Bending magnet

Abstract

An example magnet includes an assembly. The assembly includes (i) a coil assembly for conducting an electrical current to generate a magnetic field, and (ii) a support structure on which the coil assembly is asymmetrically disposed, and a ferromagnetic yoke surrounding a portion of the assembly. The ferromagnetic yoke and assembly are curved.

Description

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS

[0002] This application claims priority to U.S. Provisional Application No. 63 / 394,461, filed on August 2, 2022. The contents of U.S. Provisional Application No. 63 / 394,461 are incorporated herein by reference. Technical Field

[0003] This specification describes examples of bending magnets (eg, cosine-theta magnets) for use in a gantry of a particle therapy system. Background Art

[0004] Particle therapy systems use particle accelerators to generate particle beams for treating diseases, such as tumors. Particle therapy systems can use a gantry to direct the particle beam toward the patient from multiple angles. In some examples, the gantry includes a device that supports the radiation delivery device during treatment. The gantry includes magnetic elements to guide the particle beam to its destination. Summary of the invention

[0005] An example magnet includes an assembly. The assembly includes: (i) a coil assembly for conducting an electric current to generate a magnetic field, and (ii) a support structure on which the coil assembly is asymmetrically disposed, and a ferromagnetic yoke surrounding a portion of the assembly. The ferromagnetic yoke and the assembly are curved. As a result, the magnet may be curved. The example magnet may include one or more of the following features, alone or in combination.

[0006] The coil group may include a first coil and a second coil. The first coil and the second coil may be used to conduct current to generate a magnetic field. The first coil and the second coil may be asymmetrically arranged on a support structure in a first hemisphere of the magnet, so that a first spacing between the first coil and the second coil in a first quadrant of the magnet may be 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 may be within the first hemisphere. The coil group may include a third coil and a fourth coil. The third coil and the fourth coil may be used to conduct current to generate a magnetic field. The third coil and the fourth coil may be asymmetrically arranged on a support structure in a second hemisphere of the magnet, so 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 may be within the second hemisphere. The asymmetry of the first and second coils in the first and second quadrants, respectively, may reflect the asymmetry of the third and fourth coils in the third and fourth quadrants, respectively.

[0007] The first and third intervals may be equal. The second and fourth intervals may be equal. The first and third intervals may be smaller than the second and fourth intervals. The first and third intervals may be at an inner bending radius of the assembly, and the second and fourth intervals may be at an outer bending radius of the assembly. In the case of an embodiment comprising the first to fourth coils, the ferromagnetic yoke may include a notch adjacent to the assembly. The notch may be asymmetric in the first quadrant and the second quadrant. The asymmetry of the notch may be 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 coil group may include a fifth coil and a sixth coil. The fifth coil and the sixth coil may be used to conduct current to generate a magnetic field. The fifth coil may be disposed on a support structure in the first hemisphere. The sixth coil may be disposed on a support structure in the second hemisphere. A fifth interval between the fifth coil and an adjacent coil in the first coil or the second coil in the first quadrant may be different from a sixth interval between the fifth coil and an adjacent coil in the first or second coil in the second quadrant. A seventh interval between the sixth coil and an adjacent coil in the third or fourth coil in the third quadrant may be different from an eighth interval between the sixth coil and an adjacent coil in the third or fourth coil in the fourth quadrant. The asymmetry of the first, second, and fifth coils in the first and second quadrants may reflect the asymmetry of the third, fourth, and sixth coils in the third and fourth quadrants, respectively.

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

[0010] The coil group may include a seventh coil and an eighth coil. The seventh coil and the eighth coil may be used to conduct current to generate 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. The ninth interval between the seventh coil and the first coil, the second coil, or the fifth coil in the first quadrant may be different from the tenth interval between the seventh coil and the first coil, the second coil, or the fifth coil in the second quadrant. The eleventh interval between the eighth coil and the third coil, the fourth coil, or the sixth coil in the third quadrant may be different from the twelfth interval between the eighth coil and the third coil, the fourth coil, or the sixth coil in the fourth quadrant. The asymmetry of the first coil, the second coil, the fifth coil, and the seventh coil in the first and second quadrants may reflect the asymmetry of the third coil, the fourth coil, the sixth coil, and the eighth coil in the third and fourth quadrants, respectively.

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

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

[0013] The magnet may be or include a cosine theta magnet, wherein the current through the coil assembly has a greater concentration near the 0° or 180° position of the magnet than near the 90° or -90 / 270° position of the magnet. The coil assembly may be configured for a two-pole function. The coil assembly may be configured for a four-pole function. The coil assembly may be configured for a six-pole function. The coil assembly may include a superconducting material that makes the magnet superconducting. The magnet may include one or more magnetic shims that are movable relative to the ferromagnetic yoke to change the magnetic field generated 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 to an irradiation target. The beamline structure may include bending magnets to bend 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 an assembly including: (i) a coil assembly for conducting an electric current to generate a magnetic field, and (ii) a support structure on which the coil assembly is asymmetrically disposed, and a ferromagnetic yoke surrounding a portion of the assembly, wherein the ferromagnetic yoke and the assembly are bent, and wherein the magnet includes one or more of the aforementioned features described above.

[0015] The system may include an energy degrader, which is the only component that actively controls the energy variation of the particle beam after the particle accelerator outputs the particle beam and before the particle beam reaches the irradiation target. The beam line structure may be configured to not 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.

[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 used to block at least a portion of the particle beam before 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 beam line structure in a circular path around the irradiation target. The support structure may have a size of 6 meters or less. The size may be a diameter of the support structure. The length of the beam line structure may be 6 meters (m) or less. The length of the beam line structure may be 5 meters (m) or less. The energy of the particle beam may vary within the beam line structure by no more than 1%. The distance between the output end of the beam line structure and the isocenter containing the irradiation target may be 1.5 meters (m) or less.

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

[0018] The gantry may be an achromatic lens from the entry point of the particle beam into the gantry to the isocenter of the system treating the patient.

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

[0020] The control of the various systems or parts thereof described herein can be implemented via a computer program product, which includes instructions stored on one or more non-transitory machine-readable storage media and executable on one or more processing devices (e.g., microprocessors, application-specific integrated circuits, programming logic such as field programmable gate arrays, etc.). The systems or parts thereof described herein can be implemented as an apparatus, method, or medical system, which can include one or more processing devices and computer memory to store executable instructions to implement control of the functions described. The apparatus (e.g., magnet), system, and / or component described herein can be configured, for example, by design, construction, composition, arrangement, placement, 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 DESCRIPTION OF THE DRAWINGS

[0022] Figure 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.

[0023] Figure 2 yes Figure 1 A cut-away side view of components of a particle therapy system, including an exemplary gantry, is shown.

[0024] Figure 3 is a cross-sectional close-up side view of components included in a treatment head and an example beamline structure, which may be Figure 1 A portion of a rack is shown.

[0025] Figure 4 is a cross-sectional close-up side view of components included in a treatment head and an example beamline structure, which may be Figure 1 A portion of the rack is shown in FIG.

[0026] Figure 5 is an illustration of a front view of an example scanning magnet configured to scan a particle beam in two orthogonal dimensions.

[0027] Figure 6 is a cross-sectional close-up side view of components included in a treatment head and an example beamline structure, which may be Figure 1 A portion of the rack is shown in FIG.

[0028] Figure 7 is an illustration of a front view of an example scanning magnet configured to scan a particle beam in a single dimension.

[0029] Figure 8is an illustration of a front view of an example scanning magnet configured to scan a particle beam in a single dimension.

[0030] Fig. 9 is a cross-sectional close-up side view of components included in a treatment head and an example beamline structure, which may be Figure 1 A portion of the rack is shown in FIG.

[0031] Fig.10 is a cross-sectional close-up side view of components included in a treatment head and an example beamline structure, which may be Figure 1 A portion of the rack is shown in FIG.

[0032] Fig.11 is an illustration of a front view of an example superconducting scanning magnet configured to scan a particle beam in two orthogonal dimensions.

[0033] 12a is an illustration of a front view of an example superconducting scanning magnet configured to scan a particle beam in a single dimension; FIG. 12b is an illustration of a front view of an example superconducting scanning magnet configured to scan a particle beam in a single dimension orthogonal to the dimension of FIG. 12a.

[0034] Fig.13 is a diagram showing a perspective view of an exemplary configurable collimator that may be part of the particle therapy system of claim 1.

[0035] Fig.14 It is shown Fig.13 FIG. 1 is a diagram of a front view of a configurable collimator.

[0036] Fig.15 It is shown Fig.13 and Fig.14 Figure 2 is a perspective, partially transparent view of a configurable collimator.

[0037] Fig.16 is configured to accommodate Figure 1 A block diagram of an example treatment space of all or a portion of a particle therapy system.

[0038] Fig.17 is a graph illustrating exemplary horizontal (x) and vertical (y) particle beam envelopes produced in an exemplary gantry described herein.

[0039] Fig.18 is a graph illustrating an exemplary achromatic lattice design for a beamline of an exemplary gantry described herein.

[0040] Fig.19 is a graph showing the results produced by scanning a particle beam in the horizontal (x) and vertical (y) planes using an example gantry described herein.

[0041] Fig. 20 is a cross-sectional side view of components in an exemplary particle accelerator that may be used with the particle therapy systems described herein.

[0042] Fig.21 is a perspective view of an exemplary particle therapy system.

[0043] Fig. 22 is a perspective view of an exemplary energy degrader.

[0044] Fig.23 is a front cross-sectional view of an exemplary superconducting magnet that may be used as a scanning magnet in the particle therapy system of claim 1.

[0045] Fig.24 is a cross-sectional view of an example superconducting coil that may be used in any of the superconducting magnets described herein.

[0046] Fig.25 is a perspective view of a coil of an exemplary bending magnet.

[0047] Fig.26 is a cross-sectional view of an exemplary bending magnet having three coils in each hemisphere.

[0048] Fig. 27 is a perspective cross-sectional view of an exemplary bending magnet showing magnetic field strengths with shading, where darker shading indicates greater magnetic field strengths.

[0049] Fig.28 is a perspective view of components of an exemplary bending magnet, portions of which are shown as transparent.

[0050] Fig.29 is a perspective view of components of an exemplary bending magnet.

[0051] Fig.30 is a cross-sectional view of a portion of an exemplary bending magnet having two coils in each hemisphere.

[0052] Fig.31 is a cross-sectional view of a portion of an exemplary bending magnet having four coils in each hemisphere.

[0053] The same reference numbers in different drawings denote the same elements. DETAILED DESCRIPTION

[0054] An example particle therapy system that can accommodate a patient and an accelerator in the same space is described herein. The example system includes a particle accelerator, which can be, but is not limited to, a synchrocyclotron accelerator with low radiation leakage and small enough to fit in a standard linear accelerator (LINAC) room. The system also includes a medical rack configured to deliver a charged particle beam output from the accelerator, such as protons or ions, to treat a tumor or other condition in a patient. The rack includes a beam line structure to guide the particle beam from the accelerator to a treatment position and to deliver the particle beam to the treatment position. The beam line structure includes magnetic elements, such as one or more dipole magnets and one or more quadrupole magnets, to guide the particle beam to the treatment position. In order to be able to deliver a particle beam in the same space for treatment, particularly in a relatively small space such as a standard LINAC room, at least some of the magnetic elements in the beam line structure are configured to bend the particle beam at an angle close to or exceeding a right angle. In the example, the magnetic component is configured and arranged to bend the particle beam by 70° or more, including a 90° angle and an obtuse angle greater than 90°.

[0055] The magnetic components in the frame may include one or more magnets, such as a cosine theta bending magnet, which has the following features that enable the particle beam to be magnetically bent: a current conducting coil (or simply "coil") group including at least a first coil and a second coil, wherein the first coil and the second coil are used to conduct current to generate a magnetic field, and a non-ferromagnetic support structure ("support"), and the coil group is asymmetrically arranged on the non-ferromagnetic support structure. The assembly including the coil and the support at least partially surrounds the air core, and the particle beam passes through the air core. The ferromagnetic block or yoke at least partially surrounds the assembly and the air core, but the ferromagnetic block or yoke is solid in other aspects except for a notch in the yoke adjacent to the outer portion of the assembly. The notch defines a channel through the yoke, which can be filled with air or a vacuum. The notch affects the amount of ferromagnetic material adjacent to the coil, and therefore affects the shape of the magnetic field generated by the coil. The cosine theta bending magnet is curved or bent. The configuration of the magnets (e.g., asymmetric coil windings and asymmetric grooves in the yoke) enables the magnets to reduce distortion of the particle beam as it passes through the gantry, particularly at bends in the gantry. For example, the cross-sectional (e.g., circular) shape of the particle beam can remain substantially or completely circular as it passes through the gantry.

[0056] Fig.25 A perspective view of an exemplary coil 180 for an exemplary cosine-theta bending magnet ("magnet") that may be used as a bending magnet in an exemplary particle therapy system gantry such as those described herein is shown. However, the magnet is not limited to use in this environment. Fig.26A cross-sectional front view of an example magnet 200 is shown, which includes a coil 180, a non-ferromagnetic support 205, an air core 213, and a yoke 181 composed of iron or other ferromagnetic material.

[0057] The magnet 200 can be a dipole magnet, a quadrupole magnet, or a sextupole magnet. A dipole magnet has two poles, one north pole and one south pole. Its magnetic field lines form a closed loop, exiting from the north pole, re-entering at the south pole, and then passing through the magnet. A quadrupole magnet includes a set of four poles arranged so that in a planar multipole expansion of the magnetic field, the dipole terms cancel and the least significant term in the field equation is the quadrupole. A sextupole magnet includes six poles arranged in alternating north and south pole arrangements around an axis. The coils described herein can be layered to generate higher order field harmonics.

[0058] refer to Fig.25 and Fig.26 , the magnet 200 is an electromagnet, which includes a plurality of conductive coils 180 in each of its hemispheres 200a and 200b, wherein Fig.25 and Fig.26 , 0° to 180° corresponds to the upper hemisphere 200a, and 180° to 360° (0° again) corresponds to the lower hemisphere. Fig.25 and Fig.26 In the example of the magnet 200, the coil 180 includes three coils 201a, 201b, 201c in the upper hemisphere 200a and three coils 201d, 201e, 201f in the lower hemisphere 200b (the coil 201f is Fig.25 In some embodiments, magnet 200 may include fewer than three coils or more than three coils in each of its hemispheres. For example, a bending magnet of the type described herein may include two coils in each hemisphere (e.g., Fig.30 ), three coils in each hemisphere (e.g. Figure 25 to Figure 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 may be included in each hemisphere of the bending magnet.

[0059] refer to Fig.26, in magnet 200, lower hemisphere 200b is a mirror image of upper hemisphere 200a, which means that the structure and relative spacing of the coils in each hemisphere are the same. For example, coils 201a and 201d have the same structure and spacing relative to coils 201b and 201e, respectively. Coils 201b and 201e have the same structure and spacing relative to coils 201c and 201f, respectively. Coils 201c and 201f have the same structure and spacing relative to 90 ° and -90 ° (270 °) positions on the magnet, respectively. Therefore, the description of the upper hemisphere herein is applicable to the lower hemisphere, and vice versa.

[0060] The bending magnet coils (e.g., coil 180) can be superconducting or non-superconducting. For example, one or more or all of the coils can be made of 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 of 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 Configuration described. Due to compactness and space constraints, the ends of the coil have been configured to meet a predetermined critical strain / stress limit on the coil mandrel and a predetermined total field integral (in a non-limiting example, 3.54 Tesla-meters).

[0061] The coil is disposed (eg, wound, held, placed, arranged, or maintained) on a non-ferromagnetic support 205 ( Fig.26 ,exist Fig.25 , 28 29). The support 205 may be made of or include a non-ferromagnetic material, such as aluminum or stainless steel. The support 205 may be a single continuous or integrated structure, or the support 205 may include a plurality of separate structures that together form a support structure. The support structure 205 may have a shape that is complementary to the shape of the coil. For example, Fig.26 As shown, the support member 205 has a shape complementary to the shape of the coils 201a, 201b, 201c, 201d, 201e and 201f. The support member 205 and the coils 201a, 201b, 201c, 201d, 201e and 201f together define a substantially circular cross-section, as shown in FIG. Fig.26 As shown, it defines a space including a hollow core 213. The hollow core 213 may contain a gas such as air or a rare gas, or it may be close to a vacuum, such as 10 -5 Torr (0.0013332 Pascal) or less.

[0062] The current flowing through coils 201a, 201b, 201c, 201d, 201e and 201f generates a magnetic field that is at least partially shaped by the ferromagnetic yoke 181. Of course, the magnitude of the current also determines the shape of the magnetic field. The yoke 181 can be a solid structure made of a ferromagnetic material such as iron, such as Fig.28 and Fig.29 As shown. Fig.28 and 29 As shown, the yoke 181 may be formed of a top piece 181a and a bottom piece 181b; however, in other embodiments, the yoke 181 may be formed of a left piece and a right piece or more than two pieces. Fig.28 , portion 181c of the yoke 181 is shown as transparent, while in Fig.29 , a portion 181a of the yoke 181 is depicted in a solid form. The transparent form is for illustrating the coil 180 passing through the yoke, and does not mean that the entire or part of the yoke 181 is actually transparent.

[0063] The yoke 181 at least partially surrounds the assembly consisting of the support 205, the coils 201a, 201b, 201c, 201d, 201e, 201f and the core 205. Fig.28 and Fig.29 In the exemplary embodiment of FIG. 1 , the yoke 181 is shown in a cutaway form to show the coil 180. However, in Fig.28 and Fig.29 In the exemplary embodiment of the present invention, the yoke 181 extends to cover the entire length of the coil 180, from the end 184 (at Fig.28 Best visible in) to 185 ( Fig.29 , Fig.28 The ends 186a, 186b of the coil 180 ( Fig.28 ) remains exposed to enable magnet 200 to be connected to other magnets in the beam line and to allow the particle beam to pass through magnet 200 in the manner described herein.

[0064] In some embodiments, the yoke 181 includes a circular notch or channel adjacent to and surrounding the outer surface of the coil 181. The notches can extend 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-section can vary along the length of the yoke 181 / magnet 200. For example, Fig.26 and Fig.29 As shown ( Fig.28201c, 201d, 201e, 201f, and core 213). The lower hemisphere 200b is a mirror image of the upper hemisphere 200a, so that the notch on the lower hemisphere 200b is a mirror image of the notch on the upper hemisphere 200a. In contrast, the notch 220 on the right hemisphere 188a and the left hemisphere 188b is asymmetric to account for the curvature in the magnet 200. In the example, the asymmetry is that the notch has a larger volume / size on average in the right hemisphere 188a than in the left hemisphere 188b, resulting in more ferromagnetic material in the left hemisphere 188b than in the right hemisphere 188a. In an example, the asymmetry is that the notches in the right hemisphere 188a are shaped differently than the notches in the left hemisphere 188b, resulting in more ferromagnetic material being present in the left hemisphere 188b than in the right hemisphere 188a. In an example, the asymmetry is that the notches are, on average, closer together 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 an example, the asymmetry is that the notches are, on average, closer together and, on average, 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 In the embodiment, the notches 220a and 220b in the right hemisphere 188a are larger and closer together (in fact, they overlap) than the corresponding notches 220c and 220d in the left hemisphere 188b. In the example, asymmetry is 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 caused by notch construction, placement, quantity, 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 realize the transmission of particle beams and maintain the integrity of particle beams.

[0065] refer to Fig.25 and Fig.26 , in magnet 200, there are more conductive coils closer to the 0° / 180° position than at 90°. Therefore, when the magnet is operational, i.e., when current is conducted through the coils, the current density is greater closer to 0° / 180° than at 90°. In the example presented, there is no conductor at 90°; therefore, the current density at 90° is zero. As noted, in this example, the lower hemisphere 200b is a mirror image of the upper hemisphere 200a. Therefore, in the lower hemisphere 200b, when current is conducted through the coils, the current density closer to 0° / 180° is greater than the current density at -90° / 270°.

[0066] Another feature of the magnet 200 is that the coils 201a, 201d closer to 0° / 180° have a greater current carrying capacity than the coils 201c closer to 90° and the coils 201f closer to -90° / 270°. For example, the coils 201a, 201d each have a larger cross-sectional area than the coils 201c, 201f, respectively. Generally, in embodiments of the magnet 200, in each quadrant 210a, 210b, 210c, and 210d ( Fig.26 ) the current carrying capacity of the coil decreases from 0° / 180° to 90° and from 0° / 180° to -90° / 270°.

[0067] In this regard, 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, located between coils 201a and 201c, has a cross-sectional area that is smaller than the cross-sectional area of ​​coil 201a and larger than the cross-sectional area of ​​coil 201c. Similarly, coil 201e, located between coils 201d and 201f, has a cross-sectional area that is smaller than the cross-sectional area of ​​coil 201d and larger than the cross-sectional area 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 variable. In a constant example, for hemisphere 200a, coil 201c may have a 20% less ampacity than coil 201b; and for hemisphere 200a, coil 201b may have a 20% less ampacity than coil 201a. In a varying example, coil 201c may have a 20% less ampacity than coil 201b; and coil 201b may have a 10% less ampacity than coil 201a. The same difference in ampacity may hold true for the coil counterparts in hemisphere 200b. The 10% and 20% numbers are non-limiting examples; and the reduction in ampacity from 0° / 180° to 90° (or 0° / 180° to -90° / 270°) in the coils of the cosine-theta magnets herein may be greater or less than these numbers.

[0068] like Fig.25 and 27 As shown in Figures 29, the cosine-theta magnet 200 is curved. This feature is particularly useful for guiding a particle beam in a compact gantry of the type described herein. The cosine-theta magnet can be curved and thus direct the particle beam relative to the straight line 211 ( Fig.25 and Fig.28) is bent 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 180°, wherein the straight line 211 passes through and follows the trajectory of the unbent portion of the magnet 200. The magnet 200 may be bent at any suitable angle. The bend may be any degree in the range of 0° to 90°, any degree in the range of 0° to 180°, or any degree in the range of 70° to 180° relative to the straight line 211.

[0069] The rectangular area 212 ( Fig.26 , Fig. 27 ) (where the particle beam is constrained to travel by the magnetic field) may be distorted due to the bending of the cosine theta magnet. More specifically, the bending of the magnet may cause the cross-section (i.e., spot) of the particle beam to become elliptical rather than remaining circular. For example, the spot size (e.g., cross-sectional area) of the particle beam may have an aspect ratio of 5% or greater and may grow from 3 millimeters (mm) or 4 mm sigma to 10 mm sigma in one or more planes due to the bending. However, the magnet 200 is configured to offset this distortion and maintain the cross-section of the particle beam in region 212 in a predetermined shape, such as a circle. In order to at least partially offset this distortion, the coil groups in the magnet 200 are asymmetrically disposed on the supports 205 in each hemisphere 200a, 200b to shape the magnetic field to prevent or reduce distortion. For example, as Fig.26 As shown, in hemisphere 200a, the spacing 214 between coils 201a and 201b in quadrant 210a is different from the spacing 215 between coils 201a and 201b in quadrant 210b. Similarly, in the same hemisphere 200a, the spacing 217 between coils 210b and 210c in quadrant 210a is different from the spacing 218 between coils 210b and 210c in quadrant 210b. The same spacing difference between coils in hemisphere 200 exists in the mirror image coils in hemisphere 200b. That is, the difference in spacing between coils 201d, 201e and 201e, 201f is the same as the difference in spacing between coils 201a, 201b and 201b, 201c, respectively. In some embodiments, the difference in coil spacing in different quadrants can be different in different hemispheres. In this example, the spacing between coil pairs in quadrant 210a is greater than the spacing between the same coil pairs in quadrant 201b; however, in other embodiments, the spacing between coil pairs in quadrant 210s is less than the spacing between the same coil pairs in quadrant 201b.

[0070] like Figures 25 to 29 As shown, in magnet 200, the curvature or bend in magnet 200 produces inner surface 206 and outer surface 207, wherein the inner surface has a smaller radius of curvature or bending radius than the outer surface. The inner surface and the outer surface can refer to coils, components or magnets because the bending radius is the same or substantially the same. In this example, coils 201a, 201b, 201c, 201d, 201e and 201f are spaced apart so that the spacing between the coils is smaller on the outer surface 207 than on the inner surface 206. For example, at the outer surface 207 and quadrant 210b of magnet 200, coils 201a, 201b and 201c are closer together, they are at the inner surface 206 and in the quadrant 210a of the magnet. In this configuration, for example, spacing 215 is smaller than spacing 214, and spacing 218 is smaller than spacing 217. The spacing difference can be on the order of 1%, 2%, 3%, 4%, 5%, 6%, 7%, 8%, 9% or more. For example, interval 215 may be 1%, 2%, 3%, 4%, 5%, 6%, 7%, 8%, 9% or more smaller than interval 214; and interval 218 may be 1%, 2%, 3%, 4%, 5%, 6%, 7%, 8%, 9% or more smaller than interval 217. Any suitable interval may be used to achieve the effects described herein.

[0071] These spacing differences can be reflected in the coils located in the lower hemisphere 200b. More specifically, since coils 201d, 201e, and 201f are arranged in a mirror configuration of coils 201a, 201b, and 201c, coils 201d, 201e, and 201f will have the same spacing variations as coils 201a, 201b, and 201c, except that they will be in quadrant 201d with the same coil spacing as quadrant 201b, and in quadrant 201c with the same coil spacing as quadrant 201a.

[0072] The combination of the asymmetry of coils 180 in the right hemisphere 188a and the left hemisphere 188b, the greater current carrying capacity of coils closer to 0° (e.g., 201a, 201d) than coils closer to 90° (e.g., 201c, 201f), and the asymmetry of notches 220 in the right hemisphere 188a and the left hemisphere 188b, forms the magnetic field of magnet 220 at region 212 to (i) maintain the particle beam within region 212, and (ii) maintain the particle beam as circular or substantially circular as it travels through the length of magnet 200, including through the curved portion of magnet 200. In some examples, substantially circular can include deviations from a perfect circle 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 size 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.

[0073] Fig. 27 Shown in grayscale or color Fig.25 , Fig.26 , Fig.28 and Fig.29 206 of the upper hemisphere 200a of the example embodiment of the magnet 200. As shown, the magnetic field intensity on the inner surface 206 of the magnet 200 is greater than the magnetic field intensity on the outer surface 207 of the magnet. As a result of the Biot-Savart law solution of the curved conductor geometry, the magnetic field is larger at the inner radius. In this regard, if the magnetic field has the same intensity on all radii (e.g., from the inner surface 206 to the outer surface 207), there will be a net focusing of the particle beam in the curved plane for different particle trajectories through the magnet. This can make the particle beam over-focused, and can be compensated by reducing the magnetic field intensity on the outer radius (207) relative to the magnetic field intensity on the inner radius (206). As a result of this difference in the magnetic field, the distortion of the particle beam spot (i.e., the cross section of the particle beam) can be reduced or eliminated. As a result, when the particle beam travels through the entire length of the magnet 200, the particle beam can remain substantially circular.

[0074] In a non-limiting example, magnet 200 achieves 0.1% uniformity of magnetic (eg, dipole) field in region 212, which may be 100 millimeters (mm) by 90 mm along the beam trajectory throughout the length of the magnet.

[0075] In some embodiments, one or more magnetic shims (not shown) can be used to change the amount of ferromagnetic material in the yoke 181, and thus change the amount of ferromagnetic material in the magnet. For example, the shim can be a rod, cone, or other structure that can be controlled to move in or out of the yoke 181 to adjust the amount of ferromagnetic material in the yoke 181, thereby changing the shape of the magnetic field generated by the magnet. In some embodiments, the shim is manually controllable. In some embodiments, the shim is computer controlled. For example, each shim can be connected to a computer-controlled actuator that controls the shim to move in or out of the ferromagnetic core. The shim can be moved by manual or computer control to be completely embedded in the yoke, or completely leave the yoke. There can be one shim per magnet quadrant, for example, two, three, four, five, etc. per quadrant. In an example, one or more magnetic field sensors can detect the magnetic field generated by the magnet 200, and the shims can be controlled to change the magnetic field to a magnetic field with a target shape. In an example, one or more sensors can detect the position of the particle beam in the core 213, and the shims can be controlled to change the magnetic field to control the particle beam placement.

[0076] Fig.30 It is shown that a magnet similar to 200 (see for example Fig.26 ) is an example assembly 230 for use in a magnet of a particle beam. Assembly 230 includes a support 231 that has the same function as support 205 and can have a similar structure and composition as support 205. Assembly 230 also includes coils 232a, 232b, 232c, and 232d. The coil groups are asymmetrically arranged on support 231 in each hemisphere 235a, 235b to shape the magnetic field so as to at least partially prevent or reduce particle beam distortion, as described herein. For example, Fig.30 As shown, in hemisphere 235a, the spacing between coils 232a and 232b in quadrant 236a at the inner radius of curvature 300 of the magnet is different from (e.g., greater than) the spacing between coils 232a and 232b in quadrant 236b at the outer radius of curvature 301 of the magnet. The same spacing difference between the 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 from (e.g., greater than) the spacing between coils 232c and 232d in quadrant 236db. Component 230 and Fig.26 The difference between the components of FIG. 230 is that the component 230 includes two sets of coils instead of three sets of coils. Otherwise, the component 230 can be combined with, for example, Fig.28 and 29 The magnet structure shown has all the accompanying features configured for a dual coil rather than a triple coil design. These features include, but are not limited to, asymmetric notches and a yoke configured for a dual coil design.

[0077] Fig.31 It is shown that a magnet similar to 200 (see for example Fig.26 ) is another example assembly 240 for use in a magnet of a particle beam. Assembly 240 includes a support 241 that 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 coil groups are asymmetrically arranged on support 241 in each hemisphere 245a, 245b to shape the magnetic field so as to at least partially prevent or reduce particle beam distortion, as described herein. For example, as Fig.31 As shown, in hemisphere 245a, the spacing between coils 242a and 242b in quadrant 246a at the inner radius of curvature 303 of the magnet is different from (e.g., greater than) the spacing between coils 242a and 242b in quadrant 246b at the outer radius of curvature 304 of the magnet; the spacing between coils 242b and 242c in quadrant 246a is different from (e.g., greater than) the spacing between coils 242a and 242b in quadrant 246b; and the spacing between coils 242c and 242d in quadrant 246a is different from (e.g., greater than) the spacing between coils 242c and 242d in quadrant 246b. The same spacing differences between the coils in hemisphere 245a exist 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 inner radius of curvature of the magnet is different from (e.g., greater than) the spacing between coils 242e and 242f in quadrant 246d at the outer radius of curvature of the magnet; the spacing between coils 242f and 242g in quadrant 246c is different from (e.g., greater than) the spacing between coils 242f and 242g in quadrant 246d; and the spacing between coils 242g and 242h in quadrant 246c is different from (e.g., greater than) the spacing between coils 242g and 242h in quadrant 246d. Component 240 and Fig.26 The difference between the components of FIG. 240 is that the component 240 includes four coils instead of three coils. Otherwise, the component 240 can be combined with, for example, Fig.28 and 29 The magnet structure shown has all the additional features configured for a four-coil design instead of a three-coil design. These features include, but are not limited to, asymmetric notches and a yoke configured for a four-coil design.

[0078] In the following description of the particle beam gantry, Figures 25 to 31 A cosine-theta bending magnet of the type described or any variant thereof may be used, for example, to achieve Figure 1 , Figure 2 , Figure 3 , Figure 4 , Figure 6 , Fig. 9 , Fig.10 , Fig.16 and Fig.21 Any or all of the bending magnets described herein and variations thereof. For example, cosine-theta magnets of the type described herein having a two-pole, four-pole, and / or six-pole configuration may be used to implement a bending magnet in any particle beam gantry described herein or variations thereof.

[0079] Figure 1 An example implementation of a particle therapy system 10 of the type described above is shown, which may include Figures 25 to 31 One or more bending magnets of the type described or variations thereof. The particle therapy system 10 includes a particle accelerator 12, an example of which is described herein. In this example, the particle accelerator 12 is a synchrocyclotron having a superconducting electromagnetic configuration that produces a maximum magnetic field strength of more than 2.5 Tesla (T) or more than 3 T. In this regard, a superconductor is an element or metal alloy such as niobium-tin (Nb3Sn) that loses most, if not all, of its electrical resistance when cooled below a threshold temperature. As a result, current flows through the superconductor substantially unimpeded. Therefore, a superconducting coil is able to conduct a greater current in its superconducting state than an ordinary wire of the same size. Due to the large amount of current that a superconducting coil is able to conduct, a superconducting coil is particularly useful in particle therapy applications.

[0080] An exemplary synchrocyclotron is configured to output protons or ions as a monoenergetic particle beam having an energy level of 150 megaelectronvolts (MeV) or higher. An exemplary synchrocyclotron has a 4.5 cubic meter (m 3 ) or less and a weight of 30 tons (T) or less. Due to its size, this type of particle accelerator is referred to as "compact". However, as described herein, a synchrocyclotron or other types of particle accelerators having weights, sizes, magnetic fields, and / or energy levels different from these may be used in the particle therapy system 10.

[0081] The particle therapy system 10 also includes a gantry 14. The gantry 14 includes an annular 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 mounted to the support structure 15 and a channel 18 that guides 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 relative to a treatment position 19. The treatment position can include a system isocenter, where a patient can be positioned for treatment. In an example, the motor can move the output channel 17 along a track on the structure 15, causing the output channel 17 to rotate relative to the treatment position 19. In an example, the structure to which the output channel 17 is attached can be rotated relative to the treatment position 19, thereby causing the output channel 17 to rotate relative to the treatment position. In some embodiments, the rotation achieved by the gantry 14 allows the output channel 17 to be positioned at any angle relative to the treatment position. For example, the output channel 17 may be rotated 360°, and thus, the output channel 17 may be positioned at 0°, 90°, 270°, and back to 0° / 360° or any angle within these rotational positions.

[0082] As previously described, the beam line structure 16 is configured to direct the particle beam from the accelerator 12 to the treatment location 19. To this end, the output channel 17 includes magnetic elements to bend the particle beam toward the treatment location. As noted, the beam line structure 16 includes a channel 18 that contains magnetic elements along the beam line that direct the particle beam from the particle accelerator 12 to the output channel 17.

[0083] refer to Figure 2 and Figure 3, the channel 18 of the exemplary beam line structure 16 includes non-superconducting quadrupole magnets 21 and 22 and a bending magnet 23, which may be a superconducting dipole magnet. The quadrupole magnets 21 and 22 are configured to keep the particle beam focused and traveling in a straight or substantially straight line within the beam line structure 16, for example, deviating from a straight line by 5% or less. The quadrupole magnets 21 and 22 are configured to focus the particle beam to maintain a substantially consistent cross-sectional area of ​​the particle beam, for example, within a tolerance of ±5%. The bending magnet 23 is configured to bend the particle beam toward the output channel 17, as shown. The bending magnet 23 can be configured to bend the particle beam anywhere within a range of 20º to 80º relative to the horizontal line 24. Generally, a larger bending angle can reduce the distance between the particle accelerator 12 and the treatment position 19 or the system isocenter, thereby reducing the space required to accommodate the gantry and, therefore, reducing the size of the particle therapy system. For example, replacing the non-superconducting bending magnets 23 with superconducting bending magnets 23 that bend the particle beam by more than 80°, such as 90° or more, can further reduce the distance from the particle accelerator 12 to the support structure 15, thereby reducing the distance to the treatment position 19 and the isocenter. The bending magnets 23 can be or include Figures 25 to 31 As shown in and referenced Figures 25 to 31 Bending magnets of the type described or modifications thereof.

[0084] In some embodiments, higher order magnetic elements may be used in place of or in addition to any of the quadrupole magnets described herein. For example, the beamline structure may include one or more sextupole magnets in place of or in addition to the quadrupole magnets. The sextupole magnets may be configured to keep the particle beam focused and traveling in a straight or substantially straight line within the beamline structure 16, e.g., with a deviation from a straight line of 5% or less. The sextupole magnets may also be configured to maintain a consistent cross-sectional area of ​​the particle beam, e.g., within a tolerance of ±5%. In addition, the sextupole magnets may correct for the dispersion effects of the quadrupole magnets. Compared to the quadrupole magnets, the sextupole magnets are less sensitive to the dispersion effects of the particle beam from the axis of the ideal position defining the beam line (e.g., at Fig.26 Particles that are displaced farther (within region 212) have a greater focusing effect.

[0085] Reference again Figure 3 In this example, the channel 18 of the beamline structure 16 also includes two non-superconducting quadrupole magnets 26 and 27. The quadrupole magnets 26 and 27 are configured to keep the particle beam focused and traveling straight or substantially straight within the beamline structure 16, for example, deviating from a straight line by 5% or less. The quadrupole magnets 26 and 27 are configured to maintain a consistent cross-sectional area of ​​the particle beam, for example, within a tolerance of ±5%. As previously described, a high-order magnetic element can replace one or more quadrupole magnets to improve focusing.

[0086] The particle therapy system 10 also includes one or more scanning magnets 30 in the path of the particle beam and configured to move the particle beam through at least a portion of a beam field covering all or a portion (i.e., at least a portion) of the irradiation target. In some examples, for a given position of the compact gantry, the beam field includes a maximum (e.g., planar) range over which the particle beam can move in a plane parallel to the patient treatment area. Movement of the particle beam through the beam field results in movement through at least a portion of the irradiation target at the treatment position 19. The scanning magnet can be sized and configured to move the particle beam through a beam field having an area of ​​20 cm×20 cm or greater, but the system 10 is not limited to any particular beam field size. For example, the scanning magnet can have an aperture of 20 cm×20 cm or greater, but the scanning magnet is not limited to any particular aperture size.

[0087] The scanning magnets can be located at different locations within the particle therapy system. Figure 4 In the beam line configuration 16a shown (which is a variation of the beam line configuration 16), all scanning magnets 30a may be located in the treatment head 40a together with the energy degrader 41a and the collimator 44a (both described below) in the path of the particle beam between the output channel 17a and the treatment site. Figure 5 , example scanning magnet 43 is controllable in two dimensions (e.g., Cartesian XY dimensions) to locate particle beams in these two dimensions and to move the particle beams through at least a portion of the irradiation target. In this example, scanning magnet 43 comprises a first group 45 of two coils and a second group 46 of two coils, the first group 45 controlling the movement of the particle beams in the Cartesian X dimension of a defined coordinate system, and the second group 46 being orthogonal to the first group of the two coils and controlling the movement of the particle beams in the Cartesian Y dimension. The control of the movement of the particle beams can be achieved by changing the current passing through one or two groups of coils, thereby changing the magnetic field generated thereby. By appropriately changing the magnetic field, the magnetic field acts on the particle beams so that the particle beams move through the beam field in the X and / or Y dimensions, thereby passing through the irradiation target.

[0088] In some embodiments, there may be more than one scanning magnet. Embodiments that include multiple scanning magnets located at different points along the path of the particle beam and separated by air or structures such as magnets or beam absorbing plates may be referred to as split scanning systems. Figure 6In the beam line structure 16b shown (which is a variation of the beam line structure 30), there may be multiple (e.g., two) scanning magnets 30b1 and 30b2 between the output channel 17b and the treatment position. The scanning magnets may be located in the treatment head 40b together with the energy degrader 41b and the collimator 44b, located on the particle beam path between the output channel 17b and the treatment position. The scanning magnets may be at separate locations and separated by air or an energy degradation structure. For example, in this embodiment, the first scanning magnet 30b1 may move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions), and the second scanning magnet 30b2 may move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, the scanning magnets 30b1 and 30b2 may have the same Figure 5 The scanning magnets shown are of the same construction and operation. Each magnet 30b1 and 30b2 can partially move the beam, wherein the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0089] exist Figure 6 In the variation shown in the embodiment, scanning magnet 30b1 can move the particle beam only in one dimension (e.g., Cartesian X dimension), and scanning magnet 30b2 can move the particle beam only in one dimension (e.g., Cartesian Y dimension). As shown, one magnet 30b1 can be upstream of the other magnet 30b2 relative to the particle accelerator. The two can be separated by air or energy degradation structure as described above. Figure 7 and Figure 8 Example magnets 90 and 91 are shown, respectively, with orthogonal coils (coil 90a is orthogonal to coil 91a) to move the particle beam in different dimensions. In this example, the scanning magnet 30b1 can be Figure 7 and includes a first set of coils 90a, and the scanning magnet 30b2 may be Figure 8 and includes a second set of coils 91a orthogonal to coils 90a. Each magnet 30b1, 30b2 can partially move the beam, wherein the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0090] In some embodiments, one or more (e.g., all or less than all) scanning magnets may be located in the beam line structure. Fig. 9 The beam line structure 16c shown (which is Figure 3In a variation of the beam line structure 16 of FIG. 1 , there may be a plurality of (e.g., two) scanning magnets, including a first scanning magnet 30c1 located within the beam line structure 16c and a second scanning magnet 30c2 located outside the beam line structure in the treatment head 40c, as well as an energy degrader 41c and a collimator 44c between the output channel 17 and the treatment position. The first scanning magnet 30c1 may be located between magnetic elements included in the beam line structure 16c. For example, the first scanning magnet 30c1 may be located within the output channel 17c upstream of the bending magnet 32c relative to the particle accelerator, or as Fig. 9 As shown, the first scanning magnet 30c1 may be located upstream of the output channel 17c relative to the particle accelerator. The bending magnet 32c may be or include Figures 25 to 31 shown in and for Figures 25 to 31 In an example, the first scanning magnet 30c1 can be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions), and the second scanning magnet 30c2 can be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, the scanning magnets 30c1 and 30c2 can have the same Figure 5 The scanning magnets shown are of the same construction and operation. Each magnet 30c1 and 30c2 can partially move the beam, wherein the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0091] exist Fig. 9 In a variation of the embodiment, the first scanning magnet 30c1 can be configured to move the particle beam in only one dimension (e.g., the Cartesian X dimension), and the second scanning magnet 30c2 can be configured to move the particle beam in only one dimension (e.g., the Cartesian Y dimension). In this example, the scanning magnet 30c1 can include a first set of coils, and the scanning magnet 30c2 can include a second set of coils orthogonal to the first set of coils. Magnets 30c1 and 30c2 in this example can have similar Figure 7 and Figure 8 Each magnet 30c1 and 30c2 can be configured to partially move the beam, where the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0092] In some embodiments, all scanning magnets may be located in the beam line structure. Fig.10As shown in the separated scanning system of FIG. 1 , the first scanning magnet 30d1 and the second scanning magnet 30d2 can both be located in the beam line structure 16d. No scanning magnet can be located in the treatment head 40d, which in this example includes an energy degrader 41d and a collimator 44d. In other examples, one or more scanning magnets can also be present in the treatment head. The first scanning magnet 30d1 and the second scanning magnet 30d2 can be located between magnetic elements included in the beam line structure 16d. For example, as Fig.10 As shown, the first scanning magnet 30d1 can be located in the output channel 17d upstream of the dipole magnet 32d relative to the particle accelerator, or the first scanning magnet can be located upstream of the output channel 17d relative to the particle accelerator. The second scanning magnet 30d2 can be located upstream of the first scanning magnet 30d1 relative to the particle accelerator. Fig.10 In the example shown, the second scanning magnet 30d2 is in the beam line before the output channel 17d. The scanning magnets can be located at separate locations within the beam line structure and separated by magnets (e.g., dipole or quadrupole magnets) and / or air within the beam line structure. The separate locations can include different points or locations along the path of the particle beam or the length of the beam line structure. For example, Fig.10 As shown, the bending magnet 31d is between the first scanning magnet 30d1 and the second scanning magnet 30d1. In another example, the scanning magnet 30d1 may be moved after the bending magnet 32d so that both the bending magnets 31d and 32d are between the scanning magnets 30d1 and 30d1. In another example, both the scanning magnets 30d1 and 30d2 may be within the output channel 17d, and the bending magnets 31d and 32d may be between the scanning magnets 30d1 and 30d2. In an example, the first scanning magnet 30d1 may be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions), and the second scanning magnet 30d2 may be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, the scanning magnets 30d1 and 30d2 may have the same Figure 5 The bending magnets 23, 31d and 32d may be of the same construction and operation as the scanning magnets shown. Each magnet 30d1 and 30d2 may partially shift the particle beam, wherein the combined movement produced by the two scanning magnets produces the desired particle beam movement specified in the treatment plan. The bending magnets 23, 31d and 32d may be or include Figures 25 to 31 shown in and for Figures 25 to 31 Bending magnets of the type described or modifications thereof.

[0093] exist Fig.10In a variation of the embodiment, the first scanning magnet 30d1 can be configured to move the particle beam in only one dimension (e.g., the Cartesian X dimension), and the second scanning magnet 30d2 can be configured to move the particle beam in only one dimension (e.g., the Cartesian Y dimension). In this example, the scanning magnet 30d1 can include a first set of coils, and the scanning magnet 30d2 can include a second set of coils orthogonal to the first set of coils. Magnets 30d1 and 30d2 in this example can have similar Figure 7 and Figure 8 Each magnet 30d1 and 30d2 can be configured to partially move the beam, where the combined movement produced by the two magnets produces the desired movement specified in the treatment plan.

[0094] In some embodiments, there may be more than two scanning magnets located within the beamline structure and / or between the output end of the output channel and the treatment location. For example, there may be three or more scanning magnets located at separate locations within the beamline structure. For example, there may be three or more scanning magnets located at separate locations between the output end of the output channel and the treatment location. In each case, the scanning magnets may be arranged in series.

[0095] In some embodiments, there may be a single scanning magnet located within the beamline structure or elsewhere upstream of the output end of the output channel. Figure 2 and Figure 3 As shown, the scanning magnet 30 can be located upstream of the output channel 17 relative to the particle accelerator and at the input of the output channel 17. The scanning magnet 30 can be configured to move the particle beam in two dimensions (e.g., Cartesian X and Y dimensions). In this example, the scanning magnet 30 can have a Figure 5 The scanning magnets shown are of the same construction and operation. In this example, all particle beam motion is achieved by controlling the current flowing through one or more coils of a single scanning magnet.

[0096] In this regard, by locating all or some of the scanning magnets within the beam line structure, the size of the particle therapy system may be reduced relative to systems that perform scanning external to the gantry.

[0097] In some embodiments, one or more scanning magnets described herein may be superconducting. For example, one or more (including all) of the scanning magnets downstream of the output channel may be superconducting. For example, one or more (including all) scanning magnets within the beam line structure may be superconducting. In this regard, in the presence of high magnetic fields (such as those found in beam line structures), it may be difficult to accurately move the particle beam. Scanning using superconducting magnets can generate a magnetic field of 2.5T or greater or 3T or greater to move the particle beam, which can overcome the influence of the high magnetic field (e.g., 2.5T or greater or 3T or greater) generated by the beam line structure on the particle beam.

[0098] Fig.11 An example embodiment of a superconducting scanning magnet 92 configured to move a particle beam in two dimensions is shown, which may be used in the scanning embodiments described herein. In this example, the scanning magnet 92 may have Figure 5 The scanning magnet 43 shown in FIG. 1 is similar in structure and operation to the scanning magnet 43 shown in FIG. 1 . The superconducting magnet 92 includes a plurality of sets of high temperature superconducting coils 92a and 92b, which are similar in structure to the scanning magnet 43 shown in FIG. Figure 5 Coils 46 and 45. 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 in a cryostat 94, which maintains the superconducting magnet at a superconducting temperature, for example, above 77 Kelvin (K) or above 90K. The cryostat may include a device configured to maintain the superconducting coil at a cryogenic temperature. The cryostat may maintain the temperature by thermally isolating the superconducting coil from room temperature. This is typically performed using vacuum insulation, thermal radiation shielding, and / or superinsulation to reduce radiative heat transfer, as well as a low thermal conductivity connection between room temperature and cryogenic temperature. In some examples, liquid helium may be used to cool the coil to a superconducting temperature in the cryostat using, for example, conduction or immersion cooling. In conduction cooling, heat is transferred away from the superconducting coil using a thermal conductor. In immersion cooling, the superconducting coil may be in direct contact with a cryogen (e.g., liquid helium). In operation, current is applied to coils 92a and 92b to generate a magnetic field for scanning.

[0099] 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 implementation described herein. The superconducting magnet includes a high-temperature superconducting coil group 95a, which is configured to move a particle beam in only one dimension (e.g., a 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 in a cryostat 96, which maintains the superconducting magnet at a superconducting temperature, such as above 77 Kelvin (K). For example, liquid helium can be used to cool the coil to a superconducting temperature. Current is applied to the coil 95a to generate a magnetic field for scanning. FIG. 12b shows an example of a superconducting scanning magnet 97 configured to move a particle beam in only one dimension. The dimension is different from (e.g., orthogonal 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 group 97a, which is configured to move the particle beam in only one dimension (e.g., a Cartesian X or Y dimension). Examples of high temperature superconductors include, but are not limited to, YBCO and BSCCO. Superconducting magnet 95 is contained in a cryostat 98, which maintains the superconducting magnet at a superconducting temperature, for example, above 77 Kelvin (K). For example, liquid helium can be used to cool the coil to a superconducting temperature. Current is applied to coil 97a to generate a magnetic field for scanning.

[0100] Fig.23 A front cross-sectional view of another exemplary embodiment of a superconducting scanning magnet 150 is shown, which is configured to move a particle beam in two dimensions, which can be used in the scanning embodiments described herein. In this example, the scanning magnet 150 can be contained in, for example, a cryostat (not shown) as described above to maintain the superconducting magnet at a superconducting temperature, for example, between 30°K and 40°K in this example, but the cryostat is not limited to these temperatures. A cryocooler can be used to maintain the temperature of the cryostat at a superconducting temperature. The cryocooler includes a device for actively cooling the superconducting coil to a low temperature. The cryocooler can be controlled by a control system described herein.

[0101] exist Fig.23 , grid 151 illustrates the scanning beam aperture in Cartesian X dimension 153 and Y dimension 154, respectively. For example, grid 151 illustrates that scanning magnet 150 can move the particle beam by ±5 cm in the X dimension and ±5 cm in the Y dimension relative to reference 0,0 point 155. In other embodiments, the scanning magnet can be configured to move the particle beam over a length greater than or less than ±5 cm in the X dimension and greater than or less than ±5 cm in the Y dimension. Fig.23, multiple sets of superconducting coils 158 and 159 are wound on a non-conductive or non-superconductive material 160 to form a hole 161 containing the grid 151. The inner superconducting coil 158 can be separated from the outer superconducting coil 159 by the non-conductive or non-superconductive material 160. The superconducting coil 158 can be configured so that the magnetic field generated thereby is orthogonal to the magnetic field generated by the superconducting coil 159. And, the superconducting coil 159 can be configured so that the magnetic field generated thereby is orthogonal to the magnetic field generated by the superconducting coil 158. For example, the windings of the superconducting coils 158 and 159 can be orthogonal to each other. In some specific implementations, the magnetic fields generated by the superconducting coils 158 and 159 need not be orthogonal, but can be different, for example, at an angle of less than 90° to each other, but still able to be scanned in a grid (e.g., grid 151).

[0102] In this example, superconducting coil 158 controls the movement of particle beam in X dimension. For example, current flows through those superconducting coils to generate magnetic field. The strength of the magnetic field is proportional to the amount of current flowing through the superconducting coils. And, the strength of the magnetic field is proportional to the amount of particle beam moving in X dimension during scanning. In this example, superconducting coil 159 controls the movement of particle beam in Y dimension. For example, current flows through those superconducting coils to generate magnetic field. The strength of the magnetic field is proportional to the amount of current flowing through the superconducting coils. And, the strength of the magnetic field is proportional to the amount of particle beam moving in X dimension during scanning. Current can flow through superconducting coils 158 and 159 at the same time to generate a cumulative magnetic field that moves particle beam in X and Y dimensions. Current can flow through superconducting coils 158 and 159 at different times, so that the particle beam moves in X dimension or Y dimension at a separate time, but still reaches the target position.

[0103] An example of a non-superconducting material that may be included in scanning magnet 150 is copper; however, scanning magnet 150 is not limited to use with copper. The non-superconducting material facilitates heat dissipation during, for example, a quench of superconducting coils 158 and 159.

[0104] Fig.24 A cross-section of an example superconducting coil 165 is shown, which may be used to implement each of superconducting coils 158 and 159 and / or for Figures 25 to 31The coil described in the example bending magnet. The superconducting coil 165 includes a copper (Cu) stabilization layer 166 that surrounds or surrounds other layers of the superconducting coil 165. The superconducting coil 165 also includes a silver (Ag) cap layer 167, a rare earth barium copper oxide (ReBCO) superconducting layer 168 (or a layer of other high temperature superconducting material) adjacent to and in contact with the silver cap layer, a buffer layer stack 169 adjacent to and in contact with the ReBCO superconducting layer to prevent interdiffusion between the oxide and the metal substrate, and a substrate layer 170 adjacent to and in contact with the buffer layer stack. Examples of materials that may be included in the base layer include, but are not limited to, conductive metals such as 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 configuration different from that shown, or may include a material different from that 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 be used, such as YBCO and / or BSCCO.

[0105] Return to reference Figure 3 The output channel 17 of the beamline structure 16 includes a large aperture superconducting (or non-superconducting) bending magnet 31 arranged in series with a large aperture superconducting (or non-superconducting) bending magnet 32. Examples of large apertures include, but are not limited to, 20 cm×20 cm. The bending magnets 31 and 32 may be or include Figures 25 to 31 shown in and for Figures 25 to 31 Bending magnets of the type described or modifications thereof.

[0106] Located between the bending magnets 31 and 32 are a plurality of large aperture superconducting (or non-superconducting) quadrupole magnets 33, 34, and 35. In this example, the quadrupole magnets 33, 34, and 35 alternately include one or more focusing magnets and one or more defocusing magnets to focus and defocus the particle beam, respectively, so as to maintain a substantially uniform cross-sectional area of ​​the particle beam. In this regard, the net effect on particles passing through the alternating magnetic field gradient of the quadrupole magnets is to converge the particle beam; that is, focus. In some embodiments, the quadrupole magnet 33 includes a defocusing magnet, the quadrupole magnet 34 includes a focusing magnet, and the quadrupole magnet 35 includes a defocusing magnet. In some embodiments, the magnet 33 includes a focusing magnet, the quadrupole magnet 34 includes a defocusing magnet, and the quadrupole magnet 35 includes a focusing magnet. In some embodiments, the output channel 17 may include different numbers of quadrupole magnets in different configurations and / or different numbers of dipole magnets in different configurations. In some embodiments, the output channel 17 may include higher order magnetic elements, such as sextupole, instead of or in addition to the quadrupole magnets shown.

[0107] In some implementations, the output channel 17 is configured to bend the particle beam in the presence of a magnetic field of 2.5 T, 3 T, or greater in the beamline structure using bending magnets 31 and 32. For example, a magnetic field of 2.5 T or greater, 3 T or greater, 4 T or greater, 5 T or greater, 6 T or greater, 7 T or greater, 8 T or greater, 9 T or greater, 10 T or greater, 11 T or greater, 12 T or greater, 13 T or greater, 14 T or greater, or 15 T or greater can be generated by passing a current through one or more coils in a magnet in the beamline structure. In the presence of a magnetic field such as these, the magnetic elements in output channel 17 are configured to produce a combined total bending angle of the particle beam anywhere in the range of 90º to 170º, e.g., 90º, 95º, 100º, 105º, 110º, 115º, 120º, 125º, 130º, 135º, 140º, 145º, 150º, 155º, 160º, 165º, or 170º. 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 90º or greater than 170º (e.g., 180º or greater). Figures 1 to 3 In the embodiment of the present invention, using bending magnets 31 and 32, output channel 17 is configured to bend the particle beam at a combined total bending angle of about 150° relative to line 38. In order to achieve a bending amplitude having a value of 110° to 170°, bending magnet 31 can be configured to bend the particle beam within a range of 20° to 85° relative to line 38, and bending magnet 32 ​​can be configured to bend the particle beam within a range of 20° to 85° relative to horizontal line 38.

[0108] In some embodiments, 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 quadrupole magnets, followed by a bending magnet of the type described herein, followed by three alternating quadrupole magnets, followed by a bending magnet of the type described herein. For example, additional magnetic elements 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 longer distances. Conversely, a fewer number of magnetic structures can be used to focus the particle beam over a shorter distance, such as Figure 1 shown.

[0109] Treatment head 40( Figure 1 ) is located at the output end or outlet of the output channel 17. Figure 1In the example of , the treatment head 40 is connected to the output channel 17 and, where applicable, moves with the output channel. The treatment head 40 may or may not be considered to be part of the compact gantry. The treatment head 40 is an example of a particle beam output device. In this example, the treatment head 40 receives the particle beam from the output channel 17 and, in some embodiments, adjusts the particle beam to output to an irradiation target at a treatment position or isocenter, such as a tumor in a patient. In this regard, as noted, the output channel 17 bends the particle beam by at least 90º. Therefore, the particle beam is directed toward the treatment position or isocenter when leaving the output channel 17. In addition, as described herein, the scanning magnet 30 can move the particle beam in a plane to move the particle beam through the irradiation target.

[0110] In this regard, as previously described, the treatment head may include one or more scanning magnets. The energy degrader is downstream of the scanning magnets, and the collimator is downstream of the scanning magnets. Figure 2 and Figure 3 In the embodiment of the present invention, the degrader 41 receives the scanning or moving particle beam from the scanning magnet (one or more). In this example, the degrader 41 is mounted to the gantry 14 (via the treatment head 40) between the output channel 17 and the irradiation target at the treatment position 19. The degrader 41 is configured and controllable to change the energy of the particle beam before the particle beam reaches the irradiation target. In some embodiments, the degrader is the only component that actively controls the energy change of the particle beam before the particle beam reaches the irradiation target. In some implementations, the energy of the particle beam is not actively controllable after the particle beam is output by the particle accelerator and before the particle beam reaches the degrader. For example, in such an embodiment, the components of the gantry between the particle accelerator and the degrader do not actively control the beam energy and are not configured to actively control the beam energy. In other words, the gantry or its beam line channel is not configured to actively control the particle beam after the particle beam is output by the particle accelerator and before the particle beam reaches the degrader. In some cases, there may be some accidental changes in energy caused by movement through the beam line structure; however, these changes are not actively controlled.

[0111] refer to Figure 3 In some embodiments, a single quadrupole magnet may be used in place of quadrupoles 21, 22; a quadrupole magnet may be substituted for scanning magnet 30; scanning magnet 30 may be substituted for quadrupole magnet 35; and the treatment head may include a second scanning magnet. In this case, the scanning magnets may each scan in two dimensions or in one dimension as described herein, or one may scan in two dimensions and one may scan in one dimension.

[0112] As previously mentioned, the particle beam output by the accelerator can be monoenergetic, and the degrader is a single / only or main vehicle for changing the beam energy during the treatment of the irradiated target. Monoenergetic particle beams include particle beams with a single fixed energy level, such as 100MeV, 150MeV, 200MeV, 250MeV, etc. Monoenergetic particle beams can deviate from a fixed energy level by a predetermined amount, such as ±10%, ±5%, ±2% or ±1%, and are still considered monoenergetic. As required for switching particle beam energy during treatment, the operation of switching accelerators during treatment may produce excessive stray neutrons, resulting in the need to increase shielding and reduce beam line efficiency. Neutrons can be generated by particle accelerators and / or by magnets along the beam line structure. By using a monoenergetic particle beam during treatment and relying on a degrader to change the beam energy, the generation of stray neutrons can be reduced or minimized, and the efficiency of the beam line structure can be increased.

[0113] In an example, the degrader may include a plate that can be moved into or out of the path of the particle beam. In another example, the degrader may include a wedge that at least partially overlaps and can move within the path of the particle beam. The example 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 particle beam output with reduced energy. The more material in the particle beam path, the less energy the particle beam will have. In some embodiments, the energy absorbing structure can move over the entire beam field or only over a portion of the beam field. As noted, in some examples, for a given position of the compact gantry, the beam field includes the maximum extent to which the particle beam can move in a plane parallel to the patient treatment area.

[0114] refer to Fig. 22In one example, the degrader 48 is a range modulator that is controllable to move the structure 42 into and out of the path of the particle beam to change the energy of the particle beam and, therefore, the depth at which the dose of the particle beam will be deposited in the irradiated target. Examples of such energy absorbing structures include, but are not limited to, plates; polyhedrons, such as wedges, tetrahedrons, or toroidal polyhedrons; and curved three-dimensional shapes, such as cylinders, spheres, or cones. In this way, the degrader can cause the particle beam to deposit a radiation dose inside the irradiated target to treat a layer or column of the target. In this regard, when protons of a specific energy move through tissue, the protons ionize the atoms of the tissue and deposit a dose primarily at a predefined tissue depth corresponding to the energy. Thus, the degrader is configured to move the particle beam through the target in the Cartesian Z dimension, 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 degrader (e.g., a plate or wedge) can be configured to move during the movement (scanning) of the particle beam and track or trail the particle beam during the movement. An example degrader that tracks or traces the motion of a particle beam is described in U.S. Patent No. 10,675,487 (Zwart), entitled "High Speed ​​Energy Conversion." The contents of U.S. Patent No. 10,675,487, particularly those related to degraders that track or trace the motion of a particle beam (e.g., Figures 36 to 46 of U.S. Patent No. 10,675,487 and the accompanying description), are incorporated herein by reference.

[0115] The Bragg peak is a significant peak on the Bragg curve, which plots the energy loss of ionizing radiation during passage through tissue. The Bragg peak represents the depth at which most radiation is deposited in the tissue. For protons, the Bragg peak occurs just before the particles come to rest. Therefore, the energy of the particle beam can be changed to change the position of its Bragg peak, thereby changing the position at which most of the proton dose will be deposited in the depth of the tissue. In this regard, the particle accelerator can be a fixed energy particle accelerator. In a fixed energy particle accelerator, the particle beam always leaves the particle accelerator with the same or approximately the same energy, for example, within 10%, 5% or 1% of the expected or target energy. In a fixed energy particle accelerator, the degrader is the main vehicle or the only vehicle for changing the beam energy applied to the irradiation target in the patient's body. In some implementations, the particle accelerator described herein is configured to output a particle beam with a single energy or two or more energies in a range between about 100 MeV and about 300 MeV (for example, between 115 MeV and 250 MeV). The fixed energy output may be within this range (eg, 250 MeV), or in some examples, above or below this range.

[0116] In some embodiments, the particle accelerator is a dual energy accelerator. In a dual energy particle accelerator, a particle beam leaves the particle accelerator with one of two different energy levels (high energy level or low energy level). The terms "high" and "low" do not have specific numerical meanings, but are intended to convey relative sizes. In some implementations, the particle accelerator described herein is configured to output particle beams with two energies in a range between about 100 MeV and about 300 MeV. The high energy output and the low energy output can be values ​​within the range, or in some examples, above or below the range. The degrader described herein can be used with a dual energy particle accelerator to reduce the energy of the particle beam to below one of the two energy levels and / or to finely adjust between the two energy levels.

[0117] In the drawings, the treatment head 40 also includes a collimator 44 downstream of the degrader 41 relative to the particle accelerator (i.e., closer to the irradiation target). In the example, the collimator is a controllable structure to allow some radiation to pass to the target and block some radiation from passing to the patient. Typically, the radiation that passes is directed to the irradiation target to be treated, and the blocked radiation would otherwise hit and potentially damage healthy patient tissue. In operation, the collimator is placed in the radiation path between the output channel 17 and the irradiation target and is controlled to create an opening of appropriate size and shape to allow some radiation to pass through the opening to the irradiation target, while the rest of the structure blocks some radiation from reaching adjacent tissue.

[0118] The collimator may be configurable, for example, its aperture may be controlled and changed during treatment. The collimator may be fixed or non-changeable. For example, the collimator may have a fixed shape that cannot be changed.

[0119] In some embodiments, components of an exemplary configurable collimator include a plurality of blades that are dynamically reconfigurable during movement of a particle beam to change the shape of an edge defined by the plurality of blades. The edge is movable between at least a portion of the particle beam and a target of the particle beam such that a first portion of the particle beam on a first side of the edge is at least partially blocked by the plurality of blades and such that a second portion of the particle beam on a second side of the edge is allowed to pass to the target.

[0120] Fig.13 , 14 15 show an example embodiment of a configurable collimator 44a that can be used with the particle therapy system described herein. The collimator 44a includes brackets 113, 114, and 115 that are configured to hold and move the blades vertically and horizontally relative to the irradiation target. As shown, vertical movement includes movement in the Cartesian Z dimension 117, and horizontal movement includes movement in the Cartesian X dimension 118 (Cartesian Y dimension entering or leaving the target). Fig.13 and Fig.14 in the ). Fig.14 and Fig.15 Portions of the carrier housing are shown as transparent in order to illustrate components within the housing; however, the housing is not actually transparent.

[0121] Bracket 113 is referred to herein as a main bracket, and brackets 114 and 115 are referred to herein as secondary brackets. The secondary brackets 114, 115 are coupled to the main bracket 113, such as Figures 13 to 15 As shown. In this example, the sub-brackets 114, 115 each include a housing that is fixed to the main bracket 115 via corresponding members 118, 119. In this example, the main bracket 113 can move vertically (Z dimension) relative to the irradiation target and relative to the particle accelerator along the track 120. The vertical movement of the main bracket 113 also causes the sub-brackets to move vertically. In some embodiments, the sub-brackets move vertically in unison.

[0122] like Figures 13 to 15 As shown, each sub-bracket 114, 115 is connected to a corresponding rod or track 122, 123 along which the sub-bracket moves. More specifically, in this example, a motor 125 drives the sub-bracket 114 to move along the rod 122 toward or away from the sub-bracket 115. Similarly, in this example, a motor 126 drives the sub-bracket 115 to move along the rod 123 toward or away from the sub-bracket 114. As described herein, control of the movement of the main bracket and the sub-bracket is implemented to position the blade relative to the irradiation target. In addition, the blade itself is also configured to move in and out of the bracket, as also described herein.

[0123] like Fig.15 As shown, the motor 130 drives the vertical movement of the main bracket 113. For example, Fig.15 As shown, the lead screw 131 is coupled to a housing 132, which houses the motors 125, 126 that drive the corresponding sub-brackets 114, 115, and the housing 132 is mounted on the track 120. The lead screw 131 is coupled to the motor 130 and is driven vertically by the motor 130. That is, the motor 130 drives the lead screw 131 vertically (Cartesian Z dimension). Because the lead screw 131 is fixed to the housing 132, this movement also causes the housing 132 and therefore the sub-brackets 114, 115 to move along the track 120 toward or away from the irradiation target.

[0124] In this exemplary embodiment, seven blades 135, 136 are mounted on each sub-bracket 114, 115. Each sub-bracket can be configured to move its blades horizontally into or out of the treatment area. Using linear motors, the individual blades on each sub-bracket can be independently and linearly moved in the X dimension relative to other blades on the same sub-bracket. In some embodiments, the blades can also be configured to move in the Y dimension. In addition, the blades on one sub-bracket 114 can move independently of the blades on another sub-bracket 115. These independent movements of the blades on the sub-brackets, together with the vertical movement achieved by the main bracket, allow the blades to move into various configurations. As a result, the blades can conform to a randomly shaped treatment area in the horizontal and vertical dimensions in the horizontal and vertical directions. The size and shape of the blades can be varied to produce different configurations. For example, the size and shape can be changed to treat a single beam spot, thereby treating a single column. In some embodiments, the individual blades on each sub-bracket can be independently and linearly moved using a motor that drives a lead screw in the X dimension relative to other blades on the same sub-bracket.

[0125] The blades can be made of any suitable material that prevents or inhibits the transmission of radiation. The type of radiation used can determine what material is used in the blades. For example, if the radiation is X-rays, the blades can be made of lead. In the examples described herein, the radiation is a proton or ion beam. Therefore, different types of metals or other materials can be used for the blades. For example, the blades can be made of nickel, tungsten, lead, brass, steel, iron, or any suitable combination thereof. The height of each blade can determine the degree to which the blade inhibits the transmission of radiation.

[0126] In U.S. Patent Publication No. 2017 / 0128746 (Zwart) entitled “Adaptive aperture”, a method for Figures 13 to 15 The contents of U.S. Patent Publication No. 2017 / 0128746, particularly with respect to the description of the adaptive aperture (e.g., U.S. Patent Publication No. 2017 / 0128746 Figures 1 to 7 and accompanying description) are incorporated herein by reference.

[0127] Return to reference Figure 1As noted, the exemplary particle therapy system includes a compact gantry 14 that reduces the overall system size. In an implementation of the compact gantry 14, the diameter of the support structure 15 can be less than 6 m, less than 5 m, or less than 4 m. In an example, the diameter of the support structure 15 is 4.8 m. The length of the beamline structure can be measured from the output of the accelerator and the system isocenter, and is equal to the distance between the output of the accelerator and the system isocenter. In an implementation of the compact gantry 14, the length of the beamline structure 16 can be less than 6 m, less than 5 m, less than 4.5 m, or less than 4 m. In an example, the length of the beamline structure 16 is 4.2 m ( Figure 2 ). In this regard, the distance between the particle accelerator and the system isocenter or treatment position can be less than 6m, less than 5m, less than 4.5m, or less than 4m. In an embodiment of the compact gantry 14, the distance between the output end of the output channel 17 and the system isocenter or treatment position is 2m or less, 1.5m or less, or 1m or less. In an embodiment of the compact gantry 14, the distance between the output end of the output channel 17 and the system isocenter or treatment position is between 0.8m and 1.4m. In an example, the distance between the output end of the output channel 17 and the system isocenter or treatment position is 1.01m ( Figure 2 ). Other embodiments may have dimensions different than those listed here.

[0128] In some embodiments, the particle therapy system has a 93 square meter (m 2 ) or less or 75m 2 In some embodiments, the particle therapy system is configured to be installed in a room designed for a linear accelerator. For example, Figures 1 to 3 The components may be small enough to fit in a machine room having the following dimensions, and have dimensions to fit in a machine room having the following dimensions: a length of 25 feet (7.62 m) or less, a width of 20 feet (6.09 m) or less, and a height of 11 feet (3.35 m) or less. For example, Figures 1 to 3 The components may be small enough to fit in a machine room having the following dimensions, and have dimensions to fit in a machine room having the following dimensions: a length of 25 feet (7.62 m) or less, a width of 26 feet (7.92 m) or less, and a height of 10 feet (3.05 m) or less. For example, Figures 1 to 3The components of the particle therapy system may be small enough to fit within a LINAC room and have dimensions to fit within a LINAC room having a footprint of 26.09 feet (11 m) or less by 29.62 feet (9 m) or less, with a height of 16.40 feet (5 m) or less. However, as noted, some embodiments of the particle therapy system may have different dimensions, including but not limited to diameter, height, width, and length. In some embodiments, the ceiling of a pre-existing LINAC room may not be high enough to support a full 360° rotation of or around the rack. In such embodiments, a pit 90 ( 11 m) may be excavated below the floor of the LINAC room. Figure 1 ) to achieve rotation.

[0129] Figure 1 and Fig.16 Examples of treatment spaces 49 and 50 in which the particle therapy system 10 and variations thereof may be housed are shown. In these examples, the treatment space is implemented in a LINAC room, which may be shielded using lead or other suitable materials (e.g., concrete, boronized polyethylene, and / or steel). In this regard, particles (e.g., protons) that are produced by the particle accelerator but do not reach the irradiation target produce secondary radiation through the production of high-energy neutrons. In one example, the particle accelerator 12 and / or gantry produces 10 millisieverts or less of such neutrons per gray dose delivered by the particle beam.

[0130] The use of monoenergetic particle beams and the reliance on de-energizers outside the beamline structure enable the magnetic elements in the beamline to effectively guide the beam. More specifically, the variation of the beamline energy within the beamline increases the generation of stray neutrons, and thus increases the loss of the particle beam within the beamline, thereby reducing its efficiency. The monoenergetic particle beam used in the embodiments of the system described herein combined with the magnetic structure in the beamline can result in increased efficiency. In some cases, the reduction in the length of the beamline structure can also improve efficiency. In some embodiments, the variants of the beamline structure described herein have an efficiency 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 a particle accelerator to particles output from a beamline structure. Thus, an efficiency of 10% or more includes 10% or more of the particles output from the particle accelerator being output from the beam line structure; an efficiency of 20% or more includes 20% or more of the particles output from the particle accelerator being output from the beam line structure; an efficiency of 30% or more includes 30% or more of the particles output from the particle accelerator being output from the beam line structure; an efficiency of 40% or more includes 40% or more of the particles output from the particle accelerator being output from the beam line structure; an efficiency of 50% or more includes 50% or more of the particles output from the particle accelerator being output from the beam line structure; 50% or more of the particles are output from the beamline structure; 60% or more efficiency includes 60% or more of the particles output from the particle accelerator being output from the beamline structure; 70% or more efficiency includes 70% or more of the particles output from the particle accelerator being output from the beamline structure; 80% or more efficiency includes 80% or more of the particles output from the particle accelerator being output from the beamline structure; and 90% or more efficiency includes 90% or more of the particles output from the particle accelerator being output from the beamline structure. In an example, the particle accelerator and gantry described herein deliver more than 70% of the proton beam to the patient even at energies in the lower range of the accelerator.

[0131] The beamline efficiency of the type described herein implements a "single room" solution, in which the particle accelerator, gantry, and patient all reside in a single room, as described above. Within this room, the particle accelerator itself may include shielding, but separate compartments 60 and 61 in the room, respectively, contain the patient and particle accelerator (see Fig.16) do not need to be shielded from each other. In other words, in some embodiments, there is no electromagnetic shielding outside the particle accelerator and the gantry to separate the particle accelerator from the patient. Due to the low level of neutrons emitted by the system, shielding may not be required. In some embodiments, there may be minimal shielding between the individual compartments 60 and 61. For example, the thickness of the shield may be 30 cm or less, the thickness may be 20 cm or less, or the thickness may be 10 cm or less.

[0132] Also refer to Figure 1 , the particle therapy system 10 also includes a treatment couch 51. The treatment couch 51 is configured to move relative to a hole 53 in or through the gantry 14 to position a patient at the 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 section 55, a second section 56, and a third section 57. The first section 55 is rotatably coupled to the second section 56, and the second section 56 is rotatably coupled to the third section 57. The treatment couch 51 is coupled to the third section 57, as shown. The arm 54 is controllable to move the treatment couch 51 into and through the hole 53 to position a patient lying on the treatment couch for treatment; that is, to move the patient to the treatment position. In some embodiments, the arm 54 can position the patient with two degrees of freedom, three degrees of freedom, four degrees of freedom, five degrees of freedom, or six degrees of freedom. Examples of two degrees of freedom are forward-backward movement and side-to-side movement; examples of three degrees of freedom are forward-backward movement, side-to-side movement, and up-down movement; examples of four degrees of freedom are forward-backward movement, side-to-side movement, up-down movement, and one of pitch, yaw, or roll movement; examples of five degrees of freedom are forward-backward movement, side-to-side movement, up-down movement, and two of pitch, yaw, or roll movement; and examples of six degrees of freedom are forward-backward movement, side-to-side movement, up-down movement, pitch movement, yaw movement, and roll movement. In some embodiments, the treatment couch can be replaced by or include a couch that is at least partially tiltable or convertible into a chair, and the couch can still be controlled 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 plurality of degrees of freedom that are compatible with the patient's position. Figure 1 The configuration shown in FIG. 5 is different from that shown in FIG. For example, arm 54 may have two segments or more than three segments. Hydraulic devices, robotic devices, or both may control or achieve non-planar movement of the treatment couch.

[0133] In some embodiments, the output channel 17 may at least partially (including completely) rotate around the support structure 15, or the output channel may remain fixed on the support structure 15, and all or part of the support structure 15 may rotate around the treatment position. In some embodiments, 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 may remain stationary, thereby providing a particle beam fixed in one direction. In embodiments such as these, the treatment couch or other chair moves relative to the fixed beam during treatment. In some systems described herein, the position of the particle beam can be set by rotation of the gantry, and then, except for scanning movement across the irradiation target and movement of the treatment couch or other chair during treatment, the particle beam remains fixed. In some embodiments, a combination of gantry movement and treatment couch (or other chair movement) can be used to implement treatment. For example, the output channel can be positioned and the beam can be temporarily fixed, during which time the treatment couch moves to implement treatment. Afterwards, the output channel can be repositioned to temporarily fix the particle beam in a new position. Treatment can be implemented at a new position by moving the treatment couch. These operations can be repeated as defined in a treatment plan drafted for the particle therapy system.

[0134] The particle therapy system 10 can be an intensity modulated proton therapy (IMPT) system. The IMPT system enables spatial control of a confined proton beam that can have variable energy and / or intensity. IMPT utilizes the charged particle Bragg peak (as noted, the dose characteristic peak at the end of the particle transmission range) combined with the modulation of the particle beam variables to create a target local dose modulation to achieve the target set in the treatment plan. IMPT can involve directing the particle beam to irradiate the target at different angles and different intensities to treat the target. In some embodiments, the particle beam can scan (e.g., move) across the layers of the irradiation target, wherein each layer is treated one or more times from the same or different angles. The movement across the irradiation target to achieve scanning can be performed using the scanning magnet (one or more) described herein.

[0135] Fig.17An exemplary horizontal (x) beam envelope 63 and vertical (y) beam envelope 64 (e.g., cross section) are shown along the length 29 of the compact gantry described herein. The x and y dimensions of the beam spot cross section are determined for the quadrupole magnets 21 and 22, the dipole magnet 23, the quadrupole magnets 26 and 27, the dipole magnet 31, the quadrupole magnets 33, 34, and 35, and the dipole magnet 32. The beam dimensions are determined based on calculations of beam optics using beam parameters measured at the exit of the particle accelerator 12 and design parameters of all beam line magnets. In some embodiments, the beam spot radius at the isocenter (e.g., treatment location) is about 3 millimeters (mm) for both x and y. In some embodiments, for a 200 MeV to 230 MeV proton beam, the magnetic field at the dipole magnets in the beam line structure 16 is no greater than 4 T, and the bending radius of the beam at each dipole magnet is about 0.6 meters. In some embodiments, for 200 MeV to 230 MeV proton beams, the magnetic field at the dipole magnets in the beamline structure 16 is at least 3 T, ie, 3 T or greater. As noted, the systems described herein are not limited to these parameter values, and some embodiments may have different sizes, energies, and magnetic fields.

[0136] Chromatic aberration correction can occur in beam lines with dispersion created by including dipole magnets and multiple correctors in the dispersion region. The standard definition of an achromatic lens is a beam line with zero spatial dispersion (R16) and zero angular dispersion (R26). Fig.18 , the magnetic elements in embodiments of the compact gantry can be configured as achromatic lenses, for example, the R16 65 and R26 66 of the beam transfer matrix elements are equal to zero at the isocenter located at or near the position 67, and are equal to zero at the beam entry point of the gantry at 0m along the beam line structural length 68 (x-axis). Therefore, the gantry as a whole defines an achromatic lens from the beam entry point to the isocenter. A single magnet or combination of magnets within the gantry (which constitutes less than the entire magnet along the beam line within the gantry) is not necessarily an achromatic lens. Reducing or minimizing spatial and angular beam dispersion may be a result of pencil beam scanning techniques implemented by the particle therapy system described herein. In this regard, in some pencil beam scanning techniques, it is required that the cross-section of the particle beam is substantially circular at the isocenter. Therefore, the x and y ( Fig.18 ) plane should be close at the isocenter 67. During beam scanning, the variation of beam shape and beam diameter over the entire scanning area should be reduced or minimized, otherwise, different beam particles with different energies may fall at different positions in the curved plane. This may cause the shape of the beam and the size of the beam to be different in another plane.

[0137] Fig.19An example of beam scanning in the x-dimension 69 and the y-dimension 70 is shown. The firing of the scanning magnets allows the beam particles to be deflected to an angle proportional to the field strength of the scanning magnets. Fig.19 In the example of FIG. 1 , a beam scanning range is shown that completely covers a beam field area of ​​20 cm by 20 cm, where the beam deflection angle is approximately ±20 milliradians (mrad) and ±30 mrad from the scanning magnet. In this example, the source to isocenter distance (SAD) (i.e., the accelerator to isocenter distance) is approximately 4 meters. In some embodiments, the beam bending angle can be as large as 110° to 170° from the scanning magnet (one or more) to the exit of the output channel 17.

[0138] Return to reference Figure 1 In some embodiments, an imaging system including one or more imaging devices 99 can be mounted to the support structure 15. Imaging can be performed before and / or during treatment to identify target locations within the patient and / or to control the operation of the gantry and scanning 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.

[0139] The imaging device can be configured and controlled to rotate around the gantry 14 or rotate with the rotation of the gantry 14. In some embodiments, one or more treatment heads can rotate on a circular track located at 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 circular track and can rotate with it. In some embodiments, the treatment head and the imaging device can be mounted to different internal circumferential tracks within the gantry. For example, the treatment head can be rotatable around a circumferential track at a first radius of the support structure, and the imaging device can be rotatable around a different circumferential track at a second radius of the support structure that is different from the first radius. In some embodiments, the gantry may include different rotatable inner rings, one of which is mounted with a treatment head for rotation, and one of which is mounted with an imaging device or system for rotation.

[0140] In some embodiments, 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 therapy to improve the precision and accuracy of treatment delivery. IGRT can be used to treat tumors in active parts of the body, such as the lungs. The 2D imaging device can be rotated to enable cone beam CT imaging, including dual energy imaging for simultaneous acquisition. The imaging device may also or alternatively include an X-ray source and an image panel or a fan beam diagnostic quality CT imaging device for cone beam CT image acquisition. Optionally, one plane may include a cone beam CT imaging device and the other plane may include a fan beam diagnostic quality CT imaging device.

[0141] As described herein, an example proton therapy system scans a proton beam in three dimensions over an irradiation target in order to destroy malignant tissue. Fig. 20 A cross section of a component 75 of an example superconducting synchrocyclotron accelerator that can be used to provide a particle (e.g., proton) beam in a proton therapy system is shown. In this example, component 75 includes a superconducting magnet 77. The superconducting magnet includes superconducting coils 78 and 79. The superconducting coils are formed by a plurality of integrated conductors, each of which includes a superconducting strand, e.g., a quad or hexafil, wound around a central strand that itself can be superconducting or non-superconducting. Each superconducting coil 78, 79 is used to conduct an electric current that produces a magnetic field (B). Magnetic yokes 80, 81 or smaller pole pieces form a magnetic field in a cavity 84 where particles are accelerated. In an example, a cryostat (not shown) conducts each coil to a low temperature superconducting temperature, e.g., about 4 Kelvin (K), using liquid helium (He).

[0142] In some embodiments, the particle accelerator includes a particle source 85, such as a Penning ion source (PIG source), to provide an ionized plasma column to the cavity 84. Hydrogen or a combination of hydrogen and a rare gas is ionized to produce a plasma column. A voltage source provides a variable radio frequency (RF) voltage to the cavity 84 to accelerate particles from the plasma column in the cavity. As noted, in an example, the particle accelerator is a synchrocyclotron. Therefore, when accelerating particles in the acceleration cavity, the RF voltage sweeps across a certain frequency range to address relativistic effects on particles, such as increasing particle mass. The RF voltage drives a D-shaped plate contained in the cavity and has a frequency that sweeps downward during the acceleration cycle to address the problem of increased relativistic mass of protons and reduced magnetic field. A dummy D-shaped plate is used as a ground reference for the D-shaped plate. The magnetic field generated by passing a current through a superconducting coil, together with the scanning RF voltage, accelerates particles from the plasma column along the track in the cavity, and increases energy as the number of turns increases. The particles in the outermost track are directed to an extraction channel (not shown) and output from the synchrocyclotron as a particle beam. In a synchrocyclotron, the particle beam is pulsed so that a bunch of particles is output periodically.

[0143] The magnetic field in the cavity is shaped to cause the particles to orbit within the cavity as described above. An exemplary synchrocyclotron employs a magnetic field whose rotation angle is uniform and whose strength decreases with increasing radius. In some embodiments, the maximum magnetic field generated by the superconducting (main) coil can be in the range of 2.5 T to 20 T at the center of the cavity, which decreases with increasing radius. For example, the superconducting coils may be used to generate magnetic fields equal to or exceeding one or more of the following magnitudes: 2.5 T, 3.0 T, 3.1 T, 3.2 T, 3.3 T, 3.4 T, 3.5 T, 3.6 T, 3.7 T, 3.8 T, 3.9 T, 4.0 T, 4.1 T, 4.2 T, 4.3 T, 4.4 T, 4.5 T, 4.6 T, 4.7 T, 4.8 T, 4.9 T, 5.0 T, 5.1 T, 5.2 T, 5.3 T, 5.4 T, 5.5 T, 5.6 T, 5.7 T, 5.8 T, 5.9 T, 6.0 T, 6.1 T, 6.2 T, 6.3 T, 6.4 T, 6.5 T, 6.6 T, 6.7 T, 6.8 T, T, 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, 9.0 T, 9.1 T, 9.2 T, 9.3 T, 9.4 T, 9.5T, 9.6 T, 9.7 T, 9.8 T, 9.9 T, 10.0 T, 10.1 T, 10.2 T, 10.3 T, 10.4 T, 10.5 T, 10.6 T, 10.7 T, 10.8 T, 10.9T, 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, 18.4T, 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 may be used to generate magnetic fields outside the range of 2.5 T to 20 T or within the range of 3 T to 20 T but not specifically listed herein.

[0144] By generating a high magnetic field having a magnitude such as those described above, the bending radius of particles orbiting within cavity 84 can be reduced. As a result of the reduced bending radius, a greater number of particle tracks can be generated within a cavity of a given size. Thus, the same number of tracks can fit within a smaller cavity. Reducing the size of the cavity generally reduces the size of the particle accelerator because a smaller cavity requires a smaller yoke or pole piece and other components. In some embodiments, the size or volume of a particle accelerator can be 4 m 3 or smaller, 3m 3 or smaller, or 2m 3 or smaller.

[0145] In some embodiments, for example Fig. 20 In the embodiment shown, relatively large ferromagnetic yokes 80, 81 serve as magnetic return paths for stray magnetic fields generated by the superconducting coils. In some systems, a magnetic shield (not shown) surrounds the yokes. The return yokes and shields together serve to reduce stray magnetic fields, thereby reducing the likelihood that stray magnetic fields will adversely affect the operation of the particle accelerator.

[0146] In some embodiments, the return yoke and / or shield may be replaced or enhanced by an active return system. An example active return system includes one or more active return coils that conduct current in a direction opposite to the current through the main superconducting coils. In some embodiments, there is an active return coil for each superconducting main coil, for example, two active return coils, one for each main superconducting coil. Each active return coil may also be a superconducting coil that concentrically surrounds the outside of the corresponding main superconducting coil. In some specific implementations, the active return coil may be or include a non-superconducting coil. By using an active return system, the relatively large ferromagnetic yokes 80, 81 may be replaced with smaller and lighter pole pieces. Thus, the size and weight of the synchrocyclotron may be further reduced without sacrificing performance. An example of an active return system that may be used is described in U.S. Patent No. 8,791,656 (Zwart), entitled "Active Return System." The contents of U.S. Patent No. 8,791,656, particularly those related to return coil construction (e.g., the contents of U.S. Patent No. 8,791,656) are described in detail. Figure 2 , Figure 4 and Figure 5 and the accompanying description) are incorporated herein by reference.

[0147] 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-Lightweight Magnetically Shielded High Current Compact Cyclotron." The contents of U.S. Patent No. 8,975,836, particularly with reference to U.S. Patent No. 8,975,836, are described in detail in detail in U.S. Patent No. 8,975,836. Figure 4 , 17 The contents related to the "cyclotron 11" or "iron-free cyclotron 11" of 18 and the accompanying description are incorporated herein by reference.

[0148] In some embodiments, the synchrocyclotron used in the proton therapy system described herein can be a variable energy synchrocyclotron. In some implementations, the variable energy synchrocyclotron is configured to 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 any one of a plurality of values ​​to generate a corresponding magnetic field. For example, the current can be set to one of two values ​​to generate the previously described dual energy particle accelerator. In an example embodiment, one or more groups of superconducting coils receive a variable current to generate a variable magnetic field in the cavity. In some examples, one group of coils receives a fixed current, while one or more groups of other coils receive a variable current so that the total current received by the coil group changes. In some embodiments, all coil groups are superconducting. In some embodiments, some coil groups (e.g., coil groups for fixed current) are superconducting, while other coil groups (e.g., one or more coil groups for variable current) are non-superconducting (e.g., copper) coils.

[0149] Typically, in a variable energy synchrocyclotron, the magnitude of the magnetic field can be scaled with the magnitude of the current. Adjusting the total current of the coil within a predetermined range can generate a magnetic field that varies within a corresponding predetermined range. In some instances, continuous adjustment of the current can result in a continuous change in the magnetic field and a continuous change in the output beam energy. Alternatively, when the current applied to the coil is adjusted in a discontinuous, step-by-step manner, the magnetic field and the output beam energy also change accordingly in a discontinuous (step-by-step) manner. The step-by-step adjustment can produce the dual energy previously described. In some embodiments, the size of each step is between 10 MeV and 80 MeV. The scaling of the magnetic field to the current can allow the change in beam energy to be performed relatively accurately, thereby reducing the need for an energy degrader. An example of a variable energy synchrocyclotron that can be used in the particle therapy system described herein is described in U.S. Patent No. 9,730,308, entitled "Particle Accelerator for Producing Charged Particles with Variable Energy." The contents of U.S. Patent No. 9,730,308 are incorporated herein by reference, particularly the contents enabling operation of a synchrocyclotron at variable energy, including columns 5 to 7 of U.S. Patent No. 9,730,308 and Fig.13 and those described in the accompanying description.

[0150] In an embodiment of a particle therapy system using a variable energy synchrocyclotron, the energy of the particle beam can be controlled according to a treatment plan by changing the energy of the particle beam output by the synchrocyclotron to treat a portion of the irradiation target. In such an embodiment, a degrader may or may not be used. For example, controlling the energy of the particle beam may include setting a current in a synchrocyclotron main coil to one of a plurality of values, each value corresponding to a different energy of the particle beam output from the synchrocyclotron. The degrader may be used with a variable energy synchrocyclotron to, for example, provide additional energy variation between discrete energy levels provided by the synchrocyclotron.

[0151] The particle therapy system and its variants described herein can be used to apply ultra-high dose rates of radiation (so-called "FLASH" dose rate radiation) to irradiation targets in patients. In this regard, experimental results of radiotherapy have shown that the condition of healthy tissues subjected to radiation is improved when the therapeutic dose is delivered at an ultra-high (FLASH) dose rate. In one example, when a radiation dose of 10 to 20 grays (Gy) is delivered in pulses of less than 500 milliseconds (ms), achieving an effective dose rate of 20 to 100 grays / second (Gy / S), healthy tissues suffer less damage than when irradiated with the same dose over a longer time scale, while tumors are treated with similar effectiveness. A theory that can explain this "FLASH effect" is based on the fact that radiation damage to tissue is proportional to the oxygen supply in the tissue. In healthy tissue, ultra-high dose rates radicalize oxygen only once, as opposed to the application of doses that radicalize oxygen multiple times over a longer time scale. Using ultra-high dose rates, this can result in less damage to healthy tissue.

[0152] In some examples, as described above, ultra-high dose rate radiation may include a radiation dose exceeding 1 Gy per second for a duration less than 500 ms. In some examples, ultra-high dose rate radiation may include a radiation dose exceeding 1 Gy per second for a duration between 10 ms and 5 s. In some examples, ultra-high dose rate radiation may include a radiation dose exceeding 1 Gy per second for a duration less than 5 s.

[0153] In some examples, ultra-high dose rate radiation includes a radiation dose exceeding one of the following doses in a duration of less than 500 ms: 2 Gy / sec, 3 Gy / sec, 4 Gy / sec, 5 Gy / sec, 6 Gy / sec, 7 Gy / sec, 8 Gy / sec, 9 Gy / sec, 10 Gy / sec, 11 Gy / sec, 12 Gy / sec, 13 Gy / sec, 14 Gy / sec, 15 Gy / sec, 16 Gy / sec, 17 Gy / sec, 18 Gy / sec, 19 Gy / sec, 20 Gy / sec, 30 Gy / sec, 40 Gy / sec, 50 Gy / sec, 60 Gy / sec, 70 Gy / sec, 80 Gy / sec, 90 Gy / sec, or 100 Gy / sec. In some examples, ultra-high dose rate radiation includes a radiation dose exceeding one of the following doses: 2 Gy / second, 3 Gy / second, 4 Gy / second, 5 Gy / second, 6 Gy / second, 7 Gy / second, 8 Gy / second, 9 Gy / second, 10 Gy / second, 11 Gy / second, 12 Gy / second, 13 Gy / second, 14 Gy / second, 15 Gy / second, 16 Gy / second, 17 Gy / second, 18 Gy / second, 19 Gy / second, 20 Gy / second, 30 Gy / second, 40 Gy / second, 50 Gy / second, 60 Gy / second, 70 Gy / second, 80 Gy / second, 90 Gy / second, or 100 Gy / second over a duration between 10 ms and 5 s. In some examples, ultra-high dose rate radiation includes a radiation dose exceeding one of the following doses in a duration of less than 5 seconds: 2 Gy / second, 3 Gy / second, 4 Gy / second, 5 Gy / second, 6 Gy / second, 7 Gy / second, 8 Gy / second, 9 Gy / second, 10 Gy / second, 11 Gy / second, 12 Gy / second, 13 Gy / second, 14 Gy / second, 15 Gy / second, 16 Gy / second, 17 Gy / second, 18 Gy / second, 19 Gy / second, 20 Gy / second, 30 Gy / second, 40 Gy / second, 50 Gy / second, 60 Gy / second, 70 Gy / second, 80 Gy / second, 90 Gy / second, or 100 Gy / second.

[0154] In some examples, ultra-high dose rate radiation includes a radiation dose exceeding one or more of the following doses: 100 Gray per second, 200 Gray per second, 300 Gray per second, 400 Gray per second, or 500 Gray per second in a duration of less than 500 ms, in a duration between 10 ms and 5 s, or in a duration of less than 5 s.

[0155] In some examples, the ultra-high dose rate of radiation includes a radiation dose 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 includes a radiation dose 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 includes a radiation dose 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 includes a radiation dose between 40 Gray per second and 120 Gray per second for a period of time, such as less than 5 s. Other examples of time periods are those provided above.

[0156] In some embodiments, particle therapy systems can use ultra-high dose rate radiation (FLASH dose rate radiation) to treat a three-dimensional column of a target. These systems use pencil beam scanning to scale ultra-high dose rate delivery to a target. In some examples, pencil beam scanning includes delivering a series of small beams of particle radiation, each of which can each 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 therapy. In addition, rather than planning treatment into layers at a constant energy, the system plans treatment into columns defined by the direction of a fixed beam. The direction of the beam can be toward the surface of the target.

[0157] In some embodiments, all or part of the column is treated before the particle beam is directed along another path through the irradiation target. In some embodiments, the path through the target passes through the target in whole or in part. In an example, the particle beam can be directed along a path through the target and does not deviate from the path. When directed along the path, the energy of the particle beam is changed. The particle beam does not move as its energy changes, whereby the particle beam treats the entire or a portion of the interior of the target extending along the length of the particle beam and along the width of the beam spot. Therefore, treatment is performed in the depth direction along the longitudinal direction of the particle beam. For example, a portion of the target being treated can extend downward from the beam spot at the target surface through the entire or a portion of the interior of the target. The result is that the particle beam uses ultra-high dose rate radiation to treat a three-dimensional columnar portion of the target. In some examples, the particle beam may no longer be directed more than once along the same three-dimensional columnar portion.

[0158] In some embodiments, the irradiation target can be broken down into microvolumes. Although cubic microvolumes can be used, the microvolumes can have any suitable shape, such as a three-dimensional positon, a regular curved shape, or an irregular or amorphous shape. In this example, each microvolume is treated by delivering FLASH radiation by a column in the manner described herein. For example, the beam energy can be changed by using a degrader plate or by controlling a variable energy synchrocyclotron to change the beam energy, thereby treating the column depth of the microvolume with radiation. After a single microvolume has been treated, the next microvolume is treated, and so on, until the entire irradiation target has been treated. Treatment of the microvolumes can be performed in any suitable order or sequence.

[0159] In some implementations, particle accelerators other than synchrocyclotrons can be used in the particle therapy systems described herein. For example, a cyclotron, synchrotron, linear accelerator, etc. can replace the synchrocyclotron in the particle therapy systems described herein.

[0160] One or more scanning magnets (not shown) may be located in the particle beam path between the particle accelerator and the treatment couch. The scanning magnets may be superconducting, non-superconducting, or a combination of superconducting and non-superconducting. For example, the scanning magnets may be Figure 5 , Figure 7 , Figure 8 , Fig.11 , Fig. 12A , Fig. 12B , Fig.23 Or the type shown in a combination thereof. In some embodiments, the scanning is controlled by changing the current through one or both sets of coils to thereby change the magnetic field generated thereby. By appropriately changing the magnetic field, the particle beam can be moved through the irradiation target in the X and / or Y dimensions.

[0161] In some embodiments, the scanning magnet can be replaced by a scattering foil, and the degrader can be a range modulator. In implementations such as this, the scattering foil scatters the particle beam over the entire treatment area, and the depth of the scattered beam application is controlled by the range modulator. A configurable collimator can be kept in place to trim the edges of the scattered beam.

[0162] Fig.21 Another example particle therapy system 320 using the bending magnets described herein is shown in FIG. Fig.21In the embodiment of the present invention, the gantry 394 can be rotationally or axially connected to the treatment room floor 396 to enable controlled movement of the gantry 394 relative to the treatment room floor. In this example, the particle accelerator 10 is mounted on the gantry and the gantry can be used to rotate around the patient in the direction of arrow 321 to direct the particle beam to the patient. The gantry 394 may include an arm 397 that extends the length of the gantry 394 and reaches the treatment room floor 396. The particle accelerator 10 and the connected beamline structure 398 are rotatably mounted to the arm 397. That is, the particle accelerator 10 and the connected beamline structure 398 are connected to the end 399 of the arm 397 so that the particle accelerator 10 and the connected beamline structure 398 can be rotated at the end 399 in the direction of arrow 322. This rotation is separate from the gantry rotation described herein. The beamline structure 398 may include a rotation mechanism related to Figures 25 to 31 One or more bending magnets of the type described herein or any variation thereof. For example, the beamline structure may include Figures 25 to 31 Two bending magnets 350 and 351 of the type described, or any variation thereof, may be used to bend the particle beam by more than 90°, such as 100°, 110°, 120° or more, towards the irradiation target.

[0163] The control system 192 ( Figure 1 ) or 392( Fig.21 ) at least partially controls the operation of the example proton therapy systems described herein and the operation of all or some of their components, the control system 192 or 392 being configured to execute one or more computer program products, such as one or more computer programs tangibly embodied in one or more non-transitory machine-readable media, for execution or control of its operation by one or more data processing devices, such as programmable processors, computers, multiple computers and / or programmable logic components.

[0164] All or part of the systems described in this specification and various modifications thereof may be configured or controlled at least in part by one or more computers, such as a control system using one or more computer programs tangibly embodied in one or more information carriers, such as one or more non-transitory machine-readable storage media. The computer program may be written in any form of programming language, including compiled or interpreted languages, and it may be deployed in any form, including as a stand-alone program or as a module, component, subroutine, or other unit suitable for use in a computing environment. The computer program may be deployed to execute on one computer, or on multiple computers at one location, or distributed across multiple locations and interconnected by a network.

[0165] The 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 part of the systems and processes may be configured or controlled by dedicated logic circuitry, such as an FPGA (field programmable gate array) and / or an ASIC (application specific integrated circuit) or an embedded microprocessor limited to instrument hardware.

[0166] For example, processors suitable for executing computer programs include both general-purpose microprocessors and special-purpose microprocessors, as well as any one or more processors of any kind of digital computer. Typically, the processor will receive instructions and data from a read-only storage area or a random access storage area or both. The elements of a computer include one or more processors for executing instructions and one or more storage area devices for storing instructions and data. In general, a computer will also include one or more machine-readable storage media, or be operatively coupled to one or more machine-readable storage media to receive data from one or more machine-readable storage media, or to transfer data to one or more machine-readable storage media, or both, the machine-readable storage medium is, for example, a mass storage device for storing data, such as a magnetic disk, a magneto-optical disk, or an optical disk. Non-transitory machine-readable storage media suitable for containing computer program instructions and data include all forms of non-volatile storage areas, for example, semiconductor storage area devices, such as EPROM (erasable programmable read-only memory), EEPROM (electrically erasable programmable read-only memory), and flash memory area devices; magnetic disks, such as internal hard disks or removable disks; magneto-optical disks; and CD-ROM (compact disk read-only memory) and DVD-ROM (digital versatile disk read-only memory).

[0167] The elements of the different embodiments described can be combined to form other embodiments not previously specifically described. Elements can be excluded from the previously described system without adversely affecting their operation or the operation of the system as a whole. In addition, various individual elements can be combined into one or more individual elements to perform the functions described in this specification.

[0168] Other implementations not specifically described in this specification are also within the scope of the appended claims.

Claims

1. A magnet, comprising: An assembly comprising: (i) a coil assembly for conducting an electric current to generate a magnetic field, and (ii) a support structure on which the coil assembly is asymmetrically disposed; and A ferromagnetic yoke surrounds a portion of the assembly, the ferromagnetic yoke and the assembly being curved.

2. The magnet according to claim 1, wherein The coil assembly includes a first coil and a second coil, the first coil and the second coil are used to conduct current to generate a magnetic field, the first coil and the second coil are asymmetrically arranged on the support structure in a first hemisphere of the magnet, so that a first interval between the first coil and the second coil in a first quadrant of the magnet is different from a second interval 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; wherein the coil group comprises a third coil and a fourth coil, the third coil and the fourth coil are used to conduct current to generate a magnetic field, the third coil and the fourth coil are asymmetrically arranged on the support structure in the second hemisphere of the magnet, so that a third interval between the third coil and the fourth coil in the third quadrant of the magnet is different from a fourth interval between the third coil and the fourth coil in the fourth quadrant of the magnet, and the third quadrant and the fourth quadrant are within the second hemisphere; and The asymmetry of the first coil and the second coil in the first quadrant and the second quadrant, respectively, is a mirror image of the asymmetry of the third coil and the fourth coil in the third quadrant and the fourth quadrant, respectively.

3. The magnet according to claim 2, wherein: 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; and Wherein, the first interval and the third interval are located at an inner bending radius of the component, and the second interval and the fourth interval are located at an outer bending radius of the component.

4. The magnet according to claim 3, wherein The coil group includes a fifth coil and a sixth coil, the fifth coil and the sixth coil are used to conduct current to generate a magnetic field, the fifth coil is arranged on the support structure in the first hemisphere, and the sixth coil is arranged on the support structure in the second hemisphere; wherein a fifth interval between the fifth coil and an adjacent coil of the first coil or the second coil in the first quadrant is different from a sixth interval between the fifth coil and an adjacent coil of the first coil or the second coil in the second quadrant; wherein a seventh interval between the sixth coil and an adjacent coil of the third coil or the fourth coil in the third quadrant is different from an eighth interval between the sixth coil and an adjacent coil of the third coil or the fourth coil in the fourth quadrant; and The asymmetry of the first coil, the second coil and the fifth coil in the first quadrant and the second quadrant respectively is a mirror image of the asymmetry of the third coil, the fourth coil and the sixth coil in the third quadrant and the fourth quadrant respectively.

5. The magnet according to claim 4, wherein The fifth interval is equal to the seventh interval, the sixth interval is equal to the eighth interval, and the fifth interval and the seventh interval are smaller than the sixth interval and the eighth interval; and Wherein, the fifth interval and the seventh interval are located at the inner bending radius of the component, and the sixth interval and the eighth interval are located at the outer bending radius of the component.

6. The magnet according to claim 5, wherein The coil group includes a seventh coil and an eighth coil, the seventh coil and the eighth coil are used to conduct current to generate a magnetic field, the seventh coil is arranged on the support structure in the first hemisphere, and the eighth coil is arranged on the support structure in the second hemisphere; wherein a ninth interval between the seventh coil and an adjacent coil of the first, second or fifth coil in the first quadrant is different from a tenth interval between the seventh coil and an adjacent coil of the first, second or fifth coil in the second quadrant; wherein an eleventh interval between the eighth coil and an adjacent coil among the third coil, the fourth coil, or the sixth coil in the third quadrant is different from a twelfth interval between the eighth coil and an adjacent coil among the third coil, the fourth coil, or the sixth coil in the fourth quadrant; and The asymmetry of the first coil, the second coil, the fifth coil and the seventh coil in the first quadrant and the second quadrant respectively is a mirror image of the asymmetry of the third coil, the fourth coil, the sixth coil and the eighth coil in the third quadrant and the fourth quadrant respectively.

7. The magnet according to claim 6, wherein 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; and Wherein, the ninth interval and the eleventh interval are located at the inner bending radius of the component, and the tenth interval and the twelfth interval are located at the outer bending radius of the component.

8. The magnet according to claim 2, wherein: the ferromagnetic yoke comprising a notch adjacent the assembly, the notch being asymmetric in the first quadrant and the second quadrant, wherein the asymmetry of the notch is with respect to at least one of a size, shape, or placement of the notch; and Wherein, the asymmetry of the notch in the third quadrant and the fourth quadrant, respectively, is a mirror image of the asymmetry of the notch in the first quadrant and the second quadrant, respectively.

9. The magnet according to claim 4, wherein the ferromagnetic yoke comprising a notch adjacent the assembly, the notch being asymmetric in the first quadrant and the second quadrant, wherein the asymmetry of the notch is with respect to at least one of a size, shape, or placement of the notch; and Wherein, the asymmetry of the notch in the third quadrant and the fourth quadrant, respectively, is a mirror image of the asymmetry of the notch in the first quadrant and the second quadrant, respectively.

10. The magnet according to claim 6, wherein the ferromagnetic yoke comprising a notch adjacent the assembly, the notch being asymmetric in the first quadrant and the second quadrant, wherein the asymmetry of the notch is with respect to at least one of a size, shape, or placement of the notch; and Wherein, the asymmetry of the notch in the third quadrant and the fourth quadrant, respectively, is a mirror image of the asymmetry of the notch in the first quadrant and the second quadrant, respectively.

11. The magnet according to claim 1, wherein The ferromagnetic yoke comprises iron; and Wherein the support structure is non-ferromagnetic.

12. The magnet according to claim 1, wherein The magnet is bent by 60° or more with respect to a straight line passing through the center of an unbent portion of the magnet.

13. The magnet according to claim 1, wherein The magnet is bent by 70° or more with respect to a straight line passing through the center of an unbent portion of the magnet.

14. The magnet according to claim 1, wherein The magnet is bent by 80° or more with respect to a straight line passing through the center of an unbent portion of the magnet.

15. The magnet according to claim 1, wherein The magnet is bent by 90° or more with respect to a straight line passing through the center of an unbent portion of the magnet.

16. The magnet according to claim 1, wherein The magnet is bent by 120° or more with respect to a straight line passing through the center of an unbent portion of the magnet.

17. The magnet according to claim 1, wherein The magnet is a cosine theta magnet wherein the current through the coil assembly has a greater concentration near the 0° or 180° position of the magnet than near the 90° or -90 / 270° position of the magnet.

18. The magnet according to claim 1, wherein The coil assembly is configured for a two-pole function.

19. The magnet according to claim 1, wherein The coil assembly is configured for four-pole functionality.

20. The magnet according to claim 1, wherein The coil assembly is configured for six-pole functionality.

21. The magnet according to claim 1, wherein The coil assembly includes a superconducting material.

22. The magnet according to claim 1, further comprising: One or more magnetic spacers are movable relative to the ferromagnetic yoke to change the magnetic field generated by the magnets.

23. The magnet according to claim 1, wherein The coil assembly includes two or more groups of coils, the two or more groups of coils being 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 system comprising: a gantry including a beamline structure configured to direct a monoenergetic particle beam from an output of a particle accelerator to an irradiation target, the beamline structure including a bending magnet to bend the particle beam along a length of the beamline structure; Wherein, at least one of the bending magnets comprises the magnet according to claim 1.

25. The system of claim 24, further comprising: An energy degrader is a unique component that actively controls the energy change of the particle beam after the particle accelerator outputs the particle beam and before the particle beam reaches the irradiation target.

26. The system of claim 24, wherein: The beam line structure is configured to not 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 degrader.

27. 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. 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. The system of claim 24, further comprising: A collimator is located downstream of the gantry relative to the particle accelerator, the collimator being configured to block at least a portion of the particle beam before at least a portion of the particle beam reaches the irradiation target.

30. The system of claim 24, wherein: The gantry includes a support structure configured to move a portion of the beam line structure along a circular path around the irradiation target; and Wherein, the supporting structure has a size of 6 meters or less.

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

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

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

34. The system of claim 24, wherein: The energy of the particle beam varies by no more than 1% within the beam line structure.

35. The system of claim 24, wherein: The distance between the output end of the beam line structure and the isocenter containing the irradiation target is 1.5 meters (m) or less.

36. The system of claim 24, wherein: The beamline structure comprises an output channel comprising at least some bending magnets comprising dipole magnets arranged in series to bend the particle beam by at least 90°, the dipole magnets comprising the at least one bending magnet.

37. The system of claim 24, wherein: The beamline structure comprises an output channel comprising at least some bending magnets, the at least one bending magnet being located before the output channel in the direction of travel of the particle beam.

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

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