JPS63272335A - Magnetic resonance imaging apparatus - Google Patents

Abstract

PURPOSE:To obtain a small-scale lightweight and low cost poplar magnetic resonance imaging apparatus easy to anchor imparting no psychological oppressive sensation to a patient and capable of performing dynamic diagnosis, by constituting said apparatus of a superconductive magnet equipped with a magnetic shield of short axial length of the patient body, a conical gradient coil and a gantry rotating mechanism. CONSTITUTION:A magnetic shield 12 is longer than a superconductive coil and, further, the inner diameter of the end plate of the magnetic shield 12 is made larger than the outer diameter of the superconductive coil. A notch part X is formed to a vacuum container part not covered with the magnetic shield 12 in order to make the superconductive magnet 3 short apparently and the eliminate the oppressive sensation to a patient. By this constitution, the superconductive magnet 3 becomes short apparently. A cylindrical uniform space 6 is obtained by the superconductive coil 8 thus constituted, iron piece shims 7a, 7b and the magnetic shield 12. By using a gradient coil, the apparatus can be made short apparently. The same RF coil as a conventional one can be used.

Description

[Detailed Description of the Invention] [Object of the Invention] (Industrial Application Field) The present invention relates to magnetic resonance (MR)
Computed tomography (CT) is based on the density distribution of specific π1 nuclear spins in a specific cross section of a living subject using the eso-nance (eso-nance) phenomenon.
CT image (co
Imaging (t
The present invention relates to a magnetic resonance imaging apparatus (magina), and particularly relates to a compact magnetic resonance imaging apparatus that can become a popular type.

(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, a Z
A very uniform static magnetic field HO along the direction is generated and acted upon by a static magnetic field magnet (not shown), and a linear magnetic field gradient Gx is applied to the static magnetic field Ho by a pair of gradient coils 100A and 100B. Here, the specific atomic nucleus with respect to the static magnetic field HO has an angular frequency ω expressed by the following equation
Resonates at 0.

ω0=γH○ (1) In this equation (1), γ is the gyromagnetic ratio, which is unique to the type of atomic nucleus. Therefore, a rotating magnetic field H1 having an angular frequency ω0 that causes only specific atomic nuclei to resonate is applied to the subject P via, for example, a pair of transmitting coils 200A and 200B provided in an RF track (probe head).

In this way, the linear magnetic field gradient Gx acts selectively only on the illustrated The magnetic resonance phenomenon occurs only when the material has a certain thickness (in reality). This magnetic resonance phenomenon is caused by the above R
For example, a pair of receiving coils 30O provided in the F coil.
A, free induction damped signal (free
:nauctrondecay:rFID signal"
It is abbreviated as. ) and is used as an MR multiplied signal. By Fourier transforming this FID signal,
A single spectrum is obtained for the rotational frequency of a particular 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 cause a magnetic resonance phenomenon, the magnetic field HO has a linear inclination in the X'-axis direction (a coordinate system rotated by an angle θ from the x-axis) as shown in Figure 5. When a linear magnetic field gradient GXY is applied by a gradient coil (not shown), the isomagnetic field line E in the sliced portion S of the subject P becomes a straight line, and the rotation frequency of a specific nuclear spin on this isomagnetic field line E is expressed by the above equation (1). Hail to you.

For convenience of explanation, it is assumed that the equal magnetic fields ME are E1 to En, and the magnetic fields on these equal magnetic field lines E1 to En generate signals D1 to [)n, 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 isomagnetic field lines E1 to En passing through the slice portion S, respectively.

However, the FID signals actually observed are signals D1~
A composite FID signal is obtained by adding all Dn. Therefore,
Projection information (one-dimensional image) PD of the slice portion S onto the X' axis is obtained by Fourier transforming the composite FID signal. Next, this x' axis is rotated within the x-y plane, which is done by, for example, two pairs of gradient coils,
This is done by creating a magnetic field gradient Gxy as a composite magnetic field of the magnetic field gradients Gx and GY in the directions, and changing the composite ratio of the magnetic field gradients Gx and Gy. By rotating this magnetic field gradient Gxy, projection information in the angular direction within the xy plane is obtained in the same manner as described 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, ie, a patient 1, is placed on a bed 2. Surrounding the patient 1, an RF tocol (probe bed: high frequency transmitting/receiving coil) 5 is arranged, and furthermore, a shim coil 19 for magnetic field correction and a gradient coil 4 for generating a gradient magnetic field are arranged around the periphery thereof. All of these coil systems are housed inside the normal temperature bore 21 (usually bore inner diameter 1771) of the large whole body magnet 20. As the whole body magnet, one of a superconducting magnet, a normal conducting magnet, and a permanent magnet is used.

This whole body magnet 20 is excited and demagnetized by an excitation power source 22 via a current lead 23 (this is not necessary in the case of a permanent magnet system). In the case of superconducting magnets, the current lead 23 is usually removed after excitation because it is operated in persistent current mode and to reduce consumption of liquid helium, which is a coolant, so that a magnetic field is always generated. ing. In most magnets, the direction of this static magnetic field is usually the 24 directions shown in the figure, that is, the direction of the body axis of the patient 1.

The gradient coil 4 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.
Consisting of coils, each with an excitation power source 25, 26.2
7 is connected. These excitation power supplies 25, 26, 27
is connected to the central control bag M2B. The RF protocol 5 is composed of a transmitting coil and a receiving coil, each of which has an RF
It is connected to an oscillating device 29 and an RF receiving device 30, which are further connected to a central control device 28. The central control device 28 is connected to a display/operation panel 31, 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 32
The ball is usually 40 to 50 cm wide, and 50 pp!
High uniformity of 1 or less is required. For this reason, the whole body magnet 20 has a length of 2.4 m in the case of a superconducting method, for example.
, a huge one is required, measuring 2 m wide, 2.4 m high, and weighing 5 to 6 tons.

Even with such a large magnet, the uniformity within a 40-50 cm sphere due to the magnet alone is only a few hundred points at most.
It only becomes pm. In order to reduce this to 50 ppm or less, one shim coil for magnetic field correction is used. The patient's diagnostic site is brought into this magnetic field uniform space 32. Then, high frequencies are applied by two RF oscillators and an RF tocol 5 in a direction perpendicular to the static magnetic field 24 to excite a desired atomic nucleus, such as a hydrogen atomic nucleus, within the human body cell. Also, at the same time, GX excitation 1a'l source 25, GY excitation! 11126, a gradient magnetic field is applied in the X, Y, and Z directions by the GZ excitation power source 27 and the gradient coil 4.

The optimum RF and gradient pulse sequence is selected depending on the lesion site and image processing method.

This pulse sequence operation is controlled by central controller 28. 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 30 and input to the central controller 28. Here, the image is processed and the required human body tomographic image is displayed on the CRT of the display/operation panel 31.

By the way, the conventional magnetic resonance imaging apparatus configured as described above has the following problems.

That is, in order to realize a space with a uniform magnetic field (A>40 to 50 cm), a huge whole-body magnet is required.

For this reason, (1) the manufacturing cost of the magnet becomes high, and the system price including the diagnostic device becomes high, exceeding the purchasing power of the user;

■ Because the magnet is large and heavy, it cannot be installed in an existing diagnostic room, and the building will need to be remodeled or a new one built.

For the above reasons, the widespread use of magnetic resonance imaging apparatuses has been hindered.

(B) Patients often exhibit claustrophobia because they are completely enclosed within the very narrow magnetic bore. Furthermore, from the point of view of the diagnosing doctor, the patient cannot be observed during the diagnosis, so there is a drawback that emergency response cannot be taken in case of a sudden change in the patient's condition.

(Problems to be solved by the invention) In this way, the conventional technology requires a large magnet, which hinders its widespread use due to the increase in size and price, and also puts psychological pressure on patients. There was a problem with giving a feeling.

Therefore, the object of the present invention is to shorten the length of the static magnetic field magnet in the direction of the patient's body axis, thereby eliminating the psychological pressure on the patient and making it possible to perform dynamic diagnosis in a compact and lightweight manner.

It is an object of the present invention to provide a magnetic resonance imaging apparatus that can be of a low cost, easy to install, and popular type.

[Structure of the Invention] (Means for Solving the Problems) The magnetic resonance imaging apparatus according to the present invention is constructed as follows in order to solve the above problems and achieve the objects.

In other words, a coil system consisting of a superconducting coil that generates a static magnetic field of the required strength, and a large number of iron shims arranged on the outer surface of the liquid helium tank and the inner surface of the vacuum vessel that house this coil to correct the non-uniform magnetic field. , by generating a cylindrical uniform magnetic field space with an axis in the body axis direction, the ratio D/T of its diameter and cylinder length (thickness) T being 1.75 or more, superconducting The length of the magnet in the body axis direction is shortened, and furthermore, in the combination configuration with the magnetic shield, the room temperature bore side of the end plate of the vacuum vessel is configured with a notch, and the conical gradient coil is brought into close contact with the notch of the magnetic shield. As a result, the length in the body axis direction is further shortened.

(Function) By configuring the superconducting magnet in this way, it is possible to shorten the body axis direction and generate a magnetic resonance phenomenon.Therefore, miniaturization is realized due to a small uniform magnetic field space, and a popular configuration is achieved. Moreover, since the body axis direction is short, the psychological pressure on the patient can be reduced.

(Example) An example of the magnetic resonance imaging apparatus of the present invention will be described below with reference to FIGS. 1 to 3.

FIG. 1 is a sectional view along the body axis direction showing the configuration of this embodiment. In this embodiment, a superconducting magnet 3 that generates a static magnetic field, a gradient coil 4 that generates a gradient magnetic field, an RF
A patient 1 placed on a bed 2 is inserted into a magnet system composed of an RF coil 5 that generates a magnetic field and detects magnetic resonance signals.

A thin slice-shaped cylindrical magnetic field uniform space 6 having a magnetic field uniformity of 50 ppm or less is applied to the patient 1 .

This cylindrical magnetic field uniform space has a shape in which the ratio D/T of the diameter perpendicular to the body axis of the patient 1 and the length T in the body axis direction is 1.75 or more. For example, D=350M, T=1
It is 0trm. Such a cylindrical magnetic field is obtained by the following superconducting magnet configuration and iron piece shims 7a, 7b, and 7c.

The superconducting magnet 3 is generated by a superconducting coil 8, a liquid helium tank 10 that stores liquid helium 9 that keeps the superconducting coil 8 in a superconducting state, a vacuum container 11 that stores these and keeps them in a vacuum state, and the superconducting coil 8. It is composed of a magnetic shield 12 for shielding leakage magnetic fields.

The superconducting coil 8 is used only to generate the magnetic field strength necessary for magnetic resonance phenomena. The magnetic field strength is, for example, 0.5 Tesla or 1.5 Tesla. Uniformity is not required for the magnetic field generated by this superconducting coil 8. Therefore, a winding configuration with relatively coarse precision is sufficient. Examples include a simple solenoid coil, a notched solenoid coil, and a Helmholtz coil. There is no limitation to this winding configuration.

Further, the uniformity in the cylindrical magnetic field uniformity space 6 is adjusted by iron piece shims 7 a, 7 b, and 7 c.

The iron piece shim 7a is attached to the outer surface of the liquid helium tank 10. The iron piece shim 7b is attached to the inner surface of the vacuum container 11. The iron piece shim 7C is the vacuum container 11
It is attached to the internal normal temperature bore. Iron piece shim 7
a, 7b are placed during the manufacture of the superconducting magnet 3. The positions at which these iron piece shims 7a and 7b should be placed are determined by solving the equation for an electromagnetic circuit including a coil and a magnetic body constituted by the superconducting coil 8, iron piece shims 7a and 7b, and magnetic shield 12, and determining the position of this circuit. The positions of the iron piece shims 7 a and 7 b may be determined so that the magnetic field generated by the magnetic field has the shape of the cylindrical uniform magnetic field space 6 .

The iron piece shim 7C is used when finely adjusting the magnetic field uniformity after the superconducting magnet 3 is assembled.

The end plate of the vacuum container 11 on the side of the normal temperature bore is cut out in an inverted trapezoidal shape (notch portion X) as shown in FIG. The magnetic shield 12 is attached to the outer surface of the vacuum container 11 excluding the room temperature pores and the notch X, and the length of the magnetic shield is made to be sufficiently longer than the length of the superconducting coil 8 in order to eliminate the influence of magnetic field uniformity deterioration. It's long.

The two gradient coils 4 have their apexes facing each other and are shaped to closely fit into the notch X of the vacuum container 11, and have a substantially conical structure as shown in FIG. 2. .

As shown in FIG. 2, the gradient coil 4 includes a circular coil GZ coil 13 that provides a gradient magnetic field in the Z-axis (body axis) direction, and a saddle-shaped coil pair that provides a gradient magnetic field in the X-axis (orthogonal to the body axis) direction. It consists of a GX coil 14 and a saddle-shaped coil pair GY coil 15 that provides a gradient magnetic field in the Y-axis (body axis, orthogonal to the X-axis) direction, and these are wound around a substantially conical winding frame. The left and right gradient coils 4 have the same shape, and the left and right gradient coils 4 are configured as a pair. Here, the conical shape of the gradient coil 4 is shaped to fit tightly into the notch of the vacuum container 11.

The superconducting magnet 3, the gradient coil 4, and the RF coil 5 are covered with a decorative cover 16 and become a single unit.
This will be Gantry 17. The gantry rotation mechanism 18 shown in FIG. 3 allows the gantry 17 to be tilted at any angle from the vertical position.

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

As with conventional superconducting magnets, when trying to obtain a certain degree of uniformity, for example several hundred ppm, using only a superconducting coil in a relatively large space such as a 50aR sphere, the coil becomes long, and the length of the coil is usually close to 2 m. turn into.

Therefore, if the superconducting coil 8 of this embodiment is used only for the purpose of generating the required magnetic field strength, and the uniform space 6 is made as thin as about 10 to 20 coils in the coil length direction, that is, in the body axis direction,
The superconducting coil 8 can be made extremely short. Further, since there is no restriction on uniformity, there is no need for a conventional coil arrangement (complicated coil arrangement), and the coil can be configured with a simple solenoid or Helmholtz arrangement.

The cylindrical magnetic field uniform space 6 is at least 501) I) Ill
In order to achieve uniformity, iron piece shims 7a and 7b are used.

The iron piece shims 7a, 7b, depending on the way they are arranged, can absorb any second-order component of the non-force-magnetic field component, for example, XY.

You can erase X2 Y, XY2, etc. Normally, the non-force-magnetic field component has a large two-direction component, X.

The Y direction component is relatively small. Since the cylindrical magnetic field uniform space 6 that is now required is very thin in the Z direction, the non-force-magnetic field component in the Z direction has almost no effect.

Therefore, only the non-force-magnetic field components in the X and Y directions, which were originally small in size, remain, so the iron piece shims 7 a, 7
It is possible to adjust to below the required 501)l)m using only b. The iron piece shim 7a is connected to the liquid helium tank 10.
The iron piece shim 7b is attached to the inner surface of the vacuum vessel 11. By doing so, the components and orders of the non-force-magnetic field can be arbitrarily adjusted over a wide range.

The non-force component and order of the magnetic field generated by the superconducting coil 8 can be determined by analysis. The arrangement of the iron piece shims 7a and 7b that cancel out this non-force component/order is similarly determined by analysis. Iron piece shims 7a and 7b are attached to the positions obtained by this analysis during the manufacture of the superconducting magnet 3.

A magnetic shield 12 is attached to the outer surface of the vacuum container 11 to reduce leakage magnetic fields. At this time, magnetic shield 12
If the length of the magnetic shield 12 is about the same as that of the superconducting coil 8, or if the inner diameter of the end plate of the magnetic shield 12 is smaller than the outer diameter of the superconducting coil 8, the uniformity of the cylindrical magnetic field uniform space 6 will be significantly degraded.

Even if you do 7b, the adjustment cannot be completed.

Therefore, as shown in FIG. 1, the magnetic shield 12 is longer than the length of the superconducting coil, and the inner diameter of the end plate of the magnetic shield 12 is larger than the outer diameter of the superconducting coil. however,
With this configuration, even if the length of the superconducting coil is shortened, the magnetic shield and the vacuum vessel length will be lengthened accordingly, making it difficult to shorten the length of the superconducting magnet 3 (which is contrary to the original purpose of eliminating the feeling of pressure on the patient). I end up going against it.

Therefore, in order to shorten the apparent length of the superconducting magnet 3 and eliminate the feeling of pressure on the patient, a notch X as shown in FIG. 1 is provided in the portion of the vacuum container that is not covered with the magnetic shield 12. By doing so, the layered conductive magnet 3 appears to be shorter.

Superconducting coil 8 and iron piece shim 7a configured as above
, 7b. Although the cylindrical uniform space 6 is obtained by the magnetic shield 12, in reality, the required degree of uniformity cannot be obtained in the cylindrical uniform space depending on the manufacturing accuracy of the superconducting magnet 3 or the magnetic material conditions of the installation environment. There are cases where it is not possible. At that time, by arranging iron piece shims 7C on the inner surface of the vacuum vessel 11, the degree of uniformity can be finely adjusted to achieve the required degree of uniformity.

Next, when installing the gradient coil 4 to this device,
When using a gradient coil wound on a conventional cylindrical winding frame, the coil length becomes long (for example, 1 Tr).
L) Although the superconducting magnet is shortened by making a notch in the vacuum container as described above, there is a drawback that the gradient coil makes the overall length longer again. In order to eliminate this, a conical gradient coil 4 shown in FIG. 2 is attached to the notch X of the vacuum vessel 11.

Since the internal coil configuration of the conical gradient coil 4 is the same as that of the conventional type, its performance is equivalent to that of the conventional type. By using the gradient coil 4, the apparent length of the device can be shortened. The same RF coil as before can be used.

Next, an image processing method using this device will be described. This device basically performs diagnosis by obtaining cross-sectional images similar to those obtained with X-rays. The bed 2 is moved so that the diagnosis site is placed in the cylindrical space 6 with a uniform magnetic field. Similar to a conventional magnetic resonance imaging apparatus, an appropriate pulse sequence is created using an RF coil 5 and a gradient coil 4 to obtain a cross-sectional image within a cylindrical magnetic field uniform space 6. If you want to obtain a coronal or sagittal image of the diagnostic site, for example, if you want to diagnose the spinal cord, there are many requests.)
First, an image of the length T in the body axis direction of the cylindrical magnetic field uniform space 6 is taken. Next, the bed 2 is moved by the height and an image of length T is taken again. Repeat this operation as many times as necessary (for example, if you want to obtain an annular or sagittal cross-sectional image of length n, repeat n times),
The obtained images may be reconstructed within the central controller 28 into one coronal or sagittal image.

Furthermore, if it is desired to perform an oblique image, for example, diagnosis of the heart (this is often requested), the following procedure is performed. The diagnostic site is brought to the position of the cylindrical magnetic field uniform space 6. The gantry rotating machine #118 tilts the gantry 17, that is, the superconducting magnet 3, by an angle necessary to take an oblique image. By doing this, the cylindrical magnetic field uniform space 6 is also tilted by the same angle, and if an image is taken in this state, a required oblique image can be obtained.

Although the uniform magnetic field space 6 of this device is thinner in the body axis direction than that of conventional magnetic resonance imaging devices, by performing the above steps, various images equivalent to those obtained with conventional devices can be obtained. .

As explained above, this embodiment has the following effects.

(1) The uniform magnetic field space 6 is made thin in the body axis direction, and the uniformity of the magnetic field is achieved by iron piece shims 7 a and 7 b placed inside the superconducting magnet 3.
The superconducting coil 8 becomes shorter. Therefore, the superconducting magnet 3 can be shortened in the body axis direction. For example, 1m or less. As a result, most of the patient's body is exposed to the outside from the device, eliminating the feeling of pressure on the patient. Doctors can make diagnoses while constantly monitoring the patient's condition. Furthermore, since the arm is exposed, it is possible to perform so-called dynamic diagnosis, which can diagnose moment-by-moment changes in the lesion site by administering a contrast agent such as Gd-DTPA, further expanding the range of diagnostic methods. effective.

(21) It is significantly smaller and lighter than the conventional model.

For example, the length is less than 1m, which is less than half of the conventional model, and the weight is less than 3 tons even with the magnetic shield, so it can be easily installed in existing hospital buildings while suppressing leakage magnetic fields. There is an effect.

(3) As stated in (1) above, the superconducting coil 8 does not have to contribute to magnetic field uniformity. Therefore, the winding only needs to be wound with a coarser opening than conventional superconducting coils.
Since the superconducting wire itself needs to be short, both material costs and winding processing costs are reduced, resulting in a low-cost superconducting magnet.

(4) A space with a thin magnetic field uniform in the body axis direction is used, but by moving the bed and reconstructing the necessary number of images obtained in this uniform magnetic field space, the same coronal shape as that obtained with the conventional model can be obtained. Images of cross-sections and sagittal cross-sections can be obtained. Furthermore, by rotating the superconducting magnet using the rotation mechanism 18, an oblique image can also be obtained. Therefore, although it is a thin uniform magnetic field space, various images equivalent to those of the conventional device can be obtained.

(5) By making the end of the superconducting magnet have a notch structure and adopting the conical gradient coil 4 that fits into that part, the length of the gantry in the body axis direction appears to be even shorter, and as mentioned above ( The effect mentioned in 1) becomes even stronger.

(6) Since the iron piece shims 7a and 7b can be placed inside the superconducting magnet 3, the adjustment range and degree of freedom are expanded compared to conventional iron piece shims placed only on the inner surface of the room-temperature bore, and uniformity is achieved. It becomes easier to adjust once.

[Effects of the Invention] As described above, according to the present invention, since it is composed of a magnetically shielded superconducting magnet with a short length in the axial direction of the patient's body, a conical gradient coil, and a gantry rotation mechanism, It is possible to provide a compact, lightweight, low-cost, easy-to-install, and popular magnetic resonance imaging device that can be used for dynamic diagnosis without causing any psychological pressure.

[Brief explanation of the drawing]

Fig. 1 is a block diagram showing an embodiment of a magnetic resonance imaging apparatus according to the present invention, Fig. 2 is a block diagram showing a gradient coil of the same embodiment, and Fig. 3 is a block diagram showing a gantry rotation mechanism of the same embodiment. The block diagram, FIGS. 4 and 5 are diagrams showing the principle of magnetic resonance imaging, and FIG. 6 is a block diagram showing a conventional magnetic resonance imaging apparatus. 1... Patient, 2... Bed, 3... Superconducting magnet, 4... Gradient coil, 5... RF coil, 6... Cylindrical magnetic field uniform space, 7a, 7b, 7c.
... Iron piece shim, 8 ... Superconducting coil, 9 ... Liquid helium, 10 ... Liquid helium tank, 11 ... Vacuum container, 12 ... Magnetic shield, 13 ... GZ coil, 14 ... ...GX coil, 15...GY coil, 1
6... makeup cover, 17... gantry, 18...
・Gantry rotation mechanism, 19...Shim coil, 2o・
...Full body magnet, 21...Normal temperature bore, 22.
...Excitation power supply, 23...Current lead, 24...Patient body axis direction, 25...GX excitation single power supply, 26...GY
Excitation power supply, 27... GZ excitation power supply, 28... Central control device, 29... RF oscillation device, 30... RF receiving device, 31... Display operation panel, 32... Magnetic field uniform space . Applicant's representative Patent attorney Takehiko Suzue Figure 1 Figure 6 Commissioner of the Patent Office Kunio Ogawa 1, Indication of the case Japanese Patent Application No. 1983-274312 2, Name of the invention Magnetic Resonance Imaging Device 3, Person making the amendment Case Relationship with Patent Applicant (307) Toshiba Corporation 4, Agent uBE Building, 3-7-2 Kasumigaseki, Chiyoda-ku, Tokyo 100 Telephone 03 (502) 3181 (Main Representative) 6 Contents of Amendment (1) Permission The scope of claims is amended as shown in the attached sheet. (2) Correct "gradient coil" on page 3, line 13 of Exhibit A to "gradient coil." (3) “Grandient coil” in page 5, line 4 of Mei No. 1
is corrected as "gradient coil". (4) “Grandient coil” on page 6, line 1 of the specification
is corrected as "gradient coil". (5) "70-bubetsu" in the first line of page 12 of the specification is replaced with "
Flobhet,” he corrected. (6) Correct "gradient coil" in page 6, line 14 of Akira 101 to "gradient coil." (7) Akira [+Fu, page 7, line 9, "gradient coil" is corrected to "gradient coil." (8) M, 75J on page 11, line 11 of the specification is rl, 5
Correct it with J. (9) ri, 75J, N, 75J, 5J, page 13, line 1,
I am corrected. (10) Akari lIm page 15 line 9 “Grangen!
・Correct ``coil'' to ``gradient coil.'' (11) [Gradient coil] in Mei III Yu, page 15, line 18 is corrected to "gradient coil." (12) In the 11th Meiji edition, page 24, line 12, change ``to hang'' to ``
It will take a while,” he corrected. 2. Scope of Claims (1) A subject is placed in a static magnetic field generated by a static magnetic field magnet, a gradient r141 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. In a magnetic resonance imaging apparatus that generates a magnetic resonance phenomenon by applying a magnetic field, and performs imaging or spectroscopy of a specific atomic nucleus within a tomographic plane of the subject using image zero processing, the static magnetic field magnet is placed on the outer surface of a liquid helium tank. Using an IIA conductive magnet consisting of an iron shim and a superconducting coil placed on the inner surface of a vacuum vessel, the thickness T in the body axis direction of the subject and the diameter in a plane perpendicular to the body axis are determined. Ratio D
The superconducting coil is configured to form a cylindrical magnetic field uniform space where /T is 1.5 or more, and this superconducting coil contributes only to static magnetic field generation, and uniformity is achieved only by the iron piece shim, and the vacuum vessel The length of the cylindrical magnetic shield directly covering the vacuum vessel is longer than the length of the superconducting coil, and the inner diameter of the end plate of the magnetic shield is larger than the outer diameter of the superconducting coil. The end plate of the vacuum container that covers only the outer peripheral surface and the end plate, and is not covered with the magnetic shield, is cut out toward the normal temperature bore side of the vacuum container, and the gradient 1-coil is circular. A magnetic resonance imaging apparatus characterized in that it is wound around a winding frame and is tightly housed in a notch of the vacuum container. (2) The magnetic resonance imaging apparatus according to claim 1, wherein the static magnetic field magnet has a rotating structure for rotating around an axis perpendicular to the bed and the vertical direction. Applicant's representative Patent attorney Takehiko Suzue 1, Indication of the case Japanese Patent Application No. 1983-274312 2, Name of the invention Magnetic Resonance Imaging Device 3, Person making the amendment Relationship to the case Patent applicant (307) Toshiba Corporation 4, Agent UBE Building 5, 3-7-2 Kasumigaseki, Chiyoda-ku, Tokyo.
Date of amendment order May 24, 1986 6 Subject of amendment 7 Contents of amendment (1) In the ``Contents of amendment'' of the written procedural amendment submitted on February 17, 1988, (1) ``Claims'' is corrected to ``Claims''. (2) Same as above, in (2), "Line 13" is corrected to "Line 13 to Line 14." (3) The description in (5) above is corrected as follows. (No change in content) Note (5) “Probe bed” on page 6, line 12 of the specification
is corrected as "probe bed".

Claims (2)

[Claims] (1) Magnetic resonance phenomenon occurs by placing a subject in a static magnetic field generated by a static magnetic field magnet, 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 the magnetic resonance imaging apparatus, the static magnetic field magnet is arranged on the outer surface of the liquid helium tank and the inner surface of the vacuum container. Using a superconducting magnet consisting of iron piece shims and superconducting coils, the thickness T in the body axis direction of the above-mentioned subject,
The ratio D/T of the diameter D in the plane perpendicular to the body axis is 1.7
This superconducting coil is configured to form a cylindrical magnetic field uniform space of 5 or more, and this superconducting coil contributes only to static magnetic field generation, and uniformity is achieved only by the iron piece shim,
The length of the cylindrical magnetic shield directly covering the vacuum vessel is longer than the length of the superconducting coil, and the inner diameter of the end plate of the magnetic shield is larger than the outer diameter of the superconducting coil,
Furthermore, the magnetic shield covers only the outer peripheral surface and end plate of the vacuum container, and the end plate of the vacuum container that is not covered with the magnetic shield is cut out toward the normal temperature bore side of the vacuum container. . A magnetic resonance imaging apparatus, wherein the gradient coil is wound around a conical winding frame and is tightly housed in a notch of the vacuum container. (2) The magnetic resonance imaging apparatus according to claim 1, wherein the static magnetic field magnet has a rotation mechanism for rotating around an axis perpendicular to the bed and the vertical direction.