Radio frequency coil design method for transmitting array space coding imaging

Abstract

The invention relates to a radio frequency coil design method for transmitting array space coding imaging, and belongs to the technical field of magnetic resonance imaging. Aiming at the problems of low single-dimensional coding efficiency, strong coupling interference among coils and nonlinear phase gradient in the traditional TRASE technology, axial and radial coding coils are designed based on a target field method and a magnetic dipole method, two-dimensional radio frequency phase coding is realized through reverse current solenoid decoupling combination, and excitation and signal receiving are realized in combination with a saddle-shaped coil. According to the technical scheme, by optimizing the coil structure and the decoupling design, multi-dimensional efficient coding, high-precision phase gradient, strong anti-interference and light system weight are achieved, and the method is suitable for portable and dynamic monitoring scenes of low-field magnetic resonance.

Description

Technical Field

[0001] The present invention belongs to the technical field of magnetic resonance imaging and relates to a radio frequency coil design method for transmitting array spatial encoding imaging. Background Art

[0002] Magnetic resonance imaging (MRI), as a noninvasive medical imaging technique, plays an irreplaceable role in the clinical diagnosis of joint and brain diseases. Traditional MRI systems rely on a high-intensity main magnetic field (typically 1.5T or 3.0T) combined with a gradient coil system for spatial encoding. However, their high manufacturing cost, bulky size, and operating energy consumption limit their application in mobile healthcare and long-term monitoring scenarios. To meet the clinical needs of low-cost, portability, and continuous monitoring, low-field (0.2T-1.0T) and ultra-low-field (<0.1T) MRI technologies have recently become a research hotspot. Low-field MRI systems have a high tolerance for patient positioning errors, making them particularly suitable for patients who find it difficult to remain still, such as children and critically ill patients. They also offer excellent soft tissue contrast in joint imaging. However, the image resolution and signal-to-noise ratio (SNR) of low-field MRI are significantly lower than those of high-field systems, making it difficult to image fine structures (such as microcartilage lesions and small brain lesions). This has become a major bottleneck in its clinical adoption.

[0003] Traditional MRI spatial encoding relies on the linear gradient field generated by the main magnetic field (B0) gradient coil, which achieves k-space traversal through frequency encoding and phase encoding. However, the volume, weight and high power consumption of the gradient coil are particularly prominent in low-field systems, and its performance drops sharply in non-uniform B0 fields. To this end, Transmit Array Spatial Encoding (TRASE) technology has been proposed, which uses the phase gradient field generated by the radio frequency (RF) coil to directly perform spatial encoding without relying on the B0 gradient coil. The core principle of TRASE is to alternately transmit 180° refocusing pulses through multiple coils with different RF phase gradient directions, so that the phase of the spin echo sequence accumulates to form a k-space trajectory, thereby achieving imaging. Compared with B0 gradient encoding, TRASE technology has the following advantages: (1) the RF coil is small and lightweight and can be customized for specific anatomical sites; (2) it can still work stably in highly non-uniform B0 fields; (3) the system power consumption and manufacturing cost are significantly reduced.

[0004] However, the practical application of existing TRASE technology still faces the following challenges:

[0005] Most TRASE systems use a single RF phase gradient direction (such as axial or radial), which can only achieve one-dimensional spatial encoding. They rely on mechanical rotation or complex coil switching to achieve multidimensional imaging, resulting in low scanning efficiency and increased risk of motion artifacts.

[0006] Existing RF phase-gradient coils are often based on empirical winding designs and lack rigorous electromagnetic field optimization. For example, the phase linearity of axial gradient coils is easily distorted (error >10%) at the edge of the target region, while the field uniformity of radial gradient coils is limited by insufficient winding symmetry, leading to image artifacts.

[0007] When multiple RF coils are used in combination, the mutual inductance effect between the coils will introduce additional phase errors. Traditional decoupling methods (such as geometric orthogonality and capacitance compensation) have limited effectiveness under wide-band conditions, especially in low-field systems where sensitivity decreases, further reducing the imaging signal-to-noise ratio.

[0008] To address these issues, existing technologies attempt to improve performance by optimizing coil layout or introducing active shielding structures, but this often results in increased coil size or a sharp increase in design complexity. Therefore, a radiofrequency coil design method that balances multi-dimensional encoding efficiency, phase gradient accuracy, and coil decoupling performance is urgently needed to promote the clinical practical application of low-field TRASE technology. Summary of the Invention

[0009] In view of this, an object of the present invention is to provide a radio frequency coil design method for transmit array spatially coded imaging.

[0010] In order to achieve the above object, the present invention provides the following technical solutions:

[0011] A method for designing a radio frequency coil for transmit array spatially coded imaging comprises the following steps:

[0012] S1: Axial encoding coil design: Based on the target field method and magnetic dipole method, at a radius of R z The cylindrical surface design of the RF coil, R z is the cylindrical radius of the axial encoding coil, so that it generates a longitudinal radio frequency phase gradient field in the target area, and the Cartesian coordinate components of the target field satisfy:

[0013]

[0014] Among them, B X 、B Y 、B Z Respectively represent the radiofrequency magnetic field components along the X-axis, Y-axis, and Z-axis in the Cartesian coordinate system; L represents half the length of the target area ROI along the Z-axis; z and y are the coordinates of the target point along the Z-axis and Y-axis, respectively;

[0015] S2: Radial encoding coil design: Based on the target field method and magnetic dipole method, the radius is R r The cylindrical surface design of the RF coil, R r is the cylindrical radius of the radial encoding coil, so that it generates a transverse radio frequency phase gradient field in the target area, and the cylindrical coordinate component of the target field satisfies:

[0016]

[0017] Among them, B ρ 、B θ 、B z Respectively represent the radial, circumferential, and axial components of the radio frequency magnetic field in the cylindrical coordinate system; μ0 represents the vacuum magnetic permeability; I represents the coil input current; A represents the winding modulation amplitude of the radial encoding coil; h represents the coil turn spacing; R represents the coil cylindrical radius; ρ represents the radial distance; θ represents the azimuth angle; ψ represents the winding rotation direction angle;

[0018] S3: Coil combination and decoupling: The axial encoding coil and the radial encoding coil are combined, and decoupling is achieved by connecting a solenoid coil with reverse current in series to form a two-dimensional RF phase encoding system.

[0019] Furthermore, in S1, the radius R of the axial encoding coil z satisfy:

[0020] The design of the axial encoding coil includes: 0.8L≤R z ≤1.2L

[0021] The cylindrical surface is divided into Q grid cells with the stream function S q Describe the current distribution;

[0022] Solve the stream function S by the least squares method q The optimal distribution of B in the target area Z The phase gradient linearity error of the component is less than 5%;

[0023] The minimum curvature radius of the stream function contour line is ≥ 2mm.

[0024] Furthermore, in S2, the radius R of the radial encoding coil r Meet: 6.0cm≤R r ≤7.0cm;

[0025] The optimal range of winding modulation amplitude A is 4≤A≤7, and the turn spacing h≥3mm.

[0026] Furthermore, in S3, the decoupling design adopts any of the following methods:

[0027] Method 1: The first solenoid coil is connected in series to one end of the axial encoding coil, and the second solenoid coil is connected in series to one end of the radial encoding coil, and the current directions of the two solenoids are opposite;

[0028] Method 2: A third solenoid coil with the same direction of current is connected in series at both ends of the axial encoding coil, and a fourth solenoid coil with the same direction of current is connected in series at both ends of the radial encoding coil, and the current directions of the third and fourth solenoids are opposite.

[0029] Furthermore, in S1 and S2, the optimization of the target field is based on a mathematical model:

[0030]

[0031] The constraints are:

[0032]

[0033] Among them, Q c is the number of grids around the cylindrical surface, S q is the unit flow function.

[0034] Furthermore, the size of the target area is 8 cm × 4 cm × 4 cm, and the radius of the axial encoding coil R z =8.5cm, radial encoding coil radius R r =6.6cm.

[0035] Furthermore, the method further comprises:

[0036] integrating the combined radio frequency coil with a saddle coil for transmitting 90° excitation pulses and receiving magnetic resonance signals;

[0037] The axial encoding coil and the radial encoding coil alternately transmit 180° refocusing pulse sequences to form spin echo encoding.

[0038] Furthermore, the refocusing pulse sequence is: (CB) 2 -(CTS1) 3 -C, where:

[0039] C represents a saddle coil, TS1 represents an axial encoding coil, and B represents a radial encoding coil; the encoding direction is K c -K TS1 With K c -K B orthogonal trajectories.

[0040] Furthermore, the working field strength of the radio frequency coil is 0.2T to 1.0T, which is suitable for ultra-low field magnetic resonance imaging of wrist joints or brain tissues.

[0041] A radio frequency coil system designed by implementing the method comprises:

[0042] Saddle coil module C: arranged along the Z-axis direction, used to transmit 90° excitation pulses and receive magnetic resonance echo signals;

[0043] Axial phase gradient module CTS1: composed of the axial encoding coil, coaxially arranged along the axial direction of the saddle coil, i.e., the Z axis, for generating a Z-axis linear phase gradient;

[0044] Radial phase gradient module B: composed of the radial encoding coils, arranged orthogonally along the radial direction of the saddle coil, i.e., the Y axis, for generating a Y-axis linear phase gradient;

[0045] Decoupling module: includes a first reverse solenoid group connected in series at both ends of the axial encoding coil, and a second reverse solenoid group connected in series at both ends of the radial encoding coil, wherein:

[0046] The current direction of the first reverse solenoid group is opposite to the driving current of the axial encoding coil;

[0047] The current direction of the second reverse solenoid group is opposite to the driving current of the radial encoding coil;

[0048] The radio frequency output end of the saddle coil module is connected to the axial / radial phase gradient module through a first switching circuit;

[0049] The axial encoding coil and the radial encoding coil are alternately turned on by a second switching circuit to switch transmission in a 180° refocusing pulse sequence;

[0050] The signal receiving end of the saddle coil module is independently connected to the data acquisition unit of the magnetic resonance imaging system.

[0051] The beneficial effects of the present invention are:

[0052] (1) Through the combined design of axial and radial radio frequency coils, the present invention achieves two-dimensional spatial encoding without mechanical rotation or complex switching, significantly improving imaging efficiency while reducing image distortion caused by motion artifacts, making low-field magnetic resonance imaging adaptable to dynamic monitoring scenarios.

[0053] (2) The coil optimization design based on the target field method and magnetic dipole method ensures that the phase gradient of the radio frequency magnetic field in the target area is strictly linear and uniformly distributed, effectively overcoming the edge distortion problem of traditional coils and providing an accurate phase encoding basis for high-resolution imaging.

[0054] (3) The unique reverse current solenoid decoupling design significantly reduces the mutual inductance coupling effect between multiple coils, allowing the axial and radial encoding coils to work together without interfering with each other, ensuring the stability and reliability of signal acquisition under complex sequences.

[0055] (4) Abandoning the traditional gradient coil system, a compact radio frequency coil array is used to achieve spatial encoding, which significantly reduces the system size, weight and power consumption. It is especially suitable for portable devices and mobile medical scenarios, while reducing hardware manufacturing costs.

[0056] (5) RF phase gradient encoding has low dependence on the homogeneity of the main magnetic field and can work stably in a highly non-uniform magnetic field environment, expanding the application range of low-field MRI in complex clinical scenarios, such as bedside monitoring or intraoperative real-time imaging.

[0057] (6) The modular design of axial, radial, and saddle coils supports rapid adaptation to different anatomical sites. By adjusting coil parameters or combinations, it can be flexibly expanded to three-dimensional encoding or high-contrast imaging of specific tissues, thus improving the universality of the technology.

[0058] Other advantages, objects, and features of the present invention will be described in part in the following description and, in part, will be apparent to those skilled in the art upon examination of the following description or may be learned from practice of the present invention. The objects and other advantages of the present invention may be realized and obtained through the following description. BRIEF DESCRIPTION OF THE DRAWINGS

[0059] In order to make the purpose, technical solutions and advantages of the present invention more clear, the present invention will be described in detail below with reference to the accompanying drawings, in which:

[0060] Figure 1 This is a schematic diagram of the vector diagram group of each component of the axial (Z) axis encoding coil and the coil winding diagram of Case 1. Figure 1 (a)~ Figure 1 (d) The vector distribution diagrams of Bx, By, Bz and Btarget, respectively. Figure 1 (e) and Figure 1 (f) Stream function contour map and 3D winding diagram along the axial encoding designed by the target field method and magnetic dipole method, respectively;

[0061] Figure 2 This is a schematic diagram of the transmitting array spatial encoding coil group along the axial encoding described in Case 1. Figure 2 (a) is the RF phase gradient coil for axial encoding. Figure 2 (b) is the Bz component distribution diagram of the coil in the ZX plane when y=3cm, Figure 2 (c) is the Bz component distribution diagram, and the arrows indicate the direction of the magnetic field component;

[0062] Figure 3 The following is a line simulation result diagram of the radial RF coil (TS1) in Case 1 at y=4cm, 0cm, and -4cm on the yz plane. Figure 3 (a)~ Figure 3 (c) Phase diagram, phase gradient diagram and B1 modulus diagram of coil simulation, respectively. Figure 3 (d) Actual phase, ideal phase, and linearity at y = 0 cm;

[0063] Figure 4 This is a schematic diagram of the vector diagram group of each component of the radial (Y-axis) encoding coil and the coil winding diagram of Case 2. Figure 4 (a)~ Figure 4 (d) The vector distribution diagrams of Bx, By, Bz and Btarget, respectively. Figure 4 (e) and Figure 4 (f) Stream function contour map and 3D winding diagram along the radial encoding designed by target field method and magnetic dipole method, respectively;

[0064] Figure 5 This is a schematic diagram of the radially encoded transmitting array spatial encoding coil group described in Case 2. Figure 5 (a) RF phase gradient coil for radial encoding; Figure 5 (b) Bx component distribution diagram of the ZY plane when X = 3 cm; Figure 5 (c) is the Bx component distribution diagram, and the arrows indicate the direction of the magnetic field component;

[0065] Figure 6 This is the decoupling model diagram, which shows the decoupling model of the axial RF coil and the radial RF coil with a regular solenoid coil added at one end and both ends respectively. Figure 6 (a)~ Figure 6 (d) Model diagrams of TS1+RS1, TS2+RS2, RS1′+TS1+RS1 and RS2′+TS2+RS2, respectively. Arrows indicate the direction of current.

[0066] Figure 7 This is the parameter diagram of the decoupling coil S11. Figure 7 (a)~ Figure 7 (d) are the S11 parameters of TS1+RS1, TS2+RS2, RS1′+TS1+RS1 and RS2′+TS2+RS2 in CST simulation respectively;

[0067] Figure 8 Schematic diagram of the K-space trajectory for two-dimensional ultra-low-field magnetic resonance imaging using the transmit array spatial encoding coil group. After 2D K-space trajectory excitation using the saddle coil C, the refocusing sequence (CB)2-(CA)3-C is executed using orthogonal encoding directions (Kc-Ka) and (Kc-Kb). DETAILED DESCRIPTION

[0068] The following describes the embodiments of the present invention by means of specific examples, and those skilled in the art can easily understand other advantages and effects of the present invention from the contents disclosed in this specification. The present invention can also be implemented or applied through other different specific embodiments, and the details in this specification can also be modified or changed in various ways based on different viewpoints and applications without departing from the spirit of the present invention. It should be noted that the illustrations provided in the following embodiments are only schematic illustrations of the basic concept of the present invention, and the following embodiments and features in the embodiments can be combined with each other without conflict.

[0069] Among them, the accompanying drawings are only for illustrative purposes and represent only schematic diagrams rather than actual pictures, and should not be understood as limiting the present invention. In order to better illustrate the embodiments of the present invention, some parts of the accompanying drawings may be omitted, enlarged or reduced, and do not represent the dimensions of actual products. For those skilled in the art, it is understandable that some well-known structures and their descriptions may be omitted in the accompanying drawings.

[0070] The same or similar numbers in the drawings of the embodiments of the present invention correspond to the same or similar parts; in the description of the present invention, it should be understood that if there are terms such as "upper", "lower", "left", "right", "front", "back", etc. indicating directions or positional relationships, they are based on the directions or positional relationships shown in the drawings. They are only for the convenience of describing the present invention and simplifying the description, and do not indicate or imply that the device or element referred to must have a specific direction, be constructed and operate in a specific direction. Therefore, the terms describing the positional relationship in the drawings are only used for illustrative purposes and cannot be understood as limiting the present invention. For ordinary technicians in this field, the specific meanings of the above terms can be understood according to specific circumstances.

[0071] See also Figures 1 to 3 , is a schematic diagram of the axially encoded radio frequency phase coil of Case 1 of the present invention, as shown Figure 1 As shown, the coil design method includes: using the target field method and the magnetic dipole method to design the general structure of the coil, and then continuously optimizing each parameter to obtain the optimal coil structure. The design is divided into the following four steps:

[0072] (1) Case 1 The method of RF phase encoding coil along the axial Z direction is as follows: the elliptical cylinder with current is divided into several current-carrying circular currents. Each arc current loop can be regarded as a magnetic dipole. The magnetic field in the target area ROI (Region of Interest) is composed of the superposition of the magnetic fields generated by these unit magnetic dipoles. If the surface of any coil is divided into Q grid units with a side length of a. When a is small enough, the stream function in the element q can be approximated by the constant Sq, and the surface of the element can be approximated by a 2 The element q at the source point will generate a magnetic field at the point in the target area, which is equivalent to the magnetic field of a magnetic dipole.

[0073] (2) Case 1: The Bx, By, and Bz components of the desired magnetic field are superimposed to obtain Btarget. The Btarget vector diagram produces a uniform axial phase change along the Z axis. The coil is designed to be constrained on a cylindrical surface using the target field method and the magnetic dipole method. The following longitudinal phase gradient target field is set in the target area:

[0074]

[0075] Where B X B Y and B z They represent the magnetic field components along the X-axis, Y-axis, and Z-axis respectively, and Btarget is the matrix form of the desired magnetic field. The vector diagrams of each component are as follows Figure 1 (a)~ Figure 1 As shown in (d), the value of L can be modified based on the size of the imaging region of interest (ROI) according to the size of the human wrist joint.

[0076] (3) The radius range of coil parameters is set to 8 cm to 10 m, with a step size of 0.5 cm. By evaluating the winding distribution of each generated coil, we avoid sharp turns and windings that are too small. The coil with a radius of 8.5 cm is selected as the optimal radius. The target area is set to a rectangular block of 8 cm × 4 cm × 4 cm according to the size of the adult wrist joint. Figure 1 (e) and Figure 1 (f) Stream function contour distribution and three-dimensional winding diagram of the 8.5 cm coil;

[0077] (4) The data calculated in MATLAB, one column for Z (m) coordinates and the other for Φ (rad), are processed using the formula x = ρcosθ, y = ρsinθ to obtain the xyz coordinates and then imported into COMSOL Multiphysics to design a coil for electromagnetic field simulation, as shown below: Figure 2 (a)~ Figure 2 (c) In case 1, the axially encoded coil has a magnetic field component Bz that changes from negative to positive along the z-axis, while the magnetic field component Bx along the x-axis remains essentially unchanged. The phase of B1 can be determined by Calculation shows that φ is the RF phase and B1 is the amplitude |B1| is the RF field, and the coil encoding along the z-axis achieves the desired encoding. The gradient linearity formula is as follows:

[0078]

[0079] Where Δφ is the phase difference (the difference between the simulated or measured phase and the ideal phase φ target ), where the ideal phase is determined by the following formula:

[0080]

[0081] Where L is half the length of the ROI in the z direction. To accurately estimate the phase gradient, a second-order polynomial and a formula are fitted to the raw Bx and Bz measurements, which are used as the B1 phase gradient according to the formula:

[0082]

[0083] In the formula, the B1 phase gradient is evaluated at some equidistant points on the x-axis, and Δx i Defined as the selected x interval. For simulated phase, phase gradient, amplitude, and phase linearity, refer to Figure 3 (a)~ Figure 3 (d).

[0084] Figures 4 and 5 This is a schematic diagram of the radially encoded RF phase coil of Case 2 of the present invention, as shown in FIG. Figure 4 As shown, the coil design method includes: using the target field method and the magnetic dipole method to design the general structure of the coil, and continuously optimizing the parameters of the coil turns, radius, modulation amplitude, and turn spacing to obtain the optimal coil structure. The design is divided into the following four steps:

[0085] (1) Case 2: The coil is designed to be constrained on a cylindrical surface using the target field method and the magnetic dipole method. The following transverse phase gradient target field is set in the target area:

[0086]

[0087] B ρ is the radial component, B θ is the circumferential component, B z For the axial component, define Cartesian coordinates; (x, y, z) = (ρcosθ, ρsinθ, z), and convert these field components to Cartesian coordinates:

[0088] B X =-B θ sinθ+B ρ cosθ

[0089] B y =+B θ cosθ+B ρ sinθ

[0090]

[0091] For the selected coil rotation direction (ψ = 0), the B1 phase gradient is only in the vertical x direction (ie, Y axis), because the Bz component is uniform and By only changes as a function of x. It can also be expressed by the following expression:

[0092]

[0093] Using the trigonometric identities x = ρ·cosθ and ψ = 0, we simplify to:

[0094]

[0095] Written in consistent matrix form:

[0096]

[0097] The Bx, By, and Bz components generated by the desired magnetic field are superimposed to obtain Btarget. The Btarget vector diagram produces a uniform lateral phase change along the Y axis, as shown in Figure 4 (a)~ Figure 4 (d)

[0098] (2) The target area is set as a rectangular parallelepiped of 8cm×4cm×4cm based on the size of an adult's wrist joint. Using the target field method and the equivalent magnetic dipole method, with the goal of minimizing energy loss, the energy consumption expression generated by all magnetic dipole units is derived. Within the target area, the magnetic field generated by the coil is required to be orthogonal to each other on the X-axis and Y-axis and as close to the expected value as possible. Based on this, a mathematical model is constructed:

[0099]

[0100] (3) The least squares method is used to solve the optimal distribution of the stream function on the cylindrical surface. After obtaining the optimal value of the stream function Sq, the contour line of the stream function is drawn to obtain the coil structure. The various parameters of the radio frequency coil in Case 2 are parametrically scanned to obtain the optimal values: A = 5.5, I = 3A, h = 0.3cm, θ = 45°, where A is the coil modulation amplitude. After optimization, it is found that when A = 5.5, the coil shape is better; I is the input current. In this coil, the range of I is 1-3A. h is the distance between the two turns of the coil. This distance cannot be too large, which will cause the coil to be less compact and affect its uniformity. It cannot be too small, which will make the coils too close and affect the impedance matching. h ≥ 3mm is the most suitable. The value range of θ is a cylindrical coordinate ranging from -Nπ to Nπ, where N is the number of coil turns; Φ is the winding displacement. The coil parameter radius range is set to 6.0cm-6.8cm with a step size of 0.2cm. The optimal radius obtained by scanning is 6.6cm. The target area is set to a rectangular block of 8cm×4cm×4cm according to the size of the adult wrist joint. Figure 4 (e) and Figure 4 (f) is the stream function contour distribution and 3D winding diagram of the 6.6 cm coil.

[0101] (4) The data generated by running in MATLAB, with one column as Z (m) coordinates and the other as Φ (rad), are processed using the formula x = ρcosθ, y = ρsinθ to obtain the xyz coordinates, which are then imported into COMSOL Multiphysics to design a coil for electromagnetic field simulation. The designed coil simulation is as follows Figure 5 (a)~ Figure 5 As shown in (c), in case 2, the radially encoded coil with a radius of 6.6 cm has a magnetic field component along the y-axis that changes from negative to positive, and the magnetic field distribution of the Bx component achieves the expected obvious symmetry on the left and right sides.

[0102] See also Figures 6 and 7 The decoupling model diagram and S11 parameter diagram of the axial RF coil in Case 1 and the radial RF coil in Case 2, respectively, with a regular solenoid coil added at one end and both ends. Decoupling design was considered when combining them, that is, adding a regular solenoid coil in series, with the current directions of the two solenoid coils in series opposite. The specific decoupling method is:

[0103] Method 1: A conventional solenoid coil (RS1) is connected in series to one end of the axial RF coil (TS1), and a conventional solenoid coil (RS2) is connected in series to one end of the radial RF coil (TS2). The currents in RS1 and RS2 flow in opposite directions.

[0104] Method 2: A conventional solenoid coil (RS1 and RS1') is connected in series at both ends of the axial RF coil (TS1), and a conventional solenoid coil (RS2 and RS2') is connected in series at both ends of the radial RF coil (TS2). The currents in RS1 and RS1' flow in the same direction, and the currents in RS2 and RS2' flow in the same direction, but the currents in RS1 and RS2 flow in opposite directions.

[0105] The modeling formula of a conventional solenoid is x=rcos(2π*s),y=rsin(2π*s),z=pitch*s,s∈[0,N truns], After parametrically scanning the number of turns (N_turns) and turn spacing (pitch) of the added conventional solenoid coil in comsolmultiphysics, it was determined that the simulation effect of the coil was best when N_turns = 4 and pitch = 3cm. The created decoupling model was imported into CST to calculate the S11 parameters. The simulation results showed that both the first and second methods achieved decoupling design for the coil. TRASE, as an independent encoding mechanism, can be used in highly inhomogeneous B0 fields without the need for B1+ phase gradient to dominate the change of B0. Unlike B0 gradient coils, RF coils are physically lightweight and easy to operate, and customized arrays can be designed for each anatomical situation. Because smaller coils are more efficient, a basic TRASE imaging pulse sequence consists of an initial 90 excitation pulse to move the magnetization, It is initially in equilibrium along the z-axis, enters the transverse X, Y plane, and then generates its own radio frequency field, followed by a series of equally spaced 180 refocusing pulses (spin echo sequence). The refocusing pulses in the echo sequence are applied alternately using coils A and B, and nuclear magnetic resonance signals can be observed after each refocusing pulse. Two-dimensional radio frequency phase-encoded imaging is achieved by combining two radio frequency coils encoded in different orthogonal directions, which are composed of axially and radially encoded radio frequency coils and a saddle coil that produces a uniform phase. A 90° pulse excitation is emitted from the saddle coil, and then a 180° refocusing pulse sequence is emitted alternately from the axial radio frequency coil and the radial radio frequency coil, and the magnetic resonance echo signal is received at the saddle coil, thereby achieving two-dimensional imaging encoding in k-space. After the 2D K-space trajectory is excited using the saddle coil C, the orthogonal encoding direction K is used. c -K TS1 and K c -K B Perform a refocus sequence (CB) 2 -(CTS1) 3 -C, wherein coil A represents the axial radio frequency coil TS1 of the case 1, coil B represents the axial radio frequency coil TS2 of the case 2, and coil C represents a saddle coil that generates a uniform phase. Figure 8 Schematic diagram of the K-space trajectory for two-dimensional ultra-low-field magnetic resonance imaging using the transmitting array spatial encoding coil group.

[0106] Finally, it should be noted that the above embodiments are only used to illustrate the technical solutions of the present invention and are not limiting. Although the present invention has been described in detail with reference to the preferred embodiments, those skilled in the art should understand that the technical solutions of the present invention can be modified or replaced by equivalents without departing from the purpose and scope of the technical solutions, which should all be included in the scope of the claims of the present invention.

Claims

1. A method for designing a radio frequency coil for transmit array spatially coded imaging, characterized by: The following steps are involved: S1: Axial encoding coil design: Based on the target field method and magnetic dipole method, at a radius of R z The cylindrical surface design of the RF coil, R z is the cylindrical radius of the axial encoding coil, so that it generates a longitudinal radio frequency phase gradient field in the target area, and the Cartesian coordinate components of the target field satisfy: Among them, B X 、B Y 、B Z Respectively represent the radiofrequency magnetic field components along the X-axis, Y-axis, and Z-axis in the Cartesian coordinate system; L represents half the length of the target area ROI along the Z-axis; z and y are the coordinates of the target point along the Z-axis and Y-axis, respectively; S2: Radial encoding coil design: Based on the target field method and magnetic dipole method, the radius is R r The cylindrical surface design of the RF coil, R r is the cylindrical radius of the radial encoding coil, so that it generates a transverse radio frequency phase gradient field in the target area, and the cylindrical coordinate component of the target field satisfies: Among them, B ρ 、B θ 、B z Respectively represent the radial, circumferential, and axial components of the radio frequency magnetic field in the cylindrical coordinate system; μ0 represents the vacuum magnetic permeability; I represents the coil input current; A represents the winding modulation amplitude of the radial encoding coil; h represents the coil turn spacing; R represents the coil cylindrical radius; ρ represents the radial distance; θ represents the azimuth angle; ψ represents the winding rotation direction angle; S3: Coil combination and decoupling: The axial encoding coil and the radial encoding coil are combined, and decoupling is achieved by connecting a solenoid coil with reverse current in series to form a two-dimensional RF phase encoding system.

2. The radio frequency coil design method for transmit array spatially coded imaging according to claim 1, characterized in that: In S1, the radius R of the axial encoding coil z satisfy: The design of the axial encoding coil includes: 0.8L≤R z ≤1.2L The cylindrical surface is divided into Q grid cells with the stream function S q Describe the current distribution; Solve the stream function S by the least squares method q The optimal distribution of B in the target area Z The phase gradient linearity error of the component is less than 5%; The minimum curvature radius of the stream function contour line is ≥ 2mm.

3. The radio frequency coil design method for transmit array spatially coded imaging according to claim 1, characterized in that: In S2, the radius R of the radial encoding coil r Meet: 6.0cm≤R r ≤7.0cm; The optimal range of winding modulation amplitude A is 4≤A≤7, and the turn spacing h≥3mm.

4. The method for designing a radio frequency coil for transmit array spatially coded imaging according to claim 1, wherein: In S3, the decoupling design adopts any of the following methods: Method 1: The first solenoid coil is connected in series to one end of the axial encoding coil, and the second solenoid coil is connected in series to one end of the radial encoding coil, and the current directions of the two solenoids are opposite; Method 2: A third solenoid coil with the same direction of current is connected in series at both ends of the axial encoding coil, and a fourth solenoid coil with the same direction of current is connected in series at both ends of the radial encoding coil, and the current directions of the third and fourth solenoids are opposite.

5. The radio frequency coil design method for transmit array spatially coded imaging according to claim 1, characterized in that: In S1 and S2, the optimization of the target field is based on the mathematical model: The constraints are: Among them, Q c is the number of grids around the cylindrical surface, S q is the unit flow function.

6. The method for designing a radio frequency coil for transmit array spatially coded imaging according to claim 1, wherein: The target area has a size of 8 cm × 4 cm × 4 cm, and the axial encoding coil radius R z =8.5cm, radial encoding coil radius R r =6.6cm.

7. The radio frequency coil design method for transmit array spatially coded imaging according to claim 1, characterized in that: The method further comprises: integrating the combined radio frequency coil with a saddle coil for transmitting 90° excitation pulses and receiving magnetic resonance signals; The axial encoding coil and the radial encoding coil alternately transmit 180° refocusing pulse sequences to form spin echo encoding.

8. The radio frequency coil design method for transmit array spatially coded imaging according to claim 1, characterized in that: The refocusing pulse sequence is: (CB) 2 -(CTS1) 3 -C, where: C represents a saddle coil, TS1 represents an axial encoding coil, and B represents a radial encoding coil; the encoding direction is K c -K TS1 With K c -K B orthogonal trajectories.

9. The radio frequency coil design method for transmit array spatially coded imaging according to claim 1, characterized in that: The working field strength of the radio frequency coil is 0.2T to 1.0T, which is suitable for ultra-low field magnetic resonance imaging of wrist joints or brain tissues.

10. A radio frequency coil system designed by implementing the method according to any one of claims 1 to 9, characterized in that: include: Saddle coil module C: arranged along the Z-axis direction, used to transmit 90° excitation pulses and receive magnetic resonance echo signals; Axial phase gradient module CTS1: composed of the axial encoding coil, coaxially arranged along the axial direction of the saddle coil, i.e., the Z axis, for generating a Z-axis linear phase gradient; Radial phase gradient module B: composed of the radial encoding coils, arranged orthogonally along the radial direction of the saddle coil, i.e., the Y axis, for generating a Y-axis linear phase gradient; Decoupling module: includes a first reverse solenoid group connected in series at both ends of the axial encoding coil, and a second reverse solenoid group connected in series at both ends of the radial encoding coil, wherein: The current direction of the first reverse solenoid group is opposite to the driving current of the axial encoding coil; The current direction of the second reverse solenoid group is opposite to the driving current of the radial encoding coil; The radio frequency output end of the saddle coil module is connected to the axial / radial phase gradient module through a first switching circuit; The axial encoding coil and the radial encoding coil are alternately turned on by a second switching circuit to switch transmission in a 180° refocusing pulse sequence; The signal receiving end of the saddle coil module is independently connected to the data acquisition unit of the magnetic resonance imaging system.

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