JP2592920B2 - Superconducting magnet for magnetic resonance imaging - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION [Object of the Invention] (Industrial Application Field) The present invention relates to a specific atomic nucleus in a specific section of a living body using a magnetic resonance (MR) phenomenon. The present invention relates to a superconducting magnet of a magnetic resonance imaging apparatus for imaging a density distribution of CT as a CT image (Com-pnted Tomgram).

(Prior Art) For example, in a medical magnetic resonance imaging apparatus used for a living body diagnosis, in order to obtain a tomographic image of a specific part of a subject, which is a living body, the subject P is moved in a Z direction as shown in FIG. It is generated by the static magnetic field magnet (not shown) very uniform static magnetic field H ₀ along to act further pair of gradient coils 100A, 100B by applying a linear magnetic field gradient G _X to the static magnetic field H _0. Here, the specific nucleus with respect to the static magnetic field H ₀ resonates at an angular frequency ω ₀ represented by the following equation.

ω ₀ = γH ₀ (1) In the equation (1), γ is a gyromagnetic ratio, which is specific to the type of atomic nucleus. Accordingly Furthermore, the angular frequency omega ₀ of the rotating magnetic field H ₁ and RF coil (probe head) transmitting the example of the pair is provided in the coil 200A to resonance only certain nuclei, to act on the patient P via 200B.

In this way, the linear field gradient G _X by acting selectively only for illustrated the x-y plane moieties selected set for the Z-axis direction, is a portion on a specific slice portion S (plane to obtain a tomographic image Only has a certain thickness), the magnetic resonance phenomenon occurs. This magnetic resonance phenomenon is
For example, a pair of receiving coils 300A, 300 provided in the coil
The signal is observed as a free induction decay (hereinafter abbreviated as “FID signal”) via B and used as an MR signal. By subjecting this FID signal to Fourier transform, a single spectrum is obtained for the rotation frequency of a specific nuclear spin.

To obtain a tomographic image as a CT image, x-
Projection in multiple directions in the y-plane is required. Therefore, after exciting the slice portion S to cause a magnetic field resonance phenomenon, as shown in FIG. 8, the magnetic field H ₀ is linearly inclined in the x′-axis direction (a coordinate system rotated by an angle θ from the x-axis). When the linear magnetic field gradient G _XY is applied by a gradient coil (not shown), the isomagnetic field line E in the slice portion S of the subject P becomes a straight line. The rotation frequency of the specific nuclear spin on the isomagnetic field line E is expressed by the above equation (1).

For convenience of explanation, an equal magnetic field lines E and E ₁ to E _n, considered to result signals D ₁ to D _n which is a type of FID signals, respectively by the magnetic field on these such as magnetic field lines E ₁ to E _n. The amplitude of the signal D ₁ to D _n is proportional to the specific nuclear spin density on such field lines E ₁ to E _n each penetrating the slice portion S. However, the FID signal that is actually observed is a synthetic FID signal combined adding all signals D ₁ to D _n. Therefore, projection information (one-dimensional image) PD on the x'-axis of the slice portion S is obtained by Fourier-transforming the composite FID signal.

Next, this x'-axis is rotated in the xy plane. This is, for example, a magnetic field gradient G _XY as a composite magnetic field of the magnetic field gradients G _X , G _{Y in} the x, y directions by two pairs of gradient coils.
And changing the combined ratio of the magnetic field gradients G _X and G _Y. By the rotation of the magnetic field gradient _GXY , projection information in the angular direction on the xy plane is obtained in the same manner as described above, and a CT image is synthesized based on the information.

The above is the principle of magnetic resonance imaging. Next, as a specific example, FIG. 9 shows a conventional magnetic resonance imaging apparatus. A subject, that is, a patient 1 is placed on a bed 2. An RF coil (probe bed: high-frequency transmission / reception coil) 3 surrounding the patient 1, and a shim coil 4 for correcting a magnetic field and a gradient coil 5 for generating a gradient magnetic field are arranged on the outer periphery thereof. All of these coil systems are provided with a normal temperature bore 7 of a large static magnetic field magnet 6 (normally, the bore diameter is about
1m) housed inside. One of a superconducting magnet, a normal conducting magnet, and a permanent magnet is used as the static magnetic field magnet.

The static magnetic field magnet 6 is connected to a current lead 9 by an excitation power supply 8.
(In the case of a permanent magnet system, this is unnecessary). In the case of a superconducting magnet, the current lead 9 is usually removed after excitation in order to operate in the permanent current mode and to reduce the consumption of liquid helium as a refrigerant, so that a magnetic field is always generated. ing. Usually, the direction of the static magnetic field is the illustrated 10 directions for many magnets, that is, the direction of the body axis of the patient 1. The gradient coil 5 includes a GX coil considering a magnetic field gradient in the X-axis direction, a GY coil in the Y-axis direction, and a GZ coil in the Z-axis direction. These excitation power supplies 11, 12, 13 are connected to a central control unit 14. The RF coil 3 includes a transmission coil and a reception coil, and is connected to an RF oscillation device 15 and an RF reception device 16, respectively, and these are further connected to a central control device 14. Central control unit 14
Is connected to the display / operation panel 17 and is operated by this.

Next, the operation of the conventional magnetic resonance imaging apparatus configured as described above will be described.

In order to obtain a whole-body tomographic image of the patient 1, a uniform magnetic field 18
Is generally as large as 40 to 50 cm sphere, and high uniformity of 50 ppm or less is required. Therefore, for example, in the case of the superconducting system, the static magnetic field magnet 6 has a length of 2.4 m, a width of 2 m, a height of 2.4 m, and a weight of 5 to 6 m.
Tons and huge things are required.

Even with such a large magnet, the uniformity within a sphere of 40 to 50 cm using only the magnet is at most several hundred ppm. In order to reduce this to 50 ppm or less, a shim coil 4 for magnetic field correction is used. The diagnosis site of the patient is brought into the uniform magnetic field space 18. Then, a high frequency is applied by the RF oscillator 15 and the RF coil 3 in a direction perpendicular to the static magnetic field 10 to excite required nuclei in the human body cells, for example, hydrogen nuclei. At the same time, GX excitation power supply 11 and GY excitation power supply 1
2. A gradient magnetic field is applied in the X, Y, and Z directions by the GZ excitation power supply 13 and the gradient coil 5.

For the RF and gradient pulse sequences, an optimal method is selected depending on the lesion site and the image processing direction.

This pulse sequence operation is controlled by the central controller 14. After the application of the gradient and the RF, a magnetic resonance signal is emitted from the body of the patient 1. This signal is the RF receiver
The signal is received and amplified by 16 and input to the central control unit 14. The image is processed here, and the required human tomographic image is displayed and
It is displayed on the CRT of the operation panel 17.

FIG. 10 shows the internal structure of a conventional superconducting magnet for a magnetic resonance imaging apparatus.

The superconducting coil 19 is housed in a helium container 20 filled with liquid helium and kept at a cryogenic temperature of 4K. The outer periphery of the helium container 20 is surrounded by a double radiant heat shield plate, that is, a 20K radiant heat shield plate 21 and an 80K radiant heat shield plate 22, and shields external heat with the shield plate. To further enhance the heat insulating effect, a vacuum is maintained between the helium container and the radiation heat shield plate, and a heat insulating material formed of an alumina / polyester film is inserted. These members are housed in a vacuum vessel 23.

When the gradient coil is excited for image processing, an eddy current is generated in the radiation heat shield plates 21 and 22 made of an electrically conductive metal such as aluminum due to the pulse magnetic field.

Since a part of the energy applied to the gradient coil is used for the generation of the eddy current, the gradient pulse magnetic field which normally has to rapidly rise in 1 to 2 ms becomes dull. Further, the local magnetic field generated by the eddy current disturbs the uniform magnetic field at the center of the normal temperature bore. For this reason, the image deteriorates. Further, the gradient pulse magnetic field couples with the shim coil, which also degrades the image.

There is a technique using an active gradient coil to avoid the eddy current and the coupling problem with the shim coil. About this technology, for example, IEEE TRANS
ACTIONS ON MAGNETICS, November 1985, Vol MAG-21, No
6,2273−2275 A Finite Fourier Element Expansion Te
chnique for The Design of a Pulsed Radial Gradient
It is described in detail in System for Magnetic Resonance Imaginmg (MRI). FIG. 10 shows this structure.

A cancel gradient coil 25 is arranged concentrically outside the main gradient coil 24, thereby forming an active gradient coil 26. The active gradient coil 26 is disposed inside the shim coil 4 and substantially concentrically with the shim coil 4.

The main gradient coil 24 and the cancel gradient coil 25 have the opposite polarity of the pulse magnetic field generated.Image processing is performed inside the main gradient coil 24 by adjusting parameters such as ampere-turn and diameter of both coils. In this case, a gradient pulse intensity necessary for the above is obtained, and outside the cancel gradient coil 25, the pulse magnetic field is canceled and becomes zero. Since the pulse magnetic field becomes zero outside the active gradient coil 26, an eddy current is generated and the shim coil 4
And the image quality is improved.

(Problems to be Solved by the Invention) The conventional superconducting magnet with an active gradient coil configured as described above has the following disadvantages.

(1) In order to reduce the power consumption, it is necessary to sufficiently separate the main gradient coil 24 and the cancel gradient coil 25 from each other in the radial direction. For example, if the inner diameter of the main gradient coil is D ₁ and the inner diameter of the cancel coil is D ₂ , it is desirable that D ₁ / D ₂ <0.8. Otherwise, to cancel the external magnetic field of the main gradient coil, the magnetic field energy of the cancel gradient coil must be equal to that of the main gradient coil. That is, the coil capacity and the exciting power supply capacity are increased, and a huge gradient coil system is required.

(2) When diverting the currently used RF coil to secure a patient insertion space, it is necessary to make the inner diameter of the main gradient coil almost the same as the inner diameter of the currently used single gradient coil. In this case, the outermost diameter of the active gradient coil is 1.2 to 1.5 times that of the conventional single gradient coil, and it is not possible to store the active gradient coil in a normally used 1 mφ bore superconducting magnet. For storage, the bore diameter of the superconducting magnet must be increased. This is an increase in size of the magnet, which significantly detracts from advantages in terms of cost and installation. This has hindered the spread of MPI.

SUMMARY OF THE INVENTION It is an object of the present invention to provide a superconducting magnet for a magnetic resonance imaging apparatus with an active gradient coil which is compact and does not generate eddy currents, eliminating the above-mentioned disadvantages of the prior art.

[Configuration of the invention]

(Means for Solving the Problems) In order to achieve the above object, in the present invention, a cancel gradient coil is arranged between the outermost radiant heat shield plate (for example, 80K radiant heat shield plate) surrounding the superconducting coil and the normal temperature bore. The main gradient coil is arranged inside the room temperature bore.

(Operation) Since the external magnetic field generated by the main gradient coil is canceled by the cancel gradient coil, no eddy current is generated in the radiation heat shield plate of the superconducting magnet.

Further, since the cancel gradient coil is disposed outside the normal temperature bore, the superconducting magnet does not increase in size.

(Embodiment) An embodiment of the present invention will be described with reference to FIG.

(Configuration of Embodiment) A cancel gradient coil 25 is arranged concentrically with the superconducting coil 19 between the room temperature bore 7 and the 80K radiation heat shield plate 22 inside the superconducting magnet. In addition, room temperature bore 7
The main gradient coil 24 is arranged concentrically with the cancel gradient coil 25. An iron shim 27 is attached to the inner surface of the normal temperature bore 7.

(Operation of the Embodiment) Since the external magnetic field generated by the main gradient coil 24 is canceled by the cancel gradient coil 25, no pulse magnetic field acts on the 80K radiant heat shield plate 22 and no eddy current is generated.

Further, since the iron shim 27 is disposed between the main gradient coil and the cancel gradient coil, there is no electrical coupling between the gradient coil and the iron shim 27.

(Effects of Embodiment) (1) A high quality image can be obtained because there is no influence of eddy current.

(2) Since the main gradient coil can be configured in the same size as the conventional gradient coil, it can be housed inside the magnet bore, so that the superconducting magnet has the same size as the conventional one. Therefore, the superconducting magnet due to the active gradient coil is prevented from being enlarged. This also avoids cost increases.

(3) When a superconducting magnet having a conventional size is used, the conventional active gradient coil is housed inside the magnet bore, but if the inside diameter of the conventional active gradient coil is not reduced, it is housed inside. It was difficult. Reducing the inner diameter of the active gradient coil reduces the patient insertion space and further increases the patient's sense of claustrophobia.

If the active gradient coil of the present embodiment is used, this disadvantage is avoided because the inner diameter of the main gradient coil is the same as that of the conventional active gradient coil.

Another Embodiment 1 Another embodiment 1 of the present invention will be described with reference to FIG.

(Configuration of Another Embodiment 1) The superconducting magnet bore inner cylinder 28 disposed on the inner peripheral side of the radiant heat shield plate closest to the room temperature bore inside the superconducting magnet insulated container of the static magnetic field magnet 6 cancels the winding frame of the gradient coil 25 Also used as In this case, the cancel gradient coil 25 may be wound on the outer circumference of the superconducting magnet bore inner cylinder 28 made of FRP, or may be wound on the inner circumference side to form a coil. Other configurations are the same as those of the embodiment shown in FIG.

(Operation of Another Embodiment 1) The operation is the same as that of the embodiment shown in FIG.

(Effects of Another Embodiment 1) The winding frame of the cancel gradient coil and the supporting structure are not required, and the structure is simplified.

 Other effects are the same as those of the embodiment shown in FIG.

Second Embodiment Another embodiment 2 of the present invention will be described with reference to FIG.

(Configuration of Other Embodiment 2) Superconducting magnet, that is, 80K radiation heat shield plate 2
The cylindrical portion 2 is also used as the winding frame of the cancel gradient coil 25. In this case, the cancel gradient coil is
Wound around the inner surface of the cylindrical part of the 80K radiation heat shield plate 22.

 Other configurations are the same as those of the embodiment shown in FIG.

(Operation of Another Embodiment 2) The cancel gradient coil is cooled to the same cryogenic temperature as the radiation heat shield plate.

 Other operations are the same as those of the embodiment shown in FIG.

(Effects of Another Embodiment 2) The winding frame of the cancel gradient coil and the supporting structure are not required, and the structure is simplified. Further, since the cancel gradient coil conductor is in an extremely low temperature state, the electric resistance is reduced. As a result, the joule loss during excitation of the gradient coil is reduced, so that the excitation power supply capacity for the cancel gradient coil can be reduced. Therefore, a compact and low-cost gradient coil can be provided.

Third Embodiment Another embodiment 3 of the present invention will be described with reference to FIG.

When performing spectroscopy diagnosis and ultra-high-speed imaging, automating with a current shim coil is necessary to correct the variation in magnetic field uniformity due to in-body magnetism of a patient.

In this case, an auto shim coil must be attached to the magnet separately from the iron shim. The present embodiment is an example in this case.

(Configuration of Another Embodiment 3) In another embodiment 1 shown in FIG. 2, the current shim coil 4 (auto shim coil) is wound inside or outside the inner cylindrical portion of the outermost radiation shield plate 22.

(Operation of Other Embodiment 3) Automating can be performed by the current shim coil. Other operations are the same as those of the first embodiment.

(Effects of Other Embodiment 3) Spectroscopy and ultra-high-speed imaging can be performed by the automation. Other effects are the same as those of the first embodiment.

(Other Embodiment 4) Another embodiment 4 of the present invention will be described with reference to FIG.

(Configuration of Another Embodiment 4) The cancel gradient coil 25 and the auto shim coil 4 of the other embodiment 1 shown in FIG. 2 are integrated and configured as the same coil. At this time, the X, Y, and Z correction coils of the shim coils are shared with the cancel gradient coils G _X , G _Y , and G _Z.

(Operation of Another Embodiment 4) The operation is the same as that of the other embodiment 3.

(Effect of Other Embodiment 4) Since the cancel gradient coil and the shim coil are integrated, the structure is simplified. Other effects are the same as the effects of the other third embodiment.

(Other Embodiment 5) Another embodiment 5 of the present invention will be described with reference to FIG.

(Configuration of Another Embodiment 5) The main gradient coil 24 of the other embodiment 1 shown in FIG.
And the auto shim coil 4 are integrated into a single coil. At this time, the X, Y, and Z correction coils of the shim coil are shared with the main gradient coils G _X , G _Y , and G _Z.

(Operation of Another Embodiment 5) The operation is the same as that of the other embodiment 3.

(Effect of Other Embodiment 5) Since the main gradient coil and the shim coil are integrated, the structure is simplified. Other effects are the same as the effects of the other third embodiment.

〔The invention's effect〕

As is clear from the above description, according to the present invention, it is possible to provide a superconducting magnet for a magnetic resonance imaging apparatus which does not generate an eddy current in the heat radiation shield plate and is not large.

[Brief description of the drawings]

FIG. 1 is a block diagram showing one embodiment of a superconducting magnet of a magnetic resonance imaging apparatus according to the present invention, and FIGS.
FIGS. 4, 4, 5 and 6 are diagrams showing the construction of another embodiment, FIGS. 7 and 8 are diagrams showing the principle of magnetic resonance imaging, and FIG. 9 is a diagram showing conventional magnetic resonance imaging. FIG. 10 is a configuration diagram showing a system of the device, and FIG. 10 is a configuration diagram showing a superconducting magnet of a conventional magnetic resonance imaging device. DESCRIPTION OF SYMBOLS 1 ... Patient, 2 ... Bed 3 ... RF coil, 4 ... Shim coil 5 ... Gradient coil 6 ... Static magnetic field magnet, 7 ... Room temperature bore 8 ... Excitation power supply, 9 ... Power supply lead 10 ... Static magnetic field direction, 11, 12, 13 ... Excitation power supply 14 ... Central control unit, 15 ... RF oscillation device 16 ... RF reception device, 17 ... Display / operation panel 18 ... Uniform magnetic field space, 19 ... Superconducting coil 20 ... Helium container, 21 ... 20K Radiation shield plate 22 ... 80K Radiation shield plate 23 ... Vacuum vessel, 24 ... Main gradient coil 25 ... Cancel gradient coil 26 ... Active gradient coil 27 ... Steel shim, 28 ... Superconducting magnet bore inner cylinder 100A, 100B ... Gradient coil 200A, 200B ... Transmission coils 300A, 300B ... receiving coil H ₀ ... static field, P ... subject S ... slice portion

Claims (7)

(57) [Claims] An object is arranged in a static magnetic field generated by a superconducting coil, a gradient magnetic field generated by a gradient coil is superimposed on the static magnetic field, and an excitation rotating magnetic field is applied by an RF coil to perform magnetic resonance. In a superconducting magnet of a magnetic resonance imaging apparatus that causes a phenomenon and performs imaging or spectroscopy of a specific nucleus in the tomographic plane of the subject by image processing, the gradient coil is mainly a gradient coil and the opposite polarity to the gradient coil The gradient coil is arranged concentrically with the superconducting coil between the radiant heat shield plate closest to the room temperature bore and the room temperature bore inside the superconducting magnet insulated container. Cancel coil with gradient coil Superconducting magnet of a magnetic resonance imaging apparatus being characterized in that disposed inside a room temperature bore in mind. 2. A superconducting magnet bore inner cylinder disposed on an inner peripheral side of a radiant heat shield plate closest to a normal temperature bore inside a superconducting magnet cool container also serves as a winding frame of a cancel gradient coil. A superconducting magnet of the magnetic resonance imaging apparatus according to 1). 3. The magnetic resonance imaging apparatus according to claim 1, wherein the inner tube of the radiant heat shield plate closest to the room temperature bore inside the superconducting magnet insulated container is also used as the winding frame of the cancel gradient coil. Superconducting magnet. 4. A superconducting magnet for a magnetic resonance imaging apparatus according to claim 2, wherein said current shim coil is wound around the inner periphery or outer periphery of a radiant heat shield plate closest to a normal temperature bore inside the superconducting magnet cooling container. . 5. An integrated coil is formed by sharing the X, Y, Z correction coil of the current shim coil with the G _X , G _Y , G _Z component coil of the cancel gradient coil, and the integrated coil is formed at the position of the cancel gradient coil. 3. A superconducting magnet for a magnetic resonance imaging apparatus according to claim 2, wherein said integrated coil is arranged. 6. A current shim coils _X, Y, G X of cancellation gradient coils and Z correction coil, G _Y, to form an integrated coil by sharing the G _Z component coil this position of the main gradient coil The superconducting magnet for a magnetic resonance imaging apparatus according to claim 2, wherein an integrated coil is disposed. 7. A superconducting magnet for a magnetic resonance imaging apparatus according to claim 1, wherein an iron shim is mounted on an inner surface of the normal temperature bore.