanning can be performed using a scanning magnet as described herein.

[0102] FIG. 17 shows an exemplary horizontal (x) beam envelope 63 and vertical (y) beam envelope 64 (e.g., cross-section) along the length 29 of a compact gantry described herein. The X and Y dimensions of the beam spot cross-section are determined for magnetic quadrupoles 21 and 22, magnetic dipole 23, magnetic quadrupoles 26 and 27, magnetic dipole 31, magnetic quadrupoles 33, 34, and 35, and magnetic dipole 32. The beam size is determined based on beam optics calculations using measured beam parameters at the exit of particle accelerator 12 and design parameters of all beamline magnets. In some implementations, the beam spot radius at the isocenter (e.g., treatment position) is approximately 3 millimeters (mm) in both x and y. In some implementations, for a 200 MeV to 230 MeV proton beam, the magnetic field at the magnetic dipoles in the beamline structures 16 is 4 T or less, and the bending radius of the beam at each of the magnetic dipoles is approximately 0.6 meters. In some implementations, for a 200 MeV to 230 MeV proton beam, the magnetic field at the magnetic dipoles in the beamline structures 16 is at least 3 T (i.e., 3 T or more). As stated, the systems described herein are not limited to these parameter values, and some implementations may have different dimensions, energies, and magnetic fields.

[0103] Chromatic aberration correction can occur in the beamline with dispersion generated by including dipole magnets and multiple correctors in the dispersion region. The standard definition of achromatic is a beam transport line with zero values ​​for spatial (R16) and angular (R26) dispersion. With reference to FIG. 18, the magnetic components in the compact gantry implementation can be configured to be achromatic (e.g., both R16 65 and R26 66 of the beam transfer matrix elements equal to zero at the isocenter, which is at or near 0 m along the beamline structure length 68 (X-axis)). Reducing or minimizing spatial and angular beam dispersion can be important for the pencil beam scanning techniques implemented by the particle therapy system described herein. In this regard, some pencil beam scanning techniques require the particle beam cross section to be substantially round at the isocenter. As such, the beam spot size in both the x and y (FIG. 18) planes should be close at the isocenter 67. During beam scanning, changes to the beam shape and beam diameter across the scan area should be reduced or minimized, otherwise different beam particles of different energies may land at different locations in the bending plane, which may cause the beam shape and beam size to be different in another plane.

[0104] FIG. 19 shows an example of beam scanning in the X dimension 69 and the Y dimension 70. Firing of the scanning magnet allows the beam particles to be deflected to an angle proportional to the field strength of the scanning magnet. In the example of FIG. 19, a beam scanning range that completely covers a beam field area of ​​20 cm×20 cm is shown with a beam deflection angle of approximately ±20 milliradians (mrad) and ±30 mrad from the scanning magnet. In this example, the source-to-isocenter distance (SAD) (i.e., the distance from the accelerator to the isocenter) is approximately 4 meters. In some implementations, from the scanning magnet to the exit of the output channel 17, the beam bending angle can be as large as 110° to 170°.

[0105] 1, in some implementations, an imaging system comprised of one or more imaging devices 99 can be mounted to the support structure 15. Imaging can be performed before and / or during a procedure to identify a target location in the patient and / or control the movement of the gantry and scanning to direct the particle beam to an irradiation target in the patient. The imaging system can include one or more of a computed tomography (CT) scanner, a two-dimensional (2D) X-ray device, a magnetic resonance imaging (MRI) device, a fan-beam CT scanner, a 2D camera, a three-dimensional (3D) camera, a surface imaging device, or a cone-beam CT scanner.

[0106] The imaging devices can be configured and controlled to rotate around or with the rotation of the gantry 14. In some implementations, one or more nozzles (with the beamlines aligned and connected to the one or more nozzles) are rotatable on a ring bearing that is positioned on the inner diameter of the support structure 15. Various two-dimensional (2D) and / or three-dimensional (3D) imaging devices can also be mounted on the ring bearings and rotatable therewith. In some implementations, the nozzles and imaging devices can be mounted on different internal circumferential tracks in the gantry.

[0107] For example, the nozzle may be rotatable about a circumferential track at a first radius of the support structure and the imaging device may be rotatable about a different circumferential track at a second radius of the support structure different from the first radius. In some implementations, the gantry may include different rotatable inner rings, one of which mounts the nozzle for rotation and one of which mounts the imaging device or system for rotation.

[0108] In some implementations, two 2D imaging devices are mounted on the support structure 15 in orthogonal planes to enable 2D image-guided radiation therapy (IGRT). IGRT involves the use of imaging during radiation treatment to improve the precision and accuracy of treatment delivery. IGRT can be used to treat tumors in moving body areas (such as the lungs). The 2D imaging devices can be rotated to enable cone-beam CT imaging (including simultaneously acquired dual-energy imaging). The imaging devices can also or alternatively include an X-ray source and image panel for cone-beam CT image acquisition, or a fan-beam diagnostic-quality CT imaging device. Alternatively, one plane can include a cone-beam CT imaging device and another plane can include a fan-beam diagnostic-quality CT imaging device.

[0109] As described herein, an exemplary proton therapy system scans a proton beam in three dimensions across an irradiation target to destroy malignant tissue. FIG. 20 shows a cross section of an exemplary superconducting synchrocyclotron component 75 that may be used to provide a particle (e.g., proton) beam in a proton therapy system. In this example, the component 75 includes a superconducting magnet 77. The superconducting magnet includes superconducting coils 78 and 79. The superconducting coils are formed of multiple integrated conductors, each of which includes superconducting strands (e.g., four strands or six strands) wound around a central strand, which itself can be superconducting or non-superconducting. Each of the superconducting coils 78, 79 is for conducting an electric current that generates a magnetic field (B). Magnetic yokes 80, 81 or smaller pole pieces shape that magnetic field in a cavity 84 in which the particles are accelerated. In one example, a cryostat (not shown) uses liquid helium (He) to conductively cool each coil to a low temperature superconducting temperature (eg, approximately 4 degrees Kelvin (K)).

[0110] In some implementations, the particle accelerator includes a particle source 85 (e.g., a Penning Ion Gauge (PIG) ​​source, etc.) to provide an ionized plasma column to the cavity 84. Hydrogen gas, or a combination of hydrogen gas and noble gases, is ionized to create the plasma column. A voltage source provides a varying radio frequency (RF) voltage to the cavity 84 to accelerate particles from the plasma column within the cavity. As mentioned, in one example, the particle accelerator is a synchrocyclotron. Thus, the RF voltage is swept across a range of frequencies to account for relativistic effects on the particles (e.g., increasing particle mass, etc.) as they accelerate within the acceleration cavity. The RF voltage drives a dee plate contained within the cavity and has a frequency that is swept downward during the acceleration cycle to account for the increase in relativistic mass of the protons and the decrease in the magnetic field. A dummy dee plate acts as a ground reference for the dee plate. The magnetic field created by running a current through the superconducting coils, together with the sweeping RF voltage, causes particles from the plasma column to accelerate orbitally within the cavity and to increase in energy as the number of turns increases. Particles in the outermost orbits are directed into an extraction channel (not shown) and output from the synchrocyclotron as a particle beam. In the synchrocyclotron, the particle beam is pulsed so that a bunch of particles are periodically output.

[0111] The magnetic field within the cavity is shaped to cause the particles to move orbitally within the cavity, as described above. An exemplary synchrocyclotron uses a magnetic field that is uniform with respect to the rotation angle and falls off in strength as the radius increases. In some implementations, the maximum magnetic field produced by the superconducting (main) coils can be in the range of 2.5T to 20T at the center of the cavity, which falls off as the radius increases. For example, superconducting coils can be used to generate magnetic fields at (or exceeding) one or more of the following magnitudes: 2.5T, 3.0T, 3.1T, 3.2T, 3.3T, 3.4T, 3.5T, 3.6T, 3.7T, 3.8T, 3.9T, 4.0T, 4.1T, 4.2T, 4.3T, 4.4T, 4.5T, 4.6T, 4.7T, 4.8T, 4.9T, 5.0T, 5.1T, 5.2T, 5.3T, 5.4T, 5.5T, 5.6T, 5.7T, 5.8T, 5.9T, 6.0T, 6.1T, 6.2T, 6.3T, 6.4T, 6.5T, 6.6T, 6.7T, 6.8T, 6.9T, 7.0T, 7.1T, 7.2T, 7.3T, 7.4T, 7.5T, 7.6T, 7.7T, 7.8T, 7.9T, 8.0T, 8.1T, 8.2T, 8.3T, 8.4T, 8.5T, 8.6T, 8.7T, 8.8T, 8.9T, 9.0T, 9.1T, 9.2T, 9.3T, 9.4T, 9.5T, 9.6T, 9.7T, 9.8T, 9.9T, 9.9T, 9.1T, 9.2T, 9.4T, 9.5T, 9.6T, 9.7T, 9.8T, 9.9T, 3T, 5.4T, 5.5T, 5.6T, 5.7T, 5.8T, 5.9T, 6.0T, 6.1T, 6.2T, 6.3T, 6.4T, 6.5T, 6.6T, 6.7T, 6.8T, 6.9T, 7.0T, 7.1T, 7.2 T, 7.3T, 7.4T, 7.5T, 7.6T, 7.7T, 7.8T, 7.9T, 8.0T, 8.1T, 8.2T, 8.3T, 8.4T, 8.5T, 8.6T, 8.7T, 8.8T, 8.9T, 9.0T, 9.1T, 9.2T, 9.3T, 9.4T, 9.5T, 9.6T, 9.7T, 9.8T, 9.9T, 10.0T, 10.1T, 10.2T, 10.3T, 10.4T, 10.5T, 10.6T, 10.7T, 10.8T, 10 .9T, 11.0T, 11.1T, 11.2T, 11.3T, 11.4T, 11.5T, 11.6T, 11.7T, 11.8T, 11.9T, 12.0T, 12.1T, 12.2T, 12.3T, 12.4T, 12 .5T, 12.6T, 12.7T, 12.8T, 12.9T, 13.0T, 13.1T, 13.2T, 13.3T, 13.4T, 13.5T, 13.6T, 13.7T, 13.8T, 13.9T, 14.0T, 14 .1T, 14.2T, 14.3T, 14.4T, 14.5T, 14.6T, 14.7T, 14.8T, 14.9T, 15.0T, 15.1T, 15.2T, 15.3T, 15.4T, 15.5T, 15.6T, 15.7T, 15.8T, 15.9T, 16.0T, 16.1T, 16.2T, 16.3T, 16.4T, 16.5T, 16.6T, 16.7T, 16.8T, 16.9T, 17.0T , 17.1T, 17.2T, 17.3T, 17.4T, 17.5T, 17.6T, 17.7T, 17.8T, 17.9T, 18.0T, 18.1T, 18.2T, 18.3T, 18 0.4T, 18.5T, 18.6T, 18.7T, 18.8T, 18.9T, 19.0T, 19.1T, 19.2T, 19.3T, 19.4T, 19.5T, 19.6T, 19.7T, 19.8T, 19.9T, 20.0T, 20.1T, 20.2T, 20.3T, 20.4T, 20.5T, 20.6T, 20.7T, 20.8T, 20.9T, or more. Additionally, superconducting coils can be used in generating magnetic fields outside the range of 2.5T to 20T, or magnetic fields within the range of 3T to 20T but not specifically enumerated herein.

[0112] By generating a high magnetic field having a magnitude such as that described above, the bending radius of particles traversing within the cavity 84 can be reduced. As a result of the reduced bending radius, a greater number of particle trajectories can be created within a cavity of a given size. Thus, the same number of trajectories can be fitted into a smaller cavity. Reducing the size of the cavity generally reduces the size of the particle accelerator, since a smaller cavity requires smaller magnetic yokes or pole pieces, among other components. In some implementations, the size or volume of the particle accelerator is less than 4 m. 3 Below, 3m 3 Less than or equal to 2m 3 It is possible that:

[0113] In some implementations (such as the implementation shown in FIG. 20), relatively large ferromagnetic magnetic yokes 80, 81 act as magnetic returns for the stray magnetic fields created by the superconducting coils. In some systems, a magnetic shield (not shown) surrounds the yokes. The return yokes and shields act together to reduce the stray magnetic fields, thereby reducing the likelihood that the stray magnetic fields will adversely affect the operation of the particle accelerator.

[0114] In some implementations, the return yoke and / or shields can be replaced or augmented by an active return system. An exemplary active return system includes one or more active return coils that conduct current in a direction opposite to that through the main superconducting coils. In some implementations, there is an active return coil for each superconducting main coil, e.g., there are two active return coils, one for each main superconducting coil. Also, each active return coil can be a superconducting coil that concentrically surrounds the outside of the corresponding main superconducting coil. In some implementations, the active return coils can be or include non-superconducting coils. By using an active return system, the relatively large ferromagnetic magnetic yokes 80, 81 can be replaced with smaller and lighter pole pieces. Thus, the size and weight of the synchrocyclotron can be further reduced without sacrificing performance. An example of an active return system that can be used is described in U.S. Patent No. 8,791,656 (Zwart), entitled "Active Return System." The contents of U.S. Pat. No. 8,791,656, particularly that relating to the return coil configuration (e.g., Figures 2, 4, and 5 of U.S. Pat. No. 8,791,656 and the accompanying description), are incorporated herein by reference.

[0115] Another example of a particle accelerator that may be used in the particle therapy systems herein is described in U.S. Patent No. 8,975,836 (Bromberg), entitled "Ultra-Light Magnetically Shielded High-Current, Compact Cyclotron." The contents of U.S. Patent No. 8,975,836, particularly the contents related to "Cyclotron 11" or "Iron-Free Cyclotron 11" in Figures 4, 17, and 18 of U.S. Patent No. 8,975,836 and the accompanying description, are incorporated herein by reference.

[0116] In some implementations, the synchrocyclotron used in the proton therapy system described herein can be a variable energy synchrocyclotron. In some implementations, the variable energy synchrocyclotron is configured to vary the energy of the output particle beam by varying the magnetic field in which the particle beam is accelerated. For example, the current can be set to any of a number of values ​​to create a corresponding magnetic field. For example, the current can be set to one of two values ​​to create the previously described dual energy particle accelerator. In one example and exemplary implementation, one or more sets of superconducting coils receive a variable current to create a variable magnetic field in the cavity. In some examples, one set of coils receives a fixed current while one or more other sets of coils receive a variable current such that the total current received by the coil sets varies. In some implementations, the coils in all sets are superconducting. In some implementations, some sets of coils (e.g., sets for fixed current, etc.) are superconducting, while other sets of coils (e.g., one or more sets for variable current, etc.) are non-superconducting (e.g., copper) coils.

[0117] Generally, in a variable energy synchrocyclotron, the magnitude of the magnetic field is scalable with the magnitude of the current. Adjusting the total current of the coils in a predetermined range can generate a magnetic field that changes in a corresponding predetermined range. In some examples, continuous adjustment of the current can lead to a continuous change in the magnetic field and a continuous change in the output beam energy. Alternatively, when the current applied to the coils is adjusted in a non-continuous, stepwise manner, the magnetic field and the output beam energy also change accordingly in a non-continuous (stepwise) manner. The stepwise adjustment can create the dual energy previously described. In some implementations, each step is between 10 MeV and 80 MeV in size. The scaling of the magnetic field to the current can allow the change in the beam energy to be performed relatively precisely, thus reducing the need for an energy degrader. An example of a variable energy synchrocyclotron that may be used in the particle therapy systems described herein is described in U.S. Patent No. 9,730,308, entitled "Particle Accelerator That Produces Charged Particles Having Variable Energies." The contents of U.S. Patent No. 9,730,308 are incorporated herein by reference, and in particular the contents that enable operation of the synchrocyclotron at variable energies (including those described in columns 5-7 and FIG. 13 of U.S. Patent No. 9,730,308 and the accompanying description) are incorporated herein by reference.

[0118] In an implementation of a particle therapy system using a variable energy synchrocyclotron, controlling the energy of the particle beam to treat a portion of an irradiation target can be performed according to a treatment plan by varying the energy of the particle beam output by the synchrocyclotron. In such an implementation, an energy degrader may or may not be used. For example, controlling the energy of the particle beam can include setting a current in the synchrocyclotron main coil to one of a plurality of values, each of the plurality of values ​​corresponding to a different energy at which the particle beam is output from the synchrocyclotron. An energy degrader can be used with a variable energy synchrocyclotron to provide additional changes in energy, for example, between the discrete energy levels provided by the synchrocyclotron.

[0119] The particle therapy system and its variants described herein can be used to apply very high dose rates of radiation (so-called "FLASH" dose rates of radiation) to irradiation targets in patients. In this regard, experimental results in radiation therapy have shown improved conditions for healthy tissues receiving radiation when treatment doses are delivered at very high (FLASH) dose rates. In one example, when a dose of radiation in the range of 10 Gray (Gy) to 20 Gray is delivered in pulses of less than 500 milliseconds (ms) to reach an effective dose rate of 20 Gray per second (Gy / S) to 100 Gray per second, healthy tissue experiences less damage than when irradiated with the same dose over a longer time scale, while tumors are treated with similar efficacy. A theory that can explain this "FLASH effect" is based on the fact that radiation damage to tissue is proportional to the oxygen supply in the tissue. In healthy tissue, the very high dose rate radicalizes oxygen only once, as opposed to a dose application that radicalizes oxygen multiple times over a longer time scale. This may lead to less damage in healthy tissue using ultra-high dose rates.

[0120] In some examples, as noted above, an ultra-high dose rate of radiation can include a dose of radiation greater than 1 Gray per second for a duration of less than 500 ms. In some examples, an ultra-high dose rate of radiation can include a dose of radiation greater than 1 Gray per second for a duration between 10 ms and 5 s. In some examples, an ultra-high dose rate of radiation can include a dose of radiation greater than 1 Gray per second for a duration of less than 5 s.

[0121] In some examples, an ultra-high dose rate of radiation includes a dose of radiation that exceeds one of the following doses over a duration of less than 500 ms: 2 Gy / sec, 3 Gy / sec, 4 Gy / sec, 5 Gy / sec, 6 Gy / sec, 7 Gy / sec, 8 Gy / sec, 9 Gy / sec, 10 Gy / sec, 11 Gy / sec, 12 Gy / sec, 13 Gy / sec, 14 Gy / sec, 15 Gy / sec, 16 Gy / sec, 17 Gy / sec, 18 Gy / sec, 19 Gy / sec, 20 Gy / sec, 30 Gy / sec, 40 Gy / sec, 50 Gy / sec, 60 Gy / sec, 70 Gy / sec, 80 Gy / sec, 90 Gy / sec, or 100 Gy / sec. In some examples, an ultra-high dose rate of radiation includes a dose of radiation that exceeds one of the following doses over a duration between 10 ms and 5 s: 2 Gy / s, 3 Gy / s, 4 Gy / s, 5 Gy / s, 6 Gy / s, 7 Gy / s, 8 Gy / s, 9 Gy / s, 10 Gy / s, 11 Gy / s, 12 Gy / s, 13 Gy / s, 14 Gy / s, 15 Gy / s, 16 Gy / s, 17 Gy / s, 18 Gy / s, 19 Gy / s, 20 Gy / s, 30 Gy / s, 40 Gy / s, 50 Gy / s, 60 Gy / s, 70 Gy / s, 80 Gy / s, 90 Gy / s, or 100 Gy / s. In some examples, an ultra-high dose rate of radiation includes a dose of radiation that exceeds one of the following doses over a duration of less than 5 s: 2 Gy / s, 3 Gy / s, 4 Gy / s, 5 Gy / s, 6 Gy / s, 7 Gy / s, 8 Gy / s, 9 Gy / s, 10 Gy / s, 11 Gy / s, 12 Gy / s, 13 Gy / s, 14 Gy / s, 15 Gy / s, 16 Gy / s, 17 Gy / s, 18 Gy / s, 19 Gy / s, 20 Gy / s, 30 Gy / s, 40 Gy / s, 50 Gy / s, 60 Gy / s, 70 Gy / s, 80 Gy / s, 90 Gy / s, or 100 Gy / s.

[0122] In some examples, an ultra-high dose rate of radiation includes a dose of radiation that exceeds one or more of the following doses for a duration of less than 500 ms, for a duration between 10 ms and 5 s, or for a duration of less than 5 s: 100 Gray per second, 200 Gray per second, 300 Gray per second, 400 Gray per second, or 500 Gray per second.

[0123] In some examples, the ultra-high dose rate of radiation includes a dose of radiation between 20 Gy / sec and 100 Gy / sec for a duration of less than 500 ms. In some examples, the ultra-high dose rate of radiation includes a dose of radiation between 20 Gy / sec and 100 Gy / sec for a duration of between 10 ms and 5 s. In some examples, the ultra-high dose rate of radiation includes a dose of radiation between 20 Gy / sec and 100 Gy / sec for a duration of less than 5 s. In some examples, the ultra-high dose rate of radiation includes a dose of radiation between 40 Gy / sec and 120 Gy / sec for a predetermined period of time (e.g., less than 5 s). Other examples of time periods are provided above.

[0124] In some implementations, particle therapy systems can treat three-dimensional columns of a target using ultra-high dose rate radiation (FLASH doses of radiation). These systems use pencil beam scanning to scale ultra-high dose rate delivery to the target. In some examples, pencil beam scanning involves delivering a series of small beams of particle radiation, each of which can have a unique direction, energy, and charge. By combining the doses from these individual beams, a three-dimensional target treatment volume can be treated with radiation. Moreover, instead of organizing the treatment into layers at a constant energy, the system organizes the treatment into columns defined by the direction of the stationary beam. The beam direction can be toward the surface of the target.

[0125] In some implementations, all or a portion of the column is treated before the particle beam is directed along another path through the irradiation target. In some implementations, the path through the target is full or partial direction through the target. In one example, the particle beam can be directed along a path through the target and not deviate from that path. While being directed along that path, the energy of the particle beam is changed. The particle beam does not move as its energy is changed, and as a result, the particle beam treats all or a portion of the interior portion of the target that extends along the length of the particle beam and along the width of the beam spot. Thus, the treatment is performed in depth along the length of the beam. For example, the portion of the target that is treated can extend from the spot of the beam at the surface of the target down through all or a portion of the interior of the target. The result is that the particle beam treats a three-dimensional columnar portion of the target using an ultra-high dose rate of radiation. In some examples, the particle beam may never be re-directed along the same three-dimensional columnar portion more than once.

[0126] In some implementations, the irradiation target can be divided into micro-volumes. Although a cubic micro-volume can be used, the micro-volumes can have any suitable shape, such as a three-dimensional rectangular parallelepiped, a regular curved shape, or an irregular or amorphous shape. In this example, each micro-volume is treated through the delivery of FLASH radiation column by column in the manner described herein. For example, a column depth of the micro-volume can be treated with radiation by using an energy degrader plate to vary the beam energy or by controlling a variable energy synchrocyclotron to vary the beam energy. After each micro-volume is treated, the next micro-volume is treated and repeated until the entire irradiation target is treated. The treatment of the micro-volumes can be performed in any suitable order or sequence.

[0127] The particle therapy system described herein is capable of delivering FLASH radiation by column in the manner described in U.S. Patent Application Publication No. 2020 / 0298025, entitled "Delivery Of Radiation By Column And Generating A Treatment Plan Therefor," the contents of which are incorporated by reference herein, in particular the contents of Figures 2, 11, 12-19, 33-43B, and the accompanying description thereof, which are incorporated by reference herein.

[0128] In some implementations, particle accelerators other than synchrocyclotrons can be used in the particle therapy systems described herein. For example, a cyclotron, synchrotron, or linear accelerator, etc., can be substituted for a synchrocyclotron in the particle therapy systems described herein.

[0129] FIG. 21 shows an alternative implementation of a fixed beam particle therapy system including a particle accelerator 141 and a treatment couch 142, which may be of the type described herein. In this implementation, the particle accelerator 141 is stationary and configured and controllable to output particles as a particle beam 143 in a single fixed direction (e.g., along a single direction vector). The particle beam is monoenergetic and may be pulsed, in the case of a synchrocyclotron. When other types of particle accelerators are used, the particle beam may be continuous rather than pulsed. As shown in the figure, in this example, there are no beamline structures between the particle accelerator and the patient.

[0130] One or more scanning magnets (not shown) can be positioned in the particle beam path between the particle accelerator and the treatment couch. The scanning magnets can be superconducting, non-superconducting, or a combination of superconducting and non-superconducting. The scanning magnets can be of the type shown in FIG. 5, the type shown in FIG. 7 and FIG. 8, or a combination thereof. In some implementations, control over scanning is achieved by varying the current through one or both sets of coils, thereby varying the magnetic field produced thereby. By appropriately varying the magnetic field, the particle beam can be moved in the X and / or Y dimensions across the irradiation target.

[0131] An energy degrader 144 having a configuration as described herein is in the particle beam path between the treatment couch and the scanning magnet. In this example, the particle beam is monoenergetic. As a result, a change in the energy of the particle beam can be implemented only by changing the configuration of the energy degrader. For example, a plate or wedge can be moved to change the thickness of the material in the path of the particle beam and thus change the energy of the particle beam.

[0132] A collimator (not shown) having a configuration as described herein can be in the particle beam path between the energy degrader and the treatment couch. The collimator is used as described herein to block a portion of the particle beam from reaching non-target tissue in the patient. The collimator, the energy degrader, and the scanning magnet can be included in an output device (e.g., a nozzle, etc.) that can be mounted to a structure (e.g., a wall, etc.). The nozzle does not move relative to the particle accelerator or the treatment couch (although the components of the device contained therein can move during operation).

[0133] During operation, the treatment couch moves (e.g., with three, four, five, or six degrees of freedom) relative to the particle beam or relative to the nozzle. The treatment couch can also be reclined to reach the treatment position. Thus, instead of moving the particle accelerator or the nozzle, the treatment couch moves the patient to a position for treatment, and then the devices in the particle accelerator and the nozzle are controlled to treat the patient at that position. The patient can then be repositioned to treat the irradiation target from a different position or angle.

[0134] In some implementations, the scanning magnet can be replaced with a scattering foil and the energy degrader can be a range modulator. In such implementations, the scattering foil scatters the particle beam across the treatment area and the depth to which the scattered beam is applied is controlled by the range modulator. The configurable collimator can remain in place to trim the edges of the scattered beam.

[0135] The operation of the example proton therapy systems described herein, and all or some of their components, may be controlled (at least in part) using a control system 192 ( FIG. 1 ) configured to execute one or more computer program products (e.g., one or more computer programs tangibly embodied in one or more non-transitory machine-readable media) for execution by or to control the operation of one or more data processing devices (e.g., a programmable processor, a computer, multiple computers, and / or programmable logic components).

[0136] All or part of the systems described herein, and various modifications thereof, may be configured or controlled, at least in part, by one or more computers, such as a control system, using one or more computer programs tangibly embodied in one or more information carriers (e.g., in one or more non-transitory machine-readable storage media, etc.). The computer programs may be written in any form of programming language (including compiled or interpreted languages) and may be deployed in any form (including as a stand-alone program or as a module, part, subroutine, or other unit suitable for use in a computing environment). The computer programs may be deployed to run on one computer, or on multiple computers at one site, or distributed across multiple sites and interconnected by a network.

[0137] Actions associated with configuring or controlling the systems described herein may be performed by one or more programmable processors executing one or more computer programs to control or perform all or some of the operations described herein. All or a portion of the systems and processes may be configured or controlled by special purpose logic circuitry (e.g., FPGAs (field programmable gate arrays) and / or ASICs (application specific integrated circuits), or embedded microprocessors localized in instrument hardware, etc.).

[0138] Processors suitable for executing a computer program include, by way of example, both general purpose and special purpose microprocessors, as well as any one or more processors of any kind of digital computer. Typically, a processor will receive instructions and data from a read-only or random access storage area, or both. Elements of a computer include one or more processors for executing instructions and one or more storage area devices for storing instructions and data. Typically, a computer will also include one or more machine-readable storage media, such as a mass storage device (e.g., magnetic disk, magneto-optical disk, or optical disk, etc.) for storing data, or be operatively coupled to receive data from or transfer data to it, or both. Non-transitory machine-readable storage media suitable for embodying computer program instructions and data include all non-volatile storage areas, including, by way of example, semiconductor storage area devices, such as EPROM (erasable programmable read-only memory), EEPROM (electrically erasable programmable read-only memory), and flash storage area devices; magnetic disks, such as internal hard disks or removable disks; magneto-optical disks; and CD-ROMs (compact discs, read-only memory) and DVD-ROMs (digital versatile discs, read-only memory).

[0139] Elements of different implementations described can be combined to form other implementations not specifically described previously. Elements can be removed from previously described systems without adversely affecting their operation or the operation of the system as a whole. Moreover, various separate elements can be incorporated into one or more individual elements to perform the functions described herein.

[0140] Other implementations not specifically described herein are within the scope of the following claims. [Explanation of symbols]

[0141] 10 Particle Therapy System 12 Particle accelerator 14 Gantry 15 Support structure 16 Beamline Structure 16a Beamline structure 16b Beamline structure 16c Beamline structure 16d Beamline structure 17 Output Channels 17a Output Channel 17b Output Channel 17c Output Channel 17d Output Channel 18 Conduit 19 Treatment location 21 Non-superconducting magnetic quadrupoles 22 Non-superconducting magnetic quadrupoles 23 Superconducting magnetic dipole 24 Horizontal 26 Non-superconducting magnetic quadrupoles 27 Non-superconducting magnetic quadrupoles 29 Compact Gantry Length 30 Scanning Magnet 30a Scanning magnet 30b1 first scanning magnet 30b2 second scanning magnet 30c1 first scanning magnet 30c2 Second scanning magnet 30d1 first scanning magnet 30d2 Second scanning magnet 31 Magnetic dipole (large diameter superconducting magnetic dipole) 31d Magnetic dipole 32 Magnetic dipole (large diameter superconducting magnetic dipole) 32c Magnetic Dipole 32d Magnetic Dipole 33 Magnetic quadrupole (large diameter superconducting magnetic quadrupole) 34 Magnetic quadrupole (large diameter superconducting magnetic quadrupole) 35 Magnetic quadrupole (large diameter superconducting magnetic quadrupole) 38 horizontal line 40 Nozzles 40a nozzle 40b Nozzle 40c nozzle 40d nozzle 41 Energy Degrader 41a Energy Degrader 41b Energy Degrader 41c Energy Degrader 41d Energy Degrader 42 Structure 43 Scanning Magnet 44 Collimator 44a Collimator 44b Collimator 44c Collimator 44d Collimator 45 First set (of two coils) 46 Second set (of two coils) 48 Energy Degrader 49 Treatment Space 50 Treatment Space 51 Treatment Couch 53 Hole 54 Robot Arm 55 First Segment 56 Second Segment 57 Third Segment 60 Compartments 61 Compartments 63 Horizontal (x) beam envelope 64 Beam envelope in the vertical direction (y) 65 Beam Transfer Matrix Element R16 66 Beam Transfer Matrix Element R26 67 Isocenter 68 Beamline structure length 69 X dimension 70 Y dimension 75 Components 77 Superconducting Magnet 78 Superconducting Coil 79 Superconducting Coil 80 Magnetic Yoke 81 Magnetic Yoke 84 Cavity 85 Particle Source 90 Pit, magnet 90a 1st set coil 91 Magnet 91a Second set of coils 92 Superconducting Scanning Magnet 92a High temperature superconducting coil 92b High-temperature superconducting coil 94 Cryostat 95 Superconducting Magnets 95a High Temperature Superconducting Coil Set 96 Cryostat 97 Superconducting Scanning Magnet 97a High Temperature Superconducting Coil Set 98 Cryostat 113 Primary Carriage 114 Secondary Carriage 115 Secondary Carriage 117 Cartesian Z dimension 118 Cartesian X-dimension, corresponding member 119 Corresponding parts 120 Tracks 122 Rods, rails 123 Rods, rails 125 Motor 126 Motor 130 Motor 131 Lead screw 132 Housing 135 Leaf 136 Leaf 141 Particle Accelerator 142 Treatment Couch 143 Particle Beam 144 Energy Degrader 150 Superconducting Scanning Magnet 151 Grid 153 Cartesian X-dimension 154 Cartesian Y dimension 155 Baseline 0,0 points 158 Inner superconducting coil 159 Outer superconducting coil 160 Electrically non-conductive materials, electrically non-superconductive materials 161 Aperture 162 Electrically non-conductive materials, electrically non-superconductive materials 165 Superconducting Coil 166 Copper (Cu) stabilization layer 167 Silver (Ag) cap layer 168 Rare earth barium copper oxide (ReBCO) superconducting layer 169 Buffer Layer Stack 170 Substrate Layer 192 Control Systems

Claims

1. 1. A particle beam therapy system, comprising: a gantry including a beamline structure configured to direct a monoenergetic particle beam from an output of a particle accelerator toward an irradiation target, the beamline structure including a magnetic bending element for bending the particle beam along a length of the beamline structure; an energy degrader downstream of the beamline structure with respect to the particle accelerator, the energy degrader configured and controllable to change the energy of at least a portion of the particle beam before the particle beam reaches the irradiation target; Including, the beamline structure has an efficiency of 10% or greater; The efficiency of 10% or more includes that 10% or more of the particles output from the particle accelerator are output from the beamline structure. Particle beam therapy system.

2. 2. The particle therapy system of claim 1, wherein the energy degrader is the only mechanism for actively controlling a change in energy of the particle beam after the particle beam is output by the particle accelerator and before the particle beam reaches the irradiation target.

3. 2. The particle therapy system of claim 1, wherein the beamline structure is configured not to actively control the energy of the particle beam after the particle beam is output by the particle accelerator and before the particle beam reaches the energy degrader.

4. the magnetic bending element includes a magnet having a magnetic field of 2.5 Tesla (T) or greater; or The magnetic bending element includes a magnet having a magnetic field of 3 Tesla (T) or more.

2. A particle beam therapy system according to claim 1.

5. the gantry includes a support structure configured to move a portion of the beamline structure in a circular path around the irradiation target; The support structure has a dimension of 6 meters or less.

2. A particle beam therapy system according to claim 1.

6. The particle therapy system of claim 5 , wherein the dimension is a diameter of the support structure.

7. the length of the beamline structure is 6 meters (m) or less; or The length of the beamline structure is 5 meters (m) or less.

2. A particle beam therapy system according to claim 1.

8. 10. The particle therapy system of claim 1, wherein the energy of the particle beam does not vary by more than 1% within the beamline structure.

9. 10. The particle therapy system of claim 1, wherein a distance between an output of the beamline structure and an isocenter containing the irradiation target is 1.5 meters (m) or less.

10. 10. The particle therapy system of claim 1, further comprising a scanning system including one or more scanning magnets for moving the particle beam in at least two dimensions across at least a portion of a beam field covering at least a portion of the irradiation target.

11. The particle therapy system of claim 10 , wherein at least one of the scanning magnets includes a superconducting magnet.

12. 11. The particle therapy system of claim 10, wherein at least one of the scanning magnets is positioned downstream of the beamline structure with respect to the particle accelerator.

13. 11. The particle therapy system of claim 10, wherein at least one of the scanning magnets is positioned within the beamline structure.

14. 11. The particle beam therapy system of claim 10, wherein the one or more scanning magnets include a first scanning magnet positioned in the beamline structure and a second scanning magnet positioned downstream of the first scanning magnet with respect to the particle accelerator, the first scanning magnet being separate from the second scanning magnet.

15. the first scanning magnet is configured to move the particle beam across the at least a portion of the beam field in the at least two dimensions; 15. The particle therapy system of claim 14, wherein the second scanning magnet is configured to move the particle beam across the at least a portion of the beam field in the at least two dimensions.

16. the first scanner magnet is configured to move the particle beam across the at least a portion of the beam field in only a first dimension; the second scanning magnet is configured to move the particle beam across the at least a portion of the beam field in only a second dimension, the second dimension being different from the first dimension. The particle beam therapy system according to claim 14.

17. 15. The particle beam therapy system of claim 14, wherein the first scanning magnet is positioned between magnetic components included in the beamline structure, and the second scanning magnet is positioned downstream of the beamline structure in a particle beam output device, the particle beam output device including a nozzle.

18. 15. The particle therapy system of claim 14, wherein the first scanning magnet is positioned within the beamline structure and the second scanning magnet is positioned within the beamline structure.

19. 10. The particle therapy system of claim 1, wherein the beamline structure includes an output channel including magnetic dipoles arranged in series to bend the particle beam by at least 90 degrees.

20. 20. The particle therapy system of claim 19, wherein the magnetic dipoles include at least a first magnetic dipole and a second magnetic dipole, the first magnetic dipole and the second magnetic dipole being between a first scanning magnet and a second scanning magnet.

21. the beamline structure includes an output channel configured to bend the particle beam by at least 90° toward an irradiation target in the presence of a magnetic field of at least 3 Tesla (T); the beamline structure includes a magnetic dipole for bending the particle beam, and two or more magnetic quadrupoles or magnetic sextupoles disposed between the magnetic dipoles along a length of the beamline structure for focusing the particle beam; 2. A particle beam therapy system according to claim 1.

22. 10. The particle therapy system of claim 1, further comprising the particle accelerator, the particle accelerator being a compact particle accelerator.

23. 10. The particle therapy system of claim 1, wherein the beamline structure has a length of 6 meters or less.

24. 1. A particle beam therapy system, comprising: a particle accelerator configured to output particles as a monoenergetic particle beam; an output channel including a magnetic component configured to bend the particle beam by at least 90° in the presence of a magnetic field of at least 2.5 Tesla (T); a support structure on which the output channel is mounted for movement at least in part about an irradiation target; and a conduit for directing the particle beam to the output channel. Including, the gantry; an energy degrader downstream of the output channel relative to the particle accelerator; a scanning system including two or more scanning magnets, at least one of which is upstream of at least some of the magnetic components in the output channel relative to the particle accelerator, the two or more scanning magnets being separated by air or other magnetic components and configured to move the particle beam across at least a portion of a beam field in at least two dimensions; Including, Particle beam therapy system.

25. 25. The particle therapy system of claim 24, wherein the energy degrader is the only mechanism for actively controlling the change in energy of the particle beam after the particle beam is output by the particle accelerator and before the particle beam reaches the irradiation target.

26. 25. The particle therapy system of claim 24, wherein the gantry is configured not to actively control the energy of the particle beam after the particle beam is output by the particle accelerator and before the particle beam reaches the energy degrader.

27. the conduit and the output channel together have an efficiency of 10% or greater; The efficiency of 10% or more includes 10% or more of the particles output from the particle accelerator being output from the output channel.

25. The particle beam therapy system of claim 24.

28. 25. The particle therapy system of claim 24, wherein the gantry is less than or equal to 5 meters in length as measured from the particle accelerator.

29. The magnetic component includes: a first magnetic dipole for bending the particle beam towards an irradiation target; a second magnetic dipole upstream of the first magnetic dipole with respect to the particle accelerator, the second magnetic dipole configured to bend the particle beam toward the first magnetic dipole; Including, 25. The particle beam therapy system of claim 24.

30. 25. The particle therapy system of claim 24, wherein the gantry and the particle accelerator are in the same space and are not separated by shielding external to the particle accelerator that exceeds 30 centimeters in thickness.

31. 25. The particle therapy system of claim 24, wherein the particle therapy system has a footprint of 75 square meters or less.

32. the particle accelerator includes a synchrocyclotron; The particle therapy system is sized to fit into a vault designed for a linear accelerator (LINAC).

25. The particle beam therapy system of claim 24.

Images