Charged particle beam radiotherapy system guided by magnetic resonance

Abstract

The invention provides a magnetic resonance guided charged particle beam radiotherapy device and an electromagnetic navigation parameter debugging method for target coordinates, and belongs to the technical field of radiotherapy devices. The system comprises a charged particle beam electromagnetic navigation device which comprises multiple stages of electromagnetic navigation coils, charged particle beams sequentially pass through the multiple stages of electromagnetic navigation coils, and parallel charged particle beams consistent with a magnetic resonance main magnetic field in direction are generated; the MR imaging device comprises an MR main magnetic field coil upper group and an MR main magnetic field coil lower group, and an imaging treatment area is formed between the MR main magnetic field coil upper group and the MR main magnetic field coil lower group. The electromagnetic navigation parameter debugging method comprises the steps of constructing a target cost function, optimizing electromagnetic navigation parameters based on an optimization algorithm without gradient information, and determining and outputting optimal navigation parameters. According to the invention, the charged particle beam parallel to the MR main magnetic field can be generated, the influence of Lorentz force on the charged particle beam is greatly reduced, the difficulty of positioning a tumor target region by the charged particle beam under the guidance of MR can be reduced, and the treatment accuracy is improved.

Description

Technical Field

[0001] This invention relates to the field of radiotherapy equipment technology, and in particular to a magnetic resonance-guided charged particle beam radiotherapy system, including a magnetic resonance-guided parallel charged particle beam radiotherapy device and a method for adjusting electromagnetic navigation parameters for target coordinates. Background Technology

[0002] Cancer is a leading cause of death in humans, and more than half of cancer patients require radiation therapy. Compared to conventional cone-beam CT-guided radiotherapy, magnetic resonance (MR)-guided radiotherapy involves no imaging radiation dose and allows for MRI imaging before or during treatment, improving the resolution of soft tissues such as tumors, resulting in higher treatment precision and improved patient outcomes. This has significant clinical implications.

[0003] Currently, all MR-guided radiotherapy techniques used in clinical practice are photon radiotherapy, such as MRIdian from ViewRay in the United States and Unity from Elekta in Sweden. Compared to photon radiotherapy, charged particle radiotherapy has specific advantages, such as insensitivity to tissue heterogeneity and a steeper longitudinal dose drop, but there are currently no MR-guided charged particle radiotherapy devices available clinically. The transport of charged particle beams in the MR main magnetic field is affected by the Lorentz force, causing the beam to deviate from its original direction. This makes it difficult to accurately locate the tumor target area, leading to missed tumor irradiation or mis-irradiation of normal tissue, ultimately affecting the therapeutic effect.

[0004] Chinese patent CN111580030A discloses an MR-guided radiotherapy system where the MR main magnetic field is perpendicular to the incident particle direction. While this invention utilizes niobium-titanium alloy in the linear accelerator region to shield the magnetic field and reduce beam deflection outside the patient, it fails to address the deflection problem within the patient's body. Both United Imaging Healthcare (China) and the University of Wollongong (Australia) have designed radiotherapy systems where the MR main magnetic field is parallel to the beam's central axis for MR-guided photon radiotherapy. However, this approach is unsuitable for charged particle radiotherapy systems. In this system, only the beam along the central axis is parallel to the MR main magnetic field; beams deviating from the central axis exhibit a conical divergence. For charged particle beams, beams deviating from the central axis are still subject to Lorentz forces, resulting in a spiral beam path, which is detrimental to precise beam positioning. To address the beam localization problem of charged particle beams in the main magnetic field of MR (Mechanical Radiation) systems, the University of Wollongong's Centre for Medical Radiation Physics proposed a complex two-step method for beam parameter correction. First, based on simulations and tabulation using a uniform water phantom, a first-order approximation is obtained for each beam energy and target point. Then, Monte Carlo simulations are performed on specific patients to acquire a large amount of beam parameter and corresponding target point location data. Based on this, a neural network model is trained to refine the first-order approximation. However, this method is overly complex, requiring tabulation of each beam energy and specific patient data, and training the neural network model, making it cumbersome and impractical for clinical translation. Summary of the Invention

[0005] In view of this, to address the technical problems existing in the prior art, such as the deflection of charged particle beam trajectories leading to difficulty in accurately locating the tumor target area, resulting in missed or false radiation, and the excessive scattering of charged particle beams during transport, causing increased lateral penumbra and hindering the protection of normal tissues, this invention provides a magnetic resonance-guided charged particle beam radiotherapy device. This device generates a charged particle beam parallel to the MR main magnetic field, significantly reducing the influence of the Lorentz force on the charged particle beam. This reduces the difficulty of MR-guided charged particle beam localization to the tumor target area and improves treatment accuracy. Furthermore, by utilizing the MR imaging main magnetic field, lateral constraint of the charged particle beam is achieved, improving its lateral dose drop gradient and better protecting the normal tissues surrounding the tumor, thus reducing the probability of radiotherapy complications.

[0006] To achieve the above objectives, the present invention provides the following technical solution:

[0007] A magnetic resonance-guided charged particle beam radiotherapy device, capable of generating a charged particle beam parallel to the MR main magnetic field, comprising:

[0008] A charged particle beam generating device, which is used to generate and accelerate charged particle beams to achieve the energy range required for radiotherapy;

[0009] A charged particle beam electromagnetic navigation device, comprising multiple stages of electromagnetic navigation coils, wherein a charged particle beam sequentially passes through the multiple stages of electromagnetic navigation coils to generate a parallel charged particle beam aligned with the direction of the main magnetic field of magnetic resonance imaging; and an MR imaging device, comprising:

[0010] The MR main magnetic field coil has its central axis coincident with the central axis of the parallel charged particle beam. It includes an upper MR main magnetic field coil group and a lower MR main magnetic field coil group. An imaging therapy area is formed between the upper MR main magnetic field coil group and the lower MR main magnetic field coil group. Both the upper MR main magnetic field coil group and the lower MR main magnetic field coil group contain an imaging coil and a shielding coil. The imaging coil is close to the imaging therapy area, and the shielding coil is far away from the imaging therapy area. The current directions of the imaging coil and the shielding coil are opposite.

[0011] Gradient magnetic field coils are used to generate spatial magnetic field gradient changes, which facilitates imaging layer selection and positioning.

[0012] Body coils are used to generate radio frequency fields to excite proton spins within tissues;

[0013] The forward and backward coils are used to receive the echo signal and complete the reconstruction.

[0014] Preferably, the multi-stage electromagnetic navigation coil is a three-stage electromagnetic navigation coil arranged sequentially from top to bottom;

[0015] The first-stage electromagnetic navigation coil deflects the charged particle beam toward the target point. The second-stage electromagnetic navigation coil has a current direction opposite to that of the first-stage coil, correcting 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 influence of the MR edge magnetic field on the incident charged particle beam, ensuring that the incident direction of the charged particle beam is consistent with the direction of the main magnetic field in the MR imaging region.

[0016] Preferably, the acceleration device is a linear accelerator, a cyclotron accelerator, or a synchrotron.

[0017] Preferably, the interval between the upper group and the lower group of the MR main magnetic field coil is greater than 60 cm.

[0018] Preferably, the charged particles in the charged particle beam are electrons, protons, carbon ions, or helium ions.

[0019] On the other hand, the present invention provides a method for adjusting the electromagnetic navigation parameters of the above-mentioned magnetic resonance-guided charged particle beam radiotherapy device for the target coordinates, comprising the following steps:

[0020] Construction of the target cost function

[0021] The particle beam phase space corresponding to the electromagnetic navigation parameters is simulated based on the particle transport algorithm. The target cost function is constructed by 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 directional cost function

[0022]

[0023]

[0024] Where λ is the directional cost function The weighting coefficients are given by x0 and y0, which are the lateral coordinates of the target point. and The average transverse coordinate of all particles in the simulated particle beam based on the electromagnetic navigation parameters of the current iteration. This represents the average 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. and All are determined by electromagnetic navigation parameters Decide;

[0025] parameter and target cost function There is a functional relationship between them. The task of electromagnetic navigation parameter debugging can be described by formula (4):

[0026]

[0027] in, Here, N represents the electromagnetic navigation parameters, and N is the number of stages in the electromagnetic navigation coil.

[0028] Optimization of electromagnetic navigation parameters based on gradient-free optimization algorithms;

[0029] The optimization process does not use the objective cost function. Electromagnetic navigation parameters Instead of relying on gradient information, it randomly adjusts the gradient based on historical optimization results. And calculate the corresponding Search or explore The room for optimization;

[0030] Determine and output the optimal navigation parameters

[0031] After reaching the preset number of iterations, Electromagnetic navigation parameter with the smallest value As the optimal approach, the optimal electromagnetic navigation parameters for the target point are output.

[0032] Preferably, electromagnetic navigation parameters This includes the current direction and intensity of multi-stage electromagnetic navigation coils; and the electromagnetic navigation parameters. Vectorized encoding in and These are the normalized currents of the 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. The positive and negative signs indicate the current direction. and and These represent the normalized currents in the second and third stage electromagnetic navigation coils, respectively.

[0033] Preferably, λ is 5.0-15.0.

[0034] Preferably, the optimization algorithm without gradient information is a Bayesian black-box optimization algorithm or a meta-heuristic algorithm.

[0035] Preferably, the preset number of iterations is 300 or more.

[0036] Compared with the prior art, the present invention has the following beneficial effects:

[0037] The magnetic resonance-guided charged particle beam radiotherapy device provided by this invention can generate a charged particle beam parallel to the MR main magnetic field, which greatly reduces the influence of the Lorentz force on the charged particle beam. This reduces the difficulty of MR-guided charged particle beam localization to the tumor target area and improves treatment accuracy. By utilizing the MR imaging main magnetic field, lateral confinement of the charged particle beam is achieved, improving its lateral dose drop gradient, which can better protect the normal tissues around the tumor and reduce the probability of radiotherapy complications. Attached Figure Description

[0038] Figure 1 A perspective view of the magnetic resonance-guided charged particle beam radiotherapy device provided by the present invention;

[0039] Figure 2 A front view of the magnetic resonance-guided charged particle beam radiotherapy device provided by the present invention;

[0040] Figure 3 This is a magnetic field distribution diagram calculated using finite element method (FEM) simulation.

[0041] In the diagram, 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 coil, 6. Lower group of MR main magnetic field coil. Detailed Implementation

[0042] The technical solutions of the present invention will be clearly and completely described below with reference to the accompanying drawings of the embodiments of the present invention. Obviously, the described embodiments are only some embodiments of the present invention, and not all embodiments. All other embodiments obtained by those skilled in the art based on the embodiments of the present invention without creative effort are within the scope of protection of the present invention.

[0043] In the description of this invention, it should be noted that the terms "upper," "lower," "inner," "outer," "top / bottom," etc., indicate the orientation or positional relationship based on the orientation or positional relationship shown in the accompanying drawings. They are used only for the convenience of describing the invention and for simplifying the description, and do not indicate or imply that the device or element referred to must have a specific orientation, or be constructed and operated in a specific orientation. Therefore, they should not be construed as limitations on the invention. Furthermore, the terms "first" and "second" are used for descriptive purposes only and should not be construed as indicating or implying relative importance.

[0044] In the description of this invention, it should be noted that, unless otherwise explicitly specified and limited, the terms "installed," "equipped with," "sleeved / connected," "connected," etc., should be interpreted broadly. For example, "connection" can be 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; it can be a connection within two components. Those skilled in the art can understand the specific meaning of the above terms in this invention based on the specific circumstances.

[0045] The following problems exist with existing MR-guided radiotherapy techniques:

[0046] 1) Charged particles are greatly affected by the main magnetic field and the edge magnetic field of MR, causing the beam trajectory to deflect, making it difficult for the beam to accurately locate the tumor target area, resulting in missed or false radiation.

[0047] 2) Compared to photons, charged particles are more likely to be scattered laterally by air or human tissue. Especially for lighter charged particles, such as electrons, a large amount of scattering will occur during transport, increasing the lateral penumbra, which is not conducive to the protection of normal tissues.

[0048] This invention provides a magnetic resonance-guided charged particle beam radiotherapy device, which is capable of generating a charged particle beam parallel to the MR main magnetic field, comprising:

[0049] A charged particle beam generating device, which is used to generate and accelerate charged particle beams so that their energy reaches the energy range required for radiotherapy;

[0050] The charged particle beam electromagnetic navigation device includes multiple electromagnetic navigation coils. The charged particle beam passes through the multiple electromagnetic navigation coils in sequence to generate a parallel charged particle beam that is consistent with the direction of the main magnetic field of the magnetic resonance imaging (MRI). Unlike the off-center divergent transport of charged particle beams used in conventional radiotherapy, the parallel charged particle beam always has the same direction as the main magnetic field of the MR, regardless of whether the beam is on the isocentric axis or off-center.

[0051] MR imaging apparatus, comprising:

[0052] The MR main magnetic field coil has its central axis coincident with the central axis of the parallel charged particle beam. It includes an upper MR main magnetic field coil group and a lower MR main magnetic field coil group. An imaging therapy area is formed between the upper MR main magnetic field coil group and the lower MR main magnetic field coil group. Both the upper MR main magnetic field coil group and the lower MR main magnetic field coil group contain an imaging coil and a shielding coil. The imaging coil is close to the imaging therapy area, and the shielding coil is far away from the imaging therapy area. The current directions of the imaging coil and the shielding coil are opposite.

[0053] Gradient magnetic field coils are used to generate spatial magnetic field gradient changes, which facilitates imaging layer selection and positioning.

[0054] Body coils are used to generate radio frequency fields to excite proton spins within tissues;

[0055] The forward and backward coils are used to receive the echo signal and complete the reconstruction.

[0056] Specifically, such as Figure 1-2 As shown, this invention provides an exemplary magnetic resonance-guided charged particle beam radiotherapy device. A three-stage multi-stage electromagnetic navigation coil is used as an example for illustration. This magnetic resonance-guided charged particle beam radiotherapy device includes:

[0057] Vacuum tube 1, with a first-stage electromagnetic navigation coil 2, a second-stage electromagnetic navigation coil 3 and a third-stage electromagnetic navigation coil 4 sequentially mounted on its outside from top to bottom;

[0058] The upper group 5 and the lower group 6 of the MR main magnetic field coil are arranged at intervals 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 edge magnetic field generated by the upper group 5 of the main magnetic field coil on the beam). The upper group 5 of the MR main magnetic field coil is sleeved on the third-level electromagnetic navigation coil 4. The area between the upper group 5 and the lower group 6 of the MR main magnetic field coil forms the imaging treatment area.

[0059] A charged particle beam generating device (not shown in the figure) generates a charged particle beam, which is accelerated by an accelerator and then taken in along the extension direction of the vacuum tube to form an image in the imaging treatment area.

[0060] This device can greatly reduce the deflection or distortion of the charged particle beam's incident trajectory by the MR main magnetic field and edge 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.

[0061] In this invention, the multi-stage electromagnetic navigation coil is a three-stage electromagnetic navigation coil arranged sequentially from top to bottom;

[0062] 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, compensating 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 of the MR imaging area.

[0063] In this invention, the acceleration device is a linear accelerator, a cyclotron accelerator, or a synchrotron accelerator.

[0064] In this invention, the interval between the upper group 5 of the MR main magnetic field coil and the lower group 6 of the MR main magnetic field coil is greater than 60cm.

[0065] In this invention, the charged particles in the charged particle beam are electrons, protons, carbon ions, or helium ions.

[0066] On the other hand, the present invention provides a method for adjusting the electromagnetic navigation parameters of the above-mentioned magnetic resonance-guided charged particle beam radiotherapy device for the target coordinates, comprising the following steps:

[0067] Construction of the target cost function

[0068] 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 for recording particle information is preferably the treatment isocenter plane (instructions). Figure 2 At the rotation center (where the particle count is greater than 300), the number of particles is preferably greater than 300. A target cost function is constructed based on 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. Including location cost function and directional cost function Their calculation formulas are (1)-(3)

[0069]

[0070]

[0071]

[0072] Where λ is the directional cost function The weighting coefficients, x0 and y0 are the lateral coordinates of the target point (the lateral direction is a plane perpendicular to the incident direction of the beam, as shown in the instruction manual). Figure 2 The XY plane in the diagram has its longitudinal direction as the incident direction of the beam, as shown in the instruction manual. Figure 2 The position in the negative Z-axis direction. and The average transverse coordinate of all particles in the simulated particle beam based on the electromagnetic navigation parameters of the current iteration. It is the average angle between the incident direction of all particles in the simulated particle beam and the MR main magnetic field, based on the electromagnetic navigation parameters of the current iteration;

[0073] and All are determined by electromagnetic navigation parameters The decision, therefore the parameters and target cost function There is a functional relationship between them. The task of electromagnetic navigation parameter debugging can be described by formula (4):

[0074]

[0075] in, Here, N represents the electromagnetic navigation parameters, and N is the number of stages in the electromagnetic navigation coil.

[0076] Optimization of electromagnetic navigation parameters based on gradient-free information optimization algorithm

[0077] Because particle transport processes are random, it is difficult to determine the functional relationships. The parsing expression cannot obtain the cost function. With input parameters The gradient information. Therefore, the optimization process does not use the objective cost function. Electromagnetic navigation parameters Instead of relying on gradient information, it randomly adjusts the gradient based on historical optimization results. And calculate the corresponding Search or explore The room for optimization;

[0078] Determine and output the optimal navigation parameters

[0079] After reaching the preset number of iterations, Electromagnetic navigation parameter with the smallest value As the optimal approach, the optimal electromagnetic navigation parameters for the target point are output.

[0080] In this invention, electromagnetic navigation parameters This includes the current direction and intensity of multi-stage electromagnetic navigation coils; and the electromagnetic navigation parameters. Vectorized encoding in and These are the normalized currents of the 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. The positive and negative signs indicate the current direction. and and These represent the normalized currents in the second and third stage electromagnetic navigation coils, respectively.

[0081] In this invention, λ is 5.0-15.0.

[0082] In this invention, the optimization algorithm without gradient information is either a Bayesian black-box optimization algorithm or a meta-heuristic algorithm.

[0083] In this invention, the preset number of iterations is 300 or more.

[0084] The technical solution of the present invention will be explained in detail below with reference to specific embodiments.

[0085] Example 1

[0086] Adopting such Figure 1-2 The magnetic resonance-guided charged particle beam radiotherapy device shown uses high-energy electron beams of 70MeV, 100MeV, and 150MeV for simulation. Other energies or other types of charged particle beams, such as protons, helium ions, or carbon ions, can be generated by modifying the size of the navigation coil or the current load based on this embodiment, combined with the electromagnetic navigation parameter adjustment method of this invention, to generate a parallel charged particle beam that is consistent with the direction of the main magnetic field of the magnetic resonance, thereby achieving precise positioning and lateral constraint of the charged particle beam.

[0087] like Figure 1-2 The magnetic resonance-guided charged particle beam radiotherapy device shown includes a vacuum tube 1, a multi-stage electromagnetic navigation coil, and an MR main magnetic field coil.

[0088] 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.

[0089] The MR main magnetic field coil includes the upper group 5 of the MR main magnetic field coil and the lower group 6 of the MR main magnetic field coil;

[0090] The copper wire used for each of the above coils is 2mm wide and has a maximum current of 80 amperes.

[0091] The magnetic field strength of the MR main magnetic field B0 in the imaging treatment area is approximately 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 approximately 0.22T, and the magnetic field strength in the central area of ​​the third electromagnetic navigation coil 4 is approximately 0.028T.

[0092] Construction of the target cost function

[0093] Based on the above method, the objective cost function is constructed. 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 uses C++ code to write an extension package for linear combination of multiple magnetic field distributions.

[0094] Optimization of electromagnetic navigation parameters based on gradient-free optimization algorithms;

[0095] The optimization algorithm for the electromagnetic navigation parameter debugging method uses Bayesian black-box optimization, and the steps are as follows: ① Initial settings Calculate using formulas (1)-(3) Corresponding cost function Based on this initial point The agent model is initialized based on Gaussian process, i.e. In parameters space ① Determine the prior distribution and initialize the acquisition function; ② Consider both local exploration and global search, and select the parameters. A point in space Maximize the acquisition function in this iteration; ③ According to Perform Monte Carlo particle transport simulations to obtain... The corresponding particle beam phase space file is calculated using formulas (1)-(3). Corresponding cost function ④ Based on the new data points Update the proxy model and acquisition function; ⑤ Repeat steps ② to ④ until the maximum number of iterations is reached.

[0096] Determine and output the optimal navigation parameters

[0097] Output after reaching the preset number of iterations. 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.

[0098] In this embodiment, 25 target points are pre-defined in the isocentric plane for each of the three energy levels of the high-energy electron beam, with coordinates of (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), and (0.5,0.5). .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 in the above-described device provided by this invention have a central axis symmetric structure. In this embodiment, the target points selected are all located in the first quadrant, and the results in the other quadrants can be obtained through symmetric operations.

[0099] like Figure 3 The diagram shows the magnetic field distribution calculated using finite element method (FEM) simulation. (a) and (b) show the two-dimensional and one-dimensional magnetic field distributions of the XZ section and central axis of the imaging therapy zone, respectively (the YZ section is the same). (c) and (d) show the two-dimensional and one-dimensional magnetic field distributions of the XZ section and central axis of the first and second-level navigation coils, respectively (only one pair of navigation coils in the X direction is energized). (c) and (d) show the two-dimensional and one-dimensional magnetic field distributions of the XZ section and central axis of the third-level navigation coil, respectively (only one pair of navigation coils in the X direction is energized).

[0100] Based on the magnetic resonance-guided charged particle beam radiotherapy device and the electromagnetic navigation parameter adjustment method for target coordinates provided by this invention, the position and angular deviations 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 small and meet clinical requirements, as shown in Tables 1 and 2.

[0101] Table 1. Position and angular deviations of the parallel charged particle beam generated based on the present invention in Example 1

[0102]

[0103] Table 2. Lateral confinement effect of parallel charged particle beam generated based on the present invention in Example 1

[0104]

[0105]

[0106] Note: Reduction percentage = (width without magnetic field - width with corresponding magnetic field) / width without magnetic field; FWHM: half-width at half maximum relative to the off-axis dose curve; penumbra: width between 20% and 80% of the off-axis dose curve; B0 is the intensity of the MR main magnetic field.

[0107] In summary, the magnetic resonance-guided charged particle beam radiotherapy device and the electromagnetic navigation parameter adjustment method for target coordinates provided by this invention, through the design and coupling of multi-stage electromagnetic navigation coils and magnetic resonance main magnetic coils, combined with the electromagnetic navigation parameter adjustment method for the target, can generate a parallel charged particle beam aligned with the direction of the magnetic resonance main magnetic field. This greatly reduces the influence of the Lorentz force on the treatment beam, reduces the difficulty of MR-guided charged particle beam localization to the tumor target area, and improves treatment accuracy. By utilizing the MR imaging main magnetic field, lateral constraint on the charged particle beam is achieved, improving its lateral dose drop gradient, which can better protect the normal tissues surrounding the tumor and reduce the probability of radiotherapy complications.

[0108] The above are merely preferred embodiments of the present invention; however, the scope of protection of the present invention is not limited thereto. Any equivalent substitutions or modifications made by those skilled in the art within the scope of the technology disclosed in the present invention, based on the technical solution and its improved concept, should be covered within 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 beam of charged particles parallel to the MR main magnetic field, which includes: A charged particle beam generating device, which is used to generate and accelerate charged particle beams to achieve the energy range required for radiotherapy; A charged particle beam electromagnetic navigation device includes multiple electromagnetic navigation coils. A charged particle beam passes through the multiple electromagnetic navigation coils in sequence to generate a parallel charged particle beam that is aligned with the direction of the main magnetic field of the magnetic resonance. MR imaging apparatus, comprising: The MR main magnetic field coil has its central axis coincident with the central axis of the parallel charged particle beam. It includes an upper MR main magnetic field coil group and a lower MR main magnetic field coil group. An imaging therapy area is formed between the upper MR main magnetic field coil group and the lower MR main magnetic field coil group. Both the upper MR main magnetic field coil group and the lower MR main magnetic field coil group contain an imaging coil and a shielding coil. The imaging coil is close to the imaging therapy area, and the shielding coil is far away from the imaging therapy area. The current directions of the imaging coil and the shielding coil are opposite. Gradient magnetic field coils are used to generate spatial magnetic field gradient changes, which facilitates imaging layer selection and positioning. Body coils are used to generate radio frequency fields to excite proton spins within tissues; The forward and backward coils are used to receive the echo signal and complete the reconstruction.

2. The magnetic resonance-guided charged particle beam radiotherapy device according to claim 1, characterized in that, The multi-stage electromagnetic navigation coil consists of three stages of electromagnetic navigation coils arranged sequentially from top to bottom. The first-stage electromagnetic navigation coil deflects the charged particle beam toward the target point. The second-stage electromagnetic navigation coil has a current direction 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 further adjusts the incident direction of the charged particle beam, compensating 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 of the MR imaging region.

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

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

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 a charged particle beam are electrons, protons, carbon ions, or helium ions.

6. A method for adjusting electromagnetic navigation parameters of a magnetic resonance-guided charged particle beam radiotherapy device for target coordinates according to any one of claims 1-5, characterized in that, Includes the following steps: ① Construction of the objective cost function The particle beam phase space corresponding to the electromagnetic navigation parameters is simulated based on the particle transport algorithm. The target cost function is constructed by 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 directional cost function Where λ is the directional cost function The weighting coefficients are given, where x0 and y0 are the lateral (plane perpendicular to the beam incident direction) coordinates of the target point. and The average transverse coordinate of all particles in the simulated particle beam based on the electromagnetic navigation parameters of the current iteration. This represents the average 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. and All are determined by electromagnetic navigation parameters Decide; parameter and target cost function There is a functional relationship between them. The task of electromagnetic navigation parameter debugging can be described by formula (4): in, Here, N represents the electromagnetic navigation parameters, and N is the number of stages in the electromagnetic navigation coil. ② Optimize electromagnetic navigation parameters using an optimization algorithm without gradient information The optimization process does not use the objective cost function. Electromagnetic navigation parameters Instead of relying on gradient information, it randomly adjusts the gradient based on historical optimization results. And calculate the corresponding Search or explore The room for optimization; ③ Determine and output the optimal navigation parameters After reaching the preset number of iterations, Electromagnetic navigation parameter with the smallest value As the optimal approach, the optimal electromagnetic navigation parameters for the target point are output.

7. The method for adjusting electromagnetic navigation parameters of a magnetic resonance-guided charged particle beam radiotherapy device for target coordinates according to claim 6, characterized in that, Electromagnetic navigation parameters This includes the current direction and intensity of multi-stage electromagnetic navigation coils; and the electromagnetic navigation parameters. Vectorized encoding in and These are the normalized currents of the 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. The positive and negative signs indicate the current direction. and and These represent the normalized currents in the second and third stage electromagnetic navigation coils, respectively.

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

0.

9. The method for adjusting electromagnetic navigation parameters of a magnetic resonance-guided charged particle beam radiotherapy device for target coordinates according to claim 6, characterized in that, Optimization algorithms without gradient information are either Bayesian black-box optimization algorithms or metaheuristic algorithms.

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

Images