High uniformity 14T superconducting magnet for animal MRI

Abstract

The present invention discloses a 14T superconducting magnet with high uniformity for animal magnetic resonance imaging, belonging to the field of superconducting magnets, which includes a plurality of superconducting magnet coils, namely the first high-field Nb3Sn coil, the second high-field Nb3Sn coil, the low-field NbTi coil, the first compensation coil, the second compensation coil, the third compensation coil, the active shim coil, the first shielding coil, and the second shielding coil; the second compensation coil and the third compensation coil are symmetrically distributed along the axial direction; the first shielding coil and the second shielding coil are symmetrically distributed along the axial direction; the active shim coil includes 9 independent coils. The present invention is used in a magnetic resonance imaging system and can significantly improve the resolution and sensitivity of animal imaging, thereby improving the image quality.

Description

Technical Field

[0001] The invention belongs to the field of superconducting magnets, and in particular relates to a high-uniformity 14T superconducting magnet for animal magnetic resonance imaging. Background Art

[0002] There are many superconducting magnets for animal magnetic resonance in the prior art, but all of them are high-uniformity 14T superconducting magnets. Therefore, they cannot meet the requirements of high magnetic field (>14T), high uniformity (<5ppm), and high stability of animal magnetic resonance imaging systems. At the same time, they cannot obtain the optimal result of coil layout, cannot provide a design and evaluation method for the stability of superconducting magnet coils, and cannot calculate the maximum magnetic field density, load rate, and safety margin of each coil. At the same time, they cannot provide corresponding protection for the quenching of superconducting magnets.

[0003] In the prior art, although there are magnet systems with a capacity greater than 10 T that are applied to the human body, their uniformity cannot meet the requirements for signal sensitivity and signal-to-noise ratio, making them unable to be applied in the analysis of material structure and the study of neuronal activity information, and therefore they are not suitable for animal detection.

[0004] In addition, existing superconducting magnet coils are all connected in series and do not use a separate electrical circuit. Therefore, when one of the coils is damaged, it is impossible to excite the target magnetic field and the entire coil must be replaced, which reduces the fault tolerance and increases the risk of quenching. Summary of the invention

[0005] In order to solve the above technical problems, the present invention provides a high-uniformity 14T superconducting magnet for animal magnetic resonance imaging.

[0006] In order to achieve the above object, the present invention adopts the following technical solution:

[0007] A high-uniformity 14T superconducting magnet for animal magnetic resonance imaging comprises a plurality of superconducting magnet coils, wherein the plurality of superconducting magnet coils are respectively, from the inside to the outside, a first high-field Nb3Sn coil in the innermost layer, a second high-field Nb3Sn coil, a second low-field NbTi coil, a first compensation coil in the third layer, a second compensation coil, a third compensation coil, an active shim coil in the fourth layer, and a first shielding coil and a second shielding coil in the outermost layer; the first high-field Nb3Sn coil and the second high-field Nb3Sn coil are connected in series to form a first electrical circuit; the low-field NbTi coil, the first compensation coil, the second compensation coil, the third compensation coil, the fourth active shim coil, and the first shielding coil and the second shielding coil in the outermost layer; the first high-field Nb3Sn coil and the second high-field Nb3Sn coil are connected in series to form a first electrical circuit; the low-field NbTi coil, the first The first compensation coil, the second compensation coil, and the third compensation coil are connected in series to form a second electrical circuit; the first shielding coil and the second shielding coil are connected in series to form a third electrical circuit, and the three electrical circuits are respectively connected to the magnet power supply through current leads; the active shimming coil includes 9 groups of independent coils, each of which is equipped with 9 superconducting switches, and each group of independent coils is an independent circulation loop. When the heater of one group of independent coils is turned on through the superconducting switch, the independent coil is connected to the DC power supply, and the current setting in the group of independent coils is changed by adjusting the output current of the DC power supply.

[0008] Furthermore, a plurality of superconducting magnet coils are connected to a quench protection circuit; the quench protection circuit comprises a quench protection diode, and the quench protection diode and the superconducting magnet coil form a loop, so that the current during quench circulates and is consumed in the loop, thereby preventing the superconducting magnet coil from being damaged.

[0009] Furthermore, when the superconducting switches are all in a closed state, the independent coils of all active shim coils are connected in parallel and are respectively connected to a DC power supply.

[0010] Furthermore, the first compensation coil is arranged in the middle of the outer side of the low-field NbTi coil, and the second compensation coil and the third compensation coil are symmetrically distributed on both sides of the first compensation coil along the axial square line; the first shielding coil and the second shielding coil are symmetrically distributed along the axial square line; the second high-field Nb3Sn coil is coaxially mounted on the outside of the first high-field Nb3Sn coil, and the low-field NbTi coil is mounted on the outside of the second high-field Nb3Sn coil; the first shielding coil and the second shielding coil are respectively arranged on the outside of the second compensation coil and the third compensation coil away from the first compensation coil; the imaging area is a spherical area of ​​60 mm, located in the middle of the first high-field Nb3Sn coil.

[0011] Furthermore, the three groups of electrical circuits are excited individually by closing their respective superconducting switches.

[0012] Furthermore, the 9 groups of independent coils of the active shim coil 18 are respectively 3 groups of solenoid coil Z1, solenoid coil Z2, solenoid coil Z3 and 6 groups of saddle coil X, saddle coil Y, saddle coil ZX, saddle coil ZY, saddle coil X2Y2, and saddle coil XY; wherein, the solenoid coil Z1 is a first-order axial shim coil, the solenoid coil Z2 is a second-order axial shim coil, the solenoid coil Z3 is a third-order axial shim coil, the saddle coil X and the saddle coil Y are first-order radial shim coils, and the saddle coil ZX, the saddle coil ZY, the saddle coil X2Y2, and the saddle coil XY are second-order radial shim coils.

[0013] Furthermore, each group of independent coils is controlled by superconducting switches to form an independent closed loop, supporting independent excitation and current input; through the optimization and debugging of the electromagnetic coupling relationship of 9 groups of independent coils, the active shim coil achieves a magnetic field uniformity of 0.1ppm.

[0014] Furthermore, the quench protection circuit can achieve active quenching of the superconducting magnet coil.

[0015] Furthermore, the supporting frames of the first shielding coil and the second shielding coil are made of epoxy resin material; the supporting frame of the active uniform field coil is made of epoxy resin material; the supporting frame of the low field NbTi coil is made of aluminum alloy material, and the supporting frame of the first high field Nb3Sn coil and the second high field Nb3Sn coil is made of 304 stainless steel material.

[0016] Furthermore, the wound superconducting magnet coil is connected and fixed to two 4k flanges through embedded assembly grooves to form a 4k cold body assembly of the superconducting magnet.

[0017] Beneficial effects:

[0018] 1. The present invention sets the superconducting magnet coil as three independent electrical circuits, thereby increasing the fault tolerance of the system and preventing the system from failing to reach the target magnetic field after one coil is damaged.

[0019] 2. The present invention sets an active shim coil so that the magnetic field uniformity meets the requirements of signal sensitivity and signal-to-noise ratio, and realizes application in material structure analysis and neuron activity information research, so it is suitable for animal detection.

[0020] 3. The superconducting magnet of the present invention has the advantages of stable operation and low quenching. It adopts the principle of safety margin in superconducting magnet coil design based on the critical current characteristics of superconducting wires and empirical data, and is controlled within a certain range. BRIEF DESCRIPTION OF THE DRAWINGS

[0021] Figure 1A schematic diagram of a high-uniformity 14T superconducting magnet for animal magnetic resonance imaging of the present invention;

[0022] Figure 2 This is a safety margin analysis diagram of the high uniformity 14T superconducting magnet for animal magnetic resonance imaging of the present invention;

[0023] Figure 3a is the axial stress distribution diagram of the superconducting magnet coil;

[0024] Figure 3b is the radial stress distribution diagram of the superconducting magnet coil;

[0025] Figure 3c is the hoop stress distribution diagram of the superconducting magnet coil;

[0026] Figure 4 This is a quench protection circuit diagram of a high-uniformity 14T superconducting magnet for animal magnetic resonance imaging of the present invention;

[0027] Figure 5 It is a cross-sectional view of a high uniformity 14T superconducting magnet for animal magnetic resonance imaging of the present invention;

[0028] Figure 6 Schematic diagram of active shimming coil;

[0029] Figure 7 Schematic diagram of the support frame of the active shim coil;

[0030] Figure 8 This is the installation diagram of the superconducting magnet coil and support frame;

[0031] Figure 9a This is the main view of the 4k flange;

[0032] Figure 9b for Figure 9a AA section view in;

[0033] Fig.10 This is the circuit diagram of the active shimming coil.

[0034] Among them, the accompanying drawings are marked as: the first high-field Nb3Sn coil 1, the second high-field Nb3Sn coil 2, the low-field NbTi coil 3, the first compensation coil 4, the second compensation coil 5, the third compensation coil 6, the first shielding coil 7, the second shielding coil 8, the imaging area 9, the quench protection diode 10, the magnet power supply 11, the DC power supply 12, the superconducting switch 13, the superconducting magnet coil 14, the quench heating plate 15, the quench monitoring and controller 16, the quench heating loop switch 17, the active shim coil 18, and the 4k flange 19. DETAILED DESCRIPTION

[0035] In order to make the purpose, technical solutions and advantages of the present invention more clearly understood, the present invention is further described in detail below in conjunction with the accompanying drawings and embodiments. It should be understood that the specific embodiments described herein are only used to explain the present invention and are not intended to limit the present invention. In addition, the technical features involved in the various embodiments of the present invention described below can be combined with each other as long as they do not conflict with each other.

[0036] like Figure 1 , Figure 5 As shown, the embodiment of the present invention proposes a high uniformity 14T superconducting magnet for animal magnetic resonance imaging, including a plurality of superconducting magnet coils, namely, the first high field Nb3Sn coil 1, the second high field Nb3Sn coil 2, the low field NbTi coil 3, the first compensation coil 4, the second compensation coil 5, the third compensation coil 6, the active uniform field coil 18, the first shielding coil 7, and the second shielding coil 8. The second compensation coil 5 and the third compensation coil 6 are symmetrically distributed along the axial square line. The first shielding coil 7 and the second shielding coil 8 are also symmetrically distributed along the axial square line. The imaging area 9 is a 60mm spherical area (60mm DSV). The first high-field Nb3Sn coil 1 and the second high-field Nb3Sn coil 2 are connected in series to form a first electrical circuit; the low-field NbTi coil 3, the first compensation coil 4, the second compensation coil 5, and the third compensation coil 6 are connected in series to form a second electrical circuit; the first shielding coil 7 and the second shielding coil 8 are connected in series to form a third electrical circuit. The three electrical circuits are independent and are connected to a magnet power supply 11 (such as Figure 4 as shown).

[0037] The main magnetic field coil (including shielding coil) of the high-uniformity 14T superconducting magnet for animal magnetic resonance imaging of the present invention adopts a three-independent electrical circuit design scheme, and three independent superconducting switches need to be configured for it respectively, which can control the closing and opening of the circuits separately, so as to set different working currents for the superconducting magnet coils of different circuits. The advantage of this design is that when the manufacturing process error affects and causes the safety margin of a certain superconducting magnet coil to be reduced, the final target magnetic field can be achieved by re-adjusting the magnetic field contribution distribution of each group of superconducting magnet coils. This method may cause the uniformity of the magnetic field in the central area to change, but within a certain range, by increasing the compensation ability of the active shim coil for the uniformity and improving the operating current of the active shim coil, the uniformity problem after the magnetic field contribution is redistributed can be solved.

[0038] Different from the situation where all the main magnetic field coils are connected in series in one loop, the main magnetic field coils are divided into three electrical loops. When the magnet is excited, the superconducting switches are controlled to close separately to excite each set of coils in turn, or the three loop coils are excited simultaneously using three magnet power supplies. Each superconducting switch has a heater to control the closing and opening of the three sets of electrical loops.

[0039] The first high-field Nb3Sn coil 1 is coaxially mounted with the second high-field Nb3Sn coil 2, and the second high-field Nb3Sn coil 2 is mounted with the low-field NbTi coil 3, with a gap between the second high-field Nb3Sn coil 2 and the low-field NbTi coil 3. The first compensation coil 4 is arranged in the middle of the outside of the low-field NbTi coil 3, and the second compensation coil 5 and the third compensation coil 6 are arranged on both sides of the first compensation coil 4. The first shielding coil 7 and the second shielding coil 8 are arranged outside the second compensation coil 5 and the third compensation coil 6 away from the first compensation coil 4. The active shim coil 18 is arranged between the first compensation coil 4, the second compensation coil 5, the third compensation coil 6 and the first shielding coil 7 and the second shielding coil 8. The imaging area 9 is located in the middle of the first high-field Nb3Sn coil 1.

[0040] The central magnetic field strength of the present invention is 14 Tesla, the room temperature aperture of the magnet is 170 mm, the inner diameter of the coil is not less than 225 mm, the uniformity within the 60 mm spherical domain (DSV) is less than 2 ppm, the magnetic field stability is less than 0.05 ppm / hour, it has an active magnetic field shielding function, and a 5 Gauss stray magnetic field (axial Radial) less than 3m 2.5m, the magnet operating current is not less than 200A. Setting the current to not less than 200A can improve the utilization rate of the superconducting wire and achieve a higher field strength. 5 Gauss wire (Guass) is a magnetic field safety wire, which can reduce the radiation and impact of the magnetic field on the surrounding environment and improve the utilization rate of the site. 3m 2.5m is usually acceptable. The uniformity of 60mm DSV is 2ppm, which is the uniformity index required to achieve magnetic resonance imaging to ensure image quality. 60mm DSV is the imaging area, which can be applied to small animal scanning only if it is not less than 60mm. The inner diameter of the coil is not less than 225mm to provide a guarantee for the room temperature aperture of 170mm, and the room temperature aperture of the magnet of 170mm is guaranteed by 60mm DSV.

[0041] The magnetic field strength of the superconducting magnet is achieved by connecting the high-field Nb3Sn coil, the low-field NbTi coil, the compensation coil and the shielding coil in series to form a separate loop, and operating at zero resistance in a low-temperature 4.2K environment. Its magnetic field strength is achieved by the innermost coil, namely the first high-field Nb3Sn coil 1, the second high-field Nb3Sn coil 2, and the low-field NbTi coil 3. Its uniformity is achieved by the first compensation coil 4, the second compensation coil 5, and the third compensation coil 6 located in the middle. Its stray field control is achieved by the outermost first shielding coil 7 and the second shielding coil 8. The three groups of superconducting magnet coils that form a loop in series finally achieve a magnetic field strength of 14T through magnetic field superposition.

[0042] The active shimming coil 18 is wound by superconducting wire and works in a low temperature environment of 4.2K. There are 9 groups of shimming coils in total, namely 3 groups of solenoid coils Z1, solenoid coil Z2, solenoid coil Z3 and 6 groups of saddle coils X, saddle coil Y, saddle coil ZX, saddle coil ZY, saddle coil X2Y2, saddle coil XY, wherein the solenoid coil Z1 is a first-order axial shimming coil, the solenoid coil Z2 is a second-order axial shimming coil, the solenoid coil Z3 is a third-order axial shimming coil, the saddle coil X and saddle coil Y are first-order radial shimming coils, the saddle coil ZX, saddle coil ZY, saddle coil X2Y2, saddle coil XY are second-order radial shimming coils. The solenoid coils Z1, Z2, and Z3 are wound in the grooves of the supporting frame of the active shimming coils and are in the same plane. The saddle-shaped coil X is wound outside the plane, and the saddle-shaped coil Y, the saddle-shaped coil ZX, the saddle-shaped coil ZY, the saddle-shaped coil X2Y2, and the saddle-shaped coil XY are sequentially arranged outside the saddle-shaped coil X.

[0043] Each set of shim coils is equipped with a superconducting switch, which forms an independent closed loop through switch control, supporting separate excitation and current input. Through the optimization and debugging of the electromagnetic coupling relationship of the 9 sets of shim coils, the active shim coil 18 can further achieve a 0.1ppm magnetic field uniformity on the basis of the original 1ppm magnetic field uniformity of the magnet. The uniformity of the high-order magnetic field component can make the magnetic field uniformity between 0.05-0.1ppm. Higher uniformity can bring about an increase in signal sensitivity and signal-to-noise ratio, thereby realizing applications in material structure analysis and neuronal activity information research, which is more suitable for animals.

[0044] like Figure 6 , Figure 7 , Figure 8As shown, each superconducting magnet coil in the present invention is wound on a supporting frame, and there are a total of 5 supporting frames, among which the supporting frames of the first shielding coil 7 and the second shielding coil 8 are designed with epoxy resin frames to achieve a lightweight structure; the supporting frame of the active uniform field coil 18 is made of epoxy resin material; the supporting frame of the low-field NbTi coil 3 is made of aluminum alloy structure, and the electromagnetic stress of the first high-field Nb3Sn coil 1 and the second high-field Nb3Sn coil 2 is relatively large, and the supporting frame is made of 304 stainless steel material. Reasonable material selection can save costs, lighten and be compact. The wound superconducting magnet coil is connected to two 4k flanges 19 (such as Figure 9a , Figure 9b The 4k cold body assembly of the superconducting magnet is formed by connecting and fixing the 4k cold body assembly (as shown in the figure). Each 4k flange 19 has 4 assembly slots for fixing and assembling with each shielding coil, active shim coil 18, each compensation coil, and low-field NbTi coil 3. The first high-field Nb3Sn coil 1 and the second high-field Nb3Sn coil 2 are fixed on the support frame of the high-field Nb3Sn coil. The two ends of the support frame of the high-field Nb3Sn coil are respectively assembled with the 4k flange 19 by continuous welding. The welding surface needs to reach 10 -4 The vacuum requirement is pa, which needs to be carried out after the other superconducting magnet coils are installed.

[0045] Considering the need of superconducting magnet installation, holes are added in the support frame of the active shim coil, and the superconducting joints and superconducting switches of all superconducting magnet coils are arranged on a mounting platform, which is fixed on the outer surface of the support frame of the active shim coil, thereby achieving a more compact structure.

[0046] The present invention also provides a design method for a high-uniformity 14T superconducting magnet for animal magnetic resonance imaging, comprising the following steps:

[0047] Step 1: Combining the linear method with the nonlinear method to form a hybrid optimization algorithm, the hybrid optimization algorithm is used to establish the preliminary size of the superconducting magnet coil, and the hybrid optimization algorithm also considers the design of the shielding coil;

[0048] Step 2: Construct a hybrid objective equation and solve it using the nonlinear multi-objective programming minimization function (Fmincon) (in MATLAB). The goal of the hybrid objective equation is to minimize the volume of the superconducting magnet coil when considering the constraints, which are the magnetic field uniformity, the number of coil turns, the electromagnetic stress, the safety margin, and the maximum magnetic field strength of the coil. The hybrid objective equation is summarized based on the application requirements of the magnetic resonance superconducting magnet and is used to optimize the final design results, specifically:

[0049] Optimize minimum value: V;

[0050] (1)

[0051] Where V is the volume of the electromagnetic coil, is the average magnetic field generated by all coils, ε is the magnetic field deviation in the central area, A is the magnetic induction density of the axial magnetic field in the magnetic field uniformity area, B and C are the magnetic induction density of the radial stray field and the magnetic induction density of the axial stray field, and accordingly, I is the current of all grids, is the area of ​​each grid. Gauss is the unit of magnetic field strength, 1 Tesla is equal to 10000 Gauss. is the critical current density of the superconducting wire. In geometric modeling, the geometric shape of the superconducting magnet coil is divided into multiple small units or regions, each of which is called a grid.

[0052] The constraints are:

[0053] ;

[0054] (2)

[0055] in, is the volume of each superconducting magnet coil, i is the number of the superconducting magnet coil, is the total number of superconducting magnet coils, is the magnetic field strength generated by all coils, and is the 5Guass line stray magnetic field in the axial and radial directions, is the operating current of the superconducting magnet, is the limiting current of the conductor under the working field strength, is the critical current value at the intersection of the excitation critical characteristic line and the superconducting critical characteristic line, Indu is the inductance of the superconducting coil, and Ener is the energy of the superconducting coil. Correspondingly, η is the maximum current safety margin, is the maximum load factor of the conductor, is the maximum hoop electromagnetic stress of all coils, and is the maximum magnetic field strength and maximum magnetic field limit strength of the coil, and is the radial layer number and axial turn number of the coil. Correspondingly, tw and ww are the thickness and width of the superconducting magnet coil. r1 is the inner diameter of the coil, r2 is the outer diameter of the coil, z1 is the axial starting position of the coil, and z2 is the axial end position of the coil. is the maximum inductance allowed for all coils in series of a superconducting magnet, The maximum energy allowed for all coils of a superconducting magnet.

[0056] The present invention innovatively proposes to optimize the coil load rate and safety margin to control the risk of safe and stable operation of the magnet. At the same time, controlling the total inductance and total energy of the superconducting magnet is conducive to reducing the thermal load impact caused by the superconducting magnet quench, reducing the hot spot temperature of the coil after the magnet quenches, and improving the safety of the coil.

[0057] Step 3: Obtain the initial geometric position of each superconducting magnet coil through repeated optimization iterations.

[0058] Preferably, two types of low-temperature superconducting wires are used in the present invention, namely Nb3Sn superconducting wire and NbTi superconducting wire. When the magnet is running, the initial current of the three groups of separate loop coils is 230A. Among them, the Nb3Sn superconducting wire includes two different specifications, which are used for high-field superconducting magnet coils, and the NbTi superconducting wire also has two different specifications, which are used for low-field superconducting magnet coils, compensation coils and shielding coils.

[0059] Preferably, the present invention takes into account the coil safety margin, and the operating current through the superconducting magnet coil Critical current of superconducting wire under corresponding magnetic field The load factor of the superconducting magnet coil is obtained, and the safety margin of the superconducting magnet coil is further calculated.

[0060] The load rate of all superconducting magnet coils shall not exceed 60%, and the design safety margin shall not be less than 10%. Figure 2 As shown, the superconducting magnet is composed of 8 coils (divided into 6 groups), the superconducting magnet includes two innermost high-field Nb3Sn coils (the first high-field Nb3Sn coil 1 and the second high-field Nb3Sn coil 2), and a low-field NbTi coil 3 (i.e. Figure 2 The first NbTi coil in the middle provides a sufficiently high central magnetic field strength, and the first compensation coil 4, the second compensation coil 5, and the third compensation coil 6 (i.e. Figure 2 The second NbTi coil and the third NbTi coil, wherein the third NbTi coil includes Figure 2 The 3-A NbTi coil and the 3-B NbTi coil not shown in the figure are a pair of symmetrically distributed, corresponding to the second compensation coil 5 and the third compensation coil 6 respectively, and are used to correct the non-uniformity of the high magnetic field generated by the first high-field Nb3Sn coil 1, the second high-field Nb3Sn coil 2, and the low-field NbTi coil 3. At the same time, the first shielding coil 7 and the second shielding coil 8 (i.e. Figure 2 The 4th NbTi coil in the invention comprises Figure 2The 4-A NbTi coil and the 4-B NbTi coil not shown in the figure are a pair, symmetrically distributed, corresponding to the 1st shielding coil 7 and the 2nd shielding coil 8 respectively, which reduces the distance of the stray magnetic field and enhances the safety of the system. All the main coils (all coils except the shielding coil) and the shielding coils can be connected in series with the magnet power supply respectively.

[0061] Figure 2 The first Nb3Sn wire is used for Figure 2 The first high field Nb3Sn coil 1 in the second Nb3Sn wire is used Figure 2 The second high field Nb3Sn coil 2 in the first NbTi wire is used Figure 2 The first NbTi coil in the second NbTi wire is used Figure 2 The 2nd NbTi coil - the 4th NbTi coil.

[0062] Since the superconducting magnet design of the present invention uses four kinds of superconducting wires, these four kinds of superconducting wires have different critical current characteristics, respectively. Figure 2 The wire curves of the first and second NbTi wires and the wire curves of the first and second Nb3Sn wires are shown in Figure 1. In addition, the four wires are used for 6 sets of superconducting magnet coils. The stability of each set of superconducting magnet coils needs to be considered when designing and optimizing. The load rate and safety margin are used to measure the load rate. The load rate is the operating current of the magnet coil. and the limiting current of the superconducting wire under the corresponding magnetic field The ratio is usually no more than 60% according to experience, and the safety margin needs to be calculated first. , that is, the operating current of the magnet coil The current corresponding to the intersection of the operating current curve and the superconducting wire critical current curve The ratio of This is the safety margin, which should not be less than 10% according to experience. Figure 2 Four superconducting wire critical current curves and six sets of magnet coil operating current curves are listed, which can easily obtain the corresponding coil load rate and safety margin, thereby evaluating the stability of the superconducting magnet coil. When this evaluation method and empirical data are introduced into the magnet optimization algorithm, it is easy to obtain calculation results that meet the requirements.

[0063] After calculation, the load rate and safety margin of each coil of the 14T animal magnetic resonance superconducting magnet are shown in Table 1. The maximum load rate of the 1st NbTi coil is 55.1%, and the minimum safety margin of the 1st NbTi coil is 10.6%, which meets the design requirements and the coil is within the design safety margin.

[0064] Table 1 Maximum magnetic field and safety margin of magnet coil

[0065]

[0066] The magnetic field uniformity of a superconducting magnet is defined as the ratio of the absolute value of the difference between the maximum magnetic field value and the minimum magnetic field value within a certain range at the center of the magnet to the average magnetic field value. The uniformity of the superconducting magnet in the present invention is achieved by a high field coil, a low field coil, an active shim coil and a compensation coil, wherein there are three compensation coils, i.e., three segmented Helmholtz coils. After the final electromagnetic calculation, the uniformity of the superconducting magnet is 1ppm@60mm DSV, where ppm is the unit of uniformity and DSV is the spherical symbol, i.e., a spherical space with a diameter of 60mm, which is better than the design requirements.

[0067] The present invention has good stray field control function, which is mainly completed by a pair of shielding coils located at the outermost layer of the magnet. Radial stray field (5 Gaussian lines) controlled at 3m Within 2.5m.

[0068] like Figure 3a , Figure 3b , Figure 3c FIG. 1 is a diagram showing the electromagnetic stress distribution of the superconducting magnet coil in the present invention. Figure 3a is the axial stress distribution diagram of the superconducting magnet coil; Figure 3b is the radial stress distribution diagram of the superconducting magnet coil; Figure 3c The hoop stress distribution diagram of the superconducting magnet coil. The maximum axial, radial and hoop stresses are 69.92MPa, 7.46MPa and 225MPa respectively. With reference to the yield strength of the superconducting wire of 330MPa, the coil stress is acceptable.

[0069] like Figure 4 As shown, the superconducting magnet in the present invention has a quench protection function, which is realized by a quench protection circuit, and the quench protection circuit includes a quench protection diode 10, a magnet power supply 11, a DC power supply 12, a superconducting switch 13, a superconducting magnet coil 14, a quench heating plate 15, a quench monitoring and controller 16, and a quench heating circuit switch 17. When quenching, the current will circulate in the loop formed by the superconducting magnet coil 14 and the quench protection diode 10 and eventually consume, thereby avoiding damage to the superconducting magnet coil. At the same time, the present invention has a quench monitoring function. When a superconducting magnet coil quenches, the voltage signal rises suddenly, and the quench monitoring and controller 16 collects the voltage signal and triggers the quench heating circuit switch 17 to close, and the quench heating plate 15 close to the superconducting magnet coil 14 is heated, and causes other superconducting magnet coils to quench, thereby dispersing the heat concentration caused by the quench of a single coil, and avoiding damage to the superconducting magnet coil.

[0070] The superconducting magnet coil 14 of the present invention includes 6 groups of coils (8 coils in total), namely the 1st high field Nb3Sn coil (1st group), the 2nd high field Nb3Sn coil (2nd group), the 1st NbTi coil (3rd group, corresponding to the low field NbTi coil 3), the 2nd NbTi coil (4th group, corresponding to the 1st compensation coil 4), the 3rd NbTi coil (5th group, in pair, can be divided into the 3rd-A NbTi coil and the 3rd-B NbTi coil), the 4th NbTi coil (6th group, in pair, can be divided into the 4th-A NbTi coil and the 4th-B NbTi coil, corresponding to the 1st shielding coil 7 and the 2nd shielding coil 8 respectively). The 1st high field Nb3Sn coil corresponds to Figure 4 Nb3Sn 1, the 2nd high field Nb3Sn coil corresponds to Figure 4 The Nb3Sn 2 in the 1st NbTi coil corresponds to Figure 4 The NbTi 1 and 2 NbTi coils correspond to Figure 4 The NbTi 2nd and 3rd NbTi coils correspond to Figure 4 The NbTi 3,4 NbTi coils correspond to Figure 4 NbTi 4 in.

[0071] The quench protection diode 10, the magnet power supply 11, the superconducting switch 13, and the superconducting magnet coil 14 are connected in parallel to form a loop, and the quench monitoring and controller 16 is connected to each group of superconducting magnet coils 14. Each quench heating plate 15 is close to each group of superconducting magnet coils 14, and each quench heating plate 15 is connected in series with the DC power supply 12 and the quench heating circuit switch 17 to form a loop.

[0072] Preferably, the quench protection diode 10 includes a diode D1 and a diode D2, which is a passive protection for the magnet quench. During the excitation or operation of the superconducting magnet, the quench may occur due to instability caused by the design or process, and a large instantaneous current and voltage may be output. At this time, the quench voltage exceeds the trigger voltage of the diode D1 or the diode D2, and the diode D1 or the diode D2 is turned on. The current forms a cycle through the diode D1 or the diode D2, and is gradually lost and attenuated. This process reduces the voltage shock of the superconducting magnet.

[0073] At the same time, the present invention increases quench monitoring and leading protection functions, mainly including quench monitoring and controller 16 and quench heating plate 15 to realize. Preferably, quench heating plate 15 is 6, and its resistance is respectively R1-R6. Quench monitoring connects both ends of each coil, collects voltage signal and transmits it to controller, and when the voltage signal obtained by controller is instantaneous sharp signal, it is judged as coil quench. Quench heating plate 15 is arranged on 6 groups of corresponding superconducting magnet coils respectively, when it is judged that a certain group of superconducting magnet coil quenches, all DC power supply 12 is started to work, triggering quench heating circuit switch 17 to close, at this time all quench heating plates 15 start to work, and trigger all superconducting magnet coils to quench, disperse the energy release of quench, and avoid the temperature concentration brought by quench. There is also a situation, when magnet has an emergency, it needs human intervention to make it actively quench, and the heating function of quench heating plate 15 can be started by operating quench heating circuit switch 17 to make superconducting magnet quench. The closed-loop stable operation of the superconducting magnet depends on the superconducting switch 13. During the magnet excitation stage, the superconducting switch 13 is in a heating off state. At this time, the magnet power supply 11 supplies power to the superconducting magnet to achieve a predetermined magnetic field strength. When the superconducting magnet reaches the predetermined magnetic field strength, the superconducting switch 13 can be turned off to stop the excitation and achieve the closed loop of the superconducting magnet. At this time, the magnet power supply 11 can be removed.

[0074] The circuits of the 9 groups of independent coils of the active shim coil 18 are as follows: Fig.10 As shown. Among them, Pin 1-Pin 9 are the negative pins of the heater current of the 9 superconducting switches, Pin 12 is the positive pin of the current, Pin 10 and Pin 11 are the positive and negative pins of the current input of the shim coil, which are used to connect to the output end of the shim DC power supply. 9 groups of independent coils are distributed with 9 superconducting switches. When the superconducting switches are all in the closed state, all coils are in parallel and connected to the DC power supply through Pin 10 and Pin 11 respectively. Each independent coil is an independent circulation loop. When one of the independent coils turns on the heater through the superconducting switch, the independent coil will be connected to the DC power supply, and the current setting in the coil can be changed by adjusting the output current of the DC power supply. Through this method, different current settings in the 9 groups of independent coils can be achieved, thereby showing the best shim performance.

[0075] The present invention mainly realizes a stable and lasting high-uniformity strong magnetic field. The closed-loop operation of the magnet does not require an external power supply. It is mainly used in animal magnetic resonance imaging (MRI) systems and can achieve high-resolution and high-sensitivity magnetic resonance scanning images.

Claims

1. A high uniformity 14T superconducting magnet for animal magnetic resonance imaging, characterized in that: The invention comprises a plurality of superconducting magnet coils, which are respectively, from the inside to the outside, the first high-field Nb3Sn coil and the second high-field Nb3Sn coil of the innermost layer, the low-field NbTi coil of the second layer, the first compensation coil, the second compensation coil and the third compensation coil of the third layer, the active shimming coil of the fourth layer and the first shielding coil and the second shielding coil of the outermost layer; the first high-field Nb3Sn coil and the second high-field Nb3Sn coil are connected in series to form a first electrical circuit; the low-field NbTi coil, the first compensation coil, the second compensation coil and the third compensation coil are connected in series to form a second electrical circuit; the first shielding coil and the second shielding coil are connected in series to form a third electrical circuit, and the three groups of electrical circuits are respectively connected to the magnet power supply through current leads; the active shimming coil comprises 9 groups of independent coils, which are respectively equipped with 9 superconducting switches, and each group of independent coils is an independent circulation circuit. When a heater is turned on by the superconducting switch in one group of independent coils, the independent coils are connected to a DC power supply, and the current setting in the group of independent coils is changed by adjusting the output current of the DC power supply.

2. A high uniformity 14T superconducting magnet for animal magnetic resonance imaging according to claim 1, characterized in that: A plurality of superconducting magnet coils are connected to a quench protection circuit; the quench protection circuit comprises a quench protection diode, and the quench protection diode and the superconducting magnet coil form a loop, so that the current during the quench circulates and is consumed in the loop, thereby preventing the superconducting magnet coil from being damaged.

3. The high uniformity 14T superconducting magnet for animal magnetic resonance imaging according to claim 1, characterized in that: When the superconducting switches are all in a closed state, the independent coils of all active shim coils are connected in parallel and are respectively connected to direct current power supplies.

4. The high uniformity 14T superconducting magnet for animal magnetic resonance imaging according to claim 1, characterized in that: The first compensation coil is arranged in the middle of the outer side of the low-field NbTi coil, and the second compensation coil and the third compensation coil are symmetrically distributed on both sides of the first compensation coil along the axial square line; the first shielding coil and the second shielding coil are symmetrically distributed along the axial square line; the second high-field Nb3Sn coil is coaxially mounted on the outside of the first high-field Nb3Sn coil, and the low-field NbTi coil is mounted on the outside of the second high-field Nb3Sn coil; The first shielding coil and the second shielding coil are respectively arranged outside the second compensation coil and the third compensation coil which are far away from the first compensation coil; the imaging area is a 60mm spherical area, which is located in the middle of the first high-field Nb3Sn coil.

5. The high uniformity 14T superconducting magnet for animal magnetic resonance imaging according to claim 1, characterized in that: The three groups of electrical circuits are excited individually by closing their respective superconducting switches.

6. The high uniformity 14T superconducting magnet for animal magnetic resonance imaging according to claim 1, characterized in that: The 9 groups of independent coils of the active shim coil (18) are respectively a solenoid coil Z1, a solenoid coil Z2, a solenoid coil Z3 and a saddle coil X, a saddle coil Y, a saddle coil ZX, a saddle coil ZY, a saddle coil X2Y2, and a saddle coil XY; wherein the solenoid coil Z1 is a first-order axial shim coil, the solenoid coil Z2 is a second-order axial shim coil, the solenoid coil Z3 is a third-order axial shim coil, the saddle coil X and the saddle coil Y are first-order radial shim coils, and the saddle coil ZX, the saddle coil ZY, the saddle coil X2Y2, and the saddle coil XY are second-order radial shim coils.

7. The high uniformity 14T superconducting magnet for animal magnetic resonance imaging according to claim 6, characterized in that: Each group of independent coils is controlled by superconducting switches to form an independent closed loop, supporting independent excitation and current input; through the optimization and debugging of the electromagnetic coupling relationship of 9 groups of independent coils, the active shim coil achieves a magnetic field uniformity of 0.1ppm.

8. The high uniformity 14T superconducting magnet for animal magnetic resonance imaging according to claim 2, characterized in that: The quench protection circuit can realize active quenching of the superconducting magnet coil.

9. The high uniformity 14T superconducting magnet for animal magnetic resonance imaging according to claim 1, characterized in that: The supporting frames of the first shielding coil, the second shielding coil and the active uniform field coil are made of epoxy resin material; the supporting frames of the low field NbTi coil are made of aluminum alloy material; and the supporting frames of the first high field Nb3Sn coil and the second high field Nb3Sn coil are made of 304 stainless steel material.

10. The high uniformity 14T superconducting magnet for animal magnetic resonance imaging according to claim 1, characterized in that: The wound superconducting magnet coil is connected and fixed to two 4k flanges through embedded assembly grooves.

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