Bent Magnet

Abstract

An exemplary magnet includes an assembly including (i) a set of coils for conducting current to create a magnetic field, and (ii) a support structure on which the set of coils are asymmetrically disposed, with a ferromagnetic yoke surrounding a portion of the assembly, and the ferromagnetic yoke and assembly being curved.

Description

[Technical Field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This application claims priority to U.S. Provisional Application No. 63 / 394,461, filed August 2, 2022. The contents of U.S. Provisional Application No. 63 / 394,461 are incorporated herein by reference.

[0002] This specification describes an example of a bending magnet, such as a cosine-theta magnet, for use in a gantry in a particle therapy system. [Background technology]

[0003] Particle therapy systems use particle accelerators to generate particle beams for treating afflictions (e.g., 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 device during treatment. The gantry includes magnetic components to direct the particle beam to its destination. [Prior art documents] [Patent documents]

[0004] [Patent Document 1] U.S. Patent No. 10,675,487 [Patent Document 2] U.S. 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 Summary of the Invention [Means for solving the problem]

[0005] An exemplary magnet includes an assembly including (i) a set of coils for conducting current to create a magnetic field, and (ii) a support structure on which the set of coils are asymmetrically disposed, with a ferromagnetic yoke surrounding a portion of the assembly. The ferromagnetic yoke and assembly are curved. The magnet can be curved as a result. An exemplary magnet can include one or more of the following features, either alone or in combination:

[0006] The set of coils may include a first coil and a second coil. The first coil and the second coil may be for conducting current to create a magnetic field. The first coil and the second coil may be asymmetrically disposed on the support structure in a first hemisphere of the magnet, such that a first spacing between the first coil and the second coil in the first quadrant of the magnet is different from a second spacing between the first coil and the second coil in the second quadrant of the magnet. The first quadrant and the second quadrant may be within the first hemisphere. The set of coils may include a third coil and a fourth coil. The third coil and the fourth coil can be for conducting current to create a magnetic field. The third coil and the fourth coil can be asymmetrically disposed on the support structure in the second hemisphere of the magnet, with a third spacing between the third coil and the fourth coil in the third quadrant of the magnet being different from a fourth spacing between the third coil and the fourth coil in the fourth quadrant of the magnet. The third and fourth quadrants can be in the second hemisphere. The asymmetry of the first and second coils in the first and second quadrants can mirror the asymmetry of the third and fourth coils in the third and fourth quadrants, respectively.

[0007] The first and third spacings may be equal. The second and fourth spacings may be equal. The first and third spacings may be smaller than the second and fourth spacings. The first and third spacings may be at an inner bend radius of the assembly, and the second and fourth spacings may be at an outer bend radius of the assembly. In the case of an implementation including first through fourth coils, the ferromagnetic yoke may include a notch adjacent to the assembly. The notch may be asymmetric in the first and second quadrants. The asymmetry of the notch may relate to at least one of the size, shape, or placement of the notch. The asymmetry of the notch in the third and fourth quadrants may reflect the asymmetry of the notch in the first and second quadrants, respectively.

[0008] The set of coils may include a fifth coil and a sixth coil. The fifth coil and the sixth coil may be for conducting current to create a magnetic field. The fifth coil may be disposed on a support structure in a first hemisphere. The sixth coil may be disposed on a support structure in a second hemisphere. A fifth spacing between the fifth coil and an adjacent one of the first or second coils in the first quadrant may be different from a sixth spacing between the fifth coil and an adjacent one of the first or second coils in the second quadrant. A seventh spacing between the sixth coil and an adjacent one of the third or fourth coils in the third quadrant may be different from an eighth spacing between the sixth coil and an adjacent one of the third or fourth coils in the fourth quadrant. The asymmetry of the first, second, and fifth coils in the first and second quadrants, respectively, can reflect the asymmetry of the third, fourth, and sixth coils in the third and fourth quadrants, respectively.

[0009] The fifth and seventh intervals may be equal. The sixth and eighth intervals may be equal. The fifth and seventh intervals may be smaller than the sixth and eighth intervals. The fifth and seventh intervals may be at the inner bend radius of the assembly. The sixth and eighth intervals may be at the outer bend radius of the assembly. In the case of an implementation including first through sixth coils, the ferromagnetic yoke may include a notch adjacent to the assembly. The notch may be asymmetric in the first and second quadrants. The asymmetry of the notch may relate to at least one of the size, shape, or placement of the notch. The asymmetry of the notch in the third and fourth quadrants may reflect the asymmetry of the notch in the first and second quadrants, respectively.

[0010] The set of coils may include a seventh coil and an eighth coil. The seventh coil and the eighth coil may be for conducting current to create a magnetic field. The seventh coil may be disposed on a support structure in the first hemisphere. The eighth coil may be disposed on a support structure in the second hemisphere. A ninth spacing between the seventh coil and an adjacent one of the first, second, or fifth coils in the first quadrant may be different from a tenth spacing between the seventh coil and an adjacent one of the first, second, or fifth coils in the second quadrant. An eleventh spacing between the eighth coil and an adjacent one of the third, fourth, or sixth coils in the third quadrant may be different from a twelfth spacing between the eighth coil and an adjacent one of the third, fourth, or sixth coils in the fourth quadrant. The asymmetry of the first, second, fifth, and seventh coils in the first and second quadrants, respectively, can reflect the asymmetry of the third, fourth, sixth, and eighth coils in the third and fourth quadrants, respectively.

[0011] The ninth and eleventh intervals may be equal. The tenth and twelfth intervals may be equal. The ninth and eleventh intervals may be smaller than the tenth and twelfth intervals. The ninth and eleventh intervals may be at an inner bend radius of the assembly. The tenth and twelfth intervals may be at an outer bend radius of the assembly. In the case of an implementation including the first through eighth coils, the ferromagnetic yoke may include a notch adjacent to the assembly. The notch may be asymmetric in the first and second quadrants. The asymmetry of the notch may relate to at least one of the size, shape, or placement of the notch. The asymmetry of the notch in the third and fourth quadrants may reflect the asymmetry of the notch in the first and second quadrants, respectively.

[0012] The ferromagnetic yoke can be or include iron. The support structure can be non-ferromagnetic. The magnet can be bent by 60° or more relative to a line passing through the center of the unbent part of the magnet. The magnet can be bent by 70° or more relative to a line passing through the center of the unbent part of the magnet. The magnet can be bent by 80° or more relative to a line passing through the center of the unbent part of the magnet. The magnet can be bent by 90° or more relative to a line passing through the center of the unbent part of the magnet.

[0013] The magnet can be or include a cosine theta magnet, where the current through the set of coils has a greater concentration near the 0° or 180° location of the magnet than near the 90° or -90 / 270° location of the magnet. The set of coils can be configured for dipole functionality. The set of coils can be configured for quadrupole functionality. The set of coils can be configured for hexapole functionality. The set of coils can include superconducting material to create a superconducting magnet. The magnet can include one or more magnetic shims movable relative to the ferromagnetic yoke to change the magnetic field produced by the magnet.

[0014] An exemplary system may include 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 may include bending magnets for bending the particle beam along the length of the beamline structure. At least one of the bending magnets may be or include a magnet of the type described above, i.e., a magnet including: (i) a set of coils for conducting current to create a magnetic field; and (ii) an assembly including a support structure on which the set of coils is asymmetrically disposed; and a ferromagnetic yoke surrounding a portion of the assembly, wherein the ferromagnetic yoke and the assembly are curved, and the magnet includes one or more of the foregoing features described above.

[0015] The system can include an energy degrader that is 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 beamline structure can 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.

[0016] The at least one bending magnet may include a magnet having a magnetic field of 2.5 Tesla (T) or greater. The at least one bending magnet may include a magnet having a magnetic field of 3 Tesla (T) or greater. The system may include a collimator downstream of the gantry 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 irradiation target. The gantry may include a support structure configured to move a portion of the beamline structure in a circular path around the irradiation target. 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% within 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.

[0017] The beamline structure can include an output channel having at least some of the bending magnets. At least some of the bending magnets can include magnetic dipoles arranged in series to bend the particle beam by at least 90°. The magnetic dipoles can include at least one bending magnet. The at least one bending magnet can precede the output channel in the direction of travel of the particle beam.

[0018] The gantry can be achromatic from the entry point of the particle beam into the gantry to the isocenter of the system where the patient is treated.

[0019] Any two or more of the features described in this specification (including this Summary section) may be combined to form implementations not specifically described in this specification.

[0020] 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., a microprocessor, an application-specific integrated circuit, or programmed logic such as a field-programmable gate array, etc.). The systems, or portions thereof, described herein may be implemented as apparatuses, methods, or medical systems that may include one or more processing devices and computer memory for storing executable instructions for implementing control of the described functions. The devices (e.g., magnets), systems, and / or components described herein may be configured through, for example, design, construction, composition, arrangement, installation, programming, operation, activation, deactivation, input, and / or control.

[0021] 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 explanation of the drawings]

[0022] [Figure 1]1 is a diagram showing a partially transparent perspective view of an exemplary particle therapy system having an exemplary gantry of the type described herein. [Figure 2] FIG. 2 is a cutaway side view of components of the particle therapy system shown in FIG. 1, including an exemplary gantry. [Figure 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 exemplary 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 exemplary beamline structure. [Figure 5] FIG. 1 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 exemplary 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 exemplary 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 exemplary 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. 1 is a front view of an exemplary superconducting scanning magnet configured to scan a particle beam in a single dimension. [Figure 12b] 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. FIG. [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 is a diagram showing a 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 exemplary treatment space configured to accommodate all or a portion of the particle therapy system of FIG. 1. [Figure 17] 1 is a graph illustrating exemplary horizontal (x) and vertical (y) particle beam envelopes produced in an exemplary gantry described herein. [Figure 18] 1 is a graph illustrating an exemplary achromatic grating design for a beamline of an exemplary gantry described herein. [Figure 19] 1 is a graph illustrating the results produced by scanning a particle beam in the horizontal (x) and vertical (y) planes using an exemplary gantry described herein. [Figure 20] FIG. 1 is a cutaway side view of components within an exemplary particle accelerator that may be used with the particle therapy systems described herein. [Figure 21] FIG. 1 is a perspective view of an exemplary particle therapy system. [Figure 22] FIG. 1 is a perspective view of an exemplary energy degrader. [Figure 23] 1 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. FIG. [Figure 24] FIG. 2 is a cutaway view of an exemplary superconducting coil that may be used in any of the superconducting magnets described herein. [Figure 25]FIG. 1 is a perspective view of the coils of an exemplary bending magnet. [Figure 26] FIG. 1 is a cross-sectional view of an exemplary bending magnet with three coils in each hemisphere. [Figure 27] FIG. 10 is a perspective cross-sectional view of an exemplary bending magnet showing magnetic field strength as shading, with darker shading indicating greater magnetic field strength. [Figure 28] FIG. 1 is a perspective view of components of an exemplary bending magnet, portions of which are shown as transparent. [Figure 29] FIG. 1 is a perspective view of components of an exemplary bending magnet. [Figure 30] FIG. 1 is a cross-sectional view of a portion of an exemplary bending magnet with two coils in each hemisphere. [Figure 31] FIG. 1 is a cross-sectional view of a portion of an exemplary bending magnet having four coils in each hemisphere. DETAILED DESCRIPTION OF THE INVENTION

[0023] Like reference numbers in different figures indicate like elements.

[0024] 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 may be, but is not limited to, a synchrocyclotron, which has low radiation leakage and is small enough to fit within 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) output from the accelerator to treat tumors or other conditions in the 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) for directing the particle beam toward the treatment location. To enable delivery of the particle beam within the same space used for the procedure, particularly within 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 angles approaching or exceeding a right angle. In one example, the magnetic components are configured and arranged to bend the particle beam by 70° or more (including angles of 90° and obtuse angles greater than 90°).

[0025] The magnetic components in the gantry can include one or more magnets, such as a cosine theta bending magnet having the following characteristics that allow the particle beam to be magnetically bent: current-conducting coils (or simply "coils") including at least first and second coils (where the first and second coils are for conducting current to create a magnetic field); and a non-ferromagnetic support structure ("support") (where a set of coils is asymmetrically positioned on the non-ferromagnetic support structure). The assembly consisting of the coils and supports surrounds (at least partially) an air core through which the particle beam passes. A ferromagnetic brick or yoke surrounds (at least partially) the assembly and air core but is otherwise solid, except for cutouts in the yoke, possibly adjacent to outer parts of the assembly. The cutouts define channels through the yoke, which can be filled with air or vacuum. The cutouts affect the amount of ferromagnetic material adjacent to the coils, which in turn affects the shape of the magnetic field created by the coils. Cosine theta bending magnets are curved or bent. The configuration of the magnets (e.g., asymmetric coil windings and asymmetric cutouts in the yoke), particularly at the bends in the gantry, allows the magnets to reduce particle beam distortion during the particle beam's travel through the gantry. For example, the cross-sectional (e.g., circular) shape of the particle beam can be maintained substantially or perfectly circular during its travel through the gantry.

[0026] FIG. 25 shows a perspective view of an exemplary coil 180 for an exemplary cosine theta bending magnet ("magnet"), which can be used as a bending magnet in an exemplary particle therapy system gantry, such as that described herein. However, the magnet is not limited to use in this context. FIG. 26 shows a cross-sectional front view of an exemplary magnet 200 including coil 180, a non-ferromagnetic support 205, an air core 213, and a yoke 181 constructed from iron or other ferromagnetic material.

[0027] Magnet 200 can be a dipole magnet, quadrupole magnet, or hexapole magnet. A dipole magnet has two poles, one north and one south. Its magnetic field lines form a closed loop, emerging from the north pole, re-entering the south pole, and then passing through the body of the magnet. A quadrupole magnet includes groups of four magnetic poles, laid out so that in a planar multipole expansion of the magnetic field, the dipole terms cancel and the quadrupole is the lowest significant term in the magnetic field equation. A hexapole magnet includes six magnetic poles set in an alternating north and south pole configuration around an axis. The coils described herein can be layered to generate higher-order field harmonics.

[0028] 25 and 26, magnet 200 is an electromagnet including multiple current-conducting coils 180 in each of its hemispheres 200a and 200b, where 0° to 180° corresponds to upper hemisphere 200a and 180° to 360° (again 0°) corresponds to lower hemisphere 200a. In the example of FIGS. 25 and 26, magnet 200's coils 180 include three coils 201a, 201b, and 201c in upper hemisphere 200a and three coils 201d, 201e, and 201f in lower hemisphere 200b (coil 201f is not visible in FIG. 25). In some implementations, magnet 200 can include fewer or more than three coils in each of its hemispheres. For example, a bending magnet of the type described herein can include two coils in each hemisphere (e.g., FIG. 30), three coils in each hemisphere (e.g., FIGS. 25-29), four coils in each hemisphere (e.g., FIG. 31), five coils in each hemisphere, six coils in each hemisphere, etc. Any suitable number of coils can be included in each hemisphere of the bending magnet.

[0029] 26, in magnet 200, lower hemisphere 200b is a mirror image of upper hemisphere 200a, meaning that the configuration and relative spacing of the coils in each hemisphere are the same. For example, coils 201a and 201d have the same configuration and spacing relative to coils 201b and 201e, respectively. Coils 201b and 201e have the same configuration and spacing relative to coils 201c and 201f, respectively. Coils 201c and 201f have the same configuration and spacing relative to the 90° and -90° (270°) locations above the magnet, respectively. Therefore, any description of the upper hemisphere herein also applies to the lower hemisphere, and vice versa.

[0030] Bending magnet coils, such as coil 180, may be superconducting or non-superconducting. For example, one or more or all of the coils can be made from copper or any other suitable non-superconducting material, examples of which are described herein. One or more or all of the coils can be made from a superconducting material, examples of which are described herein. One or more or all of the coils can have a configuration as described below with respect to FIG. 24. Due to compaction and space limitations, the ends of the coils are configured to meet predefined critical strain / stress limits and a predefined total field integral over the coil mandrel (in a non-limiting example, 3.54 Tesla-meters).

[0031] The coils are disposed (e.g., wrapped, held, placed, disposed, or maintained) on non-ferromagnetic support 205 (see FIG. 26 , not shown in FIGS. 25 , 28 , or 29 ). Support 205 can be made of or include a non-ferromagnetic material such as aluminum or stainless steel. Support 205 can be a single continuous or integrated structure, or support 205 can include multiple separate structures that together make up the support structure. Support structure 205 can have a shape complementary to the shape of the coils. For example, as shown in FIG. 26 , support 205 has a shape complementary to the shape of coils 201 a, 201 b, 201 c, 201 d, 201 e, and 201 f. Together, support 205 and coils 201a, 201b, 201c, 201d, 201e, and 201f define a substantially circular cross-section, as shown in Figure 26, which defines a space that includes air core 213. Air core 213 can contain a gas, such as air or a noble gas, or it can be a vacuum (e.g., 10 -5 It is possible to approach pressures of up to 1000 kPa (0.0013332 Pascals or less).

[0032] Current flowing through coils 201a, 201b, 201c, 201d, 201e, and 201f generates a magnetic field that is shaped (at least in part) by ferromagnetic yoke 181. Naturally, the magnitude of the current also shapes the magnetic field. Yoke 181 can be a solid structure made of a ferromagnetic material, such as iron, as shown in FIGS. 28 and 29. As shown in FIGS. 28 and 29, yoke 181 can be formed from a top piece 181a and a bottom piece 181b. However, in other implementations, yoke 181 can be formed from a left piece and a right piece, or can be formed from more than two pieces. In FIG. 28, portion 181c of yoke 181 is illustrated as transparent, while in FIG. 29, portion 181a of yoke 181 is depicted as solid. The transparent form is intended to show the coil 180 through the yoke, not to indicate that all or part of the yoke 181 is actually transparent.

[0033] Yoke 181 surrounds (at least partially) the assembly consisting of support 205, coils 201a, 201b, 201c, 201d, 201e, and 201f, and core 213. In the exemplary implementation of FIGS. 28 and 29, yoke 181 is shown in cutaway form to illustrate coil 180. However, in the exemplary implementation of FIGS. 28 and 29, yoke 181 extends to cover the entire length of coil 180, from end 184 (best seen in FIG. 28) to end 185 (see FIG. 29, not shown in FIG. 28). Ends 186a, 186b ( FIG. 28) of coil 180 are left exposed to allow connection of magnet 200 to other magnetic components in the beamline and to allow a particle beam to pass through magnet 200 in the manner described herein.

[0034] In some implementations, the yoke 181 includes rounded notches or channels adjacent to the coil 180 and around the outer surface of the coil 180. The notches can run along the entire length of the yoke 181 / magnet 200 and can have the same cross-section along the entire length of the yoke 181 / magnet 200, or their cross-sections can vary along the length of the yoke 181 / magnet 200. For example, as shown in FIGS. 26 and 29 (not shown in FIG. 28), the notches 220 are rounded or semicircular and extend around the circular cross-section of the assembly consisting of the support 205, the coils 201a, 201b, 201c, 201d, 201e, 201f, and the core 213. The lower hemisphere 200b is a mirror image of the upper hemisphere 200a, such that the cutouts on the lower hemisphere 200b are a mirror image of the cutouts on the upper hemisphere 200a. In contrast, the cutouts 220 on the right and left hemispheres 188a, 188b are asymmetric to account for bending of the magnet 200. In one example, the asymmetry is such that the cutouts have a larger volume / size, on average, in the right hemisphere 188a than in the left hemisphere 188b, resulting in more ferromagnetic material being present in the left hemisphere 188b than in the right hemisphere 188a. In one example, the asymmetry is such that the cutouts are shaped differently in the right hemisphere 188a than in the left hemisphere 188b, causing more ferromagnetic material to be present in the left hemisphere 188b than in the right hemisphere 188a. In one example, the asymmetry is such that the cutouts are, on average, closer to each other in the right hemisphere 188a than in the left hemisphere 188b, resulting in more ferromagnetic material being present in the left hemisphere 188b than in the right hemisphere 188a. In one example, the asymmetry is such that the cutouts are, on average, closer to each other and larger in the right hemisphere 188a than in the left hemisphere 188b, resulting in more ferromagnetic material being present in the left hemisphere 188b than in the right hemisphere 188a.For example, in FIG. 26 , notches 220a and 220b in the right hemisphere 188a are larger and closer to each other (in fact, they overlap) than their counterparts 220c and 220d in the left hemisphere 188b. In one example, the asymmetry is such that the number of notches in the right hemisphere 188a is greater than the number of notches in the left hemisphere 188b, resulting in more ferromagnetic material in the left hemisphere 188b than in the right hemisphere 188a. Any asymmetry resulting from notch configuration, placement, number, size, shape, and / or other factors can be used to shape the magnetic field. The placement and configuration of the notches affect the magnetic field to enable particle beam transmission and maintain particle beam integrity.

[0035] 25 and 26, magnet 200 has more current-conducting coils nearer the 0° / 180° location than at 90°. Therefore, while the magnet is operating, i.e., while current is conducting through the coils, there is a greater current density nearer the 0° / 180° location than there is at 90°. In the example shown, there are no conductors at 90°. Therefore, the current density at 90° is zero. As noted, in this example, lower hemisphere 200b is a mirror image of upper hemisphere 200a. Therefore, in lower hemisphere 200b, while current is conducting through the coils, there is a greater current density nearer the 0° / 180° location than there is at -90° / 270°.

[0036] Another feature of magnet 200 is that coils 201a, 201d closer to 0° / 180° have a larger current-carrying capacity than coil 201c closer to 90° and coil 201f closer to -90° / 270°. For example, coils 201a, 201d have a larger cross-sectional area than coils 201c, 201f, respectively. Generally, in implementations of magnet 200, the current-carrying capacity of the coils decreases from 0° / 180° to 90° and from 0° / 180° to -90° / 270° in each quadrant 210a, 210b, 210c, and 210d (FIG. 26).

[0037] In this respect, quadrant 210a extends from 0° to 90°, quadrant 210b extends from 90° to 180°, quadrant 210d extends from 180° to -90° / 270°, and quadrant 210c extends from 270° to 360° / 0°. In this example, coil 201b (which is between coils 201a and 201c) has a cross-sectional area that is smaller than that of coil 201a and larger than that of coil 201c. Similarly, coil 201e (which is between coils 201d and 201f) has a cross-sectional area that is smaller than that of coil 201d and larger than that of coil 201f. The reduction in current-carrying capacity from 0° / 180° to 90° and from 0° / 180° to -90° / 270° can be constant or can vary. In a certain example, for hemisphere 200a, coil 201c can have 20% less current-carrying capacity than coil 201b, and for hemisphere 200a, coil 201b can have 20% less current-carrying capacity than coil 201a. In a varying example, coil 201c can have 20% less current-carrying capacity than coil 201b, and coil 201b can have 10% less current-carrying capacity than coil 201a. The same difference in current-carrying capacity can be maintained for the coil's counterpart in hemisphere 200b. The 10% and 20% figures are non-limiting examples of 201, and the reduction in current carrying capacity in the coils of cosine theta magnets herein from 0° / 180° to 90° (or from 0° / 180° to -90° / 270°) may be greater or less than these figures.

[0038] As shown in Figures 25 and 27-29, cosine theta magnet 200 is curved. This feature is particularly useful for steering particle beams in compact gantries of the type described herein. The cosine theta magnet can be curved by 10° or more, 20° or more, 30° or more, 40° or more, 50° or more, 60° or more, 70° or more, 80° or more, 90° or more, 100° or more, 110° or more, 120° or more, 130° or more, 140° or more, 150° or more, 160° or more, 170° or more, or even 180° relative to line 211 (Figures 25 and 28), thus bending the particle beam where line 211 passes through and follows the trajectory of unbent parts of magnet 200. Magnet 200 can be curved at any suitable angle. The bend can be any degree between 0° and 90°, any degree between 0° and 180°, or any degree between 70° and 180° relative to the straight line 211.

[0039] The homogeneity of the magnetic field in rectangular region 212 (FIGS. 26, 27) of core 213 of magnet 200 (where the particle beam is constrained to travel by the magnetic field) can be distorted by bending in the cosine theta magnet. More specifically, bending in the magnet can cause the particle beam cross-section (i.e., spot) to become elliptical rather than remain circular. For example, the particle beam spot size (e.g., cross-sectional area) can have an aspect ratio of 5% or more and grow in size from 3 millimeters (mm) or 4 mm sigma to 10 mm sigma in one or more planes due to bending. However, magnet 200 is configured to counter such distortion and maintain the particle beam cross-section in region 212 in a predefined shape, such as circular. To at least partially counter such distortion, sets of coils in magnet 200 are asymmetrically disposed on supports 205 in each hemisphere 200 a, 200 b to shape the magnetic field to prevent or reduce distortion. For example, as shown in FIG. 26 , in hemisphere 200a, spacing 214 between coil 201a and coil 201b in quadrant 210a is different from spacing 215 between coil 201a and coil 201b in quadrant 210b. Similarly, in the same hemisphere 200a, spacing 217 between coil 210b and coil 210c in quadrant 210a is different from spacing 218 between coil 210b and coil 210c in quadrant 210b. The same spacing difference between coils in hemisphere 200a exists in the mirror-image coils in hemisphere 200b. That is, the spacing difference between coils 201d, 201e and coils 201e, 201f is the same as the spacing difference between coils 201a, 201b and coils 201b, 201c, respectively. In some implementations, the coil spacing difference in different quadrants may be different in different hemispheres. In this example, the spacing between pairs of coils in quadrant 210a is greater than the spacing between the same pairs of coils in quadrant 201b.However, in other implementations, the spacing between pairs of coils in quadrant 210s can be smaller than the spacing between the same pairs of coils in quadrant 201b.

[0040] As shown in FIGS. 25-29, the curvature or bending in magnet 200 creates inner surface 206 and outer surface 207, with the inner surface having a smaller curvature or bending radius than the outer surface. The inner and outer surfaces may refer to a coil, assembly, or magnet, since they have the same or substantially the same bending radius. In this example, coils 201a, 201b, 201c, 201d, 201e, and 201f are spaced apart such that the spacing between the coils is smaller at outer surface 207 than at inner surface 206. For example, at outer surface 207 and quadrant 210b of magnet 200, coils 201a, 201b, and 201c are closer together than they are at inner surface 206 and quadrant 210a of the magnet. In this configuration, for example, space 215 is smaller than space 214, and space 218 is smaller than space 217. The difference in spacing can be on the order of 1%, 2%, 3%, 4%, 5%, 6%, 7%, 8%, 9%, or more. For example, space 215 can be 1%, 2%, 3%, 4%, 5%, 6%, 7%, 8%, 9%, or more smaller than space 214, and space 218 can be 1%, 2%, 3%, 4%, 5%, 6%, 7%, 8%, 9%, or more smaller than space 217. Any suitable spacing can be used to achieve the effects described herein.

[0041] These spacing differences can be reflected in the coils positioned in lower hemisphere 200b. More specifically, coils 201d, 201e, and 201f are arranged in a mirror image configuration of coils 201a, 201b, and 201c, and therefore will have the same spacing variations as coils 201a, 201b, and 201c, except that they will be in quadrant 201d (which will have the same coil spacing as quadrant 201b) and quadrant 201c (which will have the same coil spacing as quadrant 201a).

[0042] The combination of the asymmetry of coil 180 in right hemisphere 188a and left hemisphere 188b, the greater current-carrying capacity of the coils closer to 0° (e.g., 201a, 201d) than 90° (e.g., 201c, 201f), and the asymmetry of cutouts 220 in right hemisphere 188a and left hemisphere 188b shapes the magnetic field of magnet 200 in region 212 to (i) keep the particle beam within region 212 and (ii) keep the particle beam circular or substantially circular as it travels the length of magnet 200 (including traveling through parts of magnet 200 that curve). In some examples, substantially circular can include a deviation from perfectly circular of 9% or less, 8% or less, 7% or less, 6% or less, 5% or less, 4% or less, 3% or less, 2% or less, or 1% or less. Other factors that affect the magnitude and shape of the magnetic field include, but are not limited to, the current-carrying capacity of the coil, the size and shape of the yoke 181, and the materials from which the various components of the magnet 200 are made.

[0043] FIG. 27 illustrates the magnitude and location of the magnetic field 206 in grayscale or color for the upper hemisphere 200a for the exemplary implementations of FIGS. 25, 26, 28, and 29. As shown, the magnetic field strength is greater at the inner surface 206 of the magnet 200 than at the outer surface 207 of the magnet. The magnetic field is greater at the inner radius as a result of the solution of Biot-Savart's law for this curved conductor geometry. At this point, if the magnetic field were the same strength across all radii (e.g., from the inner surface 206 to the outer surface 207), there would be a net focusing of the beam in the bending plane, resulting in different particle trajectories through the magnet. This can overfocus the particle beam, which can be compensated for by reducing the magnetic field strength at the outer radius (207) relative to the magnetic field strength at the inner radius (206). As a result of this magnetic field difference, distortion of the particle beam spot (i.e., the cross section of the particle beam) can be reduced or eliminated. As a result, the particle beam is able to remain substantially circular as it travels through the entire length of magnet 200 .

[0044] In a non-limiting example, magnet 200 achieves 0.1% homogeneity of the magnetic field (e.g., dipole) in region 212, which can be 100 millimeters (mm) by 90 mm along the beam trajectory throughout the length of the magnet.

[0045] In some implementations, one or more magnetic shims (not shown) can be used to vary the amount of ferromagnetic material in the yoke 181 (and thus the magnet). For example, the shims can be rods, cones, or other structures that can be controlled to move into or out of the yoke 181 to adjust the amount of ferromagnetic material in the yoke 181 and thereby change the shape of the magnetic field produced by the magnet. In some implementations, the shims can be manually controlled. In some implementations, the shims are computer-controlled. For example, each shim can be connected to a computer-controlled actuator that controls the movement of the shim into or out of the ferromagnetic core. The shims can be moved to be fully embedded in or fully ejected from the yoke either through manual or computer control. There can be one shim per magnet quadrant, and there can be, for example, two, three, four, five, etc. per quadrant. In one example, one or more magnetic field sensors can detect the magnetic field produced by magnet 200, and shims can be controlled to change the magnetic field to a magnetic field having a target shape. In one example, one or more sensors can detect the location of the particle beam within core 213, and shims can be controlled to change the magnetic field to control particle beam placement.

[0046] FIG. 30 illustrates an exemplary assembly 230 that can be used in a magnet such as magnet 200 (see, e.g., FIG. 26 ). Assembly 230 includes support 231, which has the same function as support 205 and can have a similar structure and composition as support 205. Assembly 230 also includes coils 232 a, 232 b, 232 c, and 232 d. The sets of coils are asymmetrically disposed on support 231 in each hemisphere 235 a, 235 b to shape the magnetic field to at least partially prevent or reduce particle beam distortion as described herein. For example, as shown in FIG. 30 , in hemisphere 235 a, the spacing between coils 232 a and 232 b in quadrant 236 a at the magnet's inner curvature radius 300 is different (e.g., larger) than the spacing between coils 232 a and 232 b in quadrant 236 b at the magnet's outer curvature radius 301. The same spacing difference between coils in hemisphere 235a exists in the mirror-image coils in hemisphere 235b. That is, in hemisphere 235b, the spacing between coils 232c and 232d in quadrant 236c is different (e.g., larger) than the spacing between coils 232c and 232d in quadrant 236d. Assembly 230 differs from that of FIG. 26 in that assembly 230 includes two sets of coils rather than three sets of coils. Assembly 230 can otherwise be incorporated into magnet structures such as those shown in FIGS. 28 and 29, with all accompanying features configured for a two-coil design rather than a three-coil design. These features include, but are not limited to, asymmetric cutouts and magnetic yokes configured for a two-coil design.

[0047] Figure 31 shows another exemplary assembly 240 that can be used in a magnet such as magnet 200 (see, e.g., Figure 26). Assembly 240 includes support 241, which has the same function as support 205 and can have a similar structure and composition as support 205. Assembly 240 also includes coils 242a, 242b, 242c, 242d, 242e, 242f, 242g, and 242h. The sets of coils are asymmetrically disposed on support 241 in respective hemispheres 245a, 245b to shape the magnetic field to at least partially prevent or reduce particle beam distortion as described herein. For example, as shown in FIG. 31 , in hemisphere 245a, the spacing between coils 242a and 242b in quadrant 246a at the magnet's inner radius of curvature 303 is different (e.g., larger) than the spacing between coils 242a and 242b in quadrant 246b at the magnet's outer radius of curvature 304. The spacing between coils 242b and 242c in quadrant 246a is different (e.g., larger) than the spacing between coils 242a and 242b in quadrant 246b. The spacing between coils 242c and 242d in quadrant 246a is different (e.g., larger) than the spacing between coils 242c and 242d in quadrant 246b. The same spacing difference between coils in hemisphere 245a exists in the mirror image coils in hemisphere 245b. That is, in hemisphere 245b, the spacing between coils 242e and 242f in quadrant 246c at the magnet's inner radius of curvature is different (e.g., larger) than the spacing between coils 242e and 242f in quadrant 246d at the magnet's outer radius of curvature. The spacing between coils 242f and 242g in quadrant 246c is different (e.g., larger) than the spacing between coils 242f and 242g in quadrant 246d. The spacing between coils 242g and 242h in quadrant 246c is different (e.g., larger) than the spacing between coils 242g and 242h in quadrant 246d.Assembly 240 differs from that of Figure 26 in that assembly 240 includes four sets of coils instead of three sets of coils. Assembly 240 can otherwise be incorporated into magnet structures such as those shown in Figures 28 and 29, with all accompanying features configured for a four-coil design rather than a three-coil design. These features include, but are not limited to, asymmetric cutouts and magnetic yokes configured for a four-coil design.

[0048] In the following description of particle beam gantries, cosine theta bending magnets of the type described with respect to Figures 25 through 31, or any variations thereof, can implement, for example, any or all of the bending magnets and variations thereof described with respect to Figures 1, 2, 3, 4, 6, 9, 10, 16, and 21. For example, cosine theta magnets of the type described herein having dipole, quadrupole, and / or hexapole configurations can be used to implement bending magnets in any of the particle beam gantries described herein, or variations thereof.

[0049] FIG. 1 illustrates an exemplary implementation of a particle therapy system 10 of the type described above, which may include one or more bending magnets or variations thereof of the type described with respect to FIGS. 25-31. The particle therapy system 10 includes a particle accelerator 12, examples of which are described herein. 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 greater, or even 3 T or greater. In this regard, a superconductor is an element or metal alloy, such as niobium-tin (NbSn), 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.

[0050] 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 weighs 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 can also be used in particle therapy system 10.

[0051] 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 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 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 location. For example, the output channel 17 can rotate through 360°, and as such, the output channel 17 can be positioned at 0°, 90°, 270°, and back to 0° / 360°, or any angle between these rotational positions.

[0052] 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. As mentioned, 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.

[0053] 2 and 3 , the conduit 18 of the exemplary beamline structure 16 includes non-superconducting magnetic quadrupoles 21 and 22 and a bending magnet 23, which may be a superconducting dipole magnet. The magnetic quadrupoles 21 and 22 are configured to maintain the particle beam focused and traveling straight or substantially straight (e.g., within a 5% or less 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 ±5% tolerance). The bending magnet 23 is configured to bend the particle beam toward the output channel 17, as shown. The bending magnet 23 may be configured to bend the particle beam anywhere within a range of 20° to 80° relative to a horizontal direction 24. Generally, 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, in turn, reducing the size of the particle therapy system. For example, replacing the bending magnets 23 with non-superconducting bending magnets 23 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, in turn, to the treatment location 19 and isocenter. The bending magnets 23 can be or include bending magnets of the type shown in (and described with respect to) FIGS. 25-31 or variations thereof.

[0054] 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 magnetic quadrupoles. The magnetic sextupoles are configured to keep the particle beam focused and traveling straight or substantially straight (e.g., within 5% or less of straightness) 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 ±5% tolerance). Sextupoles can also compensate for chromatic effects of quadrupole magnets. Compared to magnetic quadrupoles, magnetic sextupoles have a greater focusing effect on particles displaced farther from the axis defining the ideal location of the beamline, such as in region 212 of FIG. 26 .

[0055] Referring back to FIG. 3 , in this example, conduit 18 of beamline structure 16 also includes two non-superconducting magnetic quadrupoles 26 and 27. 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 beamline structure 16. Magnetic quadrupoles 26 and 27 are configured to maintain a consistent cross-sectional area of ​​the particle beam (e.g., within a ±5% tolerance). As previously described, higher order magnetic components may be substituted for one or more of the magnetic quadrupoles to improve focusing.

[0056] 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 the maximum (e.g., planar) extent to 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 larger, 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 larger, but the scanning magnet is not limited to any particular aperture size.

[0057] The scanning magnet can be positioned at different locations within 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 within the nozzle 40a, along with the energy degrader 41a and collimator 44a (both described below), on the path of the particle beam between the output channel 17a and the treatment location. Referring to FIG. 5, an exemplary scanning magnet 43 is controllable in two dimensions (e.g., Cartesian X-Y dimensions) to position and move the particle beam across at least a portion of the irradiation target. In this example, the scanning magnet 43 includes two coils in a first set 45 and two coils in a second set 46, where the two coils in the first set 45 control particle beam movement in the Cartesian X-dimension of a defined coordinate system, and the two coils in the second set 46 are orthogonal to the two coils in the first set and control 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, therefore, across the irradiation target.

[0058] In some implementations, there may be more than one scanning magnet. An implementation including multiple scanning magnets at different points along the particle beam path and separated by air or a structure (e.g., a magnet or a beam absorbing plate) may be referred to as a split-scan system. For example, in the beamline structure 16b shown in FIG. 6 (which is a variation of the beamline structure 30), multiple (e.g., two) scanning magnets 30b1 and 30b2 may be present between the output channel 17b and the treatment position. The scanning magnets may be positioned in the nozzle 40b along with the energy degrader 41b and the collimator 44b on the particle beam path between the output channel 17b and the treatment position. The scanning magnets may be in separate locations and separated by air or an energy degrading structure. For example, in this implementation, first scanning magnet 30b1 is capable of moving the particle beam in two dimensions (e.g., Cartesian X and Y dimensions), and second scanning magnet 30b2 is capable of moving the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, 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 produces the desired movement specified in the treatment plan.

[0059] In the variation shown in the implementation of FIG. 6, scanning magnet 30b1 can move the particle beam in only one dimension (e.g., the Cartesian X dimension), and 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, scanning magnet 30b1 can be of the type shown in FIG. 7 and include a first set of coils 90a, and scanning magnet 30b2 can be of the type shown in FIG. 8 and 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.

[0060] In some implementations, one or more (e.g., all or fewer) of the scanning magnets can be positioned within the beamline structure. For example, in the beamline structure 16c shown in FIG. 9 (which is a variation of the beamline structure 16 of FIG. 3), 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 within the beamline structure 16c, and the second scanning magnet 30c2 is positioned outside the beamline structure within the nozzle 40c, along 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, first scanning magnet 30c1 can be positioned in output channel 17c, upstream of bending magnet 32c relative to the particle accelerator, or, as shown in FIG. 9, first scanning magnet 30c1 can be positioned upstream of output channel 17c relative to the particle accelerator. Bending magnet 32c can be or include a bending magnet of the type shown in (and described with respect to) FIGS. 25-31 or variations thereof. In one example, first scanning magnet 30c1 can be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions), and 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, 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 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.

[0061] In a variation of the implementation of FIG. 9, first scanning magnet 30c1 can be configured to move the particle beam in only one dimension (e.g., the Cartesian X dimension), and 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, scanning magnet 30c1 can include a first set of coils, and scanning magnet 30c2 can include a second set of coils that are orthogonal to the first set of coils. Magnets 30c1 and 30c2, in this example, can have a configuration similar to the magnets shown in FIGS. 7 and 8. Each magnet 30c1 and 30c2 can be configured to move a portion of the beam, and the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0062] In some implementations, all of the scanning magnets can be positioned within the beamline structure. As shown in the split-scan system of FIG. 10, both the first scanning magnet 30d1 and the second scanning magnet 30d2 can be positioned within the beamline structure 16d. The scanning magnets do not have to be positioned within 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 within the nozzle. The first scanning magnet 30d1 and the second scanning magnet 30d2 can be positioned between magnetic components included within the beamline structure 16d. For example, as shown in FIG. 10, the first scanning magnet 30d1 can be positioned within 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., dipole or quadrupole magnets) 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 bending magnet 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 bending magnet 32d, so that both bending magnets 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 bending magnets 31d and 32d can be between scanning magnets 30d1 and 30d2.In one example, first scanning magnet 30d1 can be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions), and 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, scanning magnets 30d1 and 30d2 can have the same construction and operation as the scanning magnets shown in FIG. 5. Each magnet 30d1 and 30d2 can move a portion of the particle beam, and the combined movement produced by the two scanning magnets produces the desired particle beam movement specified in the treatment plan. Bending magnets 23, 31d, and 32d can be or include bending magnets of the type shown in (and described with respect to) FIGS. 25-31 or variations thereof.

[0063] In a variation of the implementation of FIG. 10 , first scanning magnet 30d1 can be configured to move the particle beam in only one dimension (e.g., the Cartesian X dimension), and 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, scanning magnet 30d1 can include a first set of coils, and scanning magnet 30d2 can include a second set of coils that are orthogonal to the first set of coils. Magnets 30d1 and 30d2, in this example, can have a configuration similar to the magnets shown in FIGS. 7 and 8 . Each magnet 30d1 and 30d2 can be configured to move a portion of the beam, and the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0064] In some implementations, there can 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 can be more than two scanning magnets positioned at various separate locations in the beamline structure. For example, there can 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 can be arranged in series.

[0065] In some implementations, there may be a single scanning magnet positioned upstream of the output of an output channel or elsewhere in the beamline structure. For example, as shown in FIGS. 2 and 3, a scanning magnet 30 may be positioned upstream of output channel 17 and at the input of output channel 17 relative to the particle accelerator. Scanning magnet 30 may be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, scanning magnet 30 may have the same construction and operation as the scanning magnet shown in FIG. 5. In this example, all particle beam movement is implemented by controlling current through one or more coils of the single scanning magnet.

[0066] 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.

[0067] 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 a 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 for the generation of a magnetic field of 2.5 T or more or 3 T or more to move the particle beam, which can overcome the effect on the particle beam of the high magnetic field (e.g., 2.5 T or more or 3 T or more) created by the beamline structure.

[0068] FIG. 11 shows 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 shown 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., above 77° Kelvin (K) or 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 examples, 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 a cryogen (e.g., liquid helium). In operation, current is applied to coils 92a and 92b to generate the magnetic field used for scanning.

[0069] 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., 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 96 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 coil to the superconducting temperature. A current is applied to the coil 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 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.

[0070] 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 within 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 the control system described herein.

[0071] In FIG. 23 , grid 151 depicts a scanning beam aperture in both the Cartesian X and Y dimensions (153 and 154, respectively). For example, grid 151 indicates that scanning magnet 150 can move a 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 lengths 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 wrapped around an electrically non-conducting or electrically non-superconducting material 160 to create aperture 161 containing grid 151. Inner superconducting coil 158 can be separated from outer superconducting coil 159 by electrically non-conducting or electrically non-superconducting material 160. 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.

[0072] In this example, superconducting coils 158 control the movement of the particle beam in the X dimension. For example, currents run through the superconducting coils to create a magnetic field. The strength of the magnetic field is proportional to the amount of current running through the superconducting coils, 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, currents run through the superconducting coils to create a magnetic field. The strength of the magnetic field is proportional to the amount of current running through the superconducting coils, 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 simultaneously 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, causing the particle beam to move in the X or Y dimension at separate times but still reach a target location.

[0073] 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.

[0074] FIG. 24 shows a cross section of an exemplary superconducting coil 165 that may be used to implement each of superconducting coils 158 and 159 and / or the coils described with respect to the exemplary bending magnets of FIGS. 25-31. Superconducting coil 165 includes a copper (Cu) stabilization layer 166 that wraps around or surrounds the other layers of superconducting coil 165. 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 Ag 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 (e.g., 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 or may include different materials than those shown. For example, the copper stabilization layer may be omitted, or a material other than copper may be used. Other types of superconducting materials may also be used, such as YBCO and / or BSCCO.

[0075] Referring back to Figure 3, the output channel 17 portion of the beamline structure 16 includes a large diameter superconducting (or non-superconducting) bending magnet 31 arranged in series with a large diameter superconducting (or non-superconducting) bending magnet 32. Examples of large diameters include, but are not limited to, 20 cm x 20 cm. The bending magnets 31 and 32 can be or include bending magnets of the type shown in (and described with respect to) Figures 25-31 or variations thereof.

[0076] Positioned between bending magnet 31 and bending magnet 32 ​​are multiple large-diameter superconducting (or non-superconducting) magnetic quadrupoles 33, 34, and 35. In this example, 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 particles passing through the alternating magnetic field gradients of the magnetic quadrupoles is to cause the beam to converge (i.e., 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 channels 17 can include different numbers of magnetic quadrupoles in different configurations and / or different numbers of magnetic dipoles in different configurations. In some implementations, output channels 17 can include higher order magnetic components (e.g., hexapoles, etc.) instead of or in addition to the magnetic quadrupoles shown.

[0077] In some implementations, using bending magnets 31 and 32, 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 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, using bending magnets 31 and 32, 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 , using bending magnets 31 and 32, output channel 17 is configured to bend the particle beam at a combined total bending angle of approximately 150° relative to line 38. To achieve a bending magnitude having a value between 110° and 170°, bending magnet 31 can be configured to bend the particle beam within a range of 20° and 85° relative to line 38, and bending magnet 32 ​​can be configured to bend the particle beam within a range of 20° and 85° relative to horizontal line 38.

[0078] In some implementations, output channel 17 can include different numbers of magnetic structures in different configurations. For example, output channel 17 can include a bending magnet of the type described herein, followed by three alternating magnetic quadrupoles, followed by a bending magnet of the type described herein, followed by three alternating magnetic quadrupoles, followed by a bending magnet 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 .

[0079] 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 (e.g., 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. Additionally, as described herein, the scanning magnet 30 can move the particle beam in a plane and across the irradiation target.

[0080] In this regard, as previously described, the nozzle can 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 FIGS. 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 (via the nozzle 40) between the output channel 17 and the irradiation target at the treatment location 19. 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, 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.

[0081] 3, in some implementations, a single quadrupole magnet can be used in place of quadrupoles 21, 22, the quadrupole magnet can replace scanning magnet 30, scanning magnet 30 can replace quadrupole magnet 35, and the nozzle can include a second scanning magnet. The scanning magnets in this case can scan in two dimensions or one dimension, respectively, as described herein, or one can scan in two dimensions and one can scan in one dimension.

[0082] As previously mentioned, the particle beam output by the accelerator can be monoenergetic, and the energy degrader is the sole 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 an energy degrader 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.

[0083] 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 wedges, which at least partially overlap and can be moved within 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 present 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 the maximum range over 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.

[0084] 22 , in one example, the energy degrader 48 is a range modulator that can be controlled to move a structure 42 into (and out of) the path of the particle beam to vary the energy of the particle beam and, therefore, the depth to which the particle beam dose is deposited within 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 manner, the energy degrader can cause the particle beam to deposit a dose of radiation within the interior of the irradiation target to treat a layer or column of the target. In this regard, as protons at a particular energy travel through tissue, they ionize tissue atoms 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 the 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 particle beam movement (scanning) and to track or follow the particle beam during 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 discussion), are incorporated herein by reference.

[0085] The Bragg peak is a prominent peak on a 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 tissue. For protons, the Bragg peak occurs just before the particle comes to rest. Therefore, the energy of a particle beam can be varied to change the location of its Bragg peak and, therefore, the depth at which most of the proton dose will be deposited in 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 a 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 within 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 within a range of 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 instances, above or below that range.

[0086] In some implementations, the particle accelerator is a dual-energy accelerator. In a dual-energy particle accelerator, a particle beam exits the particle accelerator at one of two different energy levels (a high energy level or a low energy level). The terms "high" and "low" do not have a specific numerical connotation but rather are intended to convey relative magnitudes. In some implementations, the particle accelerators described herein are configured to output particle beams at two energies ranging from about 100 MeV to about 300 MeV. The high-energy output and the low-energy output can be values ​​within the range, or in some instances, can be values ​​above or below the range. The energy degraders described herein can be used with dual-energy particle accelerators 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.

[0087] In the figure, 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 controllable structure that allows some radiation to reach the target and blocks some radiation from reaching the patient. Typically, the passing radiation is directed to the irradiation target to be treated, while the blocked radiation would otherwise strike and potentially damage healthy patient tissue. In operation, the collimator is positioned in the radiation path between the output channel 17 and the irradiation target and is controlled to create an opening of an appropriate size and shape that allows some radiation to pass through the opening and reach the irradiation target, while the remainder of the structure blocks some radiation from reaching adjacent tissue.

[0088] The collimator can be configurable, e.g., its aperture can be controlled and changed during the procedure. The collimator can be fixed or non-variable, e.g., the collimator can have a fixed shape that cannot be changed.

[0089] In some implementations, an exemplary configurable collimator component includes multiple leaves that are dynamically reconfigurable during movement of the particle beam to change the 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.

[0090] 13, 14, and 15 illustrate 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 configured to hold (and move) the above-described leaves both vertically and horizontally relative to the irradiation target. As shown, vertical movement involves movement in a Cartesian Z dimension 117, and horizontal movement involves 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). Although FIGS. 14 and 15 illustrate portions of the carriage housing as transparent to show the components inside the housing, the housing is not actually transparent.

[0091] Carriage 113 is referred to herein as the primary carriage, and carriages 114 and 115 are referred to herein as secondary carriages. Secondary carriages 114, 115 are coupled to primary carriage 113 as shown in FIGS. 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 the secondary carriages to move vertically. In some implementations, the secondary carriages move vertically in unison.

[0092] 13-15 , each secondary carriage 114, 115 is connected to a corresponding rod or rail 122, 123 along which the secondary carriage moves. More specifically, in this example, motor 125 drives secondary carriage 114 to move along rod 122 toward or away from secondary carriage 115. Similarly, in this example, motor 126 drives secondary carriage 115 to move along rod 123 toward 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 well, as described herein.

[0093] As shown in FIG. 15 , motor 130 drives vertical movement of primary carriage 113. For example, as shown in FIG. 15 , lead screw 131 is coupled to housing 132, which holds motors 125, 126 that drive corresponding secondary carriages 114, 115, and housing 132 is mounted on track 120. Lead screw 131 is coupled to and driven vertically by motor 130. That is, motor 130 drives lead screw 131 in the vertical direction (Cartesian Z dimension). Because lead screw 131 is fixed to housing 132, this movement also causes housing 132 (and thus secondary carriages 114, 115) to move along track 120 either toward or away from the irradiation target.

[0094] 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, leaves can be configured to move in the Y dimension. Additionally, leaves on the secondary carriage 114 can be independently movable from leaves on other secondary carriages 115. These independent movements of leaves on the secondary carriages, along with the vertical movements enabled 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 the horizontal and vertical dimensions. The size and shape of the leaves can be varied to create different conformances. For example, the 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 that drive lead screws in the X dimension relative to other leaves on the same secondary carriage.

[0095] The leaves can be made of any suitable material that prevents or inhibits the penetration of radiation. The type of radiation used can dictate what material is used in the leaves. For example, if the radiation is x-rays, the leaves can be made of lead. In the examples described herein, the radiation is a proton beam or an ion beam. Therefore, 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 the leaf inhibits the penetration of radiation.

[0096] 13-15 are 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 those relating 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.

[0097] Referring back to FIG. 1 , as noted, an exemplary particle therapy system includes an isocentric gantry that is compact in size, reducing the overall system size. In the compact gantry 14 implementation, 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, measured from the accelerator output and the system isocenter, can be equal to the distance between the accelerator output and the system isocenter. In the compact gantry 14 implementation, 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 location can be less than 6 m, less than 5 m, less than 4.5 m, or less than 4 m. In implementations of the compact gantry 14, the distance between the output of the output channel 17 and the system isocenter or treatment position is 2 m or less, 1.5 m or less, or 1 m or less. In implementations of the compact gantry 14, the distance between the output of the output channel 17 and the system isocenter or treatment position is between 0.8 m and 14 m. In one example, the distance between the output of the output channel 17 and the system isocenter or treatment position is 1.01 m ( FIG. 2 ). Other implementations may have different dimensions than those listed here.

[0098] 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 FIGS. 1-3 can be small enough to fit into a vault having the following dimensions: length of 25 feet (7.62 m) or less, width of 20 feet (6.09 m) or less, and height of 11 feet (3.35 m) or less. For example, the components of FIGS. 1-3 can be small enough to fit into a vault having the following dimensions: length of 25 feet (7.62 m) or less, width of 26 feet (7.92 m) or less, and height of 10 feet (3.05 m) or less. For example, the components of FIGS. 1-3 may be small enough to fit within 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, and may have dimensions to fit within the vault. 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 (FIG. 1) may be dug below the floor of the LINAC vault to allow for rotation.

[0099] 1 and 16 show examples of treatment spaces 49 and 50 within which the particle therapy system 10 and its variations can be housed. The treatment spaces, in these examples, are implemented within a LINAC vault, 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 no more than 10 millisieverts of such neutrons per Gray dose delivered by the particle beam.

[0100] The use of a monoenergetic particle beam and reliance on an energy degrader outside the beamline structure allows magnetic components within the beamline to efficiently guide the beam. More specifically, changes in beam energy within the beamline increase the production of stray neutrons, which in turn increases losses of the particle beam within the beamline, thereby reducing its efficiency. The monoenergetic particle beam used in implementations of the systems described herein, combined with magnetic structures within the beamline, can lead to increased efficiency. In some cases, reducing the length of the beamline structure can also increase efficiency. In some implementations, variations in the beamline structures 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 structure that are output from the particle accelerator. Thus, an efficiency of 10% or greater includes that 10% or greater of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 20% or greater includes that 20% or greater of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 30% or greater includes that 30% or greater of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 40% or greater includes that 40% or greater of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 50% or greater includes that 50% or greater of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 60% or greater includes that 60% or greater of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 70% or greater includes that 70% or greater of the particles output from the particle accelerator are output from the beamline structure; an efficiency of 80% or greater includes that 80% or greater of the particles output from the particle accelerator are output from the beamline structure; and an efficiency of 90% or greater includes that 90% or greater of the particles output from the particle accelerator are output from the beamline structure.In one example, the particle accelerator and gantry described herein transmits over 70% of the proton beam to the patient, even at energies in the lower range of the accelerator.

[0101] The type of beamline efficiency described herein allows for a "single-room" solution, in which the particle accelerator, gantry, and patient all reside within a single vault, as described above. Within this vault, the particle accelerator itself may include shielding, but the separate compartments 60 and 61 (see FIG. 16 ) within the vault containing the patient and particle accelerator, respectively, do not need to be shielded from one another. 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.

[0102] 1 , the particle therapy system 10 also includes a treatment couch 51. The treatment couch 51 is configured to move relative to a bore 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 bore 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 forward-backward and left-right movement; examples of three degrees of freedom are forward-backward, left-right, and up-down movement; examples of four degrees of freedom are forward-backward, left-right, up-down, and one of pitch, yaw, or roll; examples of five degrees of freedom are forward-backward, left-right, up-down, and two of pitch, yaw, or roll; examples of six degrees of freedom are forward-backward, left-right, up-down, pitch, yaw, and roll. In some implementations, the treatment couch can be replaced by or include a couch that is at least partially tiltable or convertible into a chair and is still 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 configuration different from 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 movement of the treatment couch.

[0103] In some implementations, the output channel 17 can rotate at least partially (including completely) around the support structure 15, or the output channel can remain fixed on the support structure 15, and all or part of the support structure 15 can 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. 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 stationary beam during treatment. In some systems described herein, the location of the particle beam can be set through gantry rotation, after which the beam remains fixed except for scanning movement across the irradiation target, and the treatment couch or other seat moves during 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, and during that time the treatment couch can be moved to implement the treatment. Thereafter, the output channel can be repositioned to temporarily fix the beam at the new position. The treatment can be implemented at the new position via couch movement. These operations can be repeated as defined by a treatment plan designed for use with the particle therapy system.

[0104] The particle therapy system 10 may be an intensity-modulated proton therapy (IMPT) system. IMPT systems enable spatial control of a localized beam of protons, which may have variable energy and / or intensity. IMPT utilizes charged particle Bragg peaks (as described, characteristic peaks of dose at the edge of a particle's delivery range) combined with modulation of particle beam variables to generate target-local modulation of dose that achieves the objectives described in the treatment plan. IMPT may include directing a particle beam toward an irradiation target at different angles and intensities to treat the target. In some implementations, the particle beam may 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 may be performed using a scanning magnet, as described herein.

[0105] 17 shows an exemplary horizontal (x) beam envelope 63 and vertical (y) beam envelope 64 (e.g., cross-section) along the length 29 of the compact gantry described herein. The X and Y dimensions of the beam spot cross-section are determined with respect to 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 location) 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 structure 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 structure 16 is at least 3 T (i.e., 3 T or greater). As stated, the systems described herein are not limited to these parameter values, and some implementations may have different dimensions, energies, and magnetic fields.

[0106] Chromatic aberration correction can occur in a beamline with dispersion generated by including dipole magnets and multiple correctors in the dispersion region. The standard definition of achromatism is a beam transport line with zero values ​​for spatial dispersion (R16) and angular dispersion (R26). Referring to FIG. 18, the magnetic components in a compact gantry implementation can be configured to be achromatic (e.g., both R16 65 and R26 66 of the beam transfer matrix elements equal zero at the isocenter (which is at or near location 67) and at the beam entry point into the gantry at 0 m along the beamline structure length 68 (X-axis)). Thus, the gantry as a whole defines an achromat from the beam entry point to the isocenter. Individual magnets or combinations of magnets in the gantry (which constitute less than the totality of the magnets in the gantry along the beamline) need not be achromatic. Reducing or minimizing spatial and angular beam divergence can be important for pencil beam scanning techniques implemented by the particle therapy systems 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 sizes 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 entire scan area should be reduced or minimized; otherwise, different beam particles of different energies may land at different locations in the bending plane. This may cause the beam shape and beam size to differ in other planes.

[0107] 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 fully covers a 20 cm × 20 cm beam field area 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°.

[0108] 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 within the patient and / or control gantry and scanning movement to direct the particle beam to an irradiation target within 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.

[0109] 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 are rotatable on a ring bearing 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 bearing and rotatable therewith. In some implementations, the nozzles and imaging devices can be mounted on different internal circumferential tracks within the gantry. For example, the nozzles may be rotatable around a circumferential track at a first radius of the support structure, and the imaging devices may be rotatable around a different circumferential track at a second radius of the support structure that is different from the first radius. In some implementations, the gantry can include different rotatable inner rings, one of which mounts the nozzles for rotation and one of which mounts the imaging devices or systems for rotation.

[0110] In some implementations, two 2D imaging devices are mounted to 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.

[0111] As described herein, an exemplary proton therapy system scans a proton beam in three dimensions across an irradiation target to destroy malignant tissue. Figure 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, component 75 includes a superconducting magnet 77. The superconducting magnet includes superconducting coils 78 and 79. The superconducting coils are formed from multiple integrated conductors, each of which includes superconducting strands (e.g., four or six strands) wound around a central strand, which may itself be superconducting or non-superconducting. Each of 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 within a cavity 84, within which 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)).

[0112] 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 a noble gas, is ionized to create the plasma column. A voltage source provides a varying radio frequency (RF) voltage to the cavity 84, accelerating particles from the plasma column within the cavity. As noted, 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) 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 increasing relativistic mass of protons and the decreasing 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, increasing 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 packet of particles is periodically output.

[0113] The magnetic field within the cavity is shaped to cause particles to move orbitally within the cavity, as described above. Exemplary synchrocyclotrons use a magnetic field that is uniform with respect to the rotation angle and that falls off in strength with increasing radius. In some implementations, the maximum magnetic field produced by the superconducting (main) coils can be in the range of 2.5 T to 20 T at the center of the cavity, which falls off with increasing radius. For example, superconducting coils can be used in generating 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, 10.0T, 10.1T, 10.2T, 10.3T, 10.4T, 10.5T, 10.6T, 10.7 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 to generate magnetic fields outside the range of 2.5T to 20T, or magnetic fields within the range of 3T to 20T but not specifically listed herein.

[0114] By generating a high magnetic field having a magnitude such as that described above, the bending radius of particles traversing within 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 or less, or 2m 3 It is possible that:

[0115] 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 produced 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.

[0116] 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 the current through the main superconducting coils. In some implementations, there is an active return coil for each superconducting main coil, e.g., two active return coils, one for each main superconducting coil. Each active return coil can also be a superconducting coil that concentrically surrounds 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 those related to return coil configurations (e.g., Figures 2, 4, and 5 of U.S. Pat. No. 8,791,656 and the accompanying description), are incorporated herein by reference.

[0117] 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 relating to the "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.

[0118] In some implementations, the synchrocyclotron used in the proton therapy systems described herein can be a variable-energy synchrocyclotron. In some implementations, the variable-energy synchrocyclotron is configured to change the energy of the output particle beam by changing the magnetic field in which the particle beam is accelerated. For example, the current can be set to one of multiple 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 exemplary implementations, 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., sets or sets for variable current, etc.) are non-superconducting (e.g., copper) coils.

[0119] 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 over a predetermined range can generate a magnetic field that varies over a corresponding predetermined range. In some examples, continuous adjustment of the current can lead to continuous changes in the magnetic field and continuous changes in the output beam energy. Alternatively, when the current applied to the coils is adjusted in a discontinuous, stepwise manner, the magnetic field and output beam energy also change accordingly in a discontinuous (stepwise) manner. Stepwise adjustment can produce the dual energy previously described. In some implementations, each step is between 10 MeV and 80 MeV in size. Scaling the magnetic field with respect to the current can allow changes in beam energy to be implemented 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 the contents 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.

[0120] In implementations of particle therapy systems using variable-energy synchrocyclotrons, controlling the energy of the particle beam to treat a portion of the irradiation target can be performed according to a treatment plan by varying the energy of the particle beam output by the synchrocyclotron. In such implementations, 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 multiple values, each of the multiple 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 energy variation, for example, between the discrete energy levels provided by the synchrocyclotron.

[0121] The particle therapy system and its variations described herein can be used to apply ultra-high dose rates of radiation (so-called "FLASH" dose rates of radiation) to irradiation targets within a patient. In this regard, experimental results in radiation therapy have shown improved conditions for irradiated healthy tissue when treatment doses are delivered at ultra-high (FLASH) dose rates. In one example, when a dose of radiation between 10 and 20 Grays is delivered in pulses of less than 500 milliseconds (ms) to reach an effective dose rate of 20 to 100 Grays per second (Gy / s), healthy tissue experiences less damage than when the same dose is irradiated over a longer timescale, 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 within the tissue. In healthy tissue, ultra-high dose rates radicalize oxygen only once, as opposed to dose applications that radicalize oxygen multiple times over a longer timescale. This may lead to less damage in healthy tissue using ultra-high dose rates.

[0122] 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.

[0123] 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 Gray per second, 3 Gray per second, 4 Gray per second, 5 Gray per second, 6 Gray per second, 7 Gray per second, 8 Gray per second, 9 Gray per second, 10 Gray per second, 11 Gray per second, 12 Gray per second, 13 Gray per second, 14 Gray per second, 15 Gray per second, 16 Gray per second, 17 Gray per second, 18 Gray per second, 19 Gray per second, 20 Gray per second, 30 Gray per second, 40 Gray per second, 50 Gray per second, 60 Gray per second, 70 Gray per second, 80 Gray per second, 90 Gray per second, or 100 Gray per second. 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 seconds: 2 Gray per second, 3 Gray per second, 4 Gray per second, 5 Gray per second, 6 Gray per second, 7 Gray per second, 8 Gray per second, 9 Gray per second, 10 Gray per second, 11 Gray per second, 12 Gray per second, 13 Gray per second, 14 Gray per second, 15 Gray per second, 16 Gray per second, 17 Gray per second, 18 Gray per second, 19 Gray per second, 20 Gray per second, 30 Gray per second, 40 Gray per second, 50 Gray per second, 60 Gray per second, 70 Gray per second, 80 Gray per second, 90 Gray per second, or 100 Gray per second.

[0124] 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 of 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.

[0125] In some examples, the ultra-high dose rate of radiation comprises a dose of radiation between 20 Gray per second and 100 Gray per second for a duration of less than 500 ms. In some examples, the ultra-high dose rate of radiation comprises a dose of radiation between 20 Gray per second and 100 Gray per second for a duration of between 10 ms and 5 s. In some examples, the ultra-high dose rate of radiation comprises a dose of radiation between 20 Gray per second and 100 Gray per second for a duration of less than 5 s. In some examples, the ultra-high dose rate of radiation comprises a dose of radiation between 40 Gray per second and 120 Gray per second for a predetermined period of time (e.g., less than 5 s). Other examples of time periods are provided above.

[0126] In some implementations, particle therapy systems are capable of treating three-dimensional columns of targets 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 treatment into layers at a constant energy, the system organizes treatment into columns defined by the direction of the stationary beam. The beam direction may be toward the surface of the target.

[0127] 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 omnidirectional or partial 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, treatment occurs in a depth direction along the longitudinal direction of the beam. For example, the portion of the target being treated can extend from the beam spot on 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.

[0128] In some implementations, the irradiation target can be divided into microvolumes. While cubic microvolumes can be used, the microvolumes 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 microvolume is treated through the delivery of FLASH radiation column by column in the manner described herein. For example, column depths of the microvolume 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 microvolume is treated, the next microvolume is treated, and so on, until the entire irradiation target is treated. The treatment of the microvolumes can be performed in any suitable order or sequence.

[0129] 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 can be substituted for a synchrocyclotron in the particle therapy systems described herein.

[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, for example, the types shown in FIG. 5, 7, 8, 11, 12A, 12B, 23, 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] 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 irradiated is controlled by the range modulator. A configurable collimator can remain in place to trim the edges of the scattered beam.

[0132] Another exemplary particle therapy system 320 using bending magnets as described herein is shown in FIG. 21 . In FIG. 21 , a gantry 394 can be rotationally or axially connected to a treatment room floor 396, allowing controlled movement of the gantry 394 relative to the treatment room floor. In this example, a particle accelerator 10 is mounted on the gantry and can rotate with the gantry around the patient in the direction of arrow 321 to direct the particle beam toward the patient. The gantry 394 can include an arm 397 that runs the length of the gantry 394 and reaches the treatment room floor 396. The particle accelerator 10 and connected beamline structure 398 are rotatably mounted on the arm 397. That is, the particle accelerator 10 and connected beamline structure 398 are connected to an end 399 of the arm 397, allowing the particle accelerator 10 and connected beamline structure 398 to rotate at the end 399 in the direction of arrow 322. This rotation is separate from the gantry rotation described herein. The beamline structure 398 can contain one or more bending magnets of the type described with respect to Figures 25-31 or any variation thereof. For example, the beamline structure can include two bending magnets 350 and 351 of the type described with respect to Figures 25-31 or any variation thereof to bend the particle beam by 90° or more (e.g., 100°, 110°, 120°, or more) toward the irradiation target.

[0133] The operation of the example proton therapy systems described herein, and of all or some of their components, can be controlled (at least in part) using a control system 192 (FIG. 1) or 392 (FIG. 21) 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).

[0134] All or portions 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 control systems, 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.

[0135] 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.).

[0136] Processors suitable for the execution of a computer program include, by way of example, both general and special purpose microprocessors, and 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, or be operatively coupled to receive data from, transfer data to, or both, one or more machine-readable storage media, such as a mass storage device (e.g., magnetic, magneto-optical, or optical disk) for storing data. 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).

[0137] 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.

[0138] Other implementations not specifically described herein are within the scope of the following claims. [Explanation of symbols]

[0139] 10 Particle beam therapy system, particle accelerator 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, beamline structure output channels 17a Output Channel 17b output channel 17c output channel 17d output channel 18 Conduit 19 Treatment location 21 Non-superconducting magnetic quadrupole 22 Non-superconducting magnetic quadrupole 23 Superconducting magnetic dipoles, bending magnets, superconducting bending magnets, non-superconducting bending magnets 24 Horizontal 26 Non-superconducting magnetic quadrupole 27 Non-superconducting magnetic quadrupole 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), large diameter superconducting (or non-superconducting) bending magnet 31d Magnetic dipole, bending magnet 32 Magnetic dipole (large diameter superconducting magnetic dipole), large diameter superconducting (or non-superconducting) bending magnet 32c Magnetic dipoles, bending magnets 32d Magnetic dipole, bending magnet 33 Magnetic quadrupole (large diameter superconducting magnetic quadrupole) 34 Magnetic quadrupole (large diameter superconducting magnetic quadrupole) 35 Magnetic quadrupole (large diameter superconducting magnetic quadrupole), quadrupole magnet 38 horizontal line, 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 (two coils) 46 Second set (two coils) 48 Energy Degrader 49 Treatment Space 50 treatment spaces 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 Vertical (y) beam envelope 65 Beam Transfer Matrix Element R16 66 Beam Transfer Matrix Element R26 67 Isocenter, location 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 First set of coils 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 Magnet 95a High Temperature Superconducting Coil Set 96 Cryostat 97 Superconducting Scanning Magnet 97a High-Temperature Superconducting Coil Set 98 Cryostat 99 Imaging Devices 113 Primary Carriage 114 Secondary Carriage 115 Secondary Carriage 117 Cartesian Z dimension 118 Cartesian X-dimension, corresponding members 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 150 Superconducting Scanning Magnet 151 Grid 153 Cartesian X-Dimension 154 Cartesian Y-dimension 155 Standard 0,0 points 158 Inner superconducting coil 159 Outer superconducting coil 160 Electrically non-conductive materials, electrically non-superconductive materials 161 aperture 165 High-temperature superconducting tapes, superconducting coils 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 layers 192 Control Systems 180 coils 181 York 181a Upper piece 181b Bottom piece 181c Part of York 184 End 185 End 186a End 186b End 188a Right hemisphere 188b left hemisphere 200 magnets 200a upper hemisphere 200b lower hemisphere 201a Coil 201b Coil 201c coil 201d coil 201e coil 201f coil 205 Non-ferromagnetic Support 206 Inner surface, inner radius 207 outer surface, outer radius 210a quadrant 210b quadrant 210c quadrant 210d quadrant 211 straight line 212 Rectangular area 213 Air Core 214 Interval 215 interval 217 Interval 218 Interval 220 Notch 220a notch 220b notch 220c cutout 220d notch 230 Assembly 231 Support 232a Coil 232b coil 232c coil 232d coil 235a hemisphere 235b hemisphere 236a quadrant 236b quadrant 236c quadrant 236d quadrant 240 Assembly 241 Support 242a Coil 242b coil 242c coil 242d coil 242e coil 242f coil 242g coil 242h coil 245a hemisphere 245b hemisphere 246a quadrant 246b quadrant 246c quadrant 246d quadrant 300 inner curve radius 301 Outer curvature radius 303 Inner curve radius 304 Outer curvature radius 320 position, particle therapy system 321 Arrow 322 Arrow 350 bending magnet 351 Bent Magnet 392 Control Systems 394 Gantry 396 Treatment Room Floor 397 Arm 398 Beamline Structure 399 End

Claims

1. an assembly including: (i) a set of coils for conducting current to create a magnetic field; and (ii) a support structure on which the set of coils is asymmetrically disposed; a ferromagnetic yoke surrounding a portion of the assembly, the ferromagnetic yoke and the assembly being bent; Including, a magnet.

2. the set of coils includes a first coil and a second coil, the first coil and the second coil for conducting current to create a magnetic field, the first coil and the second coil asymmetrically disposed on the support structure in a first hemisphere of the magnet such that a first spacing between the first coil and the second coil in a first quadrant of the magnet is different from a second spacing between the first coil and the second coil in a second quadrant of the magnet, the first quadrant and the second quadrant being within the first hemisphere; the set of coils includes a third coil and a fourth coil, the third coil and the fourth coil for conducting current to create a magnetic field, the third coil and the fourth coil being asymmetrically disposed on the support structure in a second hemisphere of the magnet such that a third spacing between the third coil and the fourth coil in a third quadrant of the magnet is different from a fourth spacing between the third coil and the fourth coil in a fourth quadrant of the magnet, the third quadrant and the fourth quadrant being within the second hemisphere; asymmetries of the first and second coils in the first and second quadrants, respectively, mirror asymmetries of the third and fourth coils in the third and fourth quadrants, respectively; The magnet of claim 1.

3. the first interval and the third interval are equal, the second interval and the fourth interval are equal, and the first interval and the third interval are smaller than the second interval and the fourth interval; the first spacing and the third spacing are at an inside bend radius of the assembly, and the second spacing and the fourth spacing are at an outside bend radius of the assembly. The magnet according to claim 2.

4. the set of coils includes a fifth coil and a sixth coil, the fifth coil and the sixth coil for conducting current to create a magnetic field, the fifth coil being disposed on the support structure in the first hemisphere and the sixth coil being disposed on the support structure in the second hemisphere; a fifth spacing between the fifth coil and an adjacent one of the first or second coils in the first quadrant is different from a sixth spacing between the fifth coil and an adjacent one of the first or second coils in the second quadrant; a seventh spacing between the sixth coil and an adjacent one of the third or fourth coils in the third quadrant is different from an eighth spacing between the sixth coil and an adjacent one of the third or fourth coils in the fourth quadrant; asymmetries of the first, second, and fifth coils in the first and second quadrants, respectively, mirror asymmetries of the third, fourth, and sixth coils in the third and fourth quadrants, respectively; The magnet according to claim 3.

5. the fifth interval and the seventh interval are equal, the sixth interval and the eighth interval are equal, and the fifth interval and the seventh interval are smaller than the sixth interval and the eighth interval; the fifth spacing and the seventh spacing are at the inner bend radius of the assembly, and the sixth spacing and the eighth spacing are at the outer bend radius of the assembly. The magnet according to claim 4.

6. the set of coils includes a seventh coil and an eighth coil, the seventh coil and the eighth coil for conducting current to create a magnetic field, the seventh coil being disposed on the support structure in the first hemisphere and the eighth coil being disposed on the support structure in the second hemisphere; a ninth spacing between the seventh coil and an adjacent one of the first, second, or fifth coils in the first quadrant is different from a tenth spacing between the seventh coil and an adjacent one of the first, second, or fifth coils in the second quadrant; an eleventh spacing between the eighth coil and an adjacent one of the third, fourth, or sixth coils in the third quadrant is different from a twelfth spacing between the eighth coil and an adjacent one of the third, fourth, or sixth coils in the fourth quadrant; asymmetries of the first, second, fifth, and seventh coils in the first and second quadrants, respectively, mirror asymmetries of the third, fourth, sixth, and eighth coils in the third and fourth quadrants, respectively; The magnet according to claim 5.

7. the ninth interval and the eleventh interval are equal, the tenth interval and the twelfth interval are equal, and the ninth interval and the eleventh interval are smaller than the tenth interval and the twelfth interval; the ninth spacing and the eleventh spacing are at the inner bend radius of the assembly, and the tenth spacing and the twelfth spacing are at the outer bend radius of the assembly. The magnet according to claim 6.

8. the ferromagnetic yoke includes a notch adjacent the assembly, the notch being asymmetrical in the first quadrant and the second quadrant, the asymmetry of the notch relating to at least one of the size, shape, or placement of the notch; the asymmetry of the cutouts in the third and fourth quadrants mirrors the asymmetry of the cutouts in the first and second quadrants, respectively; The magnet according to claim 2.

9. the ferromagnetic yoke includes a notch adjacent the assembly, the notch being asymmetrical in the first quadrant and the second quadrant, the asymmetry of the notch relating to at least one of the size, shape, or placement of the notch; the asymmetry of the cutouts in the third and fourth quadrants mirrors the asymmetry of the cutouts in the first and second quadrants, respectively; The magnet according to claim 4.

10. the ferromagnetic yoke includes a notch adjacent the assembly, the notch being asymmetrical in the first quadrant and the second quadrant, the asymmetry of the notch relating to at least one of the size, shape, or placement of the notch; the asymmetry of the cutouts in the third and fourth quadrants mirrors the asymmetry of the cutouts in the first and second quadrants, respectively; The magnet according to claim 6.

11. the ferromagnetic yoke comprises iron; the support structure is non-ferromagnetic; The magnet of claim 1.

12. The magnet is bent at an angle of 60° or more relative to a line passing through the center of an unbent part of the magnet. The magnet of claim 1.

13. the magnet is bent at an angle of 70° or more relative to a line passing through the center of an unbent part of the magnet; The magnet of claim 1.

14. the magnet is bent at an angle of more than 80° relative to a line passing through the center of the unbent part of the magnet; The magnet of claim 1.

15. The magnet is bent by more than 90° relative to a line passing through the center of the unbent part of the magnet. The magnet of claim 1.

16. the magnet is bent at an angle of 120° or more relative to a line passing through the center of an unbent part of the magnet; The magnet of claim 1.

17. the magnet is a cosine theta magnet, in which the current through the set of coils is more concentrated near the 0° or 180° location of the magnet than near the 90° or -90 / 270° location of the magnet; The magnet of claim 1.

18. The magnet of claim 1 , wherein the set of coils is configured for dipole functionality.

19. The magnet of claim 1 , wherein the set of coils is configured for quadrupole functionality.

20. The magnet of claim 1 , wherein the set of coils is configured for hexapole functionality.

21. The magnet of claim 1 , wherein the set of coils comprises a superconducting material.

22. The magnet of claim 1 , further comprising one or more magnetic shims movable relative to the ferromagnetic yoke to vary the magnetic field produced by the magnet.

23. 10. The magnet of claim 1, wherein the coil sets include two or more coil sets configured asymmetrically with respect to a first dimension and symmetrically with respect to a second dimension, the first dimension being perpendicular to the second dimension.

24. 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 bending magnet for bending the particle beam along a length of the beamline structure; At least one of the bending magnets comprises the magnet of claim 1. system.

25. 25. The system of claim 24, further comprising an energy degrader that 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.

26. 25. The system of claim 24, 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.

27. 25. The system of claim 24, wherein the at least one bending magnet comprises a magnet having a magnetic field of 2.5 Tesla (T) or greater.

28. 25. The system of claim 24, wherein the at least one bending magnet comprises a magnet having a magnetic field of 3 Tesla (T) or greater.

29. 25. The system of claim 24, further comprising a collimator downstream of the gantry relative to the particle accelerator, the collimator for blocking at least a portion of the particle beam before the at least a portion of the particle beam reaches the irradiation target.

30. 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.

25. The system of claim 24.

31. 31. The system of claim 30, wherein the dimension is a diameter of the support structure.

32. 25. The system of claim 24, wherein the beamline structure has a length of 6 meters (m) or less.

33. 25. The system of claim 24, wherein the beamline structure has a length of 5 meters (m) or less.

34. 25. The system of claim 24, wherein the energy of the particle beam does not vary by more than 1% within the beamline structure.

35. 25. The system of claim 24, wherein the distance between the output of the beamline structure and an isocenter containing the irradiation target is 1.5 meters (m) or less.

36. 25. The system of claim 24, wherein the beamline structure includes an output channel including at least some of the bending magnets, the at least some bending magnets including magnetic dipoles arranged in series to bend the particle beam by at least 90 degrees, the magnetic dipole including the at least one bending magnet.

37. 25. The system of claim 24, wherein the beamline structure includes an output channel that includes at least some of the bending magnets, and wherein the at least one bending magnet precedes the output channel in a direction of travel of the particle beam.

38. 25. The system of claim 24, wherein the gantry is achromatic from a point of entry of the particle beam into the gantry to an isocenter of the system.

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