Magnetic resonance guided charged particle beam radiotherapy system

Abstract

Provided in the present invention are a magnetic resonance guided charged particle beam radiotherapy apparatus and an electromagnetic navigation parameter debugging method for target point coordinates, belonging to the technical field of radiotherapy apparatuses. The present invention comprises a charged particle beam electromagnetic navigation apparatus. The charged particle beam electromagnetic navigation apparatus comprises a multi-stage electromagnetic navigation coil, a charged particle beam sequentially passing through the multi-stage electromagnetic navigation coil to generate a parallel charged particle beam with a direction consistent with the direction of a magnetic resonance main magnetic field; and an MR imaging apparatus, comprising an MR main magnetic field coil upper group and an MR main magnetic field coil lower group, an imaging therapy area being formed therebetween. The electromagnetic navigation parameter debugging method of the present invention comprises constructing a target cost function, optimizing electromagnetic navigation parameters based on a gradient information-free optimization algorithm, and determining and outputting optimal navigation parameters. According to the present invention, the charged particle beam parallel to the MR main magnetic field can be generated, thereby greatly reducing the influence of Lorentz force on the charged particle beam, enabling to reduce the positioning difficulty of the charged particle beam on a tumor target area under MR guidance, and improving the therapeutic accuracy.

Description

A magnetic resonance-guided charged particle beam radiotherapy system Technical Field

[0001] The present invention relates to the technical field of radiotherapy devices, and in particular to a magnetic resonance-guided charged particle beam radiotherapy system, comprising a magnetic resonance-guided parallel charged particle beam radiotherapy device and an electromagnetic navigation parameter debugging method for target coordinates. Background Art

[0002] Cancer is a leading cause of death in humans, and more than half of cancer patients require radiation therapy (RT). Compared to conventional cone-beam CT-guided RT, magnetic resonance (MR)-guided RT eliminates the need for imaging radiation doses. MRI imaging can be performed before or during treatment, improving the resolution of soft tissues such as tumors and resulting in higher treatment accuracy, which can improve patient outcomes and has significant clinical significance.

[0003] Currently, all MR-guided radiotherapy technologies used clinically are photon radiotherapy, such as MRIdian from ViewRay (USA) and Unity from Elekta (Sweden). Charged-particle radiotherapy offers certain advantages over photon radiotherapy, such as insensitivity to tissue inhomogeneities and a steeper longitudinal dose drop-off. However, MR-guided charged-particle radiotherapy equipment is currently unavailable in clinical practice. The Lorentz force acting on the charged-particle beam during transport within the main MR magnetic field can cause the beam to deviate from its original direction. This can make it difficult to accurately target the tumor, leading to missed tumors and inadvertent irradiation of normal tissue, ultimately compromising efficacy.

[0004] Patent publication number CN111580030A discloses an MR-guided radiotherapy system in which the MR main magnetic field is perpendicular to the direction of the incident particles. While this invention utilizes niobium-titanium alloy in the linear accelerator region to shield the magnetic field, which can reduce beam deflection outside the patient's body, it cannot address the deflection of the incident particle beam inside the patient's body. Both China United Imaging and the University of Wollongong in Australia have designed radiotherapy systems in which the MR main magnetic field is parallel to the central axis of the beam for MR-guided photon radiotherapy. However, this approach is not suitable for charged particle radiotherapy systems. In this approach, only the beam on the central axis is parallel to the MR main magnetic field; beams that deviate from the central axis diverge in a conical shape. For charged particle beams, beams that deviate from the central axis are still subject to the Lorentz force, causing the beam path to exhibit a spiral motion, which is not conducive to precise beam positioning. To address the beam positioning problem of charged particle beams within the main magnetic field of an MRI machine, the University of Wollongong's Medical Radiation Physics Center proposed a complex two-step beam parameter correction method. First, a first-order approximation for each beam energy and target point is obtained based on simulation and tabulation of a uniform water phantom. Then, a Monte Carlo simulation is performed on a specific patient to obtain a large amount of beam parameter and corresponding target position data. A neural network model is trained based on this data to refine the first-order approximation. This method is overly complex, requiring tabulation and neural network training for each beam energy and specific patient, making it cumbersome and inconvenient for clinical translation.

[0005] Summary of the Invention

[0006] In view of this, in order to solve the technical problems in the prior art that the trajectory of the charged particle beam is deflected, resulting in difficulty in accurately locating the tumor target area with the beam, resulting in missed or mis-irradiated areas, and a large amount of scattering of the charged particle beam during transportation, which increases the lateral penumbra and is not conducive to the protection of normal tissue, on the one hand, the present invention provides a magnetic resonance-guided charged particle beam radiotherapy device, which can generate a charged particle beam parallel to the MR main magnetic field, greatly reducing the influence of the Lorentz force on the charged particle beam, and can reduce the difficulty of MR-guided charged particle beam positioning on the tumor target area and improve the treatment accuracy; using the MR imaging main magnetic field, lateral constraint of the charged particle beam is achieved, and its lateral dose drop gradient is improved, which can better protect the normal tissue around the tumor and reduce the probability of radiotherapy complications.

[0007] To achieve the above object, the present invention provides the following technical solutions:

[0008] A magnetic resonance guided charged particle beam radiotherapy device capable of generating a charged particle beam parallel to the MR main magnetic field, comprising:

[0009] A charged particle beam generator, which is used to generate and accelerate a charged particle beam so that its energy reaches the energy range required for radiotherapy;

[0010] A charged particle beam electromagnetic navigation device, comprising a multi-stage electromagnetic navigation coil, wherein the charged particle beam sequentially passes through the multi-stage electromagnetic navigation coil to generate a parallel charged particle beam consistent with the direction of the main magnetic field of the magnetic resonance; an MR imaging device, comprising:

[0011] An MR main magnetic field coil, wherein the central axis thereof coincides with the central axis of the parallel charged particle beam, comprises an upper group of MR main magnetic field coils and a lower group of MR main magnetic field coils, wherein an imaging treatment region is formed between the upper group of MR main magnetic field coils and the lower group of MR main magnetic field coils, wherein the upper group of MR main magnetic field coils and the lower group of MR main magnetic field coils each comprise an imaging coil and a shielding coil, wherein the imaging coil is close to the imaging treatment region and the shielding coil is far from the imaging treatment region, and current directions of the imaging coil and the shielding coil are opposite;

[0012] Gradient magnetic field coils are used to generate spatial magnetic field gradient changes to facilitate imaging layer selection and positioning;

[0013] Body coil, used to generate radiofrequency fields to excite proton spins in tissues;

[0014] The forward coil and the backward coil are used to receive the echo signal and complete the reconstruction.

[0015] Preferably, the multi-stage electromagnetic navigation coil is a three-stage electromagnetic navigation coil arranged sequentially from top to bottom;

[0016] The first-stage electromagnetic navigation coil deflects the charged particle beam toward the target. The second-stage electromagnetic navigation coil, with its current in the opposite direction to the first-stage coil, corrects the deflection angle of the charged particle beam. The third-stage electromagnetic navigation coil further adjusts the incident direction of the charged particle beam, compensating for the effects of the MR fringe magnetic field on the incident charged particle beam and aligning the incident direction of the charged particle beam with the main magnetic field in the MR imaging area.

[0017] Preferably, the acceleration device is a linear accelerator, a cyclotron or a synchrotron.

[0018] Preferably, the interval between the upper group of MR main magnetic field coils and the lower group of MR main magnetic field coils is greater than 60 cm.

[0019] Preferably, the charged particles in the charged particle beam are electrons, protons, carbon ions or helium ions.

[0020] In another aspect, the present invention provides a method for debugging electromagnetic navigation parameters of the magnetic resonance-guided charged particle beam radiotherapy device according to target coordinates, comprising the following steps:

[0021] Construction of target cost function

[0022] The target cost function is constructed by simulating the particle beam phase space corresponding to the electromagnetic navigation parameters based on the particle transport algorithm and taking the difference between the ideal particle beam phase space bombarding the target point and the simulated particle beam phase space based on the electromagnetic navigation parameters of the current iteration. It includes the location cost function and direction cost function

[0023] Where λ is the direction cost function The weight coefficients are x0 and y0, respectively, for the horizontal coordinate positions of the target point. and is the average transverse coordinate of all particles in the simulated particle beam based on the electromagnetic navigation parameters of the current iteration, is the average value of the angle between the incident direction of all particles in the particle beam simulated based on the electromagnetic navigation parameters of the current iteration and the MR main magnetic field, and Electromagnetic navigation parameters Decide;

[0024] parameter and the target cost function There is a functional relationship between The electromagnetic navigation parameter debugging task can be described as formula (4):

[0025] in, is the electromagnetic navigation parameter, N is the number of electromagnetic navigation coils;

[0026] Optimize electromagnetic navigation parameters based on the optimization algorithm without gradient information;

[0027] The optimization process does not use an objective cost function Electromagnetic navigation parameters Instead of gradient information, it is randomly modified based on historical optimization results. And calculate the corresponding Search or Explore Optimization space;

[0028] Determine and output optimal navigation parameters

[0029] After reaching the preset number of iterations, Minimum electromagnetic navigation parameter As the optimal, the optimal electromagnetic navigation parameters of the target are output.

[0030] Preferably, the electromagnetic navigation parameters Including the current direction and intensity of the multi-level electromagnetic navigation coil; the electromagnetic navigation parameters Vectorized coding, in and They are the normalized currents of the two pairs of orthogonal magnetic field coils in the first-stage electromagnetic navigation coil, normalized by the corresponding maximum load current, with values ​​between -1.0 and +1.0, where positive and negative indicate the direction of the current. and and represent the normalized currents in the second and third electromagnetic navigation coils, respectively.

[0031] Preferably, λ is 5.0-15.0.

[0032] Preferably, the optimization algorithm without gradient information is a Bayesian black box optimization algorithm or a meta-heuristic algorithm.

[0033] Preferably, the preset number of iterations is more than 300 times.

[0034] Compared with the prior art, the present invention has the following beneficial effects:

[0035] The magnetic resonance-guided charged particle beam radiotherapy device provided by the present invention can generate a charged particle beam parallel to the MR main magnetic field, greatly reducing the influence of the Lorentz force on the charged particle beam, reducing the difficulty of positioning the charged particle beam on the tumor target area under MR guidance, and improving the accuracy of treatment; utilizing the MR imaging main magnetic field to achieve lateral constraint of the charged particle beam, improve its lateral dose drop gradient, better protect the normal tissue around the tumor, and reduce the probability of radiotherapy complications. BRIEF DESCRIPTION OF THE DRAWINGS

[0036] FIG1 is a perspective view of a magnetic resonance-guided charged particle beam radiotherapy device provided by the present invention;

[0037] FIG2 is a front view of the magnetic resonance-guided charged particle beam radiotherapy device provided by the present invention;

[0038] FIG3 is a diagram of magnetic field distribution calculated using finite element simulation;

[0039] In the figure, 1. vacuum tube, 2. first-stage electromagnetic navigation coil, 3. second-stage electromagnetic navigation coil, 4. third-stage electromagnetic navigation coil, 5. upper group of MR main magnetic field coils, 6. lower group of MR main magnetic field coils. DETAILED DESCRIPTION

[0040] The technical solutions in the embodiments of the present invention will be clearly and completely described below in conjunction with the drawings in the embodiments of the present invention; it is obvious that the described embodiments are only part of the embodiments of the present invention, rather than all the embodiments. Based on the embodiments of the present invention, all other embodiments obtained by ordinary technicians in this field without making creative work are within the scope of protection of the present invention.

[0041] In the description of the present invention, it should be noted that the terms "upper," "lower," "inner," "outer," "top / bottom," and the like, indicating orientations or positional relationships, are based on the orientations or positional relationships shown in the accompanying drawings and are intended solely to facilitate and simplify the description of the present invention. They are not intended to indicate or imply that the devices or components referred to must have, be constructed, or operate in a specific orientation, and therefore should not be construed as limitations on the present invention. Furthermore, the terms "first" and "second" are used for descriptive purposes only and should not be construed as indicating or implying relative importance.

[0042] In the description of the present invention, it should be noted that, unless otherwise expressly specified or limited, the terms "installed," "provided with," "mounted / connected," and "connected" should be understood in a broad sense. For example, "connected" can mean a fixed connection, a detachable connection, or an integral connection; it can be a mechanical connection or an electrical connection; it can be a direct connection or an indirect connection through an intermediate medium, and it can be internal communication between two components. Those skilled in the art will be able to understand the specific meanings of the above terms in the present invention in specific circumstances.

[0043] The following problems exist based on the existing MR-guided radiotherapy technology:

[0044] 1) Charged particles are significantly affected by the main and marginal magnetic fields of the MR system, causing the beam trajectory to deflect, making it difficult to accurately locate the tumor target, resulting in missed or mis-irradiated areas.

[0045] 2) Compared with photons, charged particles are more likely to be laterally scattered by air or human tissue. In particular, for charged particles with lighter mass, such as electrons, a large amount of scattering will occur during the transport process, and the lateral penumbra will increase, which is not conducive to the protection of normal tissues.

[0046] The present invention provides a magnetic resonance-guided charged particle beam radiotherapy device, which can generate a charged particle beam parallel to the MR main magnetic field, and includes:

[0047] A charged particle beam generator, which is used to generate and accelerate a charged particle beam so that its energy reaches the energy range required for radiotherapy;

[0048] A charged particle beam electromagnetic navigation device includes a multi-stage electromagnetic navigation coil. The charged particle beam sequentially passes through the multi-stage electromagnetic navigation coil to generate a parallel charged particle beam aligned with the main magnetic field of the magnetic resonance imaging (MRI). Unlike conventional radiotherapy charged particle beams, which diverge and are transported at an off-center position, the parallel charged particle beam always aligns with the main magnetic field of the MRI, regardless of whether the beam is on the isocentric axis or off the central axis.

[0049] An MR imaging apparatus comprising:

[0050] An MR main magnetic field coil, wherein the central axis thereof coincides with the central axis of the parallel charged particle beam, comprises an upper group of MR main magnetic field coils and a lower group of MR main magnetic field coils, wherein an imaging treatment region is formed between the upper group of MR main magnetic field coils and the lower group of MR main magnetic field coils, wherein the upper group of MR main magnetic field coils and the lower group of MR main magnetic field coils each comprise an imaging coil and a shielding coil, wherein the imaging coil is close to the imaging treatment region and the shielding coil is far from the imaging treatment region, and current directions of the imaging coil and the shielding coil are opposite;

[0051] Gradient magnetic field coils are used to generate spatial magnetic field gradient changes to facilitate imaging layer selection and positioning;

[0052] Body coil, used to generate radiofrequency fields to excite proton spins in tissues;

[0053] The forward coil and the backward coil are used to receive the echo signal and complete the reconstruction.

[0054] Specifically, as shown in FIG1-2 , the present invention exemplarily provides a magnetic resonance-guided charged particle beam radiotherapy device. A three-stage multi-stage electromagnetic navigation coil is used as an example for explanation. The magnetic resonance-guided charged particle beam radiotherapy device includes:

[0055] A vacuum tube 1 is provided with a first-stage electromagnetic navigation coil 2, a second-stage electromagnetic navigation coil 3 and a third-stage electromagnetic navigation coil 4 in sequence from top to bottom;

[0056] The MR main magnetic field coil upper group 5 and the MR main magnetic field coil lower group 6 are spaced apart from top to bottom above the third-level electromagnetic navigation coil 4 (because the third-level electromagnetic navigation coil 4 needs to correct the influence of the fringe magnetic field generated by the main magnetic field coil upper group 5 on the beam), and the MR main magnetic field coil upper group 5 is sleeved on the third-level electromagnetic navigation coil 4, and the area between the MR main magnetic field coil upper group 5 and the MR main magnetic field coil lower group 6 forms an imaging treatment area.

[0057] The charged particle beam generating device (not shown in the figure) generates a charged particle beam, which is accelerated by the accelerator and taken in along the extension direction of the vacuum tube to form an image in the imaging treatment area.

[0058] The device can greatly reduce the deflection or distortion of the incident trajectory of the charged particle beam caused by the MR main magnetic field and fringe field, while constraining the lateral scattering of the charged particle beam, reducing the lateral penumbra, achieving precise positioning of the charged particle beam under MR guidance, and improving the accuracy of radiotherapy.

[0059] In the present invention, the multi-stage electromagnetic navigation coil is a three-stage electromagnetic navigation coil arranged in sequence from top to bottom;

[0060] The first-stage electromagnetic navigation coil 2 deflects the charged particle beam toward the target point. The current direction of the second-stage electromagnetic navigation coil 3 is opposite to that of the first-stage electromagnetic navigation coil, correcting the deflection angle of the charged particle beam. The third-stage electromagnetic navigation coil 4 further adjusts the incident direction of the charged particle beam to compensate for the influence of the MR edge magnetic field on the incident charged particle beam, so that the incident direction of the charged particle beam is consistent with the main magnetic field direction of the MR imaging area.

[0061] In the present invention, the acceleration device is a linear accelerator, a cyclotron or a synchrotron.

[0062] In the present invention, the interval between the upper group 5 of MR main magnetic field coils and the lower group 6 of MR main magnetic field coils is greater than 60 cm.

[0063] In the present invention, the charged particles in the charged particle beam are electrons, protons, carbon ions or helium ions.

[0064] In another aspect, the present invention provides a method for debugging electromagnetic navigation parameters of the magnetic resonance-guided charged particle beam radiotherapy device according to target coordinates, comprising the following steps:

[0065] Construction of target cost function

[0066] The particle beam phase space corresponding to the electromagnetic navigation parameters is simulated based on the Monte Carlo particle transport algorithm. Specifically, the phase space plane recording the particle information is preferably the treatment isocenter plane (the rotation isocenter in Figure 2 of the specification), and the number of particles is preferably greater than 300. The target cost function is constructed by the difference between the ideal particle beam phase space bombarding the target and the simulated particle beam phase space based on the electromagnetic navigation parameters of the current iteration. Including location cost function and direction cost function Their calculation formula is (1)-(3)

[0067] Where λ is the direction cost function The weight coefficients are x0 and y0, which are the horizontal coordinates of the target point (the horizontal coordinates are the plane perpendicular to the incident direction of the beam, such as the XY plane in Figure 2 of the specification, and the vertical coordinates are the incident direction of the beam, such as the negative direction of the Z axis in Figure 2 of the specification). and is the average transverse coordinate of all particles in the simulated particle beam based on the electromagnetic navigation parameters of the current iteration, is the average value of the angle between the incident directions of all particles in the simulated particle beam and the MR main magnetic field based on the electromagnetic navigation parameters of the current iteration;

[0068] and Electromagnetic navigation parameters Determine, therefore the parameters and the target cost function There is a functional relationship between The electromagnetic navigation parameter debugging task can be described as formula (4):

[0069] in, is the electromagnetic navigation parameter, N is the number of electromagnetic navigation coils;

[0070] Optimization of electromagnetic navigation parameters based on optimization algorithm without gradient information

[0071] Due to the randomness of the particle transport process, it is difficult to determine the functional relationship The analytical expression of the cost function cannot be obtained With input parameters Therefore, the optimization process does not use the objective cost function Electromagnetic navigation parameters Instead of gradient information, it is randomly modified based on historical optimization results. And calculate the corresponding Search or Explore Optimization space;

[0072] Determine and output optimal navigation parameters

[0073] After reaching the preset number of iterations, Minimum electromagnetic navigation parameter As the optimal, the optimal electromagnetic navigation parameters of the target are output.

[0074] In the present invention, the electromagnetic navigation parameters Including the current direction and intensity of the multi-level electromagnetic navigation coil; the electromagnetic navigation parameters Vectorized coding, in and They are the normalized currents of the two pairs of orthogonal magnetic field coils in the first-stage electromagnetic navigation coil, normalized by the corresponding maximum load current, with values ​​between -1.0 and +1.0, where positive and negative indicate the direction of the current. and and represent the normalized currents in the second and third electromagnetic navigation coils, respectively.

[0075] In the present invention, λ is 5.0-15.0.

[0076] In the present invention, the optimization algorithm without gradient information is a Bayesian black box optimization algorithm or a meta-heuristic algorithm.

[0077] In the present invention, the preset number of iterations is more than 300 times.

[0078] The technical solution of the present invention is explained in detail below with reference to specific embodiments.

[0079] Example 1

[0080] A magnetic resonance-guided charged particle beam radiotherapy device as shown in Figures 1-2 is used, and high-energy electron beam simulation with three energies of 70MeV, 100MeV, and 150MeV is used. Charged particle beams of other energies or types, such as protons, helium ions, or carbon ions, can be used. On the basis of this embodiment, the size or current load of the navigation coil can be modified. In combination with the electromagnetic navigation parameter debugging method of the present invention, a parallel charged particle beam consistent with the direction of the magnetic resonance main magnetic field can be generated to achieve precise positioning and lateral confinement of the charged particle beam.

[0081] The magnetic resonance guided charged particle beam radiotherapy device shown in FIG1-2 includes a vacuum tube 1, a multi-stage electromagnetic navigation coil and an MR main magnetic field coil;

[0082] The multi-stage electromagnetic navigation coil is a three-stage electromagnetic navigation coil, which includes a first-stage electromagnetic navigation coil 2, a second-stage electromagnetic navigation coil 3 and a third-stage electromagnetic navigation coil 4;

[0083] The MR main magnetic field coil includes an upper group 5 of MR main magnetic field coils and a lower group 6 of MR main magnetic field coils;

[0084] The copper wire used in the above coils is 2mm wide and has a maximum current of 80 amperes.

[0085] The magnetic field strength of the MR main magnetic field B0 in the imaging treatment area is about 0.35T, the magnetic field strength in the central area of ​​the first electromagnetic navigation coil 2 and the second electromagnetic navigation coil 3 is about 0.22T, and the magnetic field strength in the central area of ​​the third electromagnetic navigation coil 4 is about 0.028T.

[0086] Construction of target cost function

[0087] The target cost function is constructed according to the above method. The particle transport simulation method used in each iteration is preferably the Monte Carlo simulation particle transport algorithm. The Monte Carlo simulation uses the TOPAS library based on the open source Geant4 code, and the linear combination extension package of multiple magnetic field distributions is written using C++ code.

[0088] Optimize electromagnetic navigation parameters based on the optimization algorithm without gradient information;

[0089] The optimization algorithm of the electromagnetic navigation parameter debugging method uses Bayesian black box optimization, and the steps are as follows: ① Initial setting Calculate using formulas (1)-(3) The corresponding cost function Based on this initial point Initialize the surrogate model based on Gaussian process, i.e. In the parameters space The prior distribution of and initialize the acquisition function; ② Considering local exploration and global search, select parameters A point in space Maximize the acquisition function of this iteration; ③ According to Perform Monte Carlo particle transport simulation to obtain The corresponding particle beam phase space file is calculated using formulas (1)-(3) The corresponding cost function ④Based on new data points Update the proxy model and acquisition function; ⑤ Repeat steps ② to ④ until the maximum number of iterations.

[0090] Determine and output optimal navigation parameters

[0091] After reaching the preset number of iterations, output The electromagnetic navigation parameter with the smallest value is taken as the final optimization result. In this embodiment, the number of iterations is set to 500.

[0092] In this embodiment, 25 target points are preset on the isocenter plane for the three energies of high-energy electron beams, and their coordinates are (0.5, 0.5), (0.5, 1.5), (0.5, 2.5), (0.5, 3.5), (0.5, 4.5), (1.5, 0.5), (1.5, 1.5), (1.5, 2.5), (1.5, 3.5), (1.5, 4.5), (2.5, 0.5) .5), (2.5,1.5), (2.5,2.5), (2.5,3.5), (2.5,4.5), (3.5,0.5), (3.5,1.5), (3.5,2.5), (3.5,3.5), (3.5,4.5), (4.5,0.5), (4.5,1.5), (4.5,2.5), (4.5,3.5), (4.5,4.5). All components of the above-mentioned device provided by the present invention have a central axis symmetrical structure. The target points selected in this embodiment are all located in the first quadrant, and the results of the other quadrants can be obtained through symmetry operations.

[0093] As shown in Figure 3, the magnetic field distribution diagram calculated using finite element simulation, where (a) and (b) are the two-dimensional magnetic field distributions of the XZ section and the one-dimensional magnetic field distributions of the central axis of the imaging treatment area (the YZ section is the same); (c) and (d) are the two-dimensional magnetic field distributions of the XZ section and the one-dimensional magnetic field distributions of the central axis of the first and second-stage navigation coils, respectively (only the pair of navigation coils in the X direction is energized); (c) and (d) are the two-dimensional magnetic field distributions of the XZ section and the one-dimensional magnetic field distributions of the central axis of the third-stage navigation coil, respectively (only the pair of navigation coils in the X direction is energized).

[0094] Based on the magnetic resonance-guided charged particle beam radiotherapy device and the electromagnetic navigation parameter debugging method for target coordinates provided by the present invention, the position and angular deviation of the parallel charged particle beam are calculated. The average position error of the parallel charged particle beam is between 0.2 and 0.3 mm, and the average angular error is between 0.4 and 0.6 degrees. Both errors are relatively small and can meet clinical requirements, as shown in Tables 1 and 2:

[0095] Table 1. Position and angular deviations of the parallel charged particle beams generated by the present invention in Example 1

[0096] Table 2. Lateral confinement effect of parallel charged particle beams generated based on the present invention in Example 1

[0097] Note: Reduction percentage = (width in the absence of magnetic field - width in the presence of magnetic field) / width in the absence of magnetic field; FWHM: full width at half maximum of the relative off-axis dose curve; penumbra: width between 20% and 80% of the relative off-axis dose curve; B0: main MR magnetic field strength.

[0098] In summary, the magnetic resonance-guided charged particle beam radiotherapy device and the electromagnetic navigation parameter debugging method for the target target coordinates provided by the present invention, through the design and coupling of multi-level electromagnetic navigation coils and magnetic resonance main magnetic coils, combined with the electromagnetic navigation parameter debugging method for the target target, can generate a parallel charged particle beam consistent with the direction of the magnetic resonance main magnetic field, greatly reducing the influence of the Lorentz force on the treatment beam, and can reduce the difficulty of positioning the tumor target area with the MR-guided charged particle beam and improve the treatment accuracy; using the MR imaging main magnetic field, lateral constraint of the charged particle beam is achieved, and its lateral dose drop gradient is improved, which can better protect the normal tissue around the tumor and reduce the probability of radiotherapy complications.

[0099] The above are only preferred embodiments of the present invention; however, the scope of protection of the present invention is not limited thereto. Any person skilled in the art who, within the technical scope disclosed by the present invention, makes equivalent substitutions or modifications based on the technical solutions and improved concepts of the present invention shall be covered by the scope of protection of the present invention.

Claims

1. A magnetic resonance-guided parallel charged particle beam radiotherapy device, characterized in that, It can generate a charged particle beam parallel to the MR main magnetic field, and it includes: A charged particle beam generation device, which is used to generate and accelerate a charged particle beam so that its energy reaches the energy range required for radiotherapy; A charged particle beam electromagnetic navigation device, which includes multiple levels of electromagnetic navigation coils. The charged particle beam passes through the multiple levels of electromagnetic navigation coils in sequence to generate a parallel charged particle beam consistent with the direction of the MR main magnetic field; An MR imaging device, which includes: An MR main magnetic field coil, whose central axis coincides with the central axis of the parallel charged particle beam. It includes an upper group of MR main magnetic field coils and a lower group of MR main magnetic field coils. An imaging and treatment area is formed between the upper group of MR main magnetic field coils and the lower group of MR main magnetic field coils. Both the upper group of MR main magnetic field coils and the lower group of MR main magnetic field coils include imaging coils and shielding coils. The imaging coils are close to the imaging and treatment area, and the shielding coils are far from the imaging and treatment area. The current directions of the imaging coils and the shielding coils are opposite; Gradient magnetic field coils, which are used to generate spatial magnetic field gradient changes to facilitate imaging slice positioning; A body coil, which is used to generate a radiofrequency field to excite proton spins in tissues; Forward coils and backward coils, which are used to receive echo signals to complete reconstruction.

2. A magnetic resonance-guided charged particle beam radiotherapy device according to claim 1, characterized in that, The multiple levels of electromagnetic navigation coils are three levels of electromagnetic navigation coils arranged successively from top to bottom; The first-level electromagnetic navigation coil deflects the charged particle beam towards the target point. The current direction of the second-level electromagnetic navigation coil is opposite to that of the first-level electromagnetic navigation coil to correct the deflection angle of the charged particle beam. The third-level electromagnetic navigation coil further adjusts the incident direction of the charged particle beam to compensate for the influence of the MR edge magnetic field on the incident charged particle beam, so that the incident direction of the charged particle beam is consistent with the direction of the main magnetic field in the MR imaging area.

3. A magnetic resonance-guided charged particle beam radiotherapy device according to claim 1, characterized in that, The acceleration device is a linear accelerator, a cyclotron or a synchrotron.

4. A magnetic resonance-guided charged particle beam radiotherapy device according to claim 1, characterized in that, The interval between the upper group of MR main magnetic field coils and the lower group of MR main magnetic field coils is greater than 60 cm.

5. A magnetic resonance-guided charged particle beam radiotherapy device according to any one of claims 1-4, characterized in that, The charged particles in the charged particle beam are electrons, protons, carbon ions or helium ions.

6. A method for debugging electromagnetic navigation parameters of a magnetic resonance-guided charged particle beam radiotherapy device according to any one of claims 1-5 for target point coordinates, characterized in that, It includes the following steps: ① Construction of the target cost function Simulate the particle beam phase space corresponding to the electromagnetic navigation parameters based on the particle transport algorithm, and construct the target cost function based on the difference between the ideal particle beam phase space bombarding the target and the simulated particle beam phase space based on the electromagnetic navigation parameters of the current iteration. It includes a location cost function and direction cost function where λ is the direction cost function The weight coefficient, where x0 and y0 are the lateral (in the plane perpendicular to the beam incident direction) coordinate positions of the target and is the average transverse coordinate of all particles of the simulated particle beam based on the electromagnetic navigation parameters of the current iteration, is the mean value of the angles between the incident directions of all particles of the simulated particle beam and the MR main magnetic field based on the electromagnetic navigation parameters of the current iteration, and All consist of electromagnetic navigation parameters Determine; Parameter and the target cost function have a functional relationship between The electromagnetic navigation parameter debugging task can be described by Equation (4): Among them, is the electromagnetic navigation parameter, and N is the number of levels of the electromagnetic navigation coils; ② Optimize the electromagnetic navigation parameters based on an optimization algorithm without gradient information The optimization process does not use the target cost function For electromagnetic navigation parameters gradient information, but randomly correct according to historical optimization results and calculate the corresponding Search or explore The optimization space of; ③ Determine and output the optimal navigation parameters After reaching the preset number of iterations, use The electromagnetic navigation parameter with the smallest value As the optimal, output the optimal electromagnetic navigation parameters of the target point.

7. A method for debugging electromagnetic navigation parameters for target point coordinates of a magnetic resonance-guided charged particle beam radiotherapy device according to claim 6, characterized in that, Electromagnetic navigation parameters including the current directions and intensities of multi-stage electromagnetic navigation coils; vectorizing and encoding the electromagnetic navigation parameters where and are respectively the normalized currents of two pairs of orthogonal magnetic field coils in the first-stage electromagnetic navigation coil, normalized using the corresponding maximum load current, with values between -1.0 and +1.0, and the positive and negative signs indicating the current directions and and respectively represent the normalized currents in the second-stage and third-stage electromagnetic navigation coils.

8. The electromagnetic navigation parameter debugging method for the target point coordinates of a magnetic resonance-guided charged particle beam radiotherapy device according to claim 6, characterized in that, λ is 5.0 - 15.

0.

9. The electromagnetic navigation parameter debugging method for the target point coordinates of a magnetic resonance-guided charged particle beam radiotherapy device according to claim 6, characterized in that, The optimization algorithm without gradient information is a Bayesian black box optimization algorithm or a metaheuristic algorithm.

10. A method for debugging electromagnetic navigation parameters of a magnetic resonance-guided charged particle beam radiotherapy device for target point coordinates according to claim 6, characterized in that, The preset number of iterations is more than 300 times.

Images