JP2009172085A - Superconductive magnet device, magnetic resonance imaging apparatus using the same, and nuclear magnetic resonance apparatus - Google Patents

Superconductive magnet device, magnetic resonance imaging apparatus using the same, and nuclear magnetic resonance apparatus Download PDF

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JP2009172085A
JP2009172085A JP2008012558A JP2008012558A JP2009172085A JP 2009172085 A JP2009172085 A JP 2009172085A JP 2008012558 A JP2008012558 A JP 2008012558A JP 2008012558 A JP2008012558 A JP 2008012558A JP 2009172085 A JP2009172085 A JP 2009172085A
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superconducting
coil
good conductor
magnetic resonance
main coil
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JP4928477B2 (en
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Hiroto Kusaka
浩都 草加
Mitsuji Abe
充志 阿部
Tsutomu Yamamoto
勉 山本
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Hitachi Ltd
Hitachi Healthcare Manufacturing Ltd
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Hitachi Medical Corp
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    • H—ELECTRICITY
    • H01—ELECTRIC ELEMENTS
    • H01F—MAGNETS; INDUCTANCES; TRANSFORMERS; SELECTION OF MATERIALS FOR THEIR MAGNETIC PROPERTIES
    • H01F6/00—Superconducting magnets; Superconducting coils
    • H01F6/06—Coils, e.g. winding, insulating, terminating or casing arrangements therefor
    • H—ELECTRICITY
    • H01—ELECTRIC ELEMENTS
    • H01F—MAGNETS; INDUCTANCES; TRANSFORMERS; SELECTION OF MATERIALS FOR THEIR MAGNETIC PROPERTIES
    • H01F6/00—Superconducting magnets; Superconducting coils
    • H01F6/02—Quenching; Protection arrangements during quenching

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Abstract

<P>PROBLEM TO BE SOLVED: To control the voltage generated in a leakage magnetic field and a coil at the time of rapid demagnetization within an allowable range and to miniaturize a superconductive magnet device. <P>SOLUTION: The superconductive magnet device comprises a plurality of pairs of coils which are annularly formed around a center axis Z passing through a magnetic field space and which are disposed to face each other in the direction of the center axis Z, and a plurality of pairs of good conductors. The plurality of pairs of coils comprise superconductive main coils 10A and 10B and a superconductive shield coil 20 disposed with a space on the outside in the radial direction of the superconductive main coils 10A and 10B as seen from the center axis Z. The plurality of pairs of good conductors include a first good conductor 11 and a second good conductor 21. The first good conductor 11 is disposed on the outer side in the radial direction of the superconductive main coil 10A, or between a pair of superconductive main coils 10A in such a state that at least part of the first good conductor overlaps on the superconductive main coil 10A as seen from the center axis Z. The second good conductor 21 is disposed on the inner side in the radial direction of the superconductive shield coil 20. <P>COPYRIGHT: (C)2009,JPO&INPIT

Description

本発明は、超電導磁石装置、およびこれを用いた磁気共鳴イメージング装置、並びに核磁気共鳴装置に関するものである。   The present invention relates to a superconducting magnet apparatus, a magnetic resonance imaging apparatus using the same, and a nuclear magnetic resonance apparatus.

磁気共鳴イメージング装置（以下、ＭＲＩ装置と称す）は、生体の大部分を構成する水素原子核の核磁気共鳴（ＮＭＲ）現象が組織によって異なることを利用して、生体組織を画像化するものであり、共鳴の強さや、共鳴の時間的変化の速さが画像のコントラストとして現われるようになっている。ＭＲＩ装置では、高画質の画像を得るために、撮像領域に高強度で高い静磁場均一度を有する磁場を生成する必要がある。そして、この高強度の静磁場を発生させるために、超電導磁石装置が用いられている。
一般的に、超電導磁石装置は、主に磁場空間に磁場を生成する超電導主コイルと、超電導主コイルが発生する磁場が装置外へ漏洩することを抑制するシールドコイルから成り、通常運転時の漏洩磁場は、許容範囲内に抑えることが可能な構成となっている。ここで、漏洩磁場は、通常運転時以外、例えば、ヒータを用いてマグネットを強制的に急速消磁させる場合においても、同様に抑制する必要がある。
ところで、超電導磁石装置は、外部からの熱侵入を遮断する輻射シールドのように、低抵抗とされた部材を有しているため、急速消磁させる場合には、誘導電流が輻射シールドに流れて、輻射シールドが作る漏洩磁場が制御できないことがある。
このような問題を解決する方法の一例として、超電導主コイルと超電導シールドコイルの外周側に、急速消磁時の漏洩磁場を抑制するように構成した良導体リングを設置した技術が開示されている（特許文献１参照）。
特開平４−２３３７０７号公報 A magnetic resonance imaging apparatus (hereinafter referred to as an MRI apparatus) images biological tissue by utilizing the fact that the nuclear magnetic resonance (NMR) phenomenon of hydrogen nuclei constituting most of the living body varies depending on the tissue. The strength of resonance and the speed of temporal change in resonance appear as image contrast. In the MRI apparatus, in order to obtain a high-quality image, it is necessary to generate a magnetic field having high intensity and high static magnetic field uniformity in the imaging region. And in order to generate this high intensity | strength static magnetic field, the superconducting magnet apparatus is used.
In general, a superconducting magnet device is mainly composed of a superconducting main coil that generates a magnetic field in a magnetic field space and a shield coil that suppresses leakage of the magnetic field generated by the superconducting main coil to the outside of the device. The magnetic field can be suppressed within an allowable range. Here, it is necessary to suppress the leakage magnetic field in the same manner even when the magnet is forcibly rapidly demagnetized using a heater, for example, other than during normal operation.
By the way, the superconducting magnet device has a low resistance member, such as a radiation shield that blocks heat intrusion from the outside. Therefore, when rapidly demagnetizing, an induced current flows to the radiation shield, The leakage magnetic field created by the radiation shield may not be controlled.
As an example of a method for solving such a problem, a technique is disclosed in which a good conductor ring configured to suppress a leakage magnetic field at the time of rapid demagnetization is installed on the outer peripheral side of a superconducting main coil and a superconducting shield coil (patent) Reference 1).
JP-A-4-233707

ところで、ＭＲＩ装置では、超電導主コイルの径方向内周側の空間を広く利用することが望まれており、超電導主コイルの径方向内周側において、追加の良導体リングを配置することは難しくなっている。そのため、超電導主コイル付近に設置する良導体リングは、超電導主コイルの径方向外周側に配置することが一般的であり、磁気的なバランスを取るためには、超電導シールドコイル付近に設置する良導体リングも、超電導シールドコイルの径方向外周側に設置される傾向にある。   By the way, in the MRI apparatus, it is desired to widely use the space on the radially inner periphery side of the superconducting main coil, and it becomes difficult to arrange an additional good conductor ring on the radially inner periphery side of the superconducting main coil. ing. Therefore, the good conductor ring installed near the superconducting main coil is generally placed on the outer periphery in the radial direction of the superconducting main coil. In order to achieve a magnetic balance, the good conductor ring installed near the superconducting shield coil However, it tends to be installed on the radially outer peripheral side of the superconducting shield coil.

しかしながら、超電導シールドコイルの径方向外周側に良導体リングを設置すると、これを支えるための支持構造部材を設ける必要が生じてくる。そこで、良導体リングに対して支持機能を併せ持たせるように構成することが考えられるが、そのように構成すると、例えば、アルミニウム合金等からなる良導体リングでは、ステンレス鋼のような構造材に比べて強度が低いために、径方向の板厚を分厚く形成する必要が生じ、結果的に、ＭＲＩ装置が大型化するという問題を生じる。   However, if a good conductor ring is installed on the radially outer side of the superconducting shield coil, it becomes necessary to provide a support structure member for supporting this. Therefore, it is conceivable that the good conductor ring is configured to have a supporting function. However, in this case, the good conductor ring made of, for example, an aluminum alloy is compared with a structural material such as stainless steel. Since the strength is low, it is necessary to form a thick plate in the radial direction, resulting in a problem that the MRI apparatus is enlarged.

一方で、ＭＲＩ装置は、高磁場化のニーズが高まるにつれ、０．５Ｔ程度の中磁場装置から１．５Ｔ〜３Ｔ程度の高磁場装置へ取って代わるケースが増えており、装置寸法および漏洩磁場の拡大を抑えることに期待が寄せられている。   On the other hand, as the need for higher magnetic field increases, the number of cases where MRI devices are replaced by medium magnetic field devices of about 0.5T to high magnetic field devices of about 1.5T to 3T is increasing. There are expectations that the expansion of

このような観点から、本発明は、急速消磁時においても漏洩磁場および超電導主コイルや超電導シールドコイルに発生する電圧を許容範囲内に抑え、かつ小型化に寄与する超電導磁石装置、およびこれを用いた磁気共鳴イメージング装置、並びに核磁気共鳴装置を提供することを課題とする。   From such a viewpoint, the present invention suppresses the leakage magnetic field and the voltage generated in the superconducting main coil and the superconducting shield coil within the allowable range even during rapid demagnetization, and uses the superconducting magnet device that contributes to downsizing. It is an object of the present invention to provide a magnetic resonance imaging apparatus and a nuclear magnetic resonance apparatus.

前記した課題を解決するための手段として本発明は、超電導主コイルと、超電導シールドコイルとを備え、前記超電導主コイルの径方向外側、または一対の前記超電導主コイル同士の間にあって、磁場空間を通る中心軸方向から見て前記超電導主コイルに少なくとも一部が重なる状態に配置された第１の良導体と、前記超電導シールドコイルの径方向内側に配置された第２の良導体と、を含んで構成した。これにより、急速消磁時には第１，第２の良導体に誘導電流が流れ、超電導主コイルおよび超電導シールドコイルの発生電圧を抑制することができる。また、超電導シールドコイルの外周側に良導体を配置する場合に比べて、支持強度を確保することができ、装置の外径寸法を抑えることができる。   As a means for solving the above-mentioned problems, the present invention comprises a superconducting main coil and a superconducting shield coil, and is located radially outside the superconducting main coil, or between the pair of superconducting main coils, and has a magnetic field space. A first good conductor disposed so as to at least partially overlap the superconducting main coil when viewed from the direction of the central axis passing therethrough, and a second good conductor disposed radially inside the superconducting shield coil did. Thereby, at the time of rapid demagnetization, an induced current flows through the first and second good conductors, and the generated voltage of the superconducting main coil and the superconducting shield coil can be suppressed. Further, the support strength can be ensured and the outer diameter of the apparatus can be suppressed as compared with the case where a good conductor is disposed on the outer peripheral side of the superconducting shield coil.

本発明によれば、急速消磁時においても漏洩磁場および超電導主コイルや超電導シールドコイルに発生する電圧を許容範囲内に抑え、かつ小型化に寄与する超電導磁石装置、およびこれを用いた磁気共鳴イメージング装置、並びに核磁気共鳴装置が得られる。   According to the present invention, a superconducting magnet device that suppresses leakage magnetic fields and voltages generated in a superconducting main coil and a superconducting shield coil within an allowable range even during rapid demagnetization, and contributes to miniaturization, and magnetic resonance imaging using the same An apparatus and a nuclear magnetic resonance apparatus are obtained.

以下、本発明の実施形態に係る超電導磁石装置が適用されるＭＲＩ装置について図面を参照して詳細に説明する。以下では、水平磁場方式のＭＲＩ装置について説明するが、本発明はこれに限定されるものではなく、他の垂直磁場方式等にも適用することができる。
（第１実施形態）
このＭＲＩ装置は、核磁気共鳴（ＮＭＲ）現象を利用して被検体（不図示、以下同じ）の断層画像を得るものであり、図１に示すように、被検体にＮＭＲ現象を誘起してＮＭＲ信号を受信するための各種装置を収容するガントリ１、被検体を載置するベッド２、このベッド２に載置された被検体を磁場空間（図２参照）内の撮像領域Ｆへ搬送する搬送手段２ａと、ガントリ１内の各種装置を制御する電源や各種制御装置を収納した制御装置３、検出された核磁気共鳴信号を処理するコンピュータ等の処理装置４、および処理された核磁気共鳴信号に基づき断層画像を表示する表示装置５等を含み、それぞれ電源・信号線６で接続される。ガントリ１、ベッド２および搬送手段２ａは、高周波電磁波と静磁場を遮蔽する図示しないシールドルーム内に配置され、制御装置３、処理装置４および表示装置５は、シールドルーム外に配置される。 Hereinafter, an MRI apparatus to which a superconducting magnet apparatus according to an embodiment of the present invention is applied will be described in detail with reference to the drawings. Hereinafter, a horizontal magnetic field type MRI apparatus will be described. However, the present invention is not limited to this, and can be applied to other vertical magnetic field methods.
(First embodiment)
This MRI apparatus uses a nuclear magnetic resonance (NMR) phenomenon to obtain a tomographic image of a subject (not shown, the same applies hereinafter), and induces the NMR phenomenon in the subject as shown in FIG. A gantry 1 that houses various devices for receiving NMR signals, a bed 2 on which the subject is placed, and the subject placed on the bed 2 are transported to the imaging region F in the magnetic field space (see FIG. 2). Conveying means 2a, power supply for controlling various devices in gantry 1, control device 3 housing various control devices, processing device 4 such as a computer for processing detected nuclear magnetic resonance signals, and processed nuclear magnetic resonance A display device 5 that displays a tomographic image based on the signal is included, and each is connected by a power source / signal line 6. The gantry 1, the bed 2, and the transport means 2a are disposed in a shield room (not shown) that shields high-frequency electromagnetic waves and static magnetic fields, and the control device 3, the processing device 4, and the display device 5 are disposed outside the shield room.

ガントリ１内には、図２、図３に示すように、超電導主コイル１０Ａ，１０Ｂ、超電導シールドコイル２０、および本発明の特徴的構成である、環状に一体的に形成された第１，第２の良導体１１，２１を、冷媒３０とともに収容する冷却容器４０と、この冷却容器４０を覆うように形成された輻射シールド５０と、冷却容器４０および輻射シールド５０を囲繞し、内部を真空にした真空容器６０等とからなる超電導磁石装置が設けられている。
冷却容器４０に収容される冷媒３０としては、例えば液体ヘリウムなどの液化した冷媒３０が用いられており、この冷媒３０と輻射シールド５０とを冷却するための図示しない冷凍機が真空容器６０に設置されている。冷却容器４０内は、このような冷媒３０によって、例えば、４．２Ｋ程度に保たれている。 In the gantry 1, as shown in FIGS. 2 and 3, the superconducting main coils 10A and 10B, the superconducting shield coil 20, and the first and first integrally formed in an annular shape, which is a characteristic configuration of the present invention. The cooling container 40 that houses the two good conductors 11 and 21 together with the refrigerant 30, the radiation shield 50 formed so as to cover the cooling container 40, the cooling container 40 and the radiation shield 50 are surrounded, and the inside is evacuated. A superconducting magnet device including a vacuum vessel 60 and the like is provided.
As the refrigerant 30 accommodated in the cooling container 40, for example, a liquefied refrigerant 30 such as liquid helium is used, and a refrigerator (not shown) for cooling the refrigerant 30 and the radiation shield 50 is installed in the vacuum container 60. Has been. The inside of the cooling container 40 is kept at, for example, about 4.2K by such a refrigerant 30.

静磁場の発生源となる超電導主コイル１０Ａ，１０Ｂおよび超電導シールドコイル２０は、磁場空間を通る中心軸Ｚの周りにそれぞれ環状に形成され、撮像領域Ｆを挟んで中心軸Ｚ方向に相対向してそれぞれ一対配置されている。
超電導主コイル１０Ａ，１０Ｂには、一定の電流が流れており、超電導シールドコイル２０には、超電導主コイル１０Ａ，１０Ｂとは逆方向の一定電流が流れている。これらの超電導主コイル１０Ａ，１０Ｂ、超電導シールドコイル２０は、電磁力、漏洩磁場、最大経験磁場、磁場均一度、および磁場強度を許容範囲内に抑えるように、位置、形状、および設置個数の変更が可能である。
本実施形態では、一方の超電導主コイル１０Ａの径方向（中心軸Ｚに鉛直に交差する方向）外側に、超電導シールドコイル２０が間隔を隔てて対向配置されている。これにより、後記するように、超電導主コイル１０Ａおよび超電導シールドコイル２０に沿わせて設けた第１，第２の良導体１１，２１も、超電導主コイル１０Ａと超電導シールドコイル２０との間において、径方向に間隔を隔てて対向配置されている。 The superconducting main coils 10A and 10B and the superconducting shield coil 20 that are the sources of the static magnetic field are formed in an annular shape around the central axis Z passing through the magnetic field space, and face each other in the direction of the central axis Z across the imaging region F. Each pair is arranged.
A constant current flows through the superconducting main coils 10A and 10B, and a constant current in the direction opposite to the superconducting main coils 10A and 10B flows through the superconducting shield coil 20. These superconducting main coils 10A and 10B and the superconducting shield coil 20 are changed in position, shape, and number of installations so as to keep the electromagnetic force, leakage magnetic field, maximum experience magnetic field, magnetic field uniformity, and magnetic field strength within an allowable range. Is possible.
In the present embodiment, the superconducting shield coil 20 is disposed opposite to the outer side in the radial direction of the one superconducting main coil 10A (the direction perpendicular to the central axis Z) with a gap therebetween. Thereby, as described later, the first and second good conductors 11 and 21 provided along the superconducting main coil 10A and the superconducting shield coil 20 also have a diameter between the superconducting main coil 10A and the superconducting shield coil 20. Oppositely arranged at intervals in the direction.

超電導主コイル１０Ａ，１０Ｂおよび超電導シールドコイル２０は、それぞれ円筒状に形成された巻き枠４１，４２によって支持されている。巻き枠４１の外周部には、周状の支持溝４１ａ，４１ｂが形成されており、これらの支持溝４１ａ，４１ｂ内に超電導主コイル１０Ａ，１０Ｂがそれぞれ配置されるようになっている。また、巻き枠４２の外周部には、周状の支持溝４２ａが形成されており、この支持溝４２ａ内に第２の良導体２１および超電導シールドコイル２０が配置されるようになっている。   Superconducting main coils 10A and 10B and superconducting shield coil 20 are supported by winding frames 41 and 42 formed in a cylindrical shape, respectively. Circumferential support grooves 41a and 41b are formed on the outer periphery of the winding frame 41, and the superconducting main coils 10A and 10B are arranged in the support grooves 41a and 41b, respectively. Further, a circumferential support groove 42a is formed on the outer peripheral portion of the winding frame 42, and the second good conductor 21 and the superconducting shield coil 20 are arranged in the support groove 42a.

第１の良導体１１は、超電導主コイル１０Ａの径方向外側に配置されており、超電導主コイル１０Ａにおける急速消磁時の漏洩磁場を抑制する役割を成す。本実施形態では、第１の良導体１１が、図示しない絶縁部材を介して超電導主コイル１０Ａの外周面に焼き嵌めされている。つまり、超電導主コイル１０Ａは、巻き枠４１の支持溝４１ａ内に収容された状態で、その外周側から焼き嵌めされる第１の良導体１１によって、巻き枠４１の支持溝４１ａ内に押え付けられるようにして強固に保持されるようになっている。このように超電導主コイル１０Ａの径方向外側に固定される第１の良導体１１は、超電導主コイル１０Ａにおける急速消磁時の漏洩磁場を抑制するとともに、超電導主コイル１０Ａを巻き枠４１に強固に固定するための固定機能を併せ備えたものとなっている。   The first good conductor 11 is arranged on the outer side in the radial direction of the superconducting main coil 10A, and serves to suppress a leakage magnetic field during rapid demagnetization in the superconducting main coil 10A. In the present embodiment, the first good conductor 11 is shrink-fitted on the outer peripheral surface of the superconducting main coil 10A via an insulating member (not shown). That is, the superconducting main coil 10 </ b> A is pressed into the support groove 41 a of the winding frame 41 by the first good conductor 11 that is shrink-fitted from the outer peripheral side in a state of being accommodated in the support groove 41 a of the winding frame 41. In this way, it is firmly held. As described above, the first good conductor 11 fixed to the outer side in the radial direction of the superconducting main coil 10A suppresses the leakage magnetic field at the time of rapid demagnetization in the superconducting main coil 10A, and firmly fixes the superconducting main coil 10A to the winding frame 41. It is equipped with a fixing function for this purpose.

第２の良導体２１は、超電導シールドコイル２０の径方向内側に配置されており、超電導シールドコイル２０における急速消磁時の漏洩磁場を抑制する役割を成す。本実施形態では、第２の良導体２１が、巻き枠４２の支持溝４２ａの周面に接するように配置されている。第２の良導体２１の外周部には、図示しない絶縁部材を介して超電導シールドコイル２０が接するように配置されている。
本実施形態では、超電導シールドコイル２０の外周部に、リング状の保持部材２２が焼き嵌めされている。つまり、第２の良導体２１および超電導シールドコイル２０は、巻き枠４２の支持溝４２ａ内に収容された状態で、超電導シールドコイル２０の外周部に焼き嵌めされる保持部材２２によって、支持溝４２ａ内に押え付けられるようにして強固に固定されるようになっている。保持部材２２は、非磁性の構造材、例えば、ステンレス鋼等を使用することができる。 The second good conductor 21 is disposed on the radially inner side of the superconducting shield coil 20 and serves to suppress a leakage magnetic field during rapid demagnetization in the superconducting shield coil 20. In the present embodiment, the second good conductor 21 is disposed so as to contact the peripheral surface of the support groove 42 a of the winding frame 42. The superconducting shield coil 20 is disposed on the outer periphery of the second good conductor 21 via an insulating member (not shown).
In the present embodiment, a ring-shaped holding member 22 is shrink-fitted on the outer peripheral portion of the superconducting shield coil 20. In other words, the second good conductor 21 and the superconducting shield coil 20 are accommodated in the support groove 42a by the holding member 22 that is shrink-fitted to the outer periphery of the superconducting shield coil 20 while being accommodated in the support groove 42a of the winding frame 42. It is designed to be firmly fixed so that it can be pressed against. The holding member 22 can be made of a nonmagnetic structural material such as stainless steel.

本実施形態では、第１，第２の良導体１１，２１の材料として、超電導主コイル１０Ａ、超電導シールドコイル２０に使用されるコイル線材に線膨張係数が近い材料、例えば、銅が用いられている。   In the present embodiment, as the material of the first and second good conductors 11 and 21, a material having a linear expansion coefficient close to that of the coil wire used for the superconducting main coil 10A and the superconducting shield coil 20, for example, copper is used. .

巻き枠４１は、冷却容器４０の一部を構成しており、支持材４３、輻射シールド５０、支持材４４を介して真空容器６０に設置されている。巻き枠４１，４２の間には、巻き枠連結材４５が配置され、この巻き枠連結材４５によって巻き枠４１，４２が相互に連結されている。   The winding frame 41 constitutes a part of the cooling container 40 and is installed in the vacuum container 60 via the support material 43, the radiation shield 50, and the support material 44. A reel connecting member 45 is disposed between the reels 41 and 42, and the reels 41 and 42 are connected to each other by the reel connecting member 45.

支持材４３，４４は、外部からの熱侵入を防ぐために、低熱伝導の材料、例えば、ＦＲＰ（繊維強化プラスチック）等を用いる。また、巻き枠４１，４２、巻き枠連結材４５は、非磁性の構造材、例えば、ステンレス鋼等を使用することができる。   The support members 43 and 44 are made of a low heat conductive material such as FRP (fiber reinforced plastic) in order to prevent heat from entering from the outside. The reels 41 and 42 and the reel connection member 45 can be made of a non-magnetic structural material, such as stainless steel.

超電導主コイル１０Ａ，１０Ｂおよび超電導シールドコイル２０は、図４に示すような回路で接続されている。つまり、一対の超電導主コイル１０Ａ，１０Ｂ、超電導シールドコイル２０のうち、一方の超電導主コイル１０Ａ，１０Ｂ、超電導シールドコイル２０（超電導コイル群Ａ１）および他方の超電導主コイル１０Ａ，１０Ｂ、超電導シールドコイル２０（超電導コイル群Ｂ１）が、直列接続され、これらの超電導コイル群Ａ１，Ｂ１と、永久電流スイッチ１５とが、並列接続されて主回路を構成している。そして、双方向ダイオード１６，１６が、それぞれの超電導コイル群Ａ１，Ｂ１と並列に接続されており、２つに分割された保護回路を構成している。
ここで、双方向ダイオード１６，１６は、２つのダイオードがそれぞれ逆方向に並列接続されて成る。
なお、超電導主コイル１０Ａ，１０Ｂ、超電導シールドコイル２０の組み合わせ、双方向ダイオード１６の構成数、保護回路の分割数は、急速消磁時のコイル電圧、温度上昇、電磁力、漏洩磁場を考慮して、これらが許容範囲に収まるように適宜設定することができる。 The superconducting main coils 10A and 10B and the superconducting shield coil 20 are connected by a circuit as shown in FIG. That is, out of the pair of superconducting main coils 10A and 10B and the superconducting shield coil 20, one superconducting main coil 10A and 10B, the superconducting shield coil 20 (superconducting coil group A1), the other superconducting main coil 10A and 10B, and the superconducting shield coil. 20 (superconducting coil group B1) are connected in series, and these superconducting coil groups A1 and B1 and the permanent current switch 15 are connected in parallel to form a main circuit. Bidirectional diodes 16 and 16 are connected in parallel with the respective superconducting coil groups A1 and B1 to form a protection circuit divided into two.
Here, the bidirectional diodes 16 and 16 are formed by connecting two diodes in parallel in opposite directions.
Note that the combination of the superconducting main coils 10A and 10B and the superconducting shield coil 20, the number of components of the bidirectional diode 16, and the number of divisions of the protection circuit take into consideration the coil voltage, temperature rise, electromagnetic force, and leakage magnetic field during rapid demagnetization. These can be set as appropriate so that they fall within the allowable range.

このような回路において、ＭＲＩ装置の運転時には、超電導コイル群Ａ１，Ｂ１と永久電流スイッチ１５とで構成する回路を電流が流れる。つまり、双方向ダイオード１６には、電流が流れないようになっている。なお、必要に応じてダイオードの数を調整することによって、仮に、超電導コイル群Ａ１，Ｂ１の励磁時および消磁時等に、双方向ダイオード１６のダイオード順方向電圧を上回る電圧が発生したような場合にも、電流がダイオード側に分流しないようにすることができる。
また、急速消磁時には、超電導主コイル１０Ａ，１０Ｂ、超電導シールドコイル２０において、ダイオード順方向電圧を大きく上回る電圧が発生するため、双方向ダイオード１６側に電流が流れる。 In such a circuit, during operation of the MRI apparatus, a current flows through a circuit constituted by the superconducting coil groups A1 and B1 and the permanent current switch 15. That is, no current flows through the bidirectional diode 16. In the case where a voltage exceeding the diode forward voltage of the bidirectional diode 16 is generated by exciting and demagnetizing the superconducting coil groups A1 and B1 by adjusting the number of diodes as necessary. In addition, the current can be prevented from being shunted to the diode side.
Further, at the time of rapid demagnetization, a voltage that greatly exceeds the diode forward voltage is generated in the superconducting main coils 10A and 10B and the superconducting shield coil 20, so that a current flows to the bidirectional diode 16 side.

以上のようなＭＲＩ装置では、超電導主コイル１０Ａに第１の良導体１１が設けられているとともに、超電導シールドコイル２０に第２の良導体２１が設けられているので、急速消磁時には、誘導電流が第１，第２の良導体１１，２１に流れ、これによって、超電導主コイル１０Ａ，１０Ｂ、超電導シールドコイル２０の発生電圧と、漏洩磁場を許容範囲内に抑えることが可能となる。   In the MRI apparatus as described above, since the first good conductor 11 is provided in the superconducting main coil 10A and the second good conductor 21 is provided in the superconducting shield coil 20, the induced current is reduced during the rapid demagnetization. 1 and the second good conductors 11 and 21, whereby the generated voltages of the superconducting main coils 10 </ b> A and 10 </ b> B and the superconducting shield coil 20 and the leakage magnetic field can be suppressed within an allowable range.

以下では、本実施形態において得られる効果を説明する。
（１）本実施形態では、急速消磁時に、超電導主コイル１０Ａに設けられた第１の良導体１１、および超電導シールドコイル２０に設けられた第２の良導体２１に、誘導電流が流れ、超電導主コイル１０Ａ，１０Ｂ、超電導シールドコイル２０の発生電圧を抑制することができるとともに、漏洩磁場を許容範囲内に抑えることができる。
（２）第２の良導体２１は、超電導シールドコイル２０の径方向内側に配置されているので、超電導シールドコイル２０の径方向外側に良導体を設けたときのように、良導体を支持するための支持構造部材を付設したり、支持機能をもたせるべく良導体を大型化したりする必要がなくなり、超電導シールドコイル２０の径方向外周側の構造を簡単な構造とすることができて、超電導磁石装置の小型化（コンパクト化）、ひいてはＭＲＩ装置の小型化を図ることができる。
（３）第１，第２の良導体１１，２１に使用される材料として銅が使用され、銅は、その線膨張係数が、超電導主コイル１０Ａ、超電導シールドコイル２０に使用されるコイル線材の線膨張係数に近いので、超電導主コイル１０Ａおよび超電導シールドコイル２０と、第１，第２の良導体１１，２１との間に、冷却時の熱収縮差による隙間が生じ難くなっている。したがって、長期的に安定した作動を実現することができる超電導磁石装置、ＭＲＩ装置が得られる。
（４）第１の良導体１１は、超電導主コイル１０Ａの径方向外側に配置されているので、超電導主コイル１０Ａの径方向内側の空間を広くとることができ、磁場空間を通る中心軸Ｚ周りに被検体を好適に収容することができる空間を容易に形成することができる。
（５）第１の良導体１１は、超電導主コイル１０Ａの径方向外側に配置されるとともに、第２の良導体２１は、超電導シールドコイル２０の径方向内側に配置される構成であるので、超電導主コイル１０Ａと超電導シールドコイル２０との間に形成される空間を有効に利用して、この空間に、第１，第２の良導体１１，２１を好適に配置することができ、第２の良導体２１を超電導シールドコイル２０の径方向外側に配置した場合に比べて、超電導磁石装置が径方向に大型化するのを好適に阻止することができる。したがって、ＭＲＩ装置の小型化を図ることが可能となる。
（６）第２の良導体２１は、巻き枠４２の支持溝４２ａ内に支持され、超電導シールドコイル２０が、支持溝４２ａ内において、図示しない絶縁部材を介して第２の良導体２１の外周面に接しているので、支持溝４２ａ内に超電導シールドコイル２０を取り付けることによって、第２の良導体２１を超電導シールドコイル２０で好適に保持することができる。
また、超電導シールドコイル２０の外周部に焼き嵌めされた保持部材２２によって、支持溝４２ａ内に超電導シールドコイル２０および第２の良導体２１の両方を強固に固定することができる。
（７）第１の良導体１１は、図示しない絶縁部材を介して超電導主コイル１０Ａの外周面に焼き嵌めされているので、超電導主コイル１０Ａにおける急速消磁時の漏洩磁場を抑制する機能と、超電導主コイル１０Ａを巻き枠４１に強固に固定するための固定機能とを併せ備えたものとなっており、機能性、組付性に優れている。また、第１の良導体１１は、超電導主コイル１０Ａに焼き嵌めされて固定されるので、第１の良導体１１を固定するための固定部材を別途必要とせず、その分、軽量化が可能であり、また、コストを低減することができる。 Below, the effect acquired in this embodiment is demonstrated.
(1) In this embodiment, at the time of rapid demagnetization, an induced current flows through the first good conductor 11 provided in the superconducting main coil 10A and the second good conductor 21 provided in the superconducting shield coil 20, and the superconducting main coil The generated voltage of 10A, 10B and the superconducting shield coil 20 can be suppressed, and the leakage magnetic field can be suppressed within an allowable range.
(2) Since the second good conductor 21 is disposed inside the superconducting shield coil 20 in the radial direction, the support for supporting the good conductor as when the good conductor is provided outside the superconducting shield coil 20 in the radial direction. There is no need to attach a structural member or to increase the size of a good conductor so as to provide a support function, and the structure on the radially outer peripheral side of the superconducting shield coil 20 can be simplified, and the superconducting magnet device can be downsized. (Compact), and thus the size of the MRI apparatus can be reduced.
(3) Copper is used as the material used for the first and second good conductors 11 and 21, and the coefficient of linear expansion of copper is the wire of the coil wire used for the superconducting main coil 10A and the superconducting shield coil 20. Since it is close to the expansion coefficient, a gap due to a difference in thermal contraction during cooling is hardly generated between the superconducting main coil 10A and the superconducting shield coil 20 and the first and second good conductors 11 and 21. Therefore, it is possible to obtain a superconducting magnet device and an MRI device that can realize stable operation over a long period of time.
(4) Since the first good conductor 11 is disposed on the radially outer side of the superconducting main coil 10A, a space on the radially inner side of the superconducting main coil 10A can be widened, and the center axis Z around the magnetic field space can be taken. Thus, it is possible to easily form a space that can suitably accommodate the subject.
(5) Since the first good conductor 11 is arranged on the radially outer side of the superconducting main coil 10A and the second good conductor 21 is arranged on the radially inner side of the superconducting shield coil 20, the superconducting main By effectively utilizing the space formed between the coil 10A and the superconducting shield coil 20, the first and second good conductors 11 and 21 can be suitably arranged in this space, and the second good conductor 21 As compared with the case where is disposed outside the superconducting shield coil 20 in the radial direction, the superconducting magnet device can be suitably prevented from being enlarged in the radial direction. Therefore, it is possible to reduce the size of the MRI apparatus.
(6) The second good conductor 21 is supported in the support groove 42a of the winding frame 42, and the superconducting shield coil 20 is placed on the outer peripheral surface of the second good conductor 21 through an insulating member (not shown) in the support groove 42a. Therefore, the second good conductor 21 can be suitably held by the superconducting shield coil 20 by attaching the superconducting shield coil 20 in the support groove 42a.
In addition, both the superconducting shield coil 20 and the second good conductor 21 can be firmly fixed in the support groove 42 a by the holding member 22 shrink-fitted on the outer periphery of the superconducting shield coil 20.
(7) Since the first good conductor 11 is shrink-fitted on the outer peripheral surface of the superconducting main coil 10A via an insulating member (not shown), the superconducting function has a function of suppressing the leakage magnetic field at the time of rapid demagnetization in the superconducting main coil 10A. The main coil 10A is provided with a fixing function for firmly fixing the main coil 10A to the winding frame 41, and is excellent in functionality and assembly. Further, since the first good conductor 11 is fixed by being shrink-fitted to the superconducting main coil 10A, a separate fixing member for fixing the first good conductor 11 is not required, and the weight can be reduced accordingly. In addition, the cost can be reduced.

（第２実施形態）
図５は本発明の第２実施形態の超電導磁石装置が適用されるＭＲＩ装置のガントリの断面図である。この例では、第１の良導体１１’，１１’が、一対の超電導主コイル１０Ａ，１０Ａ同士の間に配置されており、中心軸Ｚ方向から見たときに、超電導主コイル１０Ａ，１０Ａに、第１の良導体１１’，１１’の少なくとも一部が重なる状態となるように配置されている。
巻き枠４１には、第１の良導体１１’を支持するための専用の支持溝１１ａが形成されている。この例では、超電導主コイル１０Ａと超電導主コイル１０Ｂとの間に位置するように支持溝１１ａが形成されており、第１の良導体１１’は、図示しない絶縁部材を介して、各支持溝１１ａに支持される。つまり、第１の良導体１１’は、前記第１実施形態のように超電導主コイル１０Ａに接するように設けられるのではなく、超電導主コイル１０Ａから中心軸Ｚ方向に離間した位置において巻き枠４１に支持されるように構成される。 (Second Embodiment)
FIG. 5 is a sectional view of a gantry of an MRI apparatus to which the superconducting magnet apparatus according to the second embodiment of the present invention is applied. In this example, the first good conductors 11 ′ and 11 ′ are disposed between the pair of superconducting main coils 10A and 10A. When viewed from the central axis Z direction, the superconducting main coils 10A and 10A are It arrange | positions so that at least one part of 1st good conductor 11 ', 11' may overlap.
The winding frame 41 is formed with a dedicated support groove 11a for supporting the first good conductor 11 ′. In this example, a support groove 11a is formed so as to be positioned between the superconducting main coil 10A and the superconducting main coil 10B, and the first good conductor 11 'is connected to each support groove 11a via an insulating member (not shown). Supported by That is, the first good conductor 11 ′ is not provided so as to be in contact with the superconducting main coil 10A as in the first embodiment, but is disposed on the winding frame 41 at a position separated from the superconducting main coil 10A in the central axis Z direction. Configured to be supported.

このように構成することによって、超電導主コイル１０Ａと第１の良導体１１’とが絶縁部材を介して接するように設けた場合に比べて、構造的制約が少なくなり、その分、低コスト化が可能となる。
また、第１の良導体１１’は、複数個に分割して構成してもよく、巻き枠４１の支持溝１１ａに組み付けた後に、溶接等によって接合するようにしてもよい。このように、第１の良導体１１’を複数個に分割して構成することによって、組み付けの自由度が高まるようになり、中心軸Ｚ方向から組み付けることができないような場合であっても、巻き枠４１に予め組み付けておく等の組付手法を採ることができるようになる。 By configuring in this way, structural restrictions are reduced compared to the case where the superconducting main coil 10A and the first good conductor 11 ′ are provided so as to be in contact with each other via an insulating member, and the cost is reduced accordingly. It becomes possible.
Further, the first good conductor 11 ′ may be divided into a plurality of parts, or may be joined by welding or the like after being assembled in the support groove 11 a of the winding frame 41. As described above, the first good conductor 11 'is divided into a plurality of parts, so that the degree of freedom of assembly is increased, and even if the assembly cannot be performed from the central axis Z direction, An assembling method such as pre-assembling the frame 41 can be employed.

（第３実施形態）
図６は本発明の第３実施形態のＭＲＩ装置に適用される超電導磁石装置を備えたガントリの断面図である。この例では、第２の良導体２１Ａ，２１Ａが、良導体線材、例えば、銅線を周方向に巻回して形成されている点が異なっている。
このように、第２の良導体２１Ａ，２１Ａを良導体線材で形成することによって、一体的に形成された第２の良導体２１，２１（図２等参照）を用いたときに比べて構造的制約が少なくなり、その分、低コスト化が可能となる。また、組み付けの自由度が高まるようになり、中心軸Ｚ方向から組み付けることに障害があるような場合にも、これをうまく避けて巻き付けることが可能となる。 (Third embodiment)
FIG. 6 is a sectional view of a gantry provided with a superconducting magnet device applied to the MRI apparatus of the third embodiment of the present invention. This example is different in that the second good conductors 21A and 21A are formed by winding a good conductor wire, for example, a copper wire in the circumferential direction.
As described above, the second good conductors 21A and 21A are formed of the good conductor wire material, so that there are structural restrictions compared to the case where the second good conductors 21 and 21 (see FIG. 2 and the like) formed integrally are used. As a result, the cost can be reduced accordingly. In addition, the degree of freedom of assembly is increased, and even when there is an obstacle in assembly from the central axis Z direction, it is possible to avoid the problem well and wind it.

ここで、前記第１〜第３実施形態で示した超電導磁石装置を適用して形成した核磁気共鳴装置（以下、ＮＭＲ装置と略称する）の構成の一例について図７を参照して説明する。なお、ここでは、第１実施形態（図２，図３）で示した超電導磁石装置を用いた例を示すが、これに限られることはなく、第２，第３実施形態で示した構成を適宜用いてもよい。また、ＮＭＲ装置に適用する場合には、冷却容器４０や真空容器６０などの形状、図７に示した、超電導主コイル１０Ａ’、超電導シールドコイル２０’の数や位置、形状などは適宜変更される。   Here, an example of the configuration of a nuclear magnetic resonance apparatus (hereinafter abbreviated as NMR apparatus) formed by applying the superconducting magnet apparatus shown in the first to third embodiments will be described with reference to FIG. Here, an example using the superconducting magnet device shown in the first embodiment (FIGS. 2 and 3) is shown, but the present invention is not limited to this, and the configuration shown in the second and third embodiments is used. You may use suitably. When applied to the NMR apparatus, the shape of the cooling container 40, the vacuum container 60, and the like, and the number, position, and shape of the superconducting main coil 10A ′ and the superconducting shield coil 20 ′ shown in FIG. The

ＮＭＲ装置は、図７に示すように、超電導磁石装置が備えた超電導主コイル１０Ａ’、超電導シールドコイル２０’によって磁場が形成される磁場空間Ｋに、検体（不図示、以下同じ）を入れるサンプル管Ｓが設置され、サンプル管Ｓに入れた検体からの核磁気共鳴信号を捉えるプローブＰ、プローブＰで捉えた核磁気共鳴信号を解析する解析手段としてのスペクトロメータＳＭ、そして超電導磁石装置やスペクトロメータＳＭの動作を制御するコンピュータ等の機器７０などで構成される。機器７０と超電導磁石装置およびスペクトロメータＳＭ、そしてスペクトロメータＳＭとプローブＰは、各々、配線Ｗを介して電気的に接続されている。また、図７に示すようなＮＭＲ装置では、冷凍機（不図示）による振動を防止するため、冷凍機を設けない構成とする場合もある。   As shown in FIG. 7, the NMR apparatus is a sample in which a specimen (not shown, the same applies hereinafter) is placed in a magnetic field space K in which a magnetic field is formed by a superconducting main coil 10A ′ and a superconducting shield coil 20 ′ provided in the superconducting magnet apparatus. A tube S is installed, a probe P for capturing a nuclear magnetic resonance signal from a specimen placed in the sample tube S, a spectrometer SM as an analysis means for analyzing the nuclear magnetic resonance signal captured by the probe P, and a superconducting magnet device or a spectrometer It comprises a device 70 such as a computer for controlling the operation of the meter SM. The device 70, the superconducting magnet device and the spectrometer SM, and the spectrometer SM and the probe P are electrically connected via a wiring W, respectively. In addition, the NMR apparatus as shown in FIG. 7 may be configured not to include a refrigerator in order to prevent vibration caused by a refrigerator (not shown).

このようなＮＭＲ装置に用いた場合にも、前記した（１）〜（７）と同様の効果を得ることができる。すなわち、主として、急速消磁時に、超電導主コイル１０Ａ’に設けられた第１の良導体１１’’、および超電導シールドコイル２０’に設けられた第２の良導体２１’に、誘導電流が流れ、超電導主コイル１０Ａ’、超電導シールドコイル２０’の発生電圧を抑制することができるとともに、漏洩磁場を許容範囲内に抑えることができる。
また、第２の良導体２１’は、超電導シールドコイル２０’の径方向内側に配置されているので、超電導シールドコイル２０’の径方向外側に良導体を設けたときのように、良導体を支持するための支持構造部材を付設したり、支持機能をもたせるべく良導体を大型化したりする必要がなくなり、超電導シールドコイル２０’の径方向外周側の構造を簡単な構造とすることができて、超電導磁石装置の小型化（コンパクト化）、ひいてはＮＭＲ装置の小型化を図ることができる。 Even when used in such an NMR apparatus, the same effects as in the above (1) to (7) can be obtained. That is, mainly during rapid demagnetization, an induced current flows through the first good conductor 11 '' provided in the superconducting main coil 10A 'and the second good conductor 21' provided in the superconducting shield coil 20 '. The generated voltage of the coil 10A ′ and the superconducting shield coil 20 ′ can be suppressed, and the leakage magnetic field can be suppressed within an allowable range.
In addition, since the second good conductor 21 'is arranged on the radially inner side of the superconducting shield coil 20', it supports the good conductor as when the good conductor is provided on the radially outer side of the superconducting shield coil 20 '. It is not necessary to attach a supporting structure member or to increase the size of a good conductor so as to provide a supporting function, and the structure on the radially outer peripheral side of the superconducting shield coil 20 'can be simplified, and the superconducting magnet device Therefore, the NMR apparatus can be reduced in size.

本発明の第１実施形態の超電導磁石装置が適用されるＭＲＩ装置の概念図である。It is a conceptual diagram of the MRI apparatus with which the superconducting magnet apparatus of 1st Embodiment of this invention is applied. 同じくガントリの断面図である。It is a sectional view of the same gantry. 同じくガントリの一部切断斜視図である。It is a partially cut perspective view of the gantry. 同じく保護回路を示す回路図である。It is a circuit diagram which similarly shows a protection circuit. 本発明の第２実施形態の超電導磁石装置が適用されるＭＲＩ装置のガントリの断面図である。It is sectional drawing of the gantry of the MRI apparatus with which the superconducting magnet apparatus of 2nd Embodiment of this invention is applied. 本発明の第３実施形態の超電導磁石装置が適用されるＭＲＩ装置のガントリの断面図である。It is sectional drawing of the gantry of the MRI apparatus with which the superconducting magnet apparatus of 3rd Embodiment of this invention is applied. 本発明の超電導磁石装置が適用されるＮＭＲ装置の概念図である。It is a conceptual diagram of the NMR apparatus to which the superconducting magnet apparatus of this invention is applied.

符号の説明Explanation of symbols

１ ガントリ
２ ベッド
２ａ 搬送手段
１０Ａ 超電導主コイル
１０Ａ’ 超電導主コイル
１１ 第１の良導体
１１’ 第１の良導体
１１ａ 支持溝
１５ 永久電流スイッチ
１６ 双方向ダイオード
２０ 超電導シールドコイル
２０’ 超電導シールドコイル
２１ 第２の良導体
２１’ 第２の良導体
２１Ａ 第２の良導体
２２ 保持部材
３０ 冷媒
４０ 冷却容器
４１，４２ 巻き枠
４１ａ，４２ａ 支持溝
４３，４４ 支持材
５０ 輻射シールド
６０ 真空容器
７０ 機器
Ａ１ 超電導コイル群
Ｂ１ 超電導コイル群
Ｆ 撮像領域
Ｚ 中心軸 1 Gantry 2 Bed 2a Conveying means 10A Superconducting main coil 10A 'Superconducting main coil 11 First good conductor 11' First good conductor 11a Support groove 15 Permanent current switch 16 Bidirectional diode 20 Superconducting shield coil 20 'Superconducting shield coil 21 Second 21 '2nd good conductor 21A 2nd good conductor 22 Holding member 30 Refrigerant 40 Cooling vessel 41, 42 Winding frame 41a, 42a Support groove 43, 44 Support material 50 Radiation shield 60 Vacuum vessel 70 Equipment A1 Superconducting coil group B1 Superconductivity Coil group F Imaging area Z Central axis

Claims (8)

磁場空間を通る中心軸周りに環状に形成され、前記中心軸方向に相対向して配置された複数対のコイルおよび複数対の良導体を有する超電導磁石装置であって、
複数対の前記コイルは、超電導主コイルと、前記中心軸方向から見て前記超電導主コイルの径方向外側に間隔を隔てて配置され、前記超電導主コイルと逆向きの電流が流れる超電導シールドコイルと、を備え、
複数対の前記良導体は、
前記超電導主コイルの径方向外側に配置され、または一対の前記超電導主コイル同士の間にあって、前記中心軸方向から見て前記超電導主コイルに少なくとも一部が重なる状態に配置された第１の良導体と、
前記超電導シールドコイルの径方向内側に配置された第２の良導体と、を含んで構成されることを特徴とする超電導磁石装置。 A superconducting magnet apparatus having a plurality of pairs of coils and a plurality of pairs of good conductors formed in an annular shape around a central axis passing through a magnetic field space and arranged to face each other in the direction of the central axis,
The plurality of pairs of coils are a superconducting main coil, a superconducting shield coil that is arranged at an interval on the radially outer side of the superconducting main coil when viewed from the central axis direction, and in which a current in a direction opposite to that of the superconducting main coil flows. With
The multiple pairs of good conductors are:
The first good conductor disposed outside the superconducting main coil in the radial direction, or disposed between the pair of superconducting main coils so as to at least partially overlap the superconducting main coil when viewed from the central axis direction. When,
A superconducting magnet device comprising: a second good conductor disposed radially inside the superconducting shield coil. 前記第２の良導体は、前記超電導シールドコイルを支持する巻き枠に設けられた支持溝に支持されおり、前記超電導シールドコイルは、前記支持溝内において、絶縁部材を介して前記第２の良導体に接していることを特徴とする請求項１に記載の超電導磁石装置。   The second good conductor is supported by a support groove provided in a winding frame that supports the superconducting shield coil, and the superconducting shield coil is connected to the second good conductor via an insulating member in the support groove. The superconducting magnet device according to claim 1, wherein the superconducting magnet device is in contact. 前記超電導主コイルは、これを支持する巻き枠に支持されており、
前記第１の良導体は、前記超電導主コイルの径方向外側に配置されており、絶縁部材を介して前記超電導主コイルに焼き嵌めされたことを特徴とする請求項１または請求項２に記載の超電導磁石装置。 The superconducting main coil is supported by a winding frame that supports the superconducting main coil,
The said 1st good conductor is arrange | positioned in the radial direction outer side of the said superconducting main coil, and it was shrink-fitted in the said superconducting main coil via the insulating member, The Claim 1 or Claim 2 characterized by the above-mentioned. Superconducting magnet device. 前記超電導シールドコイルの外周部には、リング状の保持部材が配置されており、前記保持部材は前記超電導シールドコイルに焼き嵌めされたことを特徴とする請求項２または請求項３に記載の超電導磁石装置。   The superconducting device according to claim 2 or 3, wherein a ring-shaped holding member is disposed on an outer peripheral portion of the superconducting shield coil, and the holding member is shrink-fitted to the superconducting shield coil. Magnet device. 複数対の前記良導体は、そのうちの少なくとも一対が銅製であることを特徴とする請求項１から請求項４のいずれか１項に記載の超電導磁石装置。   5. The superconducting magnet device according to claim 1, wherein at least one of the plurality of pairs of good conductors is made of copper. 複数対の前記良導体は、そのうちの少なくとも一対が良導体線材を巻回して形成されていることを特徴とする請求項１から請求項５のいずれか１項に記載の超電導磁石装置。   The superconducting magnet device according to any one of claims 1 to 5, wherein at least a pair of the good conductors is formed by winding a good conductor wire. 請求項１から請求項６のいずれか１項に記載の超電導磁石装置を備えた磁気共鳴イメージング装置であって、
被検体を載置するベッドと、このベッドに載置された前記被検体を前記磁場空間内の撮像領域へ搬送する搬送手段と、この搬送手段によって前記撮像領域に搬送された前記被検体からの核磁気共鳴号を解析する解析手段とを備えたことを特徴とする磁気共鳴イメージング装置。 A magnetic resonance imaging apparatus comprising the superconducting magnet apparatus according to any one of claims 1 to 6,
A bed on which the subject is placed, transport means for transporting the subject placed on the bed to the imaging region in the magnetic field space, and from the subject transported to the imaging region by the transport means A magnetic resonance imaging apparatus comprising: an analysis unit that analyzes a nuclear magnetic resonance signal. 請求項１から請求項６のいずれか１項に記載の超電導磁石装置を備えた核磁気共鳴装置であって、
前記磁場空間内に配置された検体からの核磁気共鳴信号を捉えるプローブと、前記プローブで捉えた信号を解析する解析手段とを備えたことを特徴とする核磁気共鳴装置。 A nuclear magnetic resonance apparatus comprising the superconducting magnet device according to any one of claims 1 to 6,
A nuclear magnetic resonance apparatus comprising: a probe that captures a nuclear magnetic resonance signal from a specimen disposed in the magnetic field space; and an analysis unit that analyzes the signal captured by the probe.
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Cited By (7)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
WO2014118390A2 (en) * 2013-02-04 2014-08-07 Siemens Plc Superconducting magnet coil arrangement
CN104183355A (en) * 2013-11-12 2014-12-03 上海联影医疗科技有限公司 Superconducting magnet system and shielding coil assembly
WO2015194576A1 (en) * 2014-06-18 2015-12-23 株式会社日立製作所 Super-conducting wire, super-conducting coil, and magnetic resonance imaging device
CN106683819A (en) * 2015-09-15 2017-05-17 上海联影医疗科技有限公司 Magnet device
WO2018020879A1 (en) * 2016-07-27 2018-02-01 株式会社日立製作所 Superconducting magnet device, magnetic resonance imaging device, mri-guided radiotherapy device, and operating room
JP2020177953A (en) * 2019-04-15 2020-10-29 三菱電機株式会社 Superconducting coil assembly and manufacturing method thereof
EP3920196A1 (en) * 2016-10-31 2021-12-08 Tokamak Energy Ltd Quench protection in superconducting magnets

Citations (3)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JPH04237105A (en) * 1991-01-22 1992-08-25 Fuji Electric Co Ltd Superconducting electromagnet
JPH04299275A (en) * 1990-11-30 1992-10-22 Philips Gloeilampenfab:Nv Magnetic resonance device
JPH06132120A (en) * 1992-10-19 1994-05-13 Mitsubishi Electric Corp Superconductive magnet apparatus

Patent Citations (3)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JPH04299275A (en) * 1990-11-30 1992-10-22 Philips Gloeilampenfab:Nv Magnetic resonance device
JPH04237105A (en) * 1991-01-22 1992-08-25 Fuji Electric Co Ltd Superconducting electromagnet
JPH06132120A (en) * 1992-10-19 1994-05-13 Mitsubishi Electric Corp Superconductive magnet apparatus

Cited By (18)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
CN105103247B (en) * 2013-02-04 2019-06-21 西门子医疗有限公司 Superconducting magnetic coil device
WO2014118390A3 (en) * 2013-02-04 2014-10-23 Siemens Plc Superconducting magnet coil arrangement
EP3252784A1 (en) * 2013-02-04 2017-12-06 Siemens Healthcare Limited Superconducting magnet coil arrangement
US10365337B2 (en) 2013-02-04 2019-07-30 Siemens Healthcare Limited Superconducting magnet coil arrangement
WO2014118390A2 (en) * 2013-02-04 2014-08-07 Siemens Plc Superconducting magnet coil arrangement
KR101874039B1 (en) 2013-02-04 2018-07-05 지멘스 헬스케어 리미티드 Superconducting magnet coil arrangement
CN105103247A (en) * 2013-02-04 2015-11-25 西门子有限公司 Superconducting magnet coil arrangement
CN104183355A (en) * 2013-11-12 2014-12-03 上海联影医疗科技有限公司 Superconducting magnet system and shielding coil assembly
JPWO2015194576A1 (en) * 2014-06-18 2017-04-20 株式会社日立製作所 Superconducting wire, superconducting coil and magnetic resonance imaging apparatus
CN106463230A (en) * 2014-06-18 2017-02-22 株式会社日立制作所 Super-conducting wire, super-conducting coil, and magnetic resonance imaging device
WO2015194576A1 (en) * 2014-06-18 2015-12-23 株式会社日立製作所 Super-conducting wire, super-conducting coil, and magnetic resonance imaging device
EP3159899A4 (en) * 2014-06-18 2018-03-21 Hitachi, Ltd. Super-conducting wire, super-conducting coil, and magnetic resonance imaging device
US10424428B2 (en) 2014-06-18 2019-09-24 Hitachi, Ltd. Super-conducting wire, super-conducting coil, and magnetic resonance imaging device
CN106683819A (en) * 2015-09-15 2017-05-17 上海联影医疗科技有限公司 Magnet device
WO2018020879A1 (en) * 2016-07-27 2018-02-01 株式会社日立製作所 Superconducting magnet device, magnetic resonance imaging device, mri-guided radiotherapy device, and operating room
US11776721B2 (en) 2016-10-31 2023-10-03 Tokamak Energy Ltd Quench protection in high-temperature superconducting magnets
EP3920196A1 (en) * 2016-10-31 2021-12-08 Tokamak Energy Ltd Quench protection in superconducting magnets
JP2020177953A (en) * 2019-04-15 2020-10-29 三菱電機株式会社 Superconducting coil assembly and manufacturing method thereof

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