JPH0268038A - Superconducting magnet of magnetic resonance imaging device - Google Patents

Abstract

PURPOSE:To obviate the generation of eddy currents in a radiant heat shielding plate and to miniaturize a superconducting magnet by installing a cancel gradient coil between the outermost peripheral radiant heat shielding plate enclosing the superconducting coil and ordinary-temp. pore and installing a main gradient coil in the ordinary-temp. core. CONSTITUTION:The cancel gradient coil 25 is disposed between the ordinary- temp. pore 7 and the 80K radint heat shielding plate 22 in the superconducting magnet concentrically with the superconducting coil 19. The main gradient coil 24 is disposed in the ordinary-temp. pore 76 concentrically with the cancel gradient coil 25. An iron shim 27 is stuck to the inside surface of the ordinary- temp. pore 7. The external magnetic field generated by the main gradient coil 24 are canceled by the gradient coil 25 and, therefore, pulse magnetic fields do not act in the 80K radiant heat shielding plate 22 and the eddy currents are not generated therein. Since the iron shim 27 is disposed between the main gradient coil and the cancel gradient coil, the electrical coupling of the gradient coil and the iron shim 27 does not arise.

Description

[Detailed Description of the Invention] [Object of the Invention] (Industrial Application Field) The present invention relates to magnetic resonance (MR)
CT images (
Imaging (Com-pnted Tomgram)
The present invention relates to a superconducting magnet for a magnetic resonance imaging device (Imaging).

(Prior Art) For example, in a medical magnetic resonance imaging apparatus used for living body diagnosis, in order to obtain a tomographic image of a specific part of a living subject, the subject P is
A very uniform static magnetic field H8 along the direction is generated and acted upon by a static magnetic field magnet (not shown), and a linear magnetic field gradient G is applied to the static magnetic field H8 by a pair of gradient coils 100A and 100B. Here, the specific atomic nucleus with respect to the static magnetic field H0 has an angular frequency ω expressed by the following equation
. It resonates with me.

ω. =γ! 1. ... ■In this ■formula, γ is the gyromagnetic ratio, which is specific to the type of atomic nucleus. Therefore, the angular frequency ω makes only specific atomic nuclei resonate. For example, a pair of transmitting coils 200A provided in the RF coil (probe head) transmit the rotating magnetic field H1 of
It acts on the subject P via 200B.

In this way, the linear magnetic field gradient GX selectively acts only on the illustrated xy plane portion selectively set in the Z-axis direction, and a specific slice portion S (a portion on a plane) from which a tomographic image is obtained is In reality, the magnetic resonance phenomenon occurs only in layers (with a certain thickness). This magnetic resonance phenomenon is caused by the above R
For example, a pair of receiving coils 300 provided within the F coil.
A, free induction damping signal (free
1induction decay: hereinafter abbreviated as "rFID signal". ) and is used as an MR multiplied signal. By Fourier transforming this FID signal, a single spectrum can be obtained for the rotational frequency of a specific nuclear spin.

To obtain a tomographic image as a CT image, x-
Projections in multiple directions within the y-plane are required. Therefore, after exciting the slice portion S to produce a magnetic resonance phenomenon, the magnetic field H0 has a linear inclination in the X' axis direction (coordinate system rotated by an angle θ from the X axis) as shown in Figure 8. When the linear magnetic field gradient GXY is applied by a gradient coil (not shown), the equal magnetic field lines E in the sliced portion S of the subject P become straight lines. The rotational frequency of the specific nuclear spin on this equal magnetic field 1jlE is given by the above equation (g).

Here, for convenience of explanation, it is assumed that the equimagnetic field lines E are E1 to En, and the magnetic fields on these equimagnetic field lines E1 to En generate signals D to Dn, which are a type of FID signal, respectively. The amplitudes of the signals D1 to Dn are proportional to the spin densities of specific atomic nuclei on the equal magnetic field 11A Et to En passing through the slice portion S, respectively. However, the FID actually observed
The signal becomes a composite FID signal that is the sum of all signals D to Dn. Therefore, projection information (one-dimensional image) PD of the slice portion S on the X' axis is obtained by Fourier transforming the composite FID signal.

Next, this X' axis is rotated within the x-y plane, for example, by creating a magnetic field gradient GX'l' as a composite magnetic field of magnetic field gradients Gx and Gy in the X+'j direction by two pairs of gradient coils. This is done by changing the composite ratio of the magnetic field gradients Gx and GY. This magnetic field gradient cx
By rotating y, projection information in the angular direction within the xy plane is obtained in the same manner as above, and a CT image is synthesized based on this information.

The above is the principle of magnetic resonance imaging. Next, as a specific example, a conventional magnetic resonance imaging apparatus is shown in FIG. A subject, that is, a patient 1 is placed on a bed 2 . An RF coil (probe bed: high frequency transmitting/receiving coil) surrounds the patient 1.3. Furthermore, a shim coil 4 for magnetic field correction and a gradient coil 5 for generating a gradient magnetic field are arranged around the outer circumference. All of these coil systems are housed inside a humid bore 7 (usually 1 m in inner diameter of the bore) of a large static magnetic field magnet 6. As the static field magnet, any one of a superconducting magnet, a normal conducting magnet, and a permanent magnet is used.

This static field magnet 6 is excited and demagnetized by an excitation power source 8 via a current lead 9 (in the case of a permanent magnet system, this is not necessary); however, in the case of a superconducting magnet, it is operated in a persistent current mode. In order to reduce the consumption of liquid helium, which is a refrigerant, the current lead 9 is usually removed after excitation.
In most magnets, the direction of this static magnetic field, which is in a state in which a magnetic field is always generated, is the 10 direction shown in the figure, that is, the direction of the body axis of the patient 1. The gradient coil 5 includes a GX coil that provides 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, and is connected to excitation power supplies 11, 12, and 13, respectively. These excitation power supplies 11, 12, 13 are connected to the central control unit 14.
It is connected to the. The RF coil 3 is composed of a transmitting coil and a receiving coil, each of which has an RF oscillator 15 and an RF
are connected to the receiving device 16, which are further connected to the central control device 1.
Connected to 4. The central control device 14 is connected to a 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, the magnetic field uniform space 18
It is usually wide as 4 (1 to 50 cm sphere) and requires high uniformity of 50 ppm or less. Therefore, for example, in the case of a superconducting type, the static field magnet 6 has a length of 2.4 m, N2 m, and a height of 2.0 m. It would require something huge, measuring 4 meters and weighing 5 to 6 tons.

Even with such a large magnet, the uniformity within the 40-501 sphere due to the magnet alone is only a few hundred pp.
It can only be m. In order to reduce this to 50 Pρm or less, a shim coil 4 for magnetic field correction is used.

The patient's diagnostic site is brought into this magnetic field uniform space 18. Then, the RF oscillator fl15 is installed in the direction perpendicular to the static magnetic field lO.
, a high frequency is applied by the RF coil 3 to excite a desired atomic nucleus, such as a hydrogen atomic nucleus, within the human body cell. At the same time, gradient magnetic fields are applied in the X, Y, and Z directions by the GX excitation power supply 11, the GY excitation power supply 12, the GZ excitation power supply 13, and the gradient coil 5.

The optimal method for this RF and gradient pulse sequence is selected depending on the lesion site and image processing direction.

This pulse sequence operation is controlled by central controller 14. After applying the gradient and RF, a magnetic resonance signal is emitted from inside the patient's body. This signal is received and amplified by the RF receiver 16 and input to the central controller 14. Here, the image is processed and the required human body tomographic image is displayed on the CRT of the display/operation panel 17.

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

The superconducting coil 19 is housed in a helium container 20 filled with liquid helium and is maintained at an extremely low temperature of 4 degrees.

The outer periphery of the helium container 20 is surrounded by double radiant heat shield plates 20, 21, 80, and 22, so that heat from the outside is shielded by the shield plates.

In order to further enhance the heat insulation effect, a vacuum is maintained between the helium container and the radiant heat shield plate, and a heat insulator made of alumina polyester film is inserted. These members are housed in a vacuum container 23.

Now, when the gradient coil is excited for image processing, the pulsed magnetic field generates eddy currents in the radiant heat shield plates 21 and 22, which are made of a metal with good electrical conductivity such as aluminum.

Since a part of the energy applied to the gradient coil is used to generate this eddy current, the gradient pulse magnetic field, which normally would have to rise rapidly in 1 to 2 cycles, becomes dull. Furthermore, the uniform magnetic field at the center of the room-temperature bore is disturbed by the local magnetic field generated by the eddy current. Therefore, one image deteriorates. Furthermore, the gradient pulse magnetic field couples with the shim coil, which also degrades the image.

There are techniques that use active gradient coils to avoid this eddy current and coupling problem with shim coils. Regarding this technology, for example, IEEE T
RANSACTIONS ON MAGNETIC5゜November 1985. Vol MAG-21
, No.6.2273 2275A Finite Fo
urier Element Expansion T
echniquefor The Design of
a Pulsed Radial Gradient
System for Magnetic Re5on
ance imaging mg (MRI). FIG. 10 shows this structure.

A cancellation gradient coil 25 is disposed concentrically outside the main gradient coil 24, thereby forming an active gradient 1-coil 26. This active gradient coil 26 is arranged inside the shim coil 4 and substantially concentrically with the shim coil 4.

The pulse magnetic fields generated by the main gradient coil 24 and the cancel gradient coil 25 have opposite polarities.
By matching parameters such as ampere, turns, and diameter of both coils, the gradient pulse intensity necessary for image processing can be obtained inside the main gradient coil 24.
It has a structure in which the pulse magnetic field is canceled and becomes zero outside the canceling gradient coil 25. Since the pulse magnetic field is zero outside the active gradient coil 26, the generation of eddy currents and coupling with the shim coil 4 due to this are eliminated, improving image quality.

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

(1) In order to reduce power consumption, it is necessary to sufficiently space the main gradient coil 24 and the canceling gradient coil 25 in the radial direction.

For example, if the main gradient coil inner diameter is Dll and the cancellation coil inner diameter is D2, o 1/ o 2 < o
-a is desirable. Otherwise, in order to cancel the external magnetic field of the main gradient coil, the magnetic field energy of the canceling gradient coil must be made equal to the main gradient coil (in other words, the coil capacity and excitation power supply capacity will increase and a huge gradient A coil system is required.

■ When reusing the RF coil currently in use to secure patient insertion space, the inner diameter of the main gradient coil must be approximately the same as the inner diameter of the single gradient coil currently in use.

In this case, the outermost diameter of the active gradient coil is 1.2 to 1.5 times that of the conventional -1 gradient i-coil, making it impossible to house the active gradient coil in the normally used 1mφ bore superconducting magnet. It becomes possible. For storage purposes, the bore diameter of the superconducting magnet must be increased. This increases the size of the magnet, which significantly impairs its advantages in terms of cost and ease of installation.

This is hindering the spread of MPI.

Therefore, the present invention eliminates the drawbacks of the above-mentioned prior art,
An object of the present invention is to provide a superconducting magnet for a magnetic resonance imaging apparatus with an active gradient coil that is compact and does not generate eddy currents.

[Structure of the invention]

(Means for solving problem a) In order to achieve the above object, the present invention.

The outermost radiant heat shield plate surrounding the superconducting coil (e.g. 8
A canceling gradient coil is placed between the radiant heat shield plate (0) and the room temperature bore, and a main gradient coil is placed inside the room temperature bore.

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

Furthermore, since the canceling gradient coil is placed outside the room-temperature bore, the superconducting magnet does not become large.

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

(Configuration of Example) A canceling gradient coil 25 is arranged concentrically with the superconducting coil 19 between the normal temperature bore 7 and the radiant heat shield plate 22 at 80 inside the superconducting magnet.

Further, a main gradient coil 24 is arranged inside the normal temperature bore 7 concentrically with the cancellation gradient coil 25. An iron shim 27 is attached to the inner surface of the room 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 pulsed magnetic field acts on the radiant heat shield plate 22 at 80, and no eddy current is generated.

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

(Effects of Example) ■High quality images can be obtained because there is no influence of eddy currents.

■ The main gradient coil can be configured with the same size as a conventional gradient coil, so it can be stored inside the magnet bore, so the superconducting magnet has the same size as a conventional one. Therefore, increasing the size of the superconducting magnet due to the use of active gradient coils can be avoided.

This also avoids cost increases.

■ Also, when using a conventional sized superconducting magnet, the conventional active gradient coil is housed inside the magnet bore, but it is not possible to house it inside unless the inner diameter of the conventional active gradient coil is made smaller. It was difficult. When the inner diameter of the active gradient coil is reduced, the patient insertion space becomes smaller, further increasing the patient's fear of opening.

If the active gradient coil of this embodiment is used, this drawback can be avoided since the main gradient coil inner diameter is the same as that of the conventional active gradient coil.

(Other Example 1) Another embodiment 1 of the present invention will be described using FIG. 2.

(Another configuration of Example 1) The superconducting magnet bore inner cylinder 28 is also used as the winding frame of the cancellation gradient coil 25. in this case.

The canceling gradient coil 25 may be wound around the outer periphery of the FRP bore inner cylinder, or may be wound around the inner periphery to form a coil. The other configurations are the same as the embodiment shown in FIG.

(Other effects of Example 1) The operation is the same as that of the embodiment shown in FIG.

(Other effects of Example 1) The winding frame and support structure of the cancellation gradient coil are unnecessary, and the structure is simplified.

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

(Other Example 2) Another embodiment 2 of the present invention will be explained using FIG. 3.

(Configuration of Other Embodiment 2) The outermost radiant heat shield plate cylindrical portion 22 of the superconducting magnet is also used as the winding frame of the canceling gradient coil 25.

In this case, the cancellation gradient coil is wound around the inner surface of the cylindrical portion of the shield plate.

The other configurations are the same as the embodiment shown in FIG.

(Other effects of Example 2) The cancellation gradient coil is cooled to the same cryogenic temperature as the radiant heat shield plate.

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

(Other effects of Embodiment 2) The winding frame and support structure of the cancellation gradient coil are unnecessary, and the structure is simplified. Furthermore, since the cancellation gradient coil conductor is brought to an extremely low temperature, its electrical resistance decreases. This reduces the Joule loss during excitation of the gradient coil, so that the excitation power supply capacity for the canceled gradient coil can be reduced. Therefore, a compact, low-cost gradient coil can be provided.

(Other Example 3) Another embodiment 3 of the present invention will be explained using FIG. 4.

When performing spectroscopy diagnosis and ultra-high-speed imaging, auto-shimming using current shim coils is required to correct for variations in magnetic field uniformity due to the patient's internal magnetism.

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

(Configuration of other embodiment 3) In other embodiment 1 shown in FIG. 2, the current shim coil 4
(auto shim coil) is wound inwardly or outwardly around the inner cylindrical portion of the outermost radiation shield plate 22.

(Other effects of Example 3) Automatic shimming is possible with the current shim coil.

Other operations are the same as in the other embodiment 1.

(Other effects of Example 3) Spectroscopy and ultra-high-speed imaging become possible by autoshimming. Other effects are the same as in the other embodiment 1.

(Other Example 4) Another embodiment 4 of the present invention will be explained using FIG. 5.

(Configuration of Other Embodiment 4) The cancellation gradient coil 25 and O1-shim coil 4 of the other embodiment 1 shown in FIG. 2 are integrated into one coil. At this time, the shim coil's X, Y, and Z correction coils cancel the gradient coil's Gx, GY
+ Shared with GZ.

(Other effects of Example 4) The operation is the same as that of the other embodiment 3.

(Other effects of Example 4) Since the cancellation gradient coil and the shim coil are integrated, the structure is simplified. Other effects are the same as those of the third embodiment.

(Other Example 5) Another embodiment 5 of the present invention will be described using FIG. 6.

(Configuration of other embodiment 5) Main gradient coil 24 of other embodiment 1 shown in FIG.
and the auto shim coil 4 are integrated and constituted as the same coil. At this time, X of the shim coil.

The Y and Z correction coils are Gx and GY of the main gradient coil.
+ Shared with GZ.

(Other effects of Example 5) The operation is the same as that of the other embodiment 3.

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

〔Effect of the invention〕

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 that does not generate eddy currents in the thermal radiation shield plate and is not large in size.

[Brief explanation of the drawing]

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. 2, 3, 4, 5, and 6 show other embodiments, respectively. 7 and 8 are diagrams showing the principle of magnetic resonance imaging, FIG. 9 is a diagram showing the system configuration of a conventional magnetic resonance imaging device, and FIG. 10 is a diagram showing the system of a conventional magnetic resonance imaging device. FIG. 2 is a configuration diagram showing a magnet. 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...・Current lead 10... Static magnetic field direction 11, 12, 13... Excitation power source 14... Central control unit 15... RF oscillator 16... RF receiver W 17... Display/operation panel 18... Uniform magnetic field space 19... Superconducting coil 20... Helium container 2
1...20 radiation shield plate 22...80 radiation shield plate 23...vacuum vessel 24...main gradient coil 25...cancel gradient coil 26...
・Active gradient coil 27...iron shim
28... Bore inner cylinder 100A, 100B... Gradient coil 200A, 200B... Transmission coil 300A, 300B... Receiving coil H0... Static magnetic field P... Subject S... Slice portion Agent Patent Attorney Nori Ken Yudo Daishimaru Ken

Claims (6)

[Claims] (1) A magnetic resonance phenomenon is produced by placing a subject in a static magnetic field generated by a superconducting coil, superimposing a gradient magnetic field generated by a gradient coil on the static magnetic field, and applying an excitation rotating magnetic field by an RF coil. In a superconducting magnet of a magnetic resonance imaging apparatus that performs imaging or spectroscopy of a specific atomic nucleus in the tomographic plane of the above-mentioned object by image processing, a gradient coil is used as a main gradient coil and a magnetic field of opposite polarity is applied to the main gradient coil. The canceling gradient coil is placed concentrically with the superconducting coil between the radiant heat shield plate closest to the room temperature bore inside the superconducting magnet cold storage container and the room temperature bore, and canceling the main gradient coil. A superconducting magnet for magnetic resonance imaging equipment, which is characterized by being placed inside a room-temperature bore concentrically with a gradient coil. (2) Claim (1) characterized in that the canceling gradient coil is wound around the inner circumference or outer circumference of the bore inner cylinder.
A superconducting magnet of the magnetic resonance imaging device described. (3) The superconducting magnet for a magnetic resonance imaging apparatus according to claim (1), wherein the canceling gradient coil is wound around the inner periphery of a radiant heat shield plate closest to the normal temperature bore inside the superconducting magnet cold storage container. (4) The superconducting magnet for a magnetic resonance imaging apparatus according to claim 2, wherein the current shim coil is wound around the inner or outer circumference of a radiant heat shield plate closest to the normal temperature bore inside the superconducting magnet cold storage container. (5) An integrated coil is formed by sharing the X, Y, and Z correction coils of the current shim coil with the Gx, GY, and Gz component coils of the canceling gradient coil, and this integrated coil is placed at the position of the canceling gradient coil. 3. A superconducting magnet for a magnetic resonance imaging apparatus according to claim 2. (6) By sharing the X, Y, and Z correction coils of the current shim coil with the Gx, GY, and Gz component coils of the main gradient coil, an integrated coil is formed, and this integrated coil is placed at the position of the main gradient coil. 3. A superconducting magnet for a magnetic resonance imaging apparatus according to claim 2, further comprising a superconducting magnet. 7. A superconducting magnet for a magnetic resonance imaging apparatus according to claim 1, characterized in that an iron shim is attached to the inner surface of the normal temperature bore.