Portable liquid helium-free MRI-guided proton therapy system

Abstract

An example system for magnetic resonance imaging (MRI) guided proton therapy includes a source configured to generate and emit a heavy ion or proton beam; and an MRI assembly including at least one segmented coil magnet comprising a plurality of windings, where the plurality of windings of the at least one segmented coil magnet include a medium or high-temperature superconductor such that the at least one segmented coil magnet does not require a cryogenic liquid for cooling. An example magnet for MRI proton therapy includes field coils and shield coils arranged in Helmholtz coil pairs, where the field coils and the shield coils include a medium or high-temperature superconductor so the magnet does not require cryogenic liquid cooling, and where the field coils and the shield coils form a central opening through which a heavy ion or proton beam is directed during proton therapy.

Description

Atty. Dkt. No.103361-624WO1 PORTABLE LIQUID HELIUM-FREE MRI-GUIDED PROTON THERAPY SYSTEM CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application claims the benefit of U.S. provisional patent application No. 63 / 621,226 filed on January 16, 2024, which is expressly incorporated herein by reference in its entirety. BACKGROUND

[0002] Proton or heavy ion particle therapy has recently become a popular treatment modality in the field of radiation therapy for cancer treatments. Compared to photon beams, heavy ion particle beams or proton beams have a characteristic called “Bragg Peak” which enables a high dose to be administered to a tumor while greatly limiting the dose to the surrounding healthy tissues (e.g., critical organs). Additionally, heavy ion particle beams or proton beams have been shown to have a higher cell-killing effect than photon beams. Proton therapy is quickly becoming a unique and important cancer treatment modality. More and more proton or heavy ion particle therapy facilities have been opened or are going to be opened for clinical practice around the world; although, the cost to build a proton or heavy ion particle therapy facility is over an order of magnitude higher than a photon therapy facility.

[0003] Despite these advantages, the superiority of proton or heavy ion particle therapy comes with implementation challenges due to the sharp dose drop-off, e.g., associated with Bragg Peak. For example, poor targeting raises the risk of missing target tissue (e.g., a tumor), putting healthy tissues and organs at risk of damage and potentially resulting in worse treatment outcomes than traditional photon therapy. Imaging guidance therefore plays an important role in proton therapy, e.g., to ensure the accuracy of targeting. The currently available imaging technology in image-guided proton therapy is mainly based on 2D orthogonal X-ray imaging. Only in some centers are in-room 3D computed tomography (CT) or on-board cone-beam CT (CBCT) imaging available; yet 3D or 4D CT and MRI images are currently limited to offline use for treatment planning. To carry out a safe and efficient treatment, higher accuracy image guidance and better imaging techniques to distinguish a tumor from healthy tissues are needed.Atty. Dkt. No.103361-624WO1 SUMMARY

[0004] In some aspects, implementations of the present disclosure include a system for magnetic resonance imaging (MRI) guided proton therapy, the system including: a source configured to generate and emit a heavy ion or proton beam; and an MRI assembly including at least one segmented coil magnet including a plurality of windings, wherein plurality of windings of the at least one segmented coil magnet include a medium or high-temperature superconductor such that the at least one segmented coil magnet does not require a cryogenic liquid for cooling.

[0005] In some aspects, implementations of the present disclosure include a system, wherein the medium or high-temperature superconductor is magnesium diboride (MgB2).

[0006] In some aspects, implementations of the present disclosure include a system, wherein the medium or high-temperature superconductor is Niobium-Tin (Nb3Sn).

[0007] In some aspects, implementations of the present disclosure include a system, wherein the windings of the at least one segmented coil magnet have a rectangular cross-section.

[0008] In some aspects, implementations of the present disclosure include a system, wherein the at least one segmented coil magnet is annularly shaped such that a central opening is formed in the at least one segmented coil magnet, wherein the proton beam is directed through the central opening toward a target.

[0009] In some aspects, implementations of the present disclosure include a system, wherein the central opening is larger than 0.7 m in diameter.

[0010] In some aspects, implementations of the present disclosure include a system, further including one or more guidance magnets configured to manipulate a direction of the proton beam after emission from the source.

[0011] In some aspects, implementations of the present disclosure include a system, wherein the at least one segmented coil magnet is configured to operate at or above 3 T.

[0012] In some aspects, implementations of the present disclosure include a system, wherein the at least one segmented coil magnet includes a cryostat that at least partially encompasses the windings.Atty. Dkt. No.103361-624WO1

[0013] In some aspects, implementations of the present disclosure include a system, wherein the windings of the at least one segmented coil magnet include: (i) a first coil, (ii) two or more first Helmholtz coil pairs, and (iii) at least one second Helmholtz coil pair, wherein the first coil and the two or more first Helmholtz coil pairs are configured as field coils, and wherein the at least one second Helmholtz coil pair is configured as a shield coil.

[0014] In some aspects, implementations of the present disclosure include a system, further including a controller configured to receive imaging data from the MRI assembly.

[0015] In some aspects, implementations of the present disclosure include a system, wherein the controller is further configured to determine a position of a tumor in a subject based on the imaging data for guidance of the heavy ion or proton beam.

[0016] In some aspects, implementations of the present disclosure include a magnet for a magnetic resonance imaging (MRI) system to be used in proton therapy, the magnet including: a plurality of field coils including a plurality of windings, wherein at least two of the plurality of field coils are configured as a Helmholtz coil pair; a plurality of shield coils, wherein at least two of the plurality of shield coils are configured as a Helmholtz coil pair, wherein the plurality of field coils and the plurality of shield coils are formed of a material that at least partially includes a medium or high-temperature superconductor, such that the magnet does not require a cryogenic liquid for cooling, and wherein the plurality of field coils and the plurality of shield coils are arranged annularly such that a central opening is formed in magnet through which a heavy ion or proton beam is directed during proton therapy.

[0017] In some aspects, implementations of the present disclosure include a magnet, wherein the medium or high-temperature superconductor is magnesium diboride (MgB2).

[0018] In some aspects, implementations of the present disclosure include a magnet, wherein the medium or high-temperature superconductor is Niobium-Tin (Nb3Sn).

[0019] In some aspects, implementations of the present disclosure include a magnet, wherein the plurality of windings of the plurality of field coils have a rectangular cross-section.

[0020] In some aspects, implementations of the present disclosure include a magnet, wherein the central opening is larger than 0.7 m in diameter.Atty. Dkt. No.103361-624WO1

[0021] In some aspects, implementations of the present disclosure include a magnet, wherein the magnet is configured to operate at or above 3 T.

[0022] In some aspects, implementations of the present disclosure include a magnet, further including a cryostat that at least partially encompasses the plurality of field coils and the plurality of shield coils.

[0023] In some aspects, implementations of the present disclosure include a magnet, wherein at least one of the field coils or at least one of the shield coils includes a segmented coil. BRIEF DESCRIPTION OF THE DRAWINGS

[0024] The components in the drawings are not necessarily to scale relative to each other. Like reference numerals designate corresponding parts throughout the several views.

[0025] FIG. 1 illustrates an example proton beam experiment.

[0026] FIG. 2 illustrates an example mechanical model of an MRI magnet geometry and stresses.

[0027] FIG. 3 illustrates an example magnetic field model of an MRI magnet.

[0028] FIG. 4 illustrates an example of energy deposition in the X-Z plane for a proton therapy system according to implementations of the present disclosure.

[0029] FIG. 5 illustrates energy deposition on the Z-axis for a proton therapy system according to implementations of the present disclosure.

[0030] FIG. 6 illustrates dose distributions in the X and Z axis for a proton therapy system according to implementations of the present disclosure.

[0031] FIG. 7A illustrates example 2D schematics of a mechanical design including a 2D view of Maxwell coil geometry in a plane, according to implementations of the present disclosure.

[0032] FIG. 7B illustrates an example schematic of a superconducting winding with copper former, stainless steel cryostat, and G10 fiber-glass epoxy mechanical support, according to a study of an example implementation of the present disclosure.

[0033] FIG. 7C illustrates an example cryocooler, according to an example implementation of the present disclosure.Atty. Dkt. No.103361-624WO1

[0034] FIG. 7D illustrates a 3D magnet geometry for an SS cryostat with copper and an MgB2winding, according to a study of an example implementation of the present disclosure.

[0035] FIG. 7E illustrates a simulated winding temperature map after 9.875 days of coolingfrom 77.3 K with thermal expansion (10-times magnified), no magnetic field. T = 4.3 K,according to a study of an example implementation of the present disclosure.

[0036] FIG. 7F illustrates a simulated winding temperature map after 4.917 days of coolingfrom 77.3 K with thermal expansion (10-times magnified), no magnetic field. T = 4.75 K,according to a study of an example implementation of the present disclosure.

[0037] FIG. 7G illustrates a simulated winding temperature map after 9.875 days of coolingfrom 77.3 K with thermal expansion (10-times magnified), magnetic field present. T = 4.75 K,according to a study of an example implementation of the present disclosure..

[0038] FIG. 7H illustrates a simulated winding temperature map after 4.917 days of coolingfrom 77.3 K with thermal expansion (10-times magnified), magnetic field present. T = 4.71 K,according to a study of an example implementation of the present disclosure.

[0039] FIG. 7I illustrates a simulated winding temperature map after 4.917 days of cooling from 77.3 K with magnetic field present, according to a study of an example implementation of the present disclosure.

[0040] FIG. 7J illustrates a pair of cryocoolers used in the temperature maps illustrated in FIGS.7E-7I.

[0041] FIG. 7K illustrates a cross section of a non-insulated rectangular MgB wire for MRIapplications, according to implementations of the present disclosure.

[0042] FIG. 7L illustrates a map of 10 ppm inhomogeneity in a study of an example implementation of the present disclosure.

[0043] FIG. 8 illustrates an example MRI-guided proton therapy system, according to implementations of the present disclosure.

[0044] FIG. 9 illustrates example configurations of magnets that can be used to implement the MRI-guided proton therapy system of FIG.8.Atty. Dkt. No.103361-624WO1 DETAILED DESCRIPTION

[0045] MRI guidance for proton therapy is a promising and desirable technology since X-ray- based imaging lacks the ability to identify soft tissue, e.g., during the patient setup, and to perform real-time on live imaging. Compared with X-ray and CBCT imaging, MRI has the following advantages: (1) soft tissue enhancement, (2) on-live treatment monitoring and assessment, (3) no adverse radiation, and (4) functional imaging features that may monitor tumor responses to the treatment. However, as mentioned above, this are currently no MRI-guided proton therapy systems available. Therefore, disclosed herein is a portable liquid helium-free MRI system that is integrated into an “MRI-guided proton therapy” with a proton beam source and gantry. The disclosed MRI-guided proton therapy system utilizes a magnet having coils fabricated from a unique high-temperature superconductor material, magnesium diboride(MgB2), which eliminates the need for a bulky and expensive liquid helium cooling system (e.g.,to cool the magnet coils). The disclosed MRI system is therefore less costly than existing MRI systems and much more compact, allowing for the acquisition of live images to guide proton beam and monitor patient treatments.

[0046] FIG. 8 is a diagram of an MRI-guided proton therapy system 800, according to some implementations. As shown, system 800 includes a source 802 configured to generate and emit a heavy ion or proton beam. In some implementations, source 802 is a synchrotron or cyclotron accelerator; however, the present disclosure is not intended to be limiting in this regard. System 800 further includes an MRI assembly 804 to enable guidance, e.g., of the heavy ion or proton beam, during therapy. In this regard, MRI assembly 804 may be configured to capture MRI images of a target 810 to determine a position of a tumor or other target area of target 810 while the heavy ion or proton beam is on or off for treatments. Notably, MRI assembly 804 may be configured to operate throughout or during a proton therapy session so that the position of a tumor or other target area of target 810 can be tracked (e.g., due to movement of a subject) and / or the direction / depth of the heavy ion or proton beam can be determined and / or adjusted. It should also be appreciated that MRI assembly 804 is configured to be movable with respect source 802. In other words, MRI assembly 804 may be connected to source 802 via a manipulatable connection system.Atty. Dkt. No.103361-624WO1

[0047] In contrast to traditional MRI systems, MRI assembly 804 is operable without (and therefore does not include) a liquid helium cooling system. In this regard, MRI assembly 804 is shown to include magnets 806 and 808, e.g., which do not require liquid helium for cooling and are therefore referred to herein as “liquid helium free” magnets, among various other components that are not illustrated but that would be known to those in the art. While not shown due to the perspective of FIG. 8, at least magnet 806 – but optionally also magnet 808 – is generally annular in shape, thereby defining a central opening through which the heavy ion or proton beam is directed towards target 810. In some implementations, magnets 806 and 808 are segmented coil magnets that include a plurality of internal windings. As used herein, a segmented coil magnet is used to refer to a magnet where multiple coil windings (“segments”) are bound together to form a coil magnet with a increased strength and / or a more uniform / homogeneous field. Accordingly, it should be understood that the “coils” referred to herein can optionally refer to segmented coil windings that are made of any number of windings and coils joined together. Optionally, the windings of magnets 806 and 808 are at least partially composed of a medium or high-temperature superconducting material. In some implementations, the medium or high-temperature superconducting material is one of magnesium diboride (MgB2) or Niobium-Tin (Nb3Sn). In some such implementations, the windings may be composed of an alloy that includes a medium or high-temperature superconducting material combined with other materials, such as copper or aluminum. As will be appreciated, utilizing a medium or high-temperature superconducting material means that the windings of magnets 806 and 808 do not require a traditional liquid helium cooling system to operate, yielding a much more compact (e.g., possibly movable) and cost-effective MRI.

[0048] Optionally, one or more guidance magnets 830 can be positioned to manipulate a direction of the proton beam after emission from the source.

[0049] With reference to FIG. 9, implementations of the present disclosure can further include magnets that can be used with the MRI assembly 804 described with reference to FIG.8. For example, the magnet 900 shown in FIG.9 can be used as either or both of the magnets 806, 808 of FIG.8.

[0050] FIG. 9 illustrates a cross section of a side view of an example magnet 900 that can be configured for proton therapy. The magnet 900 can include any number of field coils 902a, 902b,Atty. Dkt. No.103361-624WO1 although two field coils 902a, 902b are shown in FIG. 9. As shown in FIG.9, at least two of the field coils 902a, 902b can be configured as a Helmholtz coil pair 904. Each field coil 902a, 902b can be formed using one or more windings.

[0051] Still with reference to FIG.9, the magnet 900 can include any number of shield coils 906a, 906b. Two shield coils 906a, 906b can optionally be configured as a Helmholtz coil pair 908. Any or all of the field coils 902a, 902b and shield coils 906a, 906b can optionally be shaped with a rectangular cross section (as shown in FIG.9), however, the present disclosure contemplates that other shapes of field coils 902a, 902b and shield coils 906a, 906b can be used.

[0052] The field coils 902a, 902b and shield coils 906a, 906b can be arranged annularly as shown in FIG.9 so that a central opening 910 is formed in the magnet 900. The central opening 910 can be configured for a heavy ion or proton beam to be through the central opening 910 directed during proton therapy. As shown in FIG. 9, the central opening 910 can optionally be .7m in diameter or larger than 0.7 m in diameter.

[0053] Any or all of the field coils 902a, 902b and shield coils 906a, 906b can be formed from a material that at least partially includes a medium or high-temperature superconductor. Non- limiting examples of medium-temperature superconductors and high temperature superconductors that can be used in the field coils 902a, 902b and shield coils 906a, 906b include magnesium diboride (MgB2) and / or Niobium-Tin (Nb3Sn).

[0054] Optionally, any or all of the field coils 902a, 902b and shield coils 906a, 906b can be configured to not require cryogenic liquid for cooling due to the use of a medium or high- temperature superconductor. Accordingly, in some implementations the magnet 900 does not require a cryogenic liquid for cooling. Alternatively or additionally, the magnet 900 can optionally produce a field at or above 3 T for MRI imaging.

[0055] In some implementations, the magnet 900 can include a cryostat 912 configured to cool any combination of the field coils 902a, 902b and shield coils 906a, 906b. The cryostat 912 can optionally partially or completely encompass the field coils 902a, 902b and shield coils 906a, 906b.

[0056] In some implementations, system 800 includes a control system 812 that is communicably coupled to MRI assembly 804, e.g., to receive imaging data from MRI assemblyAtty. Dkt. No.103361-624WO1 804 for processing by an MRI imaging system. In some implementations, MRI imaging system automates at least portions of the image-capturing process and / or otherwise controls MRI assembly 804. For example, the MRI imaging system may generally operate MRI assembly 804 in at least two modes, e.g., to capture MRI images of target 810 and surrounding tissues / organs. In the first mode, the proton beam is off, and in the second mode, the proton beam is on and target 810 is being irradiated (e.g., the patient is being treated by MRI assembly 804). The first mode is commonly used to create or adjust a treatment plan for a patient (e.g., adaptive radiation therapy) and / or to setup MRI assembly 804 to treat the patient. The second mode is commonly used for real-time or live monitoring during treatment (e.g., while the proton beam is on). The MRI imaging system can selectively operate system 800 between these two modes. However, it should be understood and appreciated that the MRI imaging system can also be configured to operate in additional modes, and that the specific functionality of the first and second modes described above is not intended to be limiting.

[0057] Control system 812 is also shown to include an automatic quench self-protection and control system, e.g., for safety and protection of MRI assembly 804. Specifically, in some implementations, the automatic quench self-protection and control system is configured to protect MRI assembly 804 from quench and other hazards. In some such implementations, the automatic quench self-protection and control system is a dual-function system that contains two main components: a first component configured to supply currents to the superconductor coils of magnets 806, 808 to create the main magnetic field, and a second component that operates when the MRI is quenching to automatically generate a certain amount of heat based on the parameters set for the MRI running conditions. The second component may further propagate the heat to different regions of the superconducting coils (e.g., of magnets 806, 808) and to bring the superconducting coils from a superconductivity status to a normal status, e.g., to avoid the coils being burned out. In some implementations, the first component is powered by a DC source. In some implementations, the second component is powered by an AC source.

[0058] Generally, control system 812 includes a processor and memory. The processor can be a general-purpose processor, an application-specific integrated circuit (ASIC), one or more field programmable gate arrays (FPGAs), a group of processing components, or other suitable electronic processing structures. In some embodiments, the processor is configured to execute program code stored on the memory to cause control system 812 to perform one or moreAtty. Dkt. No.103361-624WO1 operations, as described below in greater detail. It will be appreciated that, in embodiments where control system 812 is part of another computing device, the components of control system 812 may be shared with, or the same as, the host device. For example, if control system 812 is implemented via a laptop computer, then control system 812 may utilize the processing circuit, processor(s), and / or memory of laptop computer to perform the functions described herein.

[0059] The memory of control system 812 can include one or more devices (e.g., memory units, memory devices, storage devices, etc.) for storing data and / or computer code for completing and / or facilitating the various processes described in the present disclosure. In some embodiments, memory includes tangible (e.g., non-transitory), computer-readable media that stores code or instructions executable by the processor. Tangible, computer-readable media refers to any physical media that is capable of providing data that causes control system 812 to operate in a particular fashion. Example tangible, computer-readable media may include, but is not limited to, volatile media, non-volatile media, removable media and non-removable media implemented in any method or technology for storage of information such as computer readable instructions, data structures, program modules or other data. Accordingly, memory can include RAM, ROM, hard drive storage, temporary storage, non-volatile memory, flash memory, optical memory, or any other suitable memory for storing software objects and / or computer instructions. Memory can include database components, object code components, script components, or any other type of information structure for supporting the various activities and information structures described in the present disclosure. Memory can be communicably connected to the processor, such as via a processing circuit, and can include computer code for executing (e.g., by the processor) one or more processes described herein.

[0060] While described above as individual components, it will be appreciated that the processor and / or memory of control system 812 can be implemented using a variety of different types and quantities of processors and memory. For example, the process may be a single processing device or multiple processing devices. Similarly, memory may be a single memory device or multiple memory devices. Additionally, in some embodiments, control system 812 may be implemented within a single computing device (e.g., one server, one housing, etc.). In other embodiments, control system 812 may be distributed across multiple devices.

[0061] Example 1:Atty. Dkt. No.103361-624WO1

[0062] Implementations of the present disclosure were studied for improving MRI-guided proton therapy using medium and high temperature superconductors. Proton therapy can be superior to photon therapy due to the known phenomenon of the Bragg peak. The Bragg peak refers to limited penetration depth in tissue

[0063] An example proton beam experiment is illustrated in FIG.1 illustrating a proton beam and example tumor. FIG.2 illustrates an example mechanical model of an MRI magnet geometry and stresses, and FIG.3 illustrates an example magnetic field model of the MRI magnet.

[0064] Monte Carlo simulations were performed for proton therapy using the magnets illustrated in FIGS.2 and 3. The simulations illustrated in FIG.4 show the energy deposition in the X-Z plane, and the simulations illustrated in FIG.5 illustrate the deposition of energy on the Z-Axis. Additionally, dose distributions are shown in FIG. 6 for the X and Z axis.

[0065] Additionally, the study included thermal and mechanical modeling of an MRI system that can be used to implement the present disclosure. The study included detailed thermal and mechanical modeling results for a 3 Tesla actively shielded whole body MRI (Magnetic Resonance Imaging) magnet consisting of coils with a square cross section of their windings. The magnet design was a segmented coil type optimized to minimize conductor length while hitting the standard field quality and OSV (Diameter of Spherical Volume) specifications as well as a standard, compact size 3 T system. It had an overall magnet length and conductor length which can lead to conduction cooled designs comparable to NbTi helium bath cooled 3 T MRI magnets. The design had a magnetic field homogeneity better than 10 ppm (part-per-million) within a OSV (Diameter of Spherical Volume) of 48 cm and the total magnet winding length of 1.37 m. A new class of MgB2 strand especially designed for MRI applications was considered as a possible candidate for winding such magnets. The 3 Tesla MRI magnet can operate at 20 K at 67 % of its critical current. Thermal modeling was focused on positioning the cryocooler's cold heads, size of the vacuum jacket as well as quality of thermal insulation with a goal of minimizing thermal gradients in the winding. The optimization process minimized the amount of superconducting material to get a field of 3 T In the middle on the magnet axis at a certain current per tum, with a required homogeneity in the chosen DSV. Constraints of the magnet parameters were specified for each optimization run. The gaps between the individual coils were kept at practical levels (at least on a level of a few mill / meters). Mechanical modeling was usedAtty. Dkt. No.103361-624WO1 to show if they are acceptable. FIG. 7A illustrates example 2D schematics of a mechanical design including a 2D view of Maxwell coil geometry in a plane containing the cylindrical symmetry axis (x) (x = coordinate in axial direction, y = coordinate in radial direction). Coil 1 and Helmholtz coil pairs (Coil 2, Coil 3, and Coil 4) are the field coils (all carrying a positive current). The Helmholtz coil pair (Coil 5) is the shielding coil (carrying a negative current). All the coils have a square winding cross-section and carry the same engineering current density. All the field coils have their winding centers positioned on a common diameter Df. The shield coils have their winding centers on diameter Ds.

[0066] FIG. 7B illustrates example schematics of the mechanical design (x = axis of rotation), including a superconducting winding with copper former, stainless steel cryostat (wall thickness = 1 cm), and G10 fiber-glass epoxy mechanical support (1 cm thick). Magnet length including cryogenics l = 1.4715 m, cryostat OD = 1.9877 m, cryostat warm bore ID = 0.77 m, d = 3 cm, g = 15 cm.

[0067] FIG. 7C illustrates an example cryocooler.

[0068] FIG. 7D illustrates a 3D magnet geometry for an SS cryostat with copper and an MgB2 winding.

[0069] Example parameters of the MRI system include: Operating temperature (K): 10; Current (A per turn): 287.2; Bmax on axis (T): 3.00; Bmax in winding (T): 6.54; Max. radial field component in the winding (T): 5.406; Max. axial field component in the winding (T): 5.4696; Volume of superconductor (m3): 0.26158; Insulated Wire (1.19 x 1.81 mm2) length (km): 109.3; Winding length (m): 1.3667; Min. winding inner radius Ri (m): 0.4375; Max. winding inner radius Ri (m): 0.4714; Winding outer radius Ro (m): 0.9447; Homogeneity (ppm): 10; DSV (m): 0.49; Self-inductance of the winding L (H): 234.2; Number of coils: 9.

[0070] Example geometry parameters of the MRI winding include: I1 (m): 0.0572219; I2 (m): 0.0621271; I3 (m): 0.0776571; I4 (m): 0.1249542; I5 (m): 0.0897296; g12 (m): 0.0724255; g23 (m): 0.1040079; g34 (m): 0.2134943; g55 (m): 1.0102704; Df (m): 1; Ds (m): 1.8;

[0071] FIG. 7E illustrates a winding temperature map after 9.875 days of cooling from 77.3 Kwith thermal expansion (10-times magnified), no magnetic field. T = 4.3 K.Atty. Dkt. No.103361-624WO1

[0072] FIG. 7F illustrates a winding temperature map after 4.917 days of cooling from 77.3 Kwith thermal expansion (10-times magnified), no magnetic field. T = 4.75 K

[0073] FIG. 7G illustrates a winding temperature map after 9.875 days of cooling from 77.3 Kwith thermal expansion (10-times magnified), magnetic field present. T = 4.75 K.

[0074] FIG. 7H illustrates a winding temperature map after 4.917 days of cooling from 77.3 Kwith thermal expansion (10-times magnified), magnetic field present. T = 4.71 K.

[0075] FIG. 7I illustrates a winding temperature map after 4.917 days of cooling from 77.3 K, magnetic field present.

[0076] FIG. 7J illustrates a pair of cryocoolers used in the temperature maps of FIGS.7E-7I.

[0077] FIG. 7K illustrates a cross section of a 1.5 mm x 1.0 mm non-insulated rectangularMgB wire for MRI applications (wire composition: 14.8% MgB , 15% Cu, 35.52% Nb, 34.68%Glidcop Al15.

[0078] FIG. 7L illustrates a map of 10 ppm inhomogeneity in DSV = 0.5 m as a 2D view with a quarter of geometry shown, and the x-coordinate in axial direction, y-coordinate in radial direction.

[0079] The study determined the following main winding parameters for 1 cryocooler after 9.875 days of cooling: Thermal strain max = -0.97235%; min = -0.97968%; Thermal von Mises stress max = 22.136 MPa; min = 0.62321 MPa; Thermal and elmg strain max = -0.96574%; min = -0.98211%; Thermal and elmg von Mises stress max = 50.524 MPa; min = 4.4278 MPa.

[0080] The study determined the following main winding parameters for 2 cryocoolers after 4.917 days of cooling: Thermal strain max = -0.97123%; min = -0.97826%; Thermal von Mises stress max = 22.757 MPa; min = 0.16658 MPa; Thermal and elmg strain max = -0.9647%; min = -0.98137%; Thermal and elmg von Mises stress max = 50.233 MPa; min = 4.236 MPa.

[0081] Finally, the only elmg strain in winding: max = 0.0127%; min = -5.8053e-3%; and the only elmg von Mises stress in winding: max = 35.5 MPa; min = 2.8 MPa.

[0082] Example 2:

[0083] In conventional radiation (photon) therapy the absorption of rays or gamma rays is intensive at the surface and decreases with distance into the subject. Thus in traditional therapy aAtty. Dkt. No.103361-624WO1 high intensity of photons must be administered using the isocentric convergence technique to allow multiple beams with the diminished intensity focus at the tumor site. As a result, much of the photon's energy causes damage to healthy tissue.

[0084] In contrast to photons, charged particles like protons are "silver bullets" whose interactions with matter are characterized by the Bragg curve. The energy loss by charged particles through matter is described by this curve which rises to a maximum (the Bragg peak) just before the end of the particle's track. The peak occurs because the cross section for particle- matter interaction increases just before the particle comes to rest.

[0085] As protons pass through tissue at velocity they continuously lose kinetic energy by inelastic Coulomb scattering against atomic electrons. The rate of energy loss, being proportionalto 1 / , rises sharply as the protons slow down, the end of their range, and form the Bragg peak.Protons also undergo repulsive non-elastic interactions with atomic nuclei and will be deflected from their original path. The product of such interactions may be secondary protons, heavier ions, neutrons, and gamma rays. These non-elastic proton-nucleus interactions, although less frequent than proton-electron ones have a stronger effect (9).

[0086] In particle therapy the beam energy is adjusted either electrically (synchrotron accelerators) or by filters (cyclotron accelerators) to ensure that the Bragg peak occurs at the tumor site. This indicates that while photon therapy is characterized by relatively high entrance and exit doses, proton therapy has not only a lower entrance dose but a negligible exit dose. It also shows that the position of the Bragg peak can be adjusted to ensure that the tumor receives the intended radiation. This allows patients to receive high radiation doses with low risk of collateral tissue damage.

[0087] To perform efficient treatments, accurate image guidance techniques that can distinguish a tumor from healthy tissue are required. Imaging is associated with target accuracy, of targeting, dosimetry, treatment outcome assessment, and prognosis. Images are used in radiation therapy during the following processes: (1) offline treatment planning - tumors and other anatomical structures are identified and the treatment is simulated on a treatment planning computer. During this process CT, MRI and even PET-CT or other type of images are commonly used; (2) imaging is applied to set up the patient in the treatment room and to convert the virtual treatment (a treatment plan) into the real treatment; (3) images are used to assess the treatmentAtty. Dkt. No.103361-624WO1 outcome: effectiveness, toxicities, and prognosis, etc. While Processes 1 and 3 are usually carried out offline, Process 2 is often performed with the imaging device attached to the treatment machine inside the treatment room to align the patient in the treatment position or during the treatment to monitor the targeting of tumors and organs. Process 2 is complex and needs to conform to the gantry and the beam delivery system. The currently available image- guided particle therapy systems are mainly based on 2D orthogonal X-ray imaging, in-room 3D computed tomography (CT) or on-board cone-beam CT (CBCT) imaging. However, 3D or 4D CT and MRI images are commonly used off-line for treatment planning (57). Imagingtechnologies such as the proton Computed Tomography (or pCT) 58, 59 or MRguided protontherapy have been proposed but are not yet available for clinical use (60). While X-ray based imaging has difficulty resolving soft-tissue the converse is true for MRIguided particle therapy. MRI guidance in photon (radiation) therapy MRgRT, which entered clinical practice not long ago, gradually became popular during the past decade (61). Currently there are several commercial available MRgRT manufacturers in the market: ViewRay (62), Elekta (63), and Varian Medical Systems / IMRIS (64).

[0088] In order to take advantage of the Bragg peak effect, the exact depth of the tumor (the target) can be used to plan the treatment. An error in the target depth of a few mm that may result in only a few percent change in the photon dose may lead to a 100% change in the proton dose (67). This emphasizes the need for direct visualization of the tumor position.

[0089] Implementations of the present disclosure enable MRI guidance to take advantage of proton over photon therapy. As described in Example 1, herein, a comprehensive and sophisticated design of accelerator, beam-line, gantry, magnets, and the imaging components can be calculated so that all sub-systems could be integrated into a unit that operates efficiently and meets the clinical requirements while reducing complexity of the resulting system.

[0090] Implementations of the present disclosure allow for the use of MRI image guidance to be combined with proton therapy. To take advantage of the Bragg peak effect the exact targeting of the tumor and positioning of the patient must figure into the treatment. The charged-particle beam has to be guided by image visualization using X-ray, CT, and / OR MRI. An error in the target depth of a few mm that may result in only a few percent change in photon dose may lead to a 100% change in the proton dose (67). Unlike MRI-guided photon therapy the directAtty. Dkt. No.103361-624WO1 interaction of the magnetic field with the charged particle beam presents a challenge such that MRI image-guided proton / particle therapy has not yet been available in clinical practice.

[0091] The use of medium and high temperature superconductors allows for improvements to rotating gantry magnets in proton therapy systems. Cooling of superconducting rotating gantry magnets is an important engineering task where cooling by liquid helium may not be feasible. Instead, some form of liquid-cryogen-free or conduction cooling can be used, as described with reference to FIGS.8 and 9 herein.

[0092] Overview of an example Proton Therapy System

[0093] Beam Energy Adjustment

[0094] Before it enters the treatment area or gantry the ion beam requires energy adjustment. Therapy cyclotrons generate a fixed-energy beam of typically 230 MeV, the needed energy variation between 60 and 230 MeV being achieved by passing the beam through an energy degrader, an absorber of variable thickness such as two opposite-facing wedges of graphite. These are followed by a magnetic analyzer consisting of a combination of dipole magnets and collimators (20). The "Energy Selection System" of Ion Beam Applications S.A. (IBA) allows tuning from 60-230MeV in less than 1 second. Although the beam loss by degradation can be as much as a factor of one hundred or more typical cyclotrons deliver sufficient beam intensity to make up for this. With regard to synchrotron sources, rapid energy variation can be achieved by extracting the ion beam at different times in the acceleration cycle (20). The synchrotron can deliver proton beams in more than 90 energy steps between 73 MeV and 222 MeV corresponding to penetration depths in water of 4 to 31 cm (29).

[0095] Gantry

[0096] A gantry is used to rigidly hold the guidance magnets of ion beam therapy. The gantry can be a massive structure. As used herein, the term “gantry” can further refer to a magnet string (5) or an entire system for guiding the ion beam therapy.

[0097] Gantry Cooling

[0098] The accelerator magnets of high energy physics are wound with high-current-carrying Rutherford cable to maintain a high ampere-turn ratio while minimizing the magnet inductance. Different considerations govern the choice of magnet design and conductor size inAtty. Dkt. No.103361-624WO1 superconducting particle beam gantries. Inductance is not an issue but magnet current has to be relatively small to minimize current-lead heat leak into the cryostat. Since the magnet heat load scales with ramp rate a large momentum acceptance (especially if it allows fixed-field operation) is beneficial for thermal design (40). Because the gantry can rotate, magnet cooling by liquid helium is not feasible. Instead all superconducting gantry systems are cryogen-free, conduction cooled by way of high conductivity links to cryocooler cold heads. Numerous cold heads may be connected in parallel to various parts of the magnet system (35). A lighter gantry could be enabled by circulating gaseous or supercritical helium through the magnets from an off-gantry- mounted cooling and pumping system (51-54) or by implementing a cryogenic oscillating heat pipe system (55).

[0099] Components of the Magnet String and Gantry

[0100] In an example ion therapy gantry (30) the beam is bent by three dipole magnets BM 1, BM 2, and BM 3, and kept in focus by six quadrupole magnets Q1 to Q6. The beam enters the gantry through a monitor PRM that monitors and records the center of the beam and its size. Three other PRMs are associated with steering magnets for beam trajectory correction and two more, PRM1 and PRM2, are located in the nozzle. Several nozzle arrangements have been used to passively or actively spread the mm -size beam over a treatment area that may be as large as30 cm 5,34 . Described in detail by (31) are: (i) uniform scanning nozzles, (ii) pencil scanningnozzles, (iii) single scattering nozzles, and (iv) double scattering nozzles.

[0101] Bending Dipoles: Momentum Acceptance

[0102] During a typical treatment session a proton beam energy may need to be varied from 70 to 250 MeV. This produces a change in rigidity of from 1.231 to 2.431 Tm requiring the field ofa 1 m radius dipole to track the beam energy and hence increase from 1.23 to 2.43 T . In theabsence of field-change a typical normal dipole will accept a beam momentum change, / , ofless than 1% ( (35), p.2). Tumors are typically scanned in layers 5 mm deep enabled by amomentum sweep of 1% ( (35), p.4), the layerscan taking of order 100 ms . A / of 1%corresponds to a beam energy change / 2 / 2%. Thus to control a 100 MeV beam(proton magnetic rigidity 1.483 Tm ) a 1 m-radius dipole would require a field change of 15mT . On this basis the average sweep rate of the scan is 150mT / s ( 20 times faster than LHC).Taking another approach, a / of 2% implies 50 scanning steps covering the energy rangeAtty. Dkt. No.103361-624WO1 from 70 to 250 MeV and a 1 -m-radius dipole field range of 1.23 to 2.43 T . At 100 ms / step this leads to an average sweep rate of 240 mT / s. From another standpoint, at 5 mm per layer a proton-beam-depth range of 10-35 cm would require 50 scanning steps, leading to the same result.

[0103] Sweep rates of hundreds of mT / s cause stability and energyloss problems for superconducting magnets. So when considering a gantry upgrade from normal-conducting to superconducting magnets these high ramp rates created a demand for achromatic bending dipoleswith wide momentum acceptances, such / 5 10%, 25%, and 40 50% [(35), p.3]. Theuse of such magnets, which enable the entire energy range to be covered in just a few steps, eases the ramp-rate requirement.

[0104] Normal-Conducting Gantry Magnets

[0105] As outlined in (20) the use of normal-conducting iron-core dipole magnets with their maximum bore field, , of about 1.8 T governs the size of most commercial gantries. At a typical 250 MeV proton beam rigidity, 2.43Tm the 1.8 T field calls for a bending radius / 1.35 m. The addition of 3.5 m for distance from the nozzle entrance to the isocenterleads to a gantry rotation radius of about 5 m. Thus proton gantries are about 10 12 m indiameter, 7 10 m long (18) and weigh about 100-200 tons (20).

[0106] Conventionally Wound Magnets

[0107] Superconductivity can enable several improvements to gantry and magnet design. (18) considered the size advantage accompanying the substitution of 2 T normal-conducting dipoles with superconducting ones with 4 T or higher fields. The smaller bend radius and smaller size were estimated to reduce the gantry diameter by about 3 m. Alonso et al. have also described a gantry based on achromatic ( 9% momentum acceptance) combined-function magnets. In thatarrangement a 60 bend was followed by a set of magnets (possibly a pair) contributing to a final150 bend. The total mass of the magnets was estimated to be less than 5 tons. The "directreplacing" of normal-conducting magnets with superconducting ones should not only lower the weight but also the cost of the gantry. Combined-function magnets possess much larger momentum acceptances than do individual dipoles and quadrupoles. In this case a fixed-field beam energy variation of 50 MeV could be accommodated, enabling an energy range of 100 toAtty. Dkt. No.103361-624WO1 250 MeV to be covered in only three steps. The superconducting gantry design study of Wan etal. (38) achieved an energy acceptance of 21% enabling beams of 150 30MeV to be bentwithout field ramping. Wan et al's compact fixed-field-alternating-gradient (FFAG) gantry, or magnet string, consisted of three groups of seven achromatic 3.2 T superconducting magnets,each group creating a 90 bend (38).

[0108] An overlaid pair of solenoidal coils tilted in opposite directions can generate a dipolar field normal to the solenoidal axis (39). Exploited in recent years by the superconducting magnet group at the Lawrence Berkeley National Laboratory (LBNL), this tilted-double-helix magnet is referred to as canted cosine theta, CCT. By suitably modifying the winding scheme quadrupoles and higher order multipoles can also be generated (18). Not only that, but a combined-function magnet can be produced by winding a CCT dipole on top of a CCT quadrupole (40). In pursuing CCT technology the LBNL group went on to introduce a new magnet concept - the alternating-gradient canted cosine theta (AGCCT) magnet 35,40 . In this magnet the inner quadrupolewinding consists of multiple sections with the current reversed between sections, e.g.5 sections for a winding designated FDFDF. LBNL's design can be characterized as fixed-field alternating- gradient since its large momentum acceptance ( 25% ) enables a large energy range to be transmitted without changing the field. The advantages claimed for gantries incorporating the new magnet design are: (i) an order of magnitude reduction in weight, (ii) possible size reduction, e.g. diameter 5 m and length 8.3 m, and (iii) rapid scanning (35). Gantry layouts described by the LBNL group are: (i) three groups of seven achromatic 3.2 T superconducting magnets (with sextupole and octupole components in the middle five of each group), with eachgroup creating a 90 bend (38), (ii) two 75 AGCCT bending magnet groups to offset the beam2.5 m "above" the exocenter followed by one 90 AGCCT group (35), (iii) one normal-conducting 45 dipole to deflect the beam "upwards" followed by a single 135 bend to guide thebeam back to the exocenter (40). As with other superconducting gantry systems the magnets are intended to be conduction cooled.

[0109] The construction and arrangement of the systems and methods as shown in the various implementations are illustrative only. Although only a few implementations have been described in detail in this disclosure, many modifications are possible (e.g., variations in sizes, dimensions, structures, shapes, and proportions of the various elements, values of parameters, mountingAtty. Dkt. No.103361-624WO1 arrangements, use of materials, colors, orientations, etc.). For example, the position of elements may be reversed or otherwise varied, and the nature or number of discrete elements or positions may be altered or varied. Accordingly, all such modifications are intended to be included within the scope of the present disclosure. The order or sequence of any process or method steps may be varied or re-sequenced according to alternative implementations. Other substitutions, modifications, changes, and omissions may be made in the design, operating conditions, and arrangement of the implementations without departing from the scope of the present disclosure.

[0110] The present disclosure contemplates methods, systems, and program products on any machine-readable media for accomplishing various operations. The implementations of the present disclosure may be implemented using existing computer processors, or by a special purpose computer processor for an appropriate system, incorporated for this or another purpose, or by a hardwired system. Implementations within the scope of the present disclosure include program products including machine-readable media for carrying or having machine-executable instructions or data structures stored thereon. Such machine-readable media can be any available media that can be accessed by a general purpose or special purpose computer or other machine with a processor. By way of example, such machine-readable media can comprise RAM, ROM, EPROM, EEPROM, CD-ROM or other optical disk storage, magnetic disk storage or other magnetic storage devices, or any other medium which can be used to carry or store desired program code in the form of machine-executable instructions or data structures, and which can be accessed by a general purpose or special purpose computer or other machine with a processor.

[0111] When information is transferred or provided over a network or another communications connection (either hardwired, wireless, or a combination of hardwired or wireless) to a machine, the machine properly views the connection as a machine-readable medium. Thus, any such connection is properly termed a machine-readable medium. Combinations of the above are also included within the scope of machine-readable media. Machine-executable instructions include, for example, instructions and data which cause a general-purpose computer, special purpose computer, or special purpose processing machines to perform a certain function or group of functions.

[0112] Although the figures show a specific order of method steps, the order of the steps may differ from what is depicted. Also, two or more steps may be performed concurrently or withAtty. Dkt. No.103361-624WO1 partial concurrence. Such variation will depend on the software and hardware systems chosen and on designer choice. All such variations are within the scope of the disclosure. Likewise, software implementations could be accomplished with standard programming techniques with rule-based logic and other logic to accomplish the various connection steps, processing steps, comparison steps and decision steps.

[0113] It is to be understood that the methods and systems are not limited to specific synthetic methods, specific components, or to particular compositions. It is also to be understood that the terminology used herein is for the purpose of describing particular implementations only and is not intended to be limiting.

[0114] As used in the specification and the appended claims, the singular forms “a,” “an” and “the” include plural referents unless the context clearly dictates otherwise. Ranges may be expressed herein as from “about” one particular value, and / or to “about” another particular value. When such a range is expressed, another implementation includes from the one particular value and / or to the other particular value. Similarly, when values are expressed as approximations, by use of the antecedent “about,” it will be understood that the particular value forms another implementation. It will be further understood that the endpoints of each of the ranges are significant both in relation to the other endpoint, and independently of the other endpoint.

[0115] “Optional” or “optionally” means that the subsequently described event or circumstance may or may not occur, and that the description includes instances where said event or circumstance occurs and instances where it does not.

[0116] Throughout the description and claims of this specification, the word “comprise” and variations of the word, such as “comprising” and “comprises,” means “including but not limited to,” and is not intended to exclude, for example, other additives, components, integers or steps. “Exemplary” means “an example of” and is not intended to convey an indication of a preferred or ideal implementation. “Such as” is not used in a restrictive sense, but for explanatory purposes.

[0117] Disclosed are components that can be used to perform the disclosed methods and systems. These and other components are disclosed herein, and it is understood that when combinations, subsets, interactions, groups, etc., of these components are disclosed, while specific reference of each various individual and collective combinations and permutation ofAtty. Dkt. No.103361-624WO1 these may not be explicitly disclosed, each is specifically contemplated and described herein, for all methods and systems. This applies to all aspects of this application including, but not limited to, steps in disclosed methods. Thus, if there are a variety of additional steps that can be performed it is understood that each of these additional steps can be performed with any specific implementation or combination of implementations of the disclosed methods.

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Claims

Atty. Dkt. No.103361-624WO1 WHAT IS CLAIMED IS:

1. A system for magnetic resonance imaging (MRI) guided proton therapy, the system comprising: a source configured to generate and emit a heavy ion or proton beam; and an MRI assembly comprising at least one segmented coil magnet comprising a plurality of windings, wherein plurality of windings of the at least one segmented coil magnet comprise a medium or high-temperature superconductor such that the at least one segmented coil magnet does not require a cryogenic liquid for cooling.

2. The system of claim 1, wherein the medium or high-temperature superconductor is magnesium diboride (MgB2).

3. The system of claim 1, wherein the medium or high-temperature superconductor is Niobium-Tin (Nb3Sn).

4. The system of claim 1, wherein the windings of the at least one segmented coil magnet have a rectangular cross-section.

5. The system of claim 1, wherein the at least one segmented coil magnet is annularly shaped such that a central opening is formed in the at least one segmented coil magnet, wherein the proton beam is directed through the central opening toward a target.

6. The system of claim 5, wherein the central opening is larger than 0.7 m in diameter.

7. The system of claim 1, further comprising one or more guidance magnets configured to manipulate a direction of the proton beam after emission from the source.

8. The system of claim 1, wherein the at least one segmented coil magnet is configured to operate at or above 3 T.

9. The system of claim 1, wherein the at least one segmented coil magnet comprises a cryostat that at least partially encompasses the windings.Atty. Dkt. No.103361-624WO1 10. The system of claim 1, wherein the windings of the at least one segmented coil magnet comprise: (i) a first coil, (ii) two or more first Helmholtz coil pairs, and (iii) at least one second Helmholtz coil pair, wherein the first coil and the two or more first Helmholtz coil pairs are configured as field coils, and wherein the at least one second Helmholtz coil pair is configured as a shield coil.

11. The system of claim 1, further comprising a controller configured to receive imaging data from the MRI assembly.

12. The system of claim 11, wherein the controller is further configured to determine a position of a tumor in a subject based on the imaging data for guidance of the heavy ion or proton beam.

13. A magnet for a magnetic resonance imaging (MRI) system to be used in proton therapy, the magnet comprising: a plurality of field coils comprising a plurality of windings, wherein at least two of the plurality of field coils are configured as a Helmholtz coil pair; a plurality of shield coils, wherein at least two of the plurality of shield coils are configured as a Helmholtz coil pair, wherein the plurality of field coils and the plurality of shield coils are formed of a material that at least partially comprises a medium or high-temperature superconductor, such that the magnet does not require a cryogenic liquid for cooling, and wherein the plurality of field coils and the plurality of shield coils are arranged annularly such that a central opening is formed in magnet through which a heavy ion or proton beam is directed during proton therapy.

14. The magnet of claim 13, wherein the medium or high-temperature superconductor is magnesium diboride (MgB2).

15. The magnet of claim 13, wherein the medium or high-temperature superconductor is Niobium-Tin (Nb3Sn).

16. The magnet of claim 13, wherein the plurality of windings of the plurality of field coils have a rectangular cross-section.Atty. Dkt. No.103361-624WO1 17. The magnet of claim 13, wherein the central opening is larger than 0.7 m in diameter.

18. The magnet of claim 13, wherein the magnet is configured to operate at or above 3 T.

19. The magnet of claim 13, further comprising a cryostat that at least partially encompasses the plurality of field coils and the plurality of shield coils.

20. The magnet of claim 13, wherein at least one of the field coils or at least one of the shield coils comprises a segmented coil.

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