JP2024035576A - Superconducting magnet device and magnetic resonance imaging device - Google Patents

Abstract

[Problem] To suppress rapid changes in the main magnetic field. A superconducting magnet device includes a superconducting coil, a thermal radiation shield, a first switch, and a compensation coil. The superconducting coil is formed of superconducting wire. The thermal radiation shield shields thermal radiation from the superconducting coil. The first switch forms a closed circuit with the superconducting coil. The compensation coil generates a magnetic field that compensates for changes in the magnetic field caused by the superconducting coil when quenching occurs in the superconducting coil. [Selection diagram] Figure 1

Description

Embodiments disclosed in this specification and the drawings relate to a superconducting magnet device and a magnetic resonance imaging device.

Conventionally, when a quench occurs in which a part of a superconducting coil included in a superconducting magnet device changes from a superconducting state to a normal conducting state, the superconducting wire is protected by attenuating the current flowing through the superconducting coil in a short period of time.

As the current flowing through the superconducting coil is attenuated in a short period of time, the main magnetic field changes rapidly, and eddy currents may be generated in the highly conductive material disposed inside the superconducting magnet device. Additionally, Lorentz forces can be generated in highly conductive materials due to eddy currents and the main magnetic field. In particular, thermal radiation shields can be subject to large Lorentz forces due to eddy currents and the main magnetic field.

However, if a structure is to withstand Lorentz force, design options will be limited. Therefore, in order to reduce the generation of eddy currents, there is a need for technology that suppresses rapid changes in the main magnetic field.

Japanese Patent Application Publication No. 11-215615

One of the problems that the embodiments disclosed in this specification and drawings attempt to solve is to suppress sudden changes in the main magnetic field. However, the problems to be solved by the embodiments disclosed in this specification and the drawings are not limited to the above problems. Problems corresponding to the effects of each configuration shown in the embodiments described later can also be positioned as other problems.

A superconducting magnet device according to an embodiment includes a superconducting coil, a thermal radiation shield, a first switch, and a compensation coil. The superconducting coil is formed of superconducting wire. The thermal radiation shield shields thermal radiation from the superconducting coil. The first switch forms a closed circuit with the superconducting coil. The compensation coil generates a magnetic field that compensates for a change in the magnetic field caused by the superconducting coil when quench occurs in the superconducting coil.

FIG. 1 is a diagram showing an example of an MRI apparatus according to the first embodiment. FIG. 2 is a diagram showing an example of the circuit configuration of the superconducting magnet device according to the first embodiment. FIG. 3 is a diagram showing an example of the arrangement of superconducting coils and compensation coils in the superconducting magnet device according to the first embodiment. FIG. 4 is a diagram showing an example of a circuit configuration of a superconducting magnet device according to the second embodiment. FIG. 5 is a diagram showing an example of a circuit configuration of a superconducting magnet device according to the third embodiment. FIG. 6 is a diagram showing an example of a circuit configuration of a superconducting magnet device according to the fourth embodiment. FIG. 7 is a diagram illustrating an example of the arrangement of superconducting coils and compensation coils in a superconducting magnet device according to Modification 1.

Hereinafter, a superconducting magnet device and a magnetic resonance imaging device according to the present embodiment will be described with reference to the drawings. In the following embodiments, parts with the same reference numerals perform similar operations, and redundant explanations will be omitted as appropriate.

(First embodiment)
FIG. 1 is a diagram showing an example of an MRI (Magnetic Resonance Imaging) apparatus 100 according to the first embodiment. As shown in FIG. 1, the MRI apparatus 100 includes a static magnetic field magnet 101, a gradient magnetic field coil 103, a gradient magnetic field power supply 105, a bed 107, a bed control circuit 109, a transmission circuit 113, a transmission coil 115, It includes a receiving coil 117, a receiving circuit 119, an imaging control circuit 121, a system control circuit 123, a memory 125, an input interface 127, a display 129, and a processing circuit 131. MRI apparatus 100 is an example of a magnetic resonance imaging apparatus.

The static field magnet 101 is a hollow, substantially cylindrical magnet. The static magnetic field magnet 101 generates a substantially uniform static magnetic field in the internal space. As the static field magnet 101, for example, a superconducting magnet or the like is used. Moreover, the static magnetic field magnet 101 has a superconducting magnet device 200 (see FIG. 2).

The gradient magnetic field coil 103 is a hollow, substantially cylindrical coil, and is arranged on the inner surface of the cylindrical cooling container. The gradient magnetic field coils 103 individually receive current supply from the gradient magnetic field power supply 105 and generate gradient magnetic fields whose magnetic field strengths vary along the X, Y, and Z axes that are orthogonal to each other. The gradient magnetic fields of the X, Y, and Z axes generated by the gradient magnetic field coil 103 form, for example, a slice selection gradient magnetic field, a phase encoding gradient magnetic field, and a frequency encoding gradient magnetic field. The slice selection gradient magnetic field is used to arbitrarily determine an imaging section. The phase encoding gradient magnetic field is used to change the phase of a magnetic resonance signal (hereinafter referred to as an MR (Magnetic Resonance) signal) depending on the spatial position. The frequency encoding gradient magnetic field is used to change the frequency of the MR signal depending on the spatial location.

The gradient magnetic field power supply 105 is a power supply device that supplies current to the gradient magnetic field coil 103 under the control of the imaging control circuit 121.

The bed 107 is a device that includes a top plate 1071 on which the subject P is placed. The bed 107 inserts the top plate 1071 on which the subject P is placed into the bore 111 under the control of the bed control circuit 109 .

The bed control circuit 109 is a circuit that controls the bed 107. The bed control circuit 109 drives the bed 107 in response to instructions from the operator via the input interface 127, thereby moving the top plate 1071 in the longitudinal direction, up and down directions, and in some cases in the left and right directions.

The transmitting circuit 113 supplies a high frequency pulse modulated at the Larmor frequency to the transmitting coil 115 under the control of the imaging control circuit 121 . For example, the transmission circuit 113 includes an oscillation section, a phase selection section, a frequency conversion section, an amplitude modulation section, an RF (Radio Frequency) amplifier, and the like. The oscillator generates an RF pulse at a resonance frequency specific to a target atomic nucleus in a static magnetic field. The phase selection section selects the phase of the RF pulse generated by the oscillation section. The frequency converter converts the frequency of the RF pulse output from the phase selector. The amplitude modulation section modulates the amplitude of the RF pulse output from the frequency conversion section, for example, according to a sinc function. The RF amplifier amplifies the RF pulse output from the amplitude modulation section and supplies the amplified RF pulse to the transmitting coil 115.

Transmission coil 115 is an RF coil placed inside gradient magnetic field coil 103. The transmitting coil 115 generates an RF pulse corresponding to a high frequency magnetic field in response to the output from the transmitting circuit 113.

The receiving coil 117 is an RF coil placed inside the gradient magnetic field coil 103. The receiving coil 117 receives the MR signal emitted from the subject P using a high frequency magnetic field. Receiving coil 117 outputs the received MR signal to receiving circuit 119. The receiving coil 117 is, for example, a coil array having one or more, typically a plurality of coil elements (hereinafter referred to as a plurality of coils). Hereinafter, in order to make the description more specific, the receiving coil 117 will be described as a coil array having a plurality of coils.

Note that although the transmitting coil 115 and the receiving coil 117 are shown as separate RF coils in FIG. 1, the transmitting coil 115 and the receiving coil 117 may be implemented as an integrated transmitting/receiving coil. The transmitting/receiving coil corresponds to the imaging region of the subject P, and is, for example, a local transmitting/receiving RF coil such as a head coil.

The receiving circuit 119 generates a digital MR signal (hereinafter referred to as MR data) based on the MR signal output from the receiving coil 117 under the control of the imaging control circuit 121 . Specifically, the receiving circuit 119 performs signal processing such as detection and filtering on the MR signal output from the receiving coil 117, and then performs analog/digital (analog/digital) processing on the data subjected to the signal processing. A/D (Analog to Digital) conversion (hereinafter referred to as A/D conversion) is performed to generate MR data. The receiving circuit 119 outputs the generated MR data to the imaging control circuit 121. For example, MR data is generated in each of the plurality of coils and output to the imaging control circuit 121 together with a tag that identifies each of the plurality of coils.

The imaging control circuit 121 controls the gradient magnetic field power supply 105, the transmitting circuit 113, the receiving circuit 119, etc. according to the imaging protocol output from the processing circuit 131, and performs imaging of the subject P. The imaging protocol has a pulse sequence depending on the type of examination. The imaging protocol includes the magnitude of the current supplied to the gradient magnetic field coil 103 by the gradient magnetic field power supply 105, the timing at which the current is supplied to the gradient magnetic field coil 103 by the gradient magnetic field power supply 105, and the timing at which the current is supplied to the gradient magnetic field coil 115 by the transmission circuit 113. The magnitude and time width of the high frequency pulse, the timing at which the high frequency pulse is supplied to the transmitting coil 115 by the transmitting circuit 113, the timing at which the receiving coil 117 receives the MR signal, etc. are defined. When the imaging control circuit 121 receives MR data from the receiving circuit 119 as a result of driving the gradient magnetic field power supply 105, the transmitting circuit 113, the receiving circuit 119, etc. to image the subject P, the imaging control circuit 121 transmits the received MR data to the processing circuit 131. Forward.

Note that the imaging control circuit 121 may collect MR data related to generation of an image showing the sensitivity distribution of the receiving coil 117 used for imaging the subject P using any imaging method. The image showing the sensitivity of the coil is expressed by complex number data. Collection of MR data related to generation of an image showing the sensitivity distribution of the receiving coil 117 is performed by the imaging control circuit 121, for example, in a pre-scan including a locator scan, prior to scanning the subject P. The imaging control circuit 121 is realized by, for example, a processor.

The word "processor" refers to, for example, a CPU, a GPU (Graphics Processing Unit), an application specific integrated circuit (ASIC), a programmable logic device (for example, a simple programmable logic device) Mable Logic Device (SPLD) , a complex programmable logic device (CPLD), and a field programmable gate array (FPGA).

The system control circuit 123 includes a processor (not shown) and memories such as a ROM (Read-Only Memory) and a RAM (Random Access Memory) as hardware resources, and controls the MRI apparatus 100 using a system control function. Specifically, the system control circuit 123 reads a system control program stored in the memory, expands it onto the memory, and controls each circuit of the MRI apparatus 100 according to the expanded system control program.

For example, the system control circuit 123 reads the imaging protocol from the memory 125 based on imaging conditions input by the operator via the input interface 127. The system control circuit 123 transmits the imaging protocol to the imaging control circuit 121 and controls imaging of the subject P. The system control circuit 123 is realized by, for example, a processor. Note that the system control circuit 123 may be incorporated into the processing circuit 131. At this time, the system control function is performed by the processing circuit 131, and the processing circuit 131 functions as a substitute for the system control circuit 123. The processor that realizes the system control circuit 123 has the same contents as described above, so a description thereof will be omitted.

The memory 125 stores various programs related to system control functions executed in the system control circuit 123, various imaging protocols, imaging conditions including a plurality of imaging parameters defining the imaging protocols, and the like. Furthermore, the memory 125 stores various functions realized by the processing circuit 131 in the form of programs executable by a computer.

Note that the memory 125 may store various data received via a communication interface (not shown). For example, the memory 125 stores information regarding an examination order for the subject P (image target region, examination purpose, etc.) received from an information processing system within a medical institution such as a radiology information system (RIS).

The memory 125 is realized by, for example, a semiconductor memory element such as a ROM, a RAM, or a flash memory, an HDD (Hard Disk Drive), an SSD (Solid State Drive), an optical disk, or the like. Further, the memory 125 may be realized by a drive device that reads and writes various information from and to a portable storage medium such as a CD (Compact Disc)-ROM drive, a DVD (Digital Versatile Disc) drive, and a flash memory. .

The input interface 127 accepts various instructions (for example, power-on instructions) and information input from the operator. The input interface 127 includes, for example, a trackball, a switch button, a mouse, a keyboard, a touchpad that performs input operations by touching the operation surface, a touchscreen that integrates a display screen and a touchpad, and a non-control device that uses an optical sensor. This is realized by a touch input circuit, a voice input circuit, etc. The input interface 127 is connected to the processing circuit 131 , converts an input operation received from an operator into an electrical signal, and outputs the electrical signal to the processing circuit 131 . Note that in this specification, the input interface 127 is not limited to one that includes physical operation parts such as a mouse and a keyboard. For example, examples of the input interface 127 include an electrical signal processing circuit that receives an electrical signal corresponding to an input operation from an external input device provided separately from the MRI apparatus 100 and outputs this electrical signal to a control circuit. It will be done.

The input interface 127 inputs the FOV of the prescan image displayed on the display 129 according to a user's instruction. Specifically, the input interface 127 inputs the FOV in the locator image displayed on the display 129 based on a range setting instruction from the user. Further, the input interface 127 inputs various imaging parameters related to scanning according to user instructions based on the examination order.

The display 129 displays various GUIs (Graphical User Interfaces), MR images generated by the processing circuit 131, etc. under the control of the processing circuit 131 or the system control circuit 123. Further, the display 129 displays various information regarding imaging parameters related to scanning, imaging, and image processing. Display 129 is implemented by a display device such as, for example, a CRT display, a liquid crystal display, an organic EL display, an LED display, a plasma display, or any other display or monitor known in the art.

The processing circuit 131 is realized by, for example, the above-mentioned processor. The processing circuit 131 includes various functions. Various functions are stored in the memory 125 in the form of programs executable by a computer. For example, the processing circuit 131 reads programs from the memory 125 and executes them to realize functions corresponding to each program. In other words, the processing circuit 131 that has read each program has various functions.

In the above description, an example has been described in which the "processor" reads out and executes a program corresponding to each function from the memory 125, but the embodiment is not limited to this. When the processor is, for example, a CPU, the processor realizes its functions by reading and executing a program stored in the memory 125. On the other hand, if the processor is an ASIC, instead of storing the program in memory 125, the functionality is built directly into the processor's circuitry as a logic circuit. Note that each processor of this embodiment is not limited to being configured as a single circuit for each processor, but may also be configured as a single processor by combining multiple independent circuits to realize its functions. good. Further, although the description has been made assuming that a single memory circuit stores programs corresponding to each processing function, a plurality of memory circuits are distributed and arranged, and the processing circuit 131 reads the corresponding programs from the individual memory circuits. It may also be a configuration.

Next, the superconducting magnet device 200 will be explained. FIG. 2 is a diagram showing an example of a circuit configuration of a superconducting magnet device 200 according to the first embodiment. The superconducting magnet device 200 is mounted on the static magnetic field magnet 101.

The superconducting magnet device 200 includes a superconducting coil 210, a superconducting coil PCS (Persistent Current Switch) 220, a protection circuit 230, and a compensation coil 240.

Superconducting coil 210 is a coil formed of superconducting wire. The superconducting coil 210 is used to generate a static magnetic field in the imaging region where the subject P is placed. When the superconducting coil 210 is cooled with a coolant such as liquid helium, it transitions to a superconducting state and has approximately zero electrical resistance. As a result, superconducting coil 210 can now flow a large current and generate a stronger magnetic field than a normal electromagnet.

The superconducting coil PCS 220 is a switch for shifting to persistent current mode. In other words, the superconducting coil PCS 220 is a switch that forms a closed circuit including the superconducting coil 210. The superconducting coil PCS 220 is an example of a first switch. For example, the superconducting coil PCS 220 is formed of superconducting wire. Moreover, the electrical connection of the superconducting coil PCS 220 is controlled by a heater.

More specifically, when the superconducting coil PCS 220 is heated by a heater, the electrical resistance increases. As a result, the superconducting coil PCS 220 stops the current. That is, the superconducting coil PCS 220 cuts off electrical connection. On the other hand, when the superconducting coil PCS 220 is not heated by a heater, it is cooled by a refrigerant, so that its electrical resistance decreases. As a result, the superconducting coil PCS 220 allows current to flow. That is, the superconducting coil PCS 220 is electrically connected to form a closed circuit. Note that the superconducting coil PCS 220 is not limited to a thermal type, but may be formed by a mechanical type.

Further, the superconducting coil PCS 220 interrupts electrical connection when exciting the superconducting coil 210. Further, the superconducting coil PCS 220 is connected to an external power source. Then, the superconducting coil PCS 220 supplies an external current to the superconducting coil 210. In this way, superconducting coil 210 is excited by receiving current from the outside.

Further, the superconducting coil PCS 220 generates a prescribed magnetic force by excitation of the superconducting coil 210, and interrupts electrical connection when the superconducting coil 210 is in a superconducting state. Thereby, the superconducting coil PCS 220 forms a closed circuit having the superconducting coil 210. The current then circulates in a closed circuit. A state in which superconducting coil 210 generates a prescribed magnetic force and current circulates in a closed circuit when superconducting coil 210 is in a superconducting state is referred to as persistent current mode.

The protection circuit 230 is a circuit that protects the superconducting coil 210 when a quench occurs. Quenching means that a part of the superconducting coil 210 changes from a superconducting state to a normal conducting state. When quenching occurs, the electrical resistance of the superconducting coil 210 increases, which generates heat due to the Joule effect. In this case, the superconducting wire forming the superconducting coil 210 may be damaged by Joule heat. Therefore, when a quench occurs, the protection circuit 230 prevents the superconducting coil 210 from being damaged by Joule heat by allowing current to flow.

More specifically, the protection circuit 230 is energized when a quench occurs. Further, the protection circuit 230 is connected in parallel with the superconducting coil 210. For example, protection circuit 230 is formed by a plurality of diodes. The protection circuit 230 shown in FIG. 2 includes a first diode 231 and a second diode 232. The first diode 231 and the second diode 232 are connected in parallel so that currents flow in different directions. That is, the first diode 231 and the second diode 232 are connected such that their anodes and cathodes are in opposite directions.

Accordingly, when current is being supplied to excite the superconducting coil 210, no current flows into the protection circuit 230. On the other hand, when a quench occurs, the voltage at both ends of the superconducting coil 210 increases. When the turn-on voltage of protection circuit 230 is exceeded, current flows into protection circuit 230.

The compensation coil 240 is a coil that generates a magnetic field that compensates for a change in the magnetic field caused by the superconducting coil 210 when a quench occurs in the superconducting coil 210. Here, when a quench occurs, current flows into the protection circuit 230, so that the current flowing into the superconducting coil 210 is attenuated. In the superconducting coil 210, the magnetic field generated by the superconducting coil 210 changes rapidly as the current passing therethrough rapidly attenuates. The compensation coil 240 suppresses changes in the magnetic field by generating a magnetic field that compensates for changes in the magnetic field caused by the superconducting coil 210.

The compensation coil 240 is a magnetic field generated by a current flowing when a quench occurs in the superconducting coil 210, and compensates for changes in the magnetic field by generating a back electromotive force in the superconducting coil 210. More specifically, compensation coil 240 is connected in series with protection circuit 230. Here, the protection circuit 230 is energized when quenching occurs. The compensation coil 240 generates a magnetic field that causes the superconducting coil 210 to generate a back electromotive force using the current supplied via the protection circuit 230 when a quench occurs.

For example, compensation coil 240 is formed of superconducting wire. Note that the compensation coil 240 is not limited to a superconducting wire, and may be formed of a highly conductive wire such as copper or aluminum.

Further, the compensation coil 240 is arranged between the superconducting coil 210 and the thermal radiation shield 260 and on the side of the magnetic field center M of the superconducting coil 210. Thermal radiation shield 260 shields the superconducting coil 210 from thermal radiation. The thermal radiation shield 260 is made of, for example, high-purity aluminum, copper, or the like. The thermal radiation shield 260 shields thermal radiation from the superconducting coil 210 by covering the superconducting coil 210.

Furthermore, the compensation coil 240 has a high coupling coefficient, such as 0.8, with respect to the superconducting coil 210. That is, the compensation coil 240 is formed such that most of the magnetic flux generated in the compensation coil 240 passes through the superconducting coil 210.

Next, the arrangement of superconducting coil 210 and compensation coil 240 will be explained. FIG. 3 is a diagram showing an example of the arrangement of the superconducting coil 210 and the compensation coil 240 in the superconducting magnet device 200 according to the first embodiment.

The superconducting coil 210 is formed by winding a superconducting wire around a winding frame 250. The winding frame 250 is formed into a substantially cylindrical shape around the bore 111 into which the subject P is inserted. By winding the superconducting wire around the winding frame 250, the superconducting coil 210 is formed into a substantially cylindrical shape around the bore 111 into which the subject P is inserted.

Furthermore, the compensation coil 240 is formed by winding a superconducting wire or a highly conductive wire around the magnetic field center M of the superconducting coil 210. That is, the compensation coil 240 is arranged so as to be in contact with the magnetic field center M side of the superconducting coil 210 in the winding frame 250. Such a structure allows the winding frame 250 to receive the hoop stress that occurs when the compensation coil 240 is energized.

Next, the circuit operation of the superconducting magnet device 200 will be explained.

The superconducting coil PCS 220 interrupts electrical connection when the superconducting coil 210 is excited. Superconducting coil 210 receives current sharing from an external power source. The superconducting coil PCS 220 forms a closed circuit including the superconducting coil 210 by generating a prescribed magnetic force by excitation of the superconducting coil 210, and by cutting off electrical connection when the superconducting coil 210 is in a superconducting state. do. This causes the superconducting coil PCS 220 to transition to persistent current mode.

When quenching occurs, the voltage across both ends of the superconducting coil 210 increases as the electrical resistance increases. When the turn-on voltage of protection circuit 230 is exceeded, current flows into protection circuit 230. Further, as the current flows into the protection circuit 230, the compensation coil 240 is energized.

Compensation coil 240 generates a magnetic field when energized. The superconducting coil 210 generates a counter electromotive force by electromagnetic induction using the magnetic field generated by the compensation coil 240. A current flows through the superconducting coil 210 due to the generated back electromotive force. The superconducting coil 210 suppresses rapid attenuation of the static magnetic field by generating a magnetic field using current. In other words, superconducting coil 210 suppresses rapid attenuation of the main magnetic field.

As described above, the superconducting magnet device 200 according to the first embodiment includes the superconducting coil 210, the thermal radiation shield 260, and the superconducting coil PCS 220 for shifting to persistent current mode. Further, the superconducting magnet device 200 includes a compensation coil 240 that generates a magnetic field that compensates for a change in the magnetic field due to the superconducting coil 210 when a quench occurs in the superconducting coil 210. Therefore, even if a quench occurs, the superconducting magnet device 200 generates a magnetic field that compensates for the change in the main magnetic field caused by the superconducting coil 210, and therefore can suppress rapid changes in the main magnetic field.

In particular, the MRI apparatus 100 generates gradient magnetic fields in order to acquire MR images. In this case, in order to reduce the influence on the image quality of the MR image, it is desirable to use a material with high purity and high conductivity for the thermal radiation shield 260. That is, it is desirable to use a relatively soft material such as high-purity aluminum or copper for the thermal radiation shield 260. The reason why we want to use a material with high conductivity is that by generating an eddy magnetic field with a somewhat long time constant, we can reduce the influence on imaging when applying gradient magnetic field pulses for imaging. However, the designer had to design with the assumption that Lorentz force would occur, making the design difficult. The superconducting magnet device 200 can be easily designed because the compensation coil 240 suppresses the thermal radiation shield 260 from being exposed to the Lorentz force.

(Second embodiment)
FIG. 4 is a diagram showing an example of a circuit configuration of a superconducting magnet device 200a according to the second embodiment. The superconducting magnet device 200a energizes the compensation coil 240 when a quench occurs.

Here, the compensation coil 240 is a coil that compensates for changes in the static magnetic field when quenching occurs. Therefore, the compensation coil 240 only needs to be energized when quenching occurs, and does not need to be energized when excited.

The excitation power source 270 is, for example, a constant current power source that supplies a current when the superconducting coil 210 is excited, and when exciting, the current flows counterclockwise. Further, the compensation coil 240 is connected in series with the first diode 231. Further, the compensation coil 240 is connected in parallel with the superconducting coil 210.

As shown in FIG. 4, the first diode 231 is connected in a direction opposite to the direction of the current that the excitation power source 270 flows during excitation. That is, the cathode side of the first diode 231 is connected with respect to the direction of current during excitation.

Compensation coil 240 is arranged in protection circuit 230 on a path through which current is applied when quench occurs. More specifically, the compensation coil 240 is connected in series with a first diode 231 connected in a direction opposite to the current direction when the superconducting coil 210 is excited. Therefore, the compensation coil 240 is not energized during excitation.

On the other hand, quenching occurs, and as the electrical resistance of the superconducting coil 210 increases, the current rapidly attenuates, and the static magnetic field generated by the superconducting coil 210 also rapidly decreases. The compensating coil 240 generates a back electromotive force at both ends of the superconducting coil 210 to suppress current attenuation. The first diode 231 energizes the compensation coil 240 when the turn-on voltage as a switch is exceeded. Therefore, the compensation coil 240 is not energized when it is excited, but is energized when quenching occurs.

As described above, in the superconducting magnet device 200a according to the second embodiment, the compensation coil 240 is connected in series to the first diode 231, which is connected in the opposite direction to the current direction when the superconducting coil 210 is excited. be done. Then, the compensation coil 240 is energized when a quench occurs. In other words, the compensation coil 240 is not energized during excitation. Therefore, the superconducting magnet device 200a can reduce power consumption.

(Third embodiment)
FIG. 5 is a diagram showing an example of a circuit configuration of a superconducting magnet device 200b according to the third embodiment.

Whether or not the compensation coil 240 is energized is switched by the quench PCS 280. Compensation coil 240 is connected in parallel to superconducting coil 210 and protection circuit 230. Further, the compensation coil 240 is connected in series with the quenching PCS 280.

The quench PCS 280 is a quench switch made of superconducting wire. The electrical connection of the quench PCS 280 is controlled by a heater. The quench PCS 280 switches whether or not to energize the compensation coil 240 in synchronization with the superconducting coil PCS 220. The quench PCS 280 is an example of a second switch. That is, the quench PCS 280 interrupts electrical connection when the superconducting coil PCS 220 does not form a closed circuit. The quench PCS 280 is electrically connected to the superconducting coil PCS 220 when the superconducting coil PCS 220 forms a closed circuit. Thereby, the compensation coil 240 is energized when a quench occurs.

Note that it is desired that the current flowing through the compensation coil 240 be zero ampere. Therefore, it is preferable that the quenching PCS 280 has an interval longer than the energization switching timing of the superconducting coil PCS 220 by about the time constant of the compensation coil 240. Furthermore, the quenching PCS 280 is not limited to one that completely cuts off current, but may be one that restricts energization to the compensation coil 240.

As described above, the superconducting magnet device 200b according to the third embodiment controls whether or not to energize the compensation coil 240 using the quench PCS 280. Even in this case, the superconducting magnet device 200b generates a magnetic field that compensates for the change in the main magnetic field caused by the superconducting coil 210 even if a quench occurs, so that a sudden change in the main magnetic field can be suppressed.

(Fourth embodiment)
FIG. 6 is a diagram showing an example of a circuit configuration of a superconducting magnet device 200c according to the fourth embodiment.

Whether or not the compensation coil 240 is energized is switched by a quench heater 290 that heats the compensation coil 240 . When the compensation coil 240 is heated by the quench heater 290, its electrical resistance increases. On the other hand, when the compensation coil 240 is not heated by the quench heater 290, the compensation coil 240 is cooled by the refrigerant, so that its electrical resistance decreases.

The quench heater 290 heats the compensation coil 240 in synchronization with the superconducting coil PCS 220. The quench heater 290 is an example of a heater. More specifically, the quench heater 290 heats the compensation coil 240 when the superconducting coil PCS 220 does not form a closed circuit. That is, the quench heater 290 cuts off the electrical connection of the superconducting coil 210. On the other hand, the quench heater 290 does not heat the compensation coil 240 when the superconducting coil PCS 220 forms a closed circuit. That is, the quench heater 290 electrically connects the superconducting coil 210. Thereby, the compensation coil 240 is energized when a quench occurs.

Note that it is desired that the current flowing through the compensation coil 240 be zero ampere. Therefore, it is preferable that the quench heater 290 has an interval that is longer than the energization switching timing of the superconducting coil PCS 220 by about the time constant of the compensation coil 240. Furthermore, the quench heater 290 is not limited to one that completely cuts off the current, but may be one that limits the energization to the compensation coil 240.

As described above, the superconducting magnet device 200c according to the fourth embodiment controls whether or not the compensation coil 240 is energized by the quench heater 290. Even in this case, the superconducting magnet device 200c generates a magnetic field that compensates for the change in the main magnetic field caused by the superconducting coil 210 even if a quench occurs, so that a sudden change in the main magnetic field can be suppressed.

(Modification 1)
The superconducting magnet device 200d according to Modification 1 has been described as having one combination of the superconducting coil 210 and the compensation coil 240, as shown in FIG. However, the superconducting magnet device 200d may have a plurality of combinations of the superconducting coil 210 and the compensation coil 240.

FIG. 7 is a diagram illustrating an example of the arrangement of superconducting coil 210 and compensation coil 240 in superconducting magnet device 200d according to Modification 1. For example, as shown in FIG. 7, the number of combinations of superconducting coils 210 and compensation coils 240 may be three. Furthermore, the number of combinations of superconducting coils 210 and compensation coils 240 may be two, or four or more.

(Modification 2)
It has been explained that the superconducting magnet devices 200, 200a, 200b, 200c, and 200d are installed in the MRI apparatus 100. However, the superconducting magnet devices 200, 200a, 200b, 200c, and 200d may be installed not only in the MRI apparatus 100 but also in other devices.
For example, the superconducting magnet devices 200, 200a, 200b, 200c, and 200d may be mounted on a semiconductor pulling device.

According to at least one embodiment described above, rapid changes in the main magnetic field can be suppressed.

Although several embodiments have been described, these embodiments are presented by way of example and are not intended to limit the scope of the invention. These embodiments can be implemented in various other forms, and various omissions, substitutions, changes, and combinations of embodiments can be made without departing from the gist of the invention. These embodiments and their modifications are included within the scope and gist of the invention as well as within the scope of the invention described in the claims and its equivalents.

100 MRI device 101 Static magnetic field magnet 200, 200a, 200b, 200c, 200d Superconducting magnet device 210 Superconducting coil 220 PCS (Persistent Current Switch) for superconducting coil
230 Protection circuit 231 First diode 232 Second diode 240 Compensation coil 250 Winding frame 260 Heat radiation shield 270 Excitation power supply 280 Quench PCS
290 Quench heater P Object M Magnetic field center

Claims (9)

A superconducting coil formed of superconducting wire,
a thermal radiation shield that blocks thermal radiation to the superconducting coil;
a first switch forming a closed circuit including the superconducting coil;
a compensation coil that generates a magnetic field that compensates for changes in the magnetic field due to the superconducting coil when quenching occurs in the superconducting coil;
A superconducting magnet device equipped with. The compensation coil compensates for a change in the magnetic field by generating a back electromotive force in the superconducting coil using a magnetic field generated by a current flowing when a quench occurs in the superconducting coil.
The superconducting magnet device according to claim 1. Further comprising a protection circuit that energizes when quenching occurs in the superconducting coil,
the compensation coil is connected in series with the protection circuit;
The superconducting magnet device according to claim 1. The protection circuit is formed by a plurality of diodes,
The compensation coil is connected in series with the diode connected in a direction opposite to the current direction when the superconducting coil is excited.
The superconducting magnet device according to claim 3. The compensation coil is disposed between the superconducting coil and the thermal radiation shield and on the side of the magnetic field center of the superconducting coil.
The superconducting magnet device according to claim 1. The superconducting coil is formed by winding the superconducting wire around a winding frame,
The compensation coil is arranged so as to be in contact with the magnetic field center side of the superconducting coil in the winding frame.
The superconducting magnet device according to claim 5. further comprising a second switch that switches whether or not to energize the compensation coil in synchronization with the first switch,
the compensation coil is connected in series with the second switch;
The superconducting magnet device according to claim 1. further comprising a heater that heats the compensation coil in synchronization with the first switch,
the heater does not overheat when the first switch forms the closed circuit;
The superconducting magnet device according to claim 1. A magnetic resonance imaging device having a superconducting magnet device,
The superconducting magnet device includes:
A superconducting coil formed of superconducting wire,
a thermal radiation shield that suppresses thermal radiation to the superconducting coil;
a first switch forming a closed circuit including the superconducting coil;
a compensation coil that suppresses a decrease in magnetic force caused by the superconducting coil when quenching occurs in the superconducting coil;
A magnetic resonance imaging device comprising:

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