A low-cost mobile ultra-low-field magnetic resonance imaging device

Abstract

This invention relates to a low-cost, mobile, ultra-low-field magnetic resonance imaging (MRI) device, belonging to the field of medical electromagnetic detection technology. The device includes a spoke-shaped magnet, gradient coils, shimming coils, and a radio frequency (RF) coil. The spoke-shaped magnet comprises upper and lower magnetic rings, or may be an assembly of upper and lower magnetic rings. The gradient coil consists of three sets of biplane coils, with the gradient magnetic field direction aligned with the main magnetic field direction, superimposed on the main magnetic field to achieve spatial imaging encoding. Multiple sets of shimming coils are used, each canceling a harmonic magnetic field of the main magnetic field. The RF coil is a transceiver solenoid coil that generates the RF excitation magnetic field and receives the magnetic resonance signal in a time-division multiplexing manner. This invention features a simple structure, small size, light weight, low cost, and high detection sensitivity. The main magnetic field B0 is less than or equal to 50 mT, enabling MRI monitoring of brain diseases and limb joints.

Description

Technical Field

[0001] The invention belongs to the technical field of medical imaging detection and relates to a low-cost mobile ultra-low field nuclear magnetic resonance imaging device. Background Art

[0002] Currently, stroke detection mainly relies on imaging equipment such as CT and MRI, but they cannot meet the needs of continuous monitoring, emergency evacuation, and community medical care. Ultra-low field nuclear magnetic resonance imaging technology is based on the non-contact and non-harmful characteristics of traditional nuclear magnetic resonance imaging technology, combined with the characteristics of lightweight and continuous monitoring after the main magnetic field is reduced. It is one of the alternative methods for stroke detection. Therefore, ultra-low field nuclear magnetic resonance technology has great application prospects.

[0003] At present, ultra-low field MRI equipment mainly uses C-type magnets or H-type magnets, and uses yoke iron to form a magnetic circuit, so the weight often exceeds 1 ton, which greatly affects the lightweight design and mobile detection of ultra-low field MRI. At the same time, the design of traditional gradient coils and shim coils adopts the target field method and residual function method, which requires the selection of target field points and the setting of target field values. Summary of the invention

[0004] In view of this, an object of the present invention is to provide a low-cost mobile ultra-low field magnetic resonance imaging device suitable for brain diseases, which has a simple structure, a small size, a light weight and reliable performance.

[0005] In order to achieve the above object, the present invention provides the following technical solutions:

[0006] A low-cost mobile ultra-low field nuclear magnetic resonance imaging device adopts a spoke-shaped magnet and a harmonic coefficient method based on a target field method to design a gradient coil and a shim coil. The device comprises a spoke-shaped magnet 1, a gradient coil, a shim coil 2 and a radio frequency coil 3;

[0007] The spoke-shaped magnet 1 comprises an upper magnetic ring 4 and a lower magnetic ring 5, which are respectively located at the upper and lower ends of the device, and a main magnetic field is generated between the two magnetic rings, which points from the center of the upper magnetic ring to the center of the lower magnetic ring, and a uniform magnetic field is generated in the geometric center area of ​​the two magnetic rings;

[0008] The gradient coils are three groups of dual-plane coils, including a Z-direction gradient coil 7, a Y-direction gradient coil 8 and an X-direction gradient coil 9; the Z-direction gradient coil 7 is located between the upper magnetic ring 4 and the lower magnetic ring 5; the Y-direction gradient coil 8 is located between the upper and lower planes of the Z-direction gradient coil 7; the X-direction gradient coil 9 is located between the upper and lower planes of the Y-direction gradient coil 8; the Z-direction gradient coil 7, the Y-direction gradient coil 8 and the X-direction gradient coil 9 generate gradient magnetic fields in the Z, Y and X directions respectively, and the directions of the gradient magnetic fields are the same as the directions of the main magnetic fields, and are superimposed on the main magnetic fields to realize spatial imaging encoding;

[0009] The shim coils 2 are multiple groups of double-plane coils, located between the upper and lower planes of the X-direction gradient coils 9; the magnetic field generated by each group of shim coils 2 offsets a harmonic magnetic field of the main magnetic field;

[0010] The radio frequency coil 3 is located between the upper and lower planes of the shim coil 2, and generates radio frequency excitation magnetic fields and receives magnetic resonance signals in a time-sharing manner.

[0011] Preferably, the upper magnetic ring 4 or the lower magnetic ring 5 is composed of magnetic bars, wherein the magnetization direction of the magnetic bars of the upper magnetic ring 4 is radially pointing toward the center of the circle, and the magnetization direction of the magnetic bars of the lower magnetic ring 5 is radially away from the center of the circle.

[0012] Preferably, the upper magnetic ring 4 can be replaced by an upper magnetic ring group, and the lower magnetic ring 5 can be replaced by a lower magnetic ring group, wherein the upper magnetic ring group is composed of multiple magnetic rings, different magnetic rings have different radii, and different magnetic rings have different sizes of magnetic rods, and the magnetization direction of the magnetic rods in the upper magnetic ring group is radially pointing to the center of the circle; the lower magnetic ring group is composed of multiple magnetic rings, the radii of the magnetic rings and the sizes of the magnetic rods of different magnetic rings are symmetrical with those of the upper magnetic ring group about a plane passing through the origin, and the magnetic ring direction of the magnetic rods in the lower magnetic ring group is radially away from the center of the circle.

[0013] Preferably, the magnetic bar is a cylindrical magnetic bar or a rectangular magnetic bar.

[0014] Preferably, the equivalent magnetic charge method is used to quickly calculate the magnetic field distribution of the main magnetic field generated by two magnetic rings or magnetic ring groups, so as to optimize the number and spacing of the magnetic bars.

[0015] Preferably, the upper and lower planes of the gradient coil are designed using a harmonic coefficient method based on a target field method, and the upper and lower planes have the same structure and the currents have the same direction.

[0016] Preferably, the shim coil 2 is designed using a harmonic coefficient method based on a target field method. The current density is first expressed as a Fourier series, which is then simplified into a series form by substituting the Biot-Savart law into the target harmonic constant coefficients to obtain the current density in the plane. Finally, the structure of the shim coil 2 is obtained by discretizing the stream function contour lines.

[0017] Preferably, the current distribution in the upper and lower planes of the shim coil 2 is the same; when l+k is an even number, the magnetic field B of the shim coil 2 is z There is an even symmetry about the z=0 plane, and the currents in the upper and lower planes are in the same direction; when l+k is an odd number, the magnetic field B of the shim coil 2 z There is an odd symmetry about the z=0 plane, and the currents in the upper and lower planes are opposite; wherein l and k represent the order and degree of the shim coil 2 respectively.

[0018] Preferably, the radio frequency coil 3 is a solenoid coil integrating transmission and reception.

[0019] Preferably, the spoke-shaped magnet 1, gradient coil, shim coil 2 and radio frequency coil 3 are made of a 3D printed integrated structure; and the supporting shell is made of non-magnetic engineering plastic.

[0020] The beneficial effects of the present invention are as follows: the present invention adopts spoke-shaped magnets and a harmonic coefficient method based on a target field method to design gradient coils and shim coils, which can overcome the disadvantages of traditional designs, and at the same time has the characteristics of lightweight, non-invasive, and continuous monitoring, and is suitable for continuous monitoring of patients with brain diseases. The device of the present invention has a simple structure, a small size, lightweight, reliable performance, and has broad application prospects.

[0021] Other advantages, objectives and features of the present invention will be described in the following description to some extent, and to some extent, will be obvious to those skilled in the art based on the following examination and study, or can be taught from the practice of the present invention. The objectives and other advantages of the present invention can be realized and obtained through the following description. BRIEF DESCRIPTION OF THE DRAWINGS

[0022] 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 in conjunction with the accompanying drawings, wherein:

[0023] Figure 1 It is an overall schematic diagram of a low-cost mobile ultra-low-field magnetic resonance imaging device;

[0024] Figure 2 is a schematic diagram of equivalent magnetic charge;

[0025] Figure 3 This is a schematic diagram of the magnet distribution of two magnetic rings of a spoke-shaped magnet. Take a magnetic ring containing 8 rectangular magnets as an example. Figure 3 (a) is the magnet distribution and polarization direction of the upper magnetic ring. Figure 3 (b) is the magnet distribution and polarization direction of the lower magnetic ring;

[0026] Figure 4 The magnetic induction intensity of a certain magnetic bar at a certain position in space is equivalent to the magnetic induction intensity of a rectangular magnetic bar along the positive direction of the X-axis in space;

[0027] Figure 5 It is a schematic diagram of the upper and lower magnetic ring groups of the spoke-shaped magnet;

[0028] Figure 6 It is a schematic diagram of the lower plane of three sets of gradient coils. Figure 6 (a) is the X-direction gradient coil, Figure 6 (b) is the Y-direction gradient coil, Figure 6 (c) is the Z-direction gradient coil;

[0029] Figure 7 is the shim coil structure, Figure 7 (a) is the z2-(x2+y2) / 2 shim coil structure, Figure 7 (b) is the xz shim coil structure, Figure 7 (c) is the x2-y2 shim coil structure, Figure 7 (d) is the 2xy shim coil structure.

[0030] Figure numerals: 1 - spoke-shaped magnet; 2 - shim coil; 3 - radio frequency coil; 4 - upper magnetic ring; 5 - lower magnetic ring; 6 - imaging target area; 7 - Z-direction gradient coil; 8 - Y-direction gradient coil; 9 - X-direction gradient coil. DETAILED DESCRIPTION

[0031] The following describes the embodiments of the present invention by 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 only illustrate the basic concept of the present invention in a schematic manner, and the following embodiments and features in the embodiments can be combined with each other without conflict.

[0032] Among them, the drawings are only used for illustrative explanations, and they only represent schematic diagrams rather than actual pictures, and should not be understood as limitations on the present invention. In order to better illustrate the embodiments of the present invention, some parts of the drawings may be omitted, enlarged or reduced, and do not represent the size of actual products. For those skilled in the art, it is understandable that some well-known structures and their descriptions in the drawings may be omitted.

[0033] 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 the terms "upper", "lower", "left", "right", "front", "rear", etc. indicate the orientation or position relationship, they are based on the orientation or position relationship shown in the drawings, which is only for the convenience of describing the present invention and simplifying the description, rather than indicating or implying that the device or element referred to must have a specific orientation, be constructed and operate in a specific orientation. Therefore, the terms describing the position 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.

[0034] Figure 1The figure shows an overall schematic diagram of a low-cost mobile ultra-low field nuclear magnetic resonance imaging device, which includes a spoke-shaped magnet 1, a gradient coil, a shim coil 2, a radio frequency coil 3, an upper magnetic ring 4, a lower magnetic ring 5, an imaging target area 6, a Z-direction gradient coil 7, a Y-direction gradient coil 8, and an X-direction gradient coil 9. The spoke-shaped magnet includes a radially magnetized upper magnetic ring and a lower magnetic ring, wherein the magnetization direction of the magnetic rod of the upper magnetic ring points to the center of the circle along the radial direction, and the magnetization direction of the magnetic rod of the lower magnetic ring points away from the center of the circle along the radial direction. The two magnetic rings are spaced a certain distance apart, so that a main magnetic field B0 is generated between the upper and lower magnetic rings from the center of the upper magnetic ring to the center of the lower magnetic ring, and a uniform magnetic field is generated in the geometric center area of ​​the two magnetic rings (i.e., the imaging target area 7). The main magnetic field B0 is less than or equal to 50 mT. The magnetic field distribution generated by the two magnetic rings is quickly calculated using the equivalent magnetic charge method to optimize the number and spacing of the magnetic rods, and the magnetic rods can be cylindrical or rectangular magnets. The magnets currently used are mainly uniformly magnetized along one magnetization direction, so the volume magnetic charge density in the magnet is always zero, and the surface magnetic charge density is σ m (r′)=M·n, where n is the normal unit vector of the surface. The surface magnetic charge density of the permanent magnet is affected by the magnitude and direction of the magnetization intensity. The relationship between the scalar magnetic potential and the surface magnetic charge density is:

[0035]

[0036] Where r is the source point, r' is the field point, and S is the boundary of the permanent magnet containing the surface magnetic charge. According to electromagnetic theory, there is the following relationship between magnetic induction intensity, magnetic field intensity and scalar magnetic potential: H = -Δφ m , B = μH. Then according to formula (1):

[0037]

[0038] Where B + , B - are the magnetic induction intensities generated by positive and negative magnetic charges, S + , S _ are the planes where the positive magnetic charge and the negative magnetic charge are located respectively. Taking a simple rectangular permanent magnet as an example, Figure 2 As shown, the magnetic induction intensity in the X direction generated by the equivalent positive magnetic charge can be obtained by formula (2):

[0039]

[0040] Where dS = dηdε. Similarly, we have:

[0041]

[0042] Combining equations (2), (3) and (4), we can get the magnetic induction intensity B x =B x+-B x- , similarly, we can get B y =B y+ -B y- , B z =B z+ -B z- , where B y+ , B y- , B z+ , B z- The calculation formula is as follows:

[0043]

[0044]

[0045] By synthesizing the magnetic field strengths in three directions, we can get the distribution of the magnetic induction intensity of the rectangular permanent magnet in space, and then get the distribution of the magnetic induction intensity of the spoke-shaped magnet ring. For the two magnetic rings of the spoke-shaped magnet, the magnetic induction intensity of the magnetic ring is equivalent to the magnetic field synthesis of 2n rectangular magnetic bars. The above formula is used to calculate the magnetic field distribution of each magnetic bar in space. Finally, all the magnetic bars are deflected according to the angle, and the magnetic fields generated are combined. For example, a magnetic ring composed of 8 rectangular magnetic bars is used. Figure 3 (a) Figure 3 (b) as shown.

[0046] Since each magnetic bar in the upper and lower magnetic rings can be approximately equivalent to uniform magnetization in a single direction, there is only equivalent surface magnetic charge. According to the integral formula of the equivalent magnetic charge method, the magnetic field of the rectangular magnetic bar in the positive direction of the X-axis in space can be calculated using equations (1) to (8). For other directions, an equivalent method can be used to map the magnetic field intensity of any rectangular magnetic bar in the rectangular magnetic bar group in space to the magnetic induction intensity of the rectangular magnetic bar in the positive direction of the X-axis, as follows: Figure 4 As shown. The magnetic induction intensity of the i-th magnetic bar at a certain position in space is equivalent to the magnetic induction intensity of the rectangular magnetic bar in space along the positive direction of the X-axis. The magnetic induction intensity of each magnetic bar in the upper and lower magnetic rings is vector-synthesized to obtain the magnetic field intensity distribution of the two magnetic rings of the spoke-shaped magnet in space. i ,y i ) is transformed into:

[0047]

[0048] θ is the angle between the axes of the two rectangular magnetic bars on the upper magnetic ring or the lower magnetic ring. The magnetic field strength of the upper magnetic ring of the spoke-shaped magnet in space can be calculated by the above method. For the lower magnetic ring, since the magnetization direction of the magnetic bar is opposite to that of the upper magnetic ring, the positive and negative charge density distribution is opposite to that of the upper magnetic ring. Therefore, when calculating the magnetic induction intensity, equations (3) to (8) should take an additional negative sign.

[0049] The spoke-shaped magnet can also be in the form of an upper and lower magnetic ring group, such as Figure 5 As shown, the upper magnetic ring group can be composed of multiple magnetic rings, the radii of each magnetic ring are different, and the sizes of magnetic rods of different magnetic rings can also be different. The magnetization direction of the magnetic rods of the upper magnetic ring group is radially pointing to the center of the circle; the lower magnetic ring group can be composed of multiple magnetic rings, the radii of the magnetic rings, the sizes of magnetic rods of different magnetic rings are symmetrical with the upper magnetic ring group about the plane passing through the origin, and the magnetization direction of the magnetic rods of the lower magnetic ring group is radially away from the center of the circle.

[0050] The gradient coil uses three sets of dual-plane (upper plane and lower plane) gradient coils. The three sets of dual-plane gradient coils include Z-direction gradient coils, Y-direction gradient coils and X-direction gradient coils. Their gradient magnetic field directions are the same as the main magnetic field directions. They are superimposed on the main magnetic field to achieve spatial imaging encoding. The dual-plane gradient coils are designed using the harmonic coefficient method based on the target field method. The design of the gradient coil can be regarded as a uniform field coil with a first-order harmonic coefficient. The upper and lower plane coils have exactly the same structure and the currents are in the same direction. The gradient coils designed using the harmonic coefficient method based on the target field method are as follows: Figure 6 (a)~ Figure 6 (c) as shown.

[0051] There are multiple groups of shim coils, which are also designed using the harmonic coefficient method based on the target field method. Each group of shim coils offsets a harmonic magnetic field of the main magnetic field. The design method is to first express the current density as a Fourier series, substitute the Biot-Savart law to simplify it into a series form, and then solve it correspondingly with the target harmonic field coefficient. Finally, the current density in the plane is obtained, and the structure of the coil can be obtained. The harmonic coefficient method based on the target field method is used to design shim coils. The relationship between the harmonic field coefficient and the harmonic current coefficient is directly established. As long as a field harmonic in the target space is specified, the corresponding current density can be directly calculated. Then, the structure of the shim coil is obtained by discretizing the stream function contour line. It is no longer necessary to select the target field point and set the target field value. It combines the advantages of both the inverse method and the forward method. The current distribution in the upper and lower planes is the same, but the current direction varies according to different situations. When l+k is an even number (odd number), the coil magnetic field Bz has even (odd) symmetry about the z=0 plane, and the currents on the upper and lower surfaces are in the same (opposite) direction. l and k represent the order and degree of the coil respectively. The typical second-order shim coil structure is as follows: Figure 7 (a)~ Figure 7 (d) as shown.

[0052] The radio frequency coil adopts an integrated transmitting and receiving solenoid coil. The advantage of the radio frequency coil is that the radio frequency magnetic field generated in the internal central area is uniform, and the radio frequency excitation magnetic field B1 is generated in a time-sharing manner to receive magnetic resonance signals.

[0053] Finally, it should be noted that the above embodiments are only used to illustrate the technical solution of the present invention rather than to limit it. 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 solution of the present invention can be modified or replaced by equivalents without departing from the purpose and scope of the technical solution, which should be included in the scope of the claims of the present invention.

Claims

1. A low-cost mobile ultra-low field nuclear magnetic resonance imaging device, characterized in that: It comprises a spoke-shaped magnet (1), a gradient coil, a shim coil (2) and a radio frequency coil (3); The spoke-shaped magnet (1) comprises an upper magnetic ring (4) and a lower magnetic ring (5), wherein the upper magnetic ring and the lower magnetic ring are respectively located at the upper and lower ends of the device; the upper magnetic ring (4) and the lower magnetic ring (5) are both composed of magnetic bars, wherein the magnetization direction of the magnetic bar of the upper magnetic ring (4) is radially pointing to the center of the circle, and the magnetization direction of the magnetic bar of the lower magnetic ring (5) is radially away from the center of the circle; The gradient coils are three groups of dual-plane coils, including a Z-direction gradient coil (7), a Y-direction gradient coil (8) and an X-direction gradient coil (9); the Z-direction gradient coil (7) is located between the upper magnetic ring (4) and the lower magnetic ring (5); the Y-direction gradient coil (8) is located between the upper and lower planes of the Z-direction gradient coil (7); and the X-direction gradient coil (9) is located between the upper and lower planes of the Y-direction gradient coil (8); the Z-direction gradient coil (7), the Y-direction gradient coil (8) and the X-direction gradient coil (9) respectively generate gradient magnetic fields in the Z, Y and X directions, and the directions of the gradient magnetic fields are the same as the directions of the main magnetic fields, and are superimposed on the main magnetic fields to achieve spatial imaging encoding; The shim coils (2) are multiple groups of double-plane coils, located between the upper and lower planes of the X-direction gradient coils (9); the magnetic field generated by each group of shim coils (2) offsets a harmonic magnetic field of the main magnetic field; The shim coil (2) is designed by using a harmonic coefficient method based on a target field method. The current density is first expressed as a Fourier series, which is then simplified into a series form by substituting the Biot-Savart law into the series form. The shim coil (2) is then solved correspondingly with the target harmonic constant coefficient, that is, a field harmonic in the target space is specified to directly calculate the corresponding current density. Finally, the structure of the shim coil (2) is obtained by discretizing using stream function contour lines. The current distribution in the upper and lower planes of the shim coil (2) is the same; when l+k is an even number, the magnetic field B of the shim coil (2) z There is an even symmetry about the z=0 plane, and the currents in the upper and lower planes are in the same direction; when l+k is an odd number, the magnetic field B of the shim coil (2) z There is an odd symmetry about the z=0 plane, and the currents in the upper and lower planes are opposite; wherein l and k represent the order and degree of the shim coil (2) respectively; The radio frequency coil (3) is located between the upper and lower planes of the shim coil (2), and generates a radio frequency excitation magnetic field and receives magnetic resonance signals in a time-sharing manner.

2. A low-cost mobile ultra-low field nuclear magnetic resonance imaging device according to claim 1, characterized in that: The upper magnetic ring (4) can be replaced by an upper magnetic ring group, and the lower magnetic ring (5) can be replaced by a lower magnetic ring group; wherein the upper magnetic ring group is composed of a plurality of magnetic rings, different magnetic rings have different radii, and different magnetic rings have different sizes of magnetic rods, and the magnetization direction of the magnetic rods of the upper magnetic ring group is radially pointing to the center of the circle; the lower magnetic ring group is composed of a plurality of magnetic rings, the radii of the magnetic rings and the sizes of the magnetic rods of different magnetic rings are symmetrical with the upper magnetic ring group about a plane passing through the origin, and the magnetic ring direction of the magnetic rods of the lower magnetic ring group is radially away from the center of the circle.

3. A low-cost mobile ultra-low field nuclear magnetic resonance imaging device according to claim 1 or 2, characterized in that: The magnetic bar is a cylindrical magnetic bar or a rectangular magnetic bar.

4. A low-cost mobile ultra-low field nuclear magnetic resonance imaging device according to any one of claims 1 or 2, characterized in that: The equivalent magnetic charge method is used to calculate the magnetic field distribution of the main magnetic field in order to optimize the number and spacing of the magnetic bars.

5. The low-cost mobile ultra-low field nuclear magnetic resonance imaging device according to claim 1, characterized in that: The upper and lower planes of the gradient coil are designed using the harmonic coefficient method based on the target field method. The upper and lower planes have the same structure and the current has the same direction.

6. The low-cost mobile ultra-low field nuclear magnetic resonance imaging device according to claim 1, characterized in that: The radio frequency coil (3) is a solenoid coil with integrated transmission and reception.

7. The low-cost mobile ultra-low field nuclear magnetic resonance imaging device according to claim 1, characterized in that: The spoke-shaped magnet (1), gradient coil, shim coil (2) and radio frequency coil (3) adopt a 3D printing integrated structure; the supporting shell adopts non-magnetic engineering plastic.

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