WO2017090620A1 - Magnetic resonance imaging apparatus - Google Patents

Magnetic resonance imaging apparatus Download PDF

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Publication number
WO2017090620A1
WO2017090620A1 PCT/JP2016/084625 JP2016084625W WO2017090620A1 WO 2017090620 A1 WO2017090620 A1 WO 2017090620A1 JP 2016084625 W JP2016084625 W JP 2016084625W WO 2017090620 A1 WO2017090620 A1 WO 2017090620A1
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Prior art keywords
magnetic field
coil
current
field strength
change
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PCT/JP2016/084625
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French (fr)
Japanese (ja)
Inventor
慎一 浦山
秀直 福山
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国立大学法人京都大学
株式会社Ｋｙｏｔｏ Ｆｕｔｕｒｅ Ｍｅｄｉｃａｌ Ｉｎｓｔｒｕｍｅｎｔｓ
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Publication of WO2017090620A1 publication Critical patent/WO2017090620A1/en

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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/05—Detecting, measuring or recording for diagnosis by means of electric currents or magnetic fields; Measuring using microwaves or radio waves 
    • A61B5/055—Detecting, measuring or recording for diagnosis by means of electric currents or magnetic fields; Measuring using microwaves or radio waves  involving electronic [EMR] or nuclear [NMR] magnetic resonance, e.g. magnetic resonance imaging
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/381—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using electromagnets
    • G01R33/3815—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using electromagnets with superconducting coils, e.g. power supply therefor
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/387—Compensation of inhomogeneities
    • G01R33/3875—Compensation of inhomogeneities using correction coil assemblies, e.g. active shimming

Definitions

  • the present invention relates to a magnetic resonance imaging apparatus, and more particularly to a magnetic resonance imaging apparatus capable of setting a plurality of magnetic field strengths.
  • MRI nuclear magnetic resonance imaging
  • a superconducting magnet that generates a strong magnetic field may be used in order to improve image accuracy.
  • a low-temperature superconducting material may be used as a superconducting coil in a magnet.
  • liquid helium is used as the refrigerant to maintain the superconducting coil at the superconducting temperature.
  • liquid helium is generally difficult to obtain, and particularly in recent years, it has become more difficult to obtain due to rising prices. Further, when liquid helium is used as a refrigerant, the entire apparatus tends to be large and the equipment cost tends to increase due to storage and cooling equipment.
  • a low-temperature superconducting material that requires liquid helium as a refrigerant such as Japanese Unexamined Patent Application Publication No. 2009-183372 (Patent Document 1) and Japanese Unexamined Patent Application Publication No. 2015-167576 (Patent Document 2).
  • Patent Document 1 Japanese Unexamined Patent Application Publication No. 2009-183372
  • Patent Document 2 Japanese Unexamined Patent Application Publication No. 2015-167576
  • Patent Document 1 Japanese Unexamined Patent Application Publication No. 2009-183372
  • Patent Document 2 Japanese Unexamined Patent Application Publication No. 2015-167576
  • JP 2009-183372 A Japanese Patent Laying-Open No. 2015-167576
  • MRI apparatuses have different static magnetic field strengths depending on their applications. For example, MRI apparatuses having a static magnetic field of 1.5T, 3T, 7T, etc. are known. In order to deal with all uses, it is necessary to provide individual MRI apparatuses according to each use, so that a large installation space is required and the cost of introducing the equipment becomes high. Therefore, it is desired to develop an MRI apparatus that can generate a plurality of magnetic field strengths with a single facility.
  • changing the magnetic field strength can be realized by changing the current flowing through the superconducting coil in the magnet.
  • the current switch is operated to supply current to the superconducting coil from the outside, or to extract current from the superconducting coil. Is required.
  • heat is generated due to increase / decrease in current due to the operation of the current switch, liquid helium is consumed (evaporated). Therefore, the cost increases due to the replenishment of liquid helium.
  • devices using low-temperature superconducting materials are often designed to operate just below the liquid helium temperature when designing each component in the device. There is a possibility of connection. For this reason, it is difficult to change the magnetic field in the apparatus using the low temperature superconducting material by a person other than an engineer having expertise.
  • the inventor of the present application has focused on the possibility that the magnetic field strength generated in the superconducting coil can be changed by adjusting the power supply device in an MRI apparatus using a high-temperature superconducting material.
  • the present invention has been made to solve the above-described problems, and an object thereof is to provide a magnetic resonance imaging (MRI) apparatus that can be used by switching a plurality of magnetic field strengths.
  • MRI magnetic resonance imaging
  • the magnetic resonance imaging apparatus can set a plurality of magnetic field strengths, and includes a superconducting coil, a DC power supply device, and a shim coil.
  • the superconducting coil includes a high-temperature superconducting material, and generates a static magnetic field to which the subject is exposed.
  • the DC power supply device supplies an output current to the superconducting coil.
  • the DC power supply device is configured to be able to set a plurality of magnetic field strengths by adjusting the magnitude of the output current supplied to the superconducting coil.
  • the shim coil is configured to be able to correct the magnetic field distribution generated by the superconducting coil. The magnetic field distribution is corrected by supplying a current corresponding to each of the plurality of magnetic field strengths to the shim coil.
  • the magnetic resonance imaging apparatus further includes a transmission coil for generating an excitation pulse of a predetermined frequency and a reception coil for receiving a magnetic resonance signal from the subject.
  • the plurality of magnetic field strengths include a first magnetic field strength and a second magnetic field strength.
  • the transmit coil is configured to resonate at both a first frequency corresponding to the first magnetic field strength and a second frequency corresponding to the second magnetic field strength.
  • the first and second magnetic field strengths are 1.5T and 3T, respectively.
  • the first frequency is 64 MHz and the second frequency is 128 MHz.
  • the current supplied to the shim coil is determined based on a principal component analysis of a change in magnetic field strength at each of a plurality of magnetic field strengths.
  • the magnetic resonance imaging apparatus further includes a current measurement device for measuring a drive current supplied to the superconducting coil.
  • the direct current power source adjusts the output current based on the fluctuation of the drive current measured by the current measuring device.
  • an MRI apparatus that can be used by switching a plurality of magnetic field strengths.
  • FIG. 1 is a functional block diagram for explaining an outline of a magnetic resonance imaging (MRI) apparatus 10 according to the present embodiment.
  • MRI magnetic resonance imaging
  • the MRI apparatus 10 includes a main body device 100, a DC power supply device 200, a control device 300, a display unit 310, and an input unit 320.
  • Main device 100 includes a superconducting magnet 110, a gradient coil 120, an RF transmitter coil 130, an RF receiver coil 140, and a shim coil 150.
  • Superconducting magnet 110 includes a high-temperature superconducting coil 115 formed of a high-temperature superconducting material (for example, a bismuth-based superconducting material).
  • the high-temperature superconducting coil 115, the gradient magnetic field coil 120, the RF transmitting coil 130, and the shim coil 150 of the main body device 100 have a generally cylindrical shape.
  • a magnetic field is generated in the cavity (tunnel).
  • the subject 170 is inserted into the tunnel of the cylindrical main body device 100 while lying on the examination table 160.
  • the RF receiving coil 140 is installed so as to cover the examination target site in the subject 170.
  • the RF receiving coil 140 is shown as an example in which the head is inspected.
  • the RF receiving coil 140 covers the trunk. Installed.
  • the high temperature superconducting coil 115 is a coil for generating a spatially and temporally uniform static magnetic field in the tunnel.
  • the high temperature superconducting coil 115 is configured to be able to generate a magnetic field (for example, 1.5T, 3T, 7T) having a strength corresponding to the magnitude of the current supplied from the DC power supply device 200. Due to the static magnetic field generated by the high-temperature superconducting coil 115, the nuclear spins of the hydrogen nuclei at the site to be examined of the subject 170 can be aligned in a certain direction.
  • the shim coil 150 is a coil for generating a magnetic field for correcting the static magnetic field generated by the high-temperature superconducting coil 115.
  • Shim coil 150 is formed of, for example, a normal conductive material.
  • the magnetic field uniformity of the static magnetic field in the MRI apparatus 10 needs to be suppressed to 10 ppm or less in the imaging region as a standard.
  • the static magnetic field generated by the high-temperature superconducting coil 115 is distorted due to the influence of each device included in the MRI apparatus 10 and structures such as reinforcing bars of a building where the apparatus is installed.
  • the shim coil 150 may be configured by a plurality of coils that generate different magnetic field distributions in order to adjust different generated magnetic field intensity.
  • the gradient magnetic field coil 120 (or gradient magnetic field coil) is a coil for forming a gradient magnetic field that linearly changes spatially. This gradient magnetic field can spatially and linearly change the frequency of the signal emitted by the hydrogen nuclei at the site to be examined. Therefore, position information can be added to the received signal received by the RF receiving coil 140.
  • the RF transmission coil 130 is a coil for transmitting an RF pulse signal having a predetermined frequency to the subject 170.
  • the RF pulse signal When the RF pulse signal is irradiated to the inspection target part, the hydrogen nuclei of the inspection target part are excited by the energy given by the RF pulse. When the RF pulse signal is stopped, the hydrogen nuclei return from the excited state. A signal that is observed when returning from this excited state is received by the RF receiving coil 140.
  • the received signals from the respective positions received by the RF receiving coil 140 have different phases. Therefore, by appropriately adjusting the gradient magnetic field to be applied and the frequency of the RF pulse signal, the position of the hydrogen nucleus that has emitted the signal can be specified from the obtained reception signal. By arranging the received signals in a two-dimensional or three-dimensional manner, the site to be examined can be imaged.
  • DC power supply device 200 supplies a current for generating a static magnetic field by high-temperature superconducting coil 115.
  • a direct current supplied from the direct current power supply device 200 to the high temperature superconducting coil 115 is measured by a current sensor 210.
  • the current output from the DC power supply device 200 can be adjusted by a signal from the control device 300.
  • Control device 300 receives the current value measured by current sensor 210 and performs feedback control so that the output current from DC power supply device 200 is constant. Thereby, a stable static magnetic field can be generated.
  • the control device 300 receives information from the user input from the input unit 320.
  • the information from the user includes, for example, information on the magnetic field strength to be used, information on the subject 170, and the like.
  • the control device 300 controls the current to be supplied from the DC power supply device 200 to the high temperature superconducting coil 115 based on the magnetic field strength information.
  • the control device 300 adjusts the excitation current for the gradient magnetic field coil 120 and the shim coil 150. In addition, the control device 300 outputs an RF pulse signal to the RF transmission coil 130 and receives a reception signal received by the RF reception coil 140 with respect to the RF pulse signal. Based on the received signal, the control device 300 images the cross section of the examination target region and displays it on the display unit 310.
  • the refrigerant does not need to be liquid helium, and a cheaper and easily available refrigerant such as liquid nitrogen can be used.
  • FIG. 2 is a diagram for explaining the characteristics of each magnetic field strength of the MRI apparatus currently used in the medical field.
  • As the magnetic field strength a case of 1.5T, 3T, and 7T, which is often used in the medical field, will be described as an example.
  • S / N ratio suitability for each inspection object is shown, A means “very suitable”, B is “suitable”, C is “slightly unsuitable”, D means “unsuitable”.
  • the S / N ratio is improved as the magnetic field strength increases. That is, as the magnetic field strength is increased, the obtained image quality (resolution, contrast) is improved. Therefore, an MRI apparatus having a high magnetic field strength is suitable for imaging fine parts such as the head (in the brain) and limbs, or imaging metabolic information.
  • the inhomogeneous magnetic field in the living body increases in proportion to the magnetic field strength, which is disadvantageous for imaging of the trunk including abdominal organs and the like that have many oxygen molecules that cause the inhomogeneous magnetic field. It is easy to become. For such a part, an MRI apparatus having a relatively small magnetic field strength is more advantageous.
  • the MRI apparatus since the MRI apparatus generates a strong magnetic field, it may not be used when there is a metal such as a pacemaker or implant in the body. This is particularly problematic when the strength of the generated magnetic field is increased. For this reason, it is preferable to use an MRI with a small magnetic field strength in the case of emergency response in which the presence or absence of metal in the body cannot be known in advance.
  • changing the magnetic field strength can be realized by changing the current flowing through the superconducting coil in the magnet.
  • a switch called a permanent current switch is operated to supply current to the superconducting coil from the outside. It is necessary to extract current from the superconducting coil.
  • liquid helium is consumed (evaporated). For this reason, if the cost increases due to replenishment of liquid helium, or if the temperature rises rapidly, so-called “quenching” in which evaporation of liquid helium occurs explosively may occur.
  • devices using low-temperature superconducting materials are often designed to operate just below the liquid helium temperature when designing each component in the device. There is a possibility of connection. Therefore, the magnetic field change in the apparatus using the low-temperature superconducting material involves a risk, and it is difficult to perform it by a person other than an engineer having specialized skills.
  • the MRI apparatus 10 when the high-temperature superconducting coil 115 is used as in the MRI apparatus 10 according to the present embodiment shown in FIG. 1, it is not necessary to use liquid helium. However, the superconducting material cannot be brought into a superconducting state, for example, at the connection portion of the superconducting material. As a result, a minute current loss may occur. Therefore, the MRI apparatus 10 according to the present embodiment has a configuration in which a uniform static magnetic field is stably generated by providing the DC power supply device 200 and supplying power to the high-temperature superconducting coil 115. .
  • the current supplied from the DC power supply apparatus 200 to the high-temperature superconducting coil 115 is changed by the setting from the input unit 320, thereby changing the magnetic field intensity that varies depending on one MRI apparatus 10. Can be generated.
  • the RF transmission coil 130 is configured to resonate at a plurality of frequencies corresponding to the magnetic field strength used.
  • the RF transmission coil 130 has a frequency corresponding to a magnetic field strength of 1.5T and a frequency corresponding to a magnetic field strength of 3T. It is configured to resonate at 128 MHz. In this way, by designing the RF transmission coil 130 so as to resonate at a plurality of frequencies, it is possible to suppress replacement of the RF transmission coil in response to a change in magnetic field strength.
  • a shim coil 150 adapted to each magnetic field strength is provided.
  • the shim coil 150 includes a plurality of shim coils, and is configured to correct a static magnetic field by a combination of the plurality of shim coils. Therefore, some shim coils may be used in common at different magnetic field strengths.
  • the high-temperature superconducting coil 115 is used in the superconducting magnet 110 so that liquid helium is not used, and a DC power supply apparatus 200 for maintaining a static magnetic field.
  • a DC power supply apparatus 200 for maintaining a static magnetic field.
  • the RF transmission coil so as to resonate at respective frequencies suitable for a plurality of magnetic field strengths to be used.
  • the magnetic field strength can be changed by using a superconducting coil made of high-temperature superconducting material, and changing the current of the DC power supply, so that the magnetic field strength can be changed compared to the case of a general superconducting magnet. Can be reduced, and the magnetic field strength can be easily changed. Therefore, in each medical institution, for example, the magnetic field strength can be frequently changed, for example, the magnetic field strength is changed every day of the week.
  • a single MRI apparatus can generate a plurality of magnetic field strengths. However, if the magnetic field strength differs, the degree and distribution of the inhomogeneous magnetic field changes. It is necessary to set the shim coil.
  • the magnetic field intensity is variable as in the present embodiment, it cannot be sufficiently corrected with a fixed correction magnetic field such as an iron shim, and the correction magnetic field is also variable using a shim coil. It will be necessary. In this case, due to the large magnetic field correction, the resistance value of the shim coil itself increases and the amount of current flowing through the shim coil increases, resulting in an increase in the amount of heat generated by the shim coil.
  • the generated heat is cooled by the refrigerant of the cooling device, but it is preferable to concentrate the heat generation portion locally from the viewpoint of the installation of the cooling pipe and the cooling efficiency.
  • the shim coil to be used is a “high heat generation shim coil” for generating a strong correction magnetic field and a “low heat generation shim coil” for generating a weak correction magnetic field for fine adjustment. By dividing, the cooling efficiency of the shim coil is improved.
  • the design may be performed according to the flowchart of FIG. 3 in a state where the installation of the MRI apparatus at the desired installation location is completed. is there.
  • the magnetic field strengths at a plurality of measurement locations in the examination region are measured a plurality of times. Measurement is performed (step S100).
  • the measured magnetic field strength at this time is represented as B b, n .
  • n represents the number of measurements (1 ⁇ n ⁇ N)
  • N represents the total number of measurements.
  • B is an M-dimensional vector
  • M represents the total number of measurement points.
  • step S110 the measured B b, n and the magnetic field strength change ⁇ B b, n from the average magnetic field are calculated as in the following equation (1) (step S110).
  • ⁇ B b, n B b, n ⁇ ( ⁇ b B b, n ) / N (1) Then, the obtained magnetic field strength change ⁇ B b, n is considered as a set of N points on the M-dimensional space, and principal component analysis is performed (step S120).
  • step S130 it is determined whether or not the cumulative contribution ratio of the Kth principal component obtained by the principal component analysis is equal to or less than a predetermined threshold value T1 (step S130).
  • T1 a predetermined threshold value
  • the shim coil is designed as a high heat generation shim coil (step S140).
  • the cumulative contribution ratio of the K-th principal component is greater than threshold value T1 (NO in step S130)
  • the cumulative contribution ratio of the K-th principal component is next equal to or less than predetermined threshold value T2 (T1 ⁇ T2). ) Is determined (step S150).
  • step S150 If the cumulative contribution rate is equal to or less than threshold value T2 (YES in step S150), a relatively low current may be passed through the shim coil in order to form the magnetic field distribution represented by the main component. Designed as a system shim coil. On the other hand, if the cumulative contribution rate is greater than threshold value T2 (NO in step S150), it can be determined that there is no need to correct the magnetic field distribution represented by the principal component.
  • a shim coil for forming a magnetic field distribution represented by each principal component is a high heat generation shim coil or a low heat generation.
  • Design and classify into system shim coils As a coil design method based on the obtained main component, a known method can be used.
  • the shim coils By designing the shim coils according to such a procedure, the number of shim coils required can be reduced and the heat generation points can be concentrated locally. Therefore, the cooling system for cooling the shim coil can be simplified and the cooling efficiency can be improved.
  • the inhomogeneity and instability of the magnetic field generated by the superconducting magnet for forming the variable magnetic field is caused by the change in the magnetic field strength ⁇ B f resulting from the change in the static magnetic field strength, Space-time of four magnetic field strength changes: magnetic field strength change ⁇ B c due to instability of magnetic field, magnetic field strength change ⁇ B r due to long-period ripple noise immediately after magnetic field change, and magnetic field strength change ⁇ B e due to shielding current It can be modeled by adding the functions.
  • the change in magnetic field strength ⁇ B f resulting from the change in the static magnetic field strength is a change in the magnetic field strength caused by the internal structure of the magnet and the surrounding environment such as a magnet manufacturing error and deformation due to electromagnetic force when the magnetic field strength is changed. . For this reason, no time fluctuation occurs while the magnetic field strength is fixed.
  • the change in magnetic field strength ⁇ B c caused by the instability of the drive current is a change in magnetic field strength caused by the fluctuation of the drive current due to the instability of the DC power supply device that supplies the drive current of the magnet.
  • spatial pattern of the magnetic field strength change .DELTA.B c is constant in principle, for the size and orientation (sign) is correlated to the variation of the drive current.
  • Field strength change .DELTA.B r due to the long cycle ripple noise begins immediately after the completion of the magnetic field change process, a magnetic field intensity change monotonously attenuated while oscillating in a specific period determined from the characteristics of the magnet.
  • This magnetic field strength change ⁇ B r is a function of time and space, and varies according to temporal and spatial positions.
  • the magnetic field intensity change ⁇ B e caused by the shielding current is a magnetic field intensity change that occurs in a direction to suppress the generated static magnetic field, and has a characteristic that decreases monotonously at a very slow rate immediately after the start of the magnetic field change process. Yes.
  • This magnetic field strength change ⁇ B e is also a function of time and space, and fluctuates according to temporal and spatial positions.
  • the magnetic field intensity change ⁇ B r and the magnetic field intensity change ⁇ B e are reduced to some extent by adjusting the impedance of the DC power supply device to be adapted to the magnet or by appropriately adjusting the current when changing the magnetic field. Is possible.
  • the magnetic field strength change ⁇ B f , the magnetic field strength change ⁇ B r, and the magnetic field strength change ⁇ B e are relatively highly reproducible, and therefore can be estimated by analyzing data of temporal and spatial changes of the magnetic field measured in advance. It is.
  • the magnetic field strength changes .DELTA.B r it can be estimated by identifying the oscillation frequency of the ripple. Furthermore, the magnetic field strength change ⁇ B e can be estimated by performing the magnetic field measurement for a long time because the decay rate is sufficiently slow and the change is large compared to the magnetic field strength change ⁇ B r .
  • each of the current values I f , I c , I r (t), and I e (t) is a vector including the same number of elements as the shim coil.
  • the current values I f , I r (t), I It is possible to calculate e (t) in advance. Since the current value I c can be calculated from the drive current measured by the current sensor, the current values I f , I r (t), I e (t) obtained in advance and the current calculated from the drive current by summing the values I c, can be determined current required for each shim coil.
  • FIG. 4 is a flowchart showing the above procedure.
  • steps S200 to S240 are previously processed off-line in advance at the time of adjustment at the time of equipment introduction, and the obtained current values I f , I c , I r (t), I e (T) is stored in a storage unit included in the control device.
  • Step S250 is processed online based on the stored current values I f , I r (t), I e (t) and the current value I c determined from the drive current measured by the current sensor. .
  • step S200 the temporal change B b (t) of the magnetic field strength and the temporal change I c (t) of the drive current when the magnetic field strength b is changed are measured, and the measurement data is stored in the storage unit in the control device.
  • the stored magnetic field strength temporal change B b (t) data is Fourier-transformed to estimate the magnetic field strength change ⁇ B r caused by ripple noise (step S210).
  • step S220 the data obtained by separating the magnetic field intensity change ⁇ B r (t) estimated in step S210 from the stored data of the temporal change B b (t) of the magnetic field intensity is subjected to smoothing processing, and the magnetic field after convergence is obtained.
  • the distribution is estimated as the magnetic field strength change ⁇ B f and the fluctuation until convergence is estimated as the magnetic field strength change ⁇ B e (t).
  • step S240 each of the magnetic field strength changes ⁇ B r (t), ⁇ B f , ⁇ B e (t), and ⁇ B c is decomposed with the main components at the time of designing the shim coil, whereby current values I r (t), I f 1 , I e (t), and I c are calculated, and the obtained current value is stored in the storage unit.
  • the steps so far are performed offline.
  • the stored current values I r (t), I f , I e (t) and the drive current actually supplied to the superconducting magnet measured by the current sensor are summed.
  • 10 MRI apparatus 100 main unit, 110 superconducting magnet, 115 high temperature superconducting coil, 120 gradient magnetic field coil, 130 RF transmitting coil, 140 RF receiving coil, 150 shim coil, 160 testing table, 170 subject, 200 DC power supply, 210 current sensor, 300 control device, 310 display unit, 320 input unit.

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Abstract

An MRI apparatus (10) includes: a superconducting coil (115) that includes a high-temperature superconducting material and that generates a static magnetic field to which a subject (170) is exposed; and a DC power supply device (200) that supplies a direct current to the superconducting coil (115). The DC power supply device (200) is configured to allow setting of a plurality of magnetic-field intensities by adjusting the amplitude of the output current that is supplied to the superconducting coil (115).

Description

磁気共鳴イメージング装置Magnetic resonance imaging system
　本発明は、磁気共鳴イメージング装置に関し、より特定的には、複数の磁場強度の設定が可能な磁気共鳴イメージング装置に関する。 The present invention relates to a magnetic resonance imaging apparatus, and more particularly to a magnetic resonance imaging apparatus capable of setting a plurality of magnetic field strengths.
　医療分野などにおいて、核磁気共鳴現象を利用して生体内の情報を画像化し、画像化された人体の断面画像に基づいて検査や診断を行なう核磁気共鳴イメージング（Magnetic　Resonance　Imaging：ＭＲＩ）装置が知られている。このＭＲＩ装置においては、画像の精度を向上させるために、強磁場を発生させる超電導磁石が用いられる場合がある。 In the medical field, etc., there is a nuclear magnetic resonance imaging (MRI) apparatus that uses nuclear magnetic resonance phenomena to image in vivo information, and performs examinations and diagnoses based on the cross-sectional images of the human body. Are known. In this MRI apparatus, a superconducting magnet that generates a strong magnetic field may be used in order to improve image accuracy.
　多くのＭＲＩ装置においては、磁石内の超電導コイルとして低温超電導材料が用いられる場合がある。この場合、超電導コイルを超電導温度に維持するために、液体ヘリウムが冷媒として用いられる。 In many MRI apparatuses, a low-temperature superconducting material may be used as a superconducting coil in a magnet. In this case, liquid helium is used as the refrigerant to maintain the superconducting coil at the superconducting temperature.
　しかしながら、液体ヘリウムは一般的に入手が困難であり、特に近年は価格高騰により、一層入手が困難となりつつある。また、液体ヘリウムを冷媒として用いる場合、その貯蔵や冷却設備のために、装置全体が大型化し設備コストが増大する傾向にある。 However, liquid helium is generally difficult to obtain, and particularly in recent years, it has become more difficult to obtain due to rising prices. Further, when liquid helium is used as a refrigerant, the entire apparatus tends to be large and the equipment cost tends to increase due to storage and cooling equipment.
　このような課題に対して、たとえば特開２００９－１８３３７２号公報（特許文献１）や特開２０１５－１６７５７６号公報（特許文献２）のように、液体ヘリウムを冷媒とする必要のある低温超電導材料に代えて、低温超電導材料よりも臨界温度が高い高温超電導材料を用いる技術が提案されている。このような高温超電導材料を用いることで、液体ヘリウムの使用を抑制することができる。 For such a problem, a low-temperature superconducting material that requires liquid helium as a refrigerant, such as Japanese Unexamined Patent Application Publication No. 2009-183372 (Patent Document 1) and Japanese Unexamined Patent Application Publication No. 2015-167576 (Patent Document 2). Instead, a technique using a high-temperature superconducting material having a critical temperature higher than that of the low-temperature superconducting material has been proposed. By using such a high-temperature superconducting material, the use of liquid helium can be suppressed.
特開２００９－１８３３７２号公報JP 2009-183372 A 特開２０１５－１６７５７６号公報Japanese Patent Laying-Open No. 2015-167576
　従来のＭＲＩ装置は、その用途によって必要とされる静磁場の強度が異なっており、たとえば、静磁場が１．５Ｔ，３Ｔ，７ＴなどのＭＲＩ装置が知られている。すべての用途に対応する場合には、各用途に応じた個別のＭＲＩ装置を設けることが必要となるため、広い設置スペースが必要となり、また設備導入の費用が高額となってしまう。そのため、１台の設備で複数の磁場強度を発生させることのできるＭＲＩ装置の開発が望まれている。 Conventional MRI apparatuses have different static magnetic field strengths depending on their applications. For example, MRI apparatuses having a static magnetic field of 1.5T, 3T, 7T, etc. are known. In order to deal with all uses, it is necessary to provide individual MRI apparatuses according to each use, so that a large installation space is required and the cost of introducing the equipment becomes high. Therefore, it is desired to develop an MRI apparatus that can generate a plurality of magnetic field strengths with a single facility.
　一般的に、磁場強度を変更するには、磁石内の超電導コイルに流れる電流を変化させることで実現可能である。しかしながら、特に液体ヘリウムを用いる低温超電導材料においては、超電導コイル内を流れる電流を変更する場合には、電流スイッチを動作させて超電導コイルに外部から電流を供給したり、超電導コイルから電流を取り出すことが必要となる。このとき、電流スイッチの動作による電流の増減によって発熱が生じるため、液体ヘリウムが消費（蒸発）されてしまう。そのため、液体ヘリウムの補充等によるコストの増加となってしまう。 Generally, changing the magnetic field strength can be realized by changing the current flowing through the superconducting coil in the magnet. However, especially in low-temperature superconducting materials using liquid helium, when changing the current flowing in the superconducting coil, the current switch is operated to supply current to the superconducting coil from the outside, or to extract current from the superconducting coil. Is required. At this time, since heat is generated due to increase / decrease in current due to the operation of the current switch, liquid helium is consumed (evaporated). Therefore, the cost increases due to the replenishment of liquid helium.
　また、低温超電導材料を用いた装置では、装置内の各部品の設計時に液体ヘリウム温度ぎりぎりで動作するように設計されていることが多いため、磁場変更時に大きな温度変動が生じると機器の破損につながる可能性がある。そのため、低温超電導材料を用いた装置における磁場変更は、専門技術を有する技術者以外では行なうことが困難である。 In addition, devices using low-temperature superconducting materials are often designed to operate just below the liquid helium temperature when designing each component in the device. There is a possibility of connection. For this reason, it is difficult to change the magnetic field in the apparatus using the low temperature superconducting material by a person other than an engineer having expertise.
　一方で、上記の特許文献に記載されているような高温超電導材料を用いた場合、たとえば超電導材料の接続部分等において超電導状態にできない状況が発生し得るため、微小の電流損失が生じ得ることが知られている。そして、高温超電導材料を用いる場合には、この電流損失を補うために、高い安定性を有する電源装置が設けられて超電導コイルに対して電源供給がなされることが一般的である。 On the other hand, when a high-temperature superconducting material as described in the above patent document is used, for example, a situation where the superconducting material cannot be brought into a superconducting state may occur, so that a minute current loss may occur. Are known. When a high-temperature superconducting material is used, in order to compensate for this current loss, a power supply device having high stability is generally provided and power is supplied to the superconducting coil.
　本願発明者は、高温超電導材料を用いたＭＲＩ装置において、当該電源装置を調節することによって、超電導コイルで発生される磁場強度を変更できる可能性があることに着目した。 The inventor of the present application has focused on the possibility that the magnetic field strength generated in the superconducting coil can be changed by adjusting the power supply device in an MRI apparatus using a high-temperature superconducting material.
　本発明は、上記のような課題を解決するためになされたものであって、その目的は、複数の磁場強度を切換えて使用できる磁気共鳴イメージング（ＭＲＩ）装置を提供することである。 The present invention has been made to solve the above-described problems, and an object thereof is to provide a magnetic resonance imaging (MRI) apparatus that can be used by switching a plurality of magnetic field strengths.
　本開示に係る磁気共鳴イメージング装置は、複数の磁場強度を設定することが可能であって、超電導コイルと、直流電源装置と、シムコイルとを備える。超電導コイルは、高温超電導材を含んで構成され、被検体が曝露する静磁場を発生する。直流電源装置は、超電導コイルに出力電流を供給する。直流電源装置は、超電導コイルへ供給する出力電流の大きさを調整することによって、複数の磁場強度を設定することが可能となるように構成される。シムコイルは、超電導コイルにより生成された磁場分布を補正可能に構成される。磁場分布は、複数の磁場強度の各々に対応した電流をシムコイルに供給することによって補正される。 The magnetic resonance imaging apparatus according to the present disclosure can set a plurality of magnetic field strengths, and includes a superconducting coil, a DC power supply device, and a shim coil. The superconducting coil includes a high-temperature superconducting material, and generates a static magnetic field to which the subject is exposed. The DC power supply device supplies an output current to the superconducting coil. The DC power supply device is configured to be able to set a plurality of magnetic field strengths by adjusting the magnitude of the output current supplied to the superconducting coil. The shim coil is configured to be able to correct the magnetic field distribution generated by the superconducting coil. The magnetic field distribution is corrected by supplying a current corresponding to each of the plurality of magnetic field strengths to the shim coil.
　磁気共鳴イメージング装置は、所定周波数の励磁パルスを発生させるための送信コイルと、被検体からの磁気共鳴信号を受信するための受信コイルをさらに備える。複数の磁場強度は、第１の磁場強度および第２の磁場強度を含む。送信コイルは、第１の磁場強度に対応する第１の周波数、および第２の磁場強度に対応する第２の周波数の双方で共振するように構成される。 The magnetic resonance imaging apparatus further includes a transmission coil for generating an excitation pulse of a predetermined frequency and a reception coil for receiving a magnetic resonance signal from the subject. The plurality of magnetic field strengths include a first magnetic field strength and a second magnetic field strength. The transmit coil is configured to resonate at both a first frequency corresponding to the first magnetic field strength and a second frequency corresponding to the second magnetic field strength.
　好ましくは、第１および第２の磁場強度は、それぞれ１．５Ｔおよび３Ｔである。そして、第１の周波数は６４ＭＨｚであり、第２の周波数は１２８ＭＨｚである。 Preferably, the first and second magnetic field strengths are 1.5T and 3T, respectively. The first frequency is 64 MHz and the second frequency is 128 MHz.
　好ましくは、シムコイルに供給される電流は、複数の磁場強度の各々における磁場強度変化の主成分分析に基づいて決定される。 Preferably, the current supplied to the shim coil is determined based on a principal component analysis of a change in magnetic field strength at each of a plurality of magnetic field strengths.
　好ましくは、磁気共鳴イメージング装置は、超電導コイルに供給される駆動電流を計測するための電流計測装置をさらに備える。直流電源は、電流計測装置によって計測される駆動電流の変動に基づいて出力電流を調整する。 Preferably, the magnetic resonance imaging apparatus further includes a current measurement device for measuring a drive current supplied to the superconducting coil. The direct current power source adjusts the output current based on the fluctuation of the drive current measured by the current measuring device.
　本発明によれば、複数の磁場強度を切換えて使用することができるＭＲＩ装置を提供することが可能となる。 According to the present invention, it is possible to provide an MRI apparatus that can be used by switching a plurality of magnetic field strengths.
本実施の形態に従うＭＲＩ装置の概略を説明するための機能ブロック図である。It is a functional block diagram for demonstrating the outline of the MRI apparatus according to this Embodiment. 異なる磁場強度のＭＲＩ装置における特徴を説明するための図である。It is a figure for demonstrating the characteristic in the MRI apparatus of a different magnetic field strength. 本実施の形態に従うＭＲＩ装置用のシムコイルの設計手法を説明するためのフローチャートである。It is a flowchart for demonstrating the design method of the shim coil for MRI apparatuses according to this Embodiment. 本実施の形態に従うＭＲＩ装置用のシムコイルへの電流供給制御を説明するためのフローチャートである。It is a flowchart for demonstrating the electric current supply control to the shim coil for MRI apparatuses according to this Embodiment.
　以下、本発明の実施の形態について、図面を参照しながら詳細に説明する。なお、図中同一または相当部分には同一符号を付してその説明は繰り返さない。 Hereinafter, embodiments of the present invention will be described in detail with reference to the drawings. In the drawings, the same or corresponding parts are denoted by the same reference numerals and description thereof will not be repeated.
　（ＭＲＩ装置の説明）
　図１は、本実施の形態に従う磁気共鳴イメージング（ＭＲＩ）装置１０の概略を説明するための機能ブロック図である。 (Description of MRI system)
FIG. 1 is a functional block diagram for explaining an outline of a magnetic resonance imaging (MRI) apparatus 10 according to the present embodiment.
　図１を参照して、ＭＲＩ装置１０は、本体装置１００と、直流電源装置２００と、制御装置３００と、表示部３１０と、入力部３２０とを備える。本体装置１００は、超電導マグネット１１０と、傾斜磁場コイル１２０と、ＲＦ送信コイル１３０と、ＲＦ受信コイル１４０と、シムコイル１５０とを含む。また、超電導マグネット１１０は、高温超電導材料（たとえばビスマス系超電導材料）で形成された高温超電導コイル１１５を含む。 Referring to FIG. 1, the MRI apparatus 10 includes a main body device 100, a DC power supply device 200, a control device 300, a display unit 310, and an input unit 320. Main device 100 includes a superconducting magnet 110, a gradient coil 120, an RF transmitter coil 130, an RF receiver coil 140, and a shim coil 150. Superconducting magnet 110 includes a high-temperature superconducting coil 115 formed of a high-temperature superconducting material (for example, a bismuth-based superconducting material).
　本実施の形態においては、ＭＲＩ装置１０は、いわゆるトンネル型のＭＲＩ装置である場合を例として説明する。本体装置１００の高温超電導コイル１１５、傾斜磁場コイル１２０、ＲＦ送信コイル１３０、シムコイル１５０は、概略的には円筒型の形状を有している。高温超電導コイル１１５、傾斜磁場コイル１２０、およびシムコイル１５０が励磁されると、空洞部（トンネル）内に磁場が発生する。 In the present embodiment, the case where the MRI apparatus 10 is a so-called tunnel type MRI apparatus will be described as an example. The high-temperature superconducting coil 115, the gradient magnetic field coil 120, the RF transmitting coil 130, and the shim coil 150 of the main body device 100 have a generally cylindrical shape. When the high temperature superconducting coil 115, the gradient magnetic field coil 120, and the shim coil 150 are excited, a magnetic field is generated in the cavity (tunnel).
　被検者１７０は、検査用テーブル１６０上に横たえられた状態で、円筒型の本体装置１００のトンネル内に挿入される。ＲＦ受信コイル１４０は、被検者１７０における検査対象部位を覆うように設置される。図１においては、ＲＦ受信コイル１４０は、頭部を検査する場合を例として示されているが、たとえば、検査対象部位が胴部の場合には、ＲＦ受信コイル１４０は胴部を覆うように設置される。 The subject 170 is inserted into the tunnel of the cylindrical main body device 100 while lying on the examination table 160. The RF receiving coil 140 is installed so as to cover the examination target site in the subject 170. In FIG. 1, the RF receiving coil 140 is shown as an example in which the head is inspected. For example, when the inspection target part is the trunk, the RF receiving coil 140 covers the trunk. Installed.
　高温超電導コイル１１５は、トンネル内に空間的および時間的に均一な静磁場を生成するためのコイルである。高温超電導コイル１１５は、直流電源装置２００から供給される電流の大きさに応じた強度の磁場（たとえば、１．５Ｔ，３Ｔ，７Ｔ）を生成することが可能に構成されている。高温超電導コイル１１５で生じる静磁場によって、被検者１７０の検査対象部位の水素原子核の持つ核スピンを一定方向に整列させることができる。 The high temperature superconducting coil 115 is a coil for generating a spatially and temporally uniform static magnetic field in the tunnel. The high temperature superconducting coil 115 is configured to be able to generate a magnetic field (for example, 1.5T, 3T, 7T) having a strength corresponding to the magnitude of the current supplied from the DC power supply device 200. Due to the static magnetic field generated by the high-temperature superconducting coil 115, the nuclear spins of the hydrogen nuclei at the site to be examined of the subject 170 can be aligned in a certain direction.
　シムコイル１５０は、高温超電導コイル１１５によって生成される静磁場を補正するための磁場を生成するためのコイルである。シムコイル１５０は、たとえば常電導材料で形成される。 The shim coil 150 is a coil for generating a magnetic field for correcting the static magnetic field generated by the high-temperature superconducting coil 115. Shim coil 150 is formed of, for example, a normal conductive material.
　ＭＲＩ装置１０における静磁場の磁場均一性は、標準的には撮像領域内で１０ｐｐｍ以下に抑える必要がある。しかしながら、高温超電導コイル１１５により生成される静磁場は、ＭＲＩ装置１０に含まれる各機器や、当該装置が設置される建物の鉄筋等の構造物の影響等により少なからず歪みが生じてしまうため、一般的には、高温超電導コイル１１５のみでは均一した静磁場を得ることは困難である。そのため、シムコイル１５０を用いて、これらの影響を相殺するような磁場を形成することによって、トンネル内に発生する静磁場を均一化する。 The magnetic field uniformity of the static magnetic field in the MRI apparatus 10 needs to be suppressed to 10 ppm or less in the imaging region as a standard. However, the static magnetic field generated by the high-temperature superconducting coil 115 is distorted due to the influence of each device included in the MRI apparatus 10 and structures such as reinforcing bars of a building where the apparatus is installed. In general, it is difficult to obtain a uniform static magnetic field only with the high-temperature superconducting coil 115. Therefore, the static magnetic field generated in the tunnel is made uniform by forming a magnetic field that cancels these influences using the shim coil 150.
　なお、上記のような各影響要素によって生じる磁場の強度は一定ではないため、発生する異なる磁場強度を調整するために、シムコイル１５０は異なる磁場分布を生成する複数のコイルで構成される場合がある。 In addition, since the intensity of the magnetic field generated by each of the influence factors as described above is not constant, the shim coil 150 may be configured by a plurality of coils that generate different magnetic field distributions in order to adjust different generated magnetic field intensity. .
　傾斜磁場コイル１２０（または勾配磁場コイル）は、空間的に線形に変化する勾配磁場を形成するためのコイルである。この傾斜磁場によって、検査対象部位の水素原子核が出す信号の周波数を空間的に線形に変化させることができる。したがって、ＲＦ受信コイル１４０によって受信される受信信号に対して位置情報を付加することができる。 The gradient magnetic field coil 120 (or gradient magnetic field coil) is a coil for forming a gradient magnetic field that linearly changes spatially. This gradient magnetic field can spatially and linearly change the frequency of the signal emitted by the hydrogen nuclei at the site to be examined. Therefore, position information can be added to the received signal received by the RF receiving coil 140.
　ＲＦ送信コイル１３０は、被検者１７０に対して所定の周波数のＲＦパルス信号を送信するためのコイルである。検査対象部位に対してＲＦパルス信号が照射されると、ＲＦパルスにより与えられるエネルギによって検査対象部位の水素原子核が励起される。そして、ＲＦパルス信号が停止されると水素原子核が励起状態から復帰する。この励起状態から復帰する際に観測される信号がＲＦ受信コイル１４０にて受信される。 The RF transmission coil 130 is a coil for transmitting an RF pulse signal having a predetermined frequency to the subject 170. When the RF pulse signal is irradiated to the inspection target part, the hydrogen nuclei of the inspection target part are excited by the energy given by the RF pulse. When the RF pulse signal is stopped, the hydrogen nuclei return from the excited state. A signal that is observed when returning from this excited state is received by the RF receiving coil 140.
　上記の傾斜磁場コイル１２０によって生成される傾斜磁場のために、ＲＦ受信コイル１４０で受信される各位置からの受信信号はその位相が異なったものとなる。そのため、印加する傾斜磁場とＲＦパルス信号の周波数とを適切に調整することで、得られた受信信号から当該信号を放出した水素原子核の位置を特定することができる。この受信信号を平面的あるいは三次元的に配列することによって、検査対象部位を画像化することができる。 Because of the gradient magnetic field generated by the gradient magnetic field coil 120, the received signals from the respective positions received by the RF receiving coil 140 have different phases. Therefore, by appropriately adjusting the gradient magnetic field to be applied and the frequency of the RF pulse signal, the position of the hydrogen nucleus that has emitted the signal can be specified from the obtained reception signal. By arranging the received signals in a two-dimensional or three-dimensional manner, the site to be examined can be imaged.
　直流電源装置２００は、高温超電導コイル１１５により静磁場を生成させるための電流を供給する。直流電源装置２００から高温超電導コイル１１５に供給される直流電流は電流センサ２１０によって測定される。なお、直流電源装置２００から出力される電流は、制御装置３００からの信号により調整可能である。制御装置３００は、電流センサ２１０で計測された電流値を受け、直流電源装置２００からの出力電流が一定となるようにフィードバック制御を行なう。これにより、安定した静磁場を発生することができる。 DC power supply device 200 supplies a current for generating a static magnetic field by high-temperature superconducting coil 115. A direct current supplied from the direct current power supply device 200 to the high temperature superconducting coil 115 is measured by a current sensor 210. Note that the current output from the DC power supply device 200 can be adjusted by a signal from the control device 300. Control device 300 receives the current value measured by current sensor 210 and performs feedback control so that the output current from DC power supply device 200 is constant. Thereby, a stable static magnetic field can be generated.
　制御装置３００は、入力部３２０から入力されるユーザからの情報を受ける。ユーザからの情報には、たとえば使用する磁場強度の情報、被検者１７０に関する情報などが含まれる。制御装置３００は、磁場強度の情報に基づいて直流電源装置２００から高温超電導コイル１１５に供給するための電流を制御する。 The control device 300 receives information from the user input from the input unit 320. The information from the user includes, for example, information on the magnetic field strength to be used, information on the subject 170, and the like. The control device 300 controls the current to be supplied from the DC power supply device 200 to the high temperature superconducting coil 115 based on the magnetic field strength information.
　制御装置３００は、傾斜磁場コイル１２０およびシムコイル１５０についての励磁電流を調整する。また、制御装置３００は、ＲＦ送信コイル１３０に対してＲＦパルス信号を出力するとともに、当該ＲＦパルス信号に対してＲＦ受信コイル１４０で受信された受信信号を受ける。制御装置３００は、この受信信号に基づいて、検査対象部位の断面を画像化し、表示部３１０に表示する。 The control device 300 adjusts the excitation current for the gradient magnetic field coil 120 and the shim coil 150. In addition, the control device 300 outputs an RF pulse signal to the RF transmission coil 130 and receives a reception signal received by the RF reception coil 140 with respect to the RF pulse signal. Based on the received signal, the control device 300 images the cross section of the examination target region and displays it on the display unit 310.
　なお、高温超電導コイルにおいては上記のように電流損失のために少なからず発熱が生じ得るため、図１には記載されていないが、高温超電導コイルを冷却するための冷却装置が設けられる。この場合、コイルが高温超電導材料であるため、冷媒は液体ヘリウムである必要はなく、たとえば液体窒素などのより安価で入手しやすい冷媒を用いることができる。 In the high temperature superconducting coil, heat generation may occur due to current loss as described above. Therefore, although not shown in FIG. 1, a cooling device for cooling the high temperature superconducting coil is provided. In this case, since the coil is a high-temperature superconducting material, the refrigerant does not need to be liquid helium, and a cheaper and easily available refrigerant such as liquid nitrogen can be used.
　図２は、現在医療分野で用いられているＭＲＩ装置の、各磁場強度における特徴を説明するための図である。磁場強度としては、医療分野で多く用いられる１．５Ｔ，３Ｔ，７Ｔの場合を例として説明する。また、特徴としては、Ｓ／Ｎ比、各検査対象に対する適否が示されており、Ａは「非常に適している」を意味し、Ｂは「適している」、Ｃは「やや不適」、Ｄは「不適」を意味している。 FIG. 2 is a diagram for explaining the characteristics of each magnetic field strength of the MRI apparatus currently used in the medical field. As the magnetic field strength, a case of 1.5T, 3T, and 7T, which is often used in the medical field, will be described as an example. Further, as characteristics, S / N ratio, suitability for each inspection object is shown, A means “very suitable”, B is “suitable”, C is “slightly unsuitable”, D means “unsuitable”.
　Ｓ／Ｎ比については、磁場強度が大きくなるほど良くなっている。すなわち、磁場強度を大きくするほど、得られる画像品質（解像度，コントラスト）が向上する。そのため、磁場強度の高いＭＲＩ装置は、頭部（脳内）や手足のような細かい部分を撮像する場合、あるいは、代謝情報を撮像する場合に適している。 The S / N ratio is improved as the magnetic field strength increases. That is, as the magnetic field strength is increased, the obtained image quality (resolution, contrast) is improved. Therefore, an MRI apparatus having a high magnetic field strength is suitable for imaging fine parts such as the head (in the brain) and limbs, or imaging metabolic information.
　一方で、磁場強度の大きいＭＲＩでは、磁場強度に比例して生体内の不均一磁場が大きくなるため、不均一磁場の原因となる酸素分子が多い腹部臓器などを含む体幹部の撮像には不利になりやすい。このような部位には、磁場強度が比較的小さいＭＲＩ装置の方が有利になる。 On the other hand, in MRI with a high magnetic field strength, the inhomogeneous magnetic field in the living body increases in proportion to the magnetic field strength, which is disadvantageous for imaging of the trunk including abdominal organs and the like that have many oxygen molecules that cause the inhomogeneous magnetic field. It is easy to become. For such a part, an MRI apparatus having a relatively small magnetic field strength is more advantageous.
　なお、ＭＲＩ装置は強磁場を生成するため、たとえば体内にペースメーカやインプラントなどの金属がある場合には使用できない場合が生じ得る。これは、生成される磁場強度が大きくなる場合に特に問題になりやすい。そのため、体内の金属の有無を事前に知ることができない救急対応の場合などには、磁場強度の小さいＭＲＩを使用することが好ましい。 Note that since the MRI apparatus generates a strong magnetic field, it may not be used when there is a metal such as a pacemaker or implant in the body. This is particularly problematic when the strength of the generated magnetic field is increased. For this reason, it is preferable to use an MRI with a small magnetic field strength in the case of emergency response in which the presence or absence of metal in the body cannot be known in advance.
　このように、医療分野においては、高分解能となる磁場強度の大きいＭＲＩ装置のみが求められるのではなく、依然として磁場強度の小さいＭＲＩ装置に対するニーズも多い。そのため、検査対象部位や検査状況に応じて、異なる磁場強度を用いて撮像変化することが必要となる。 Thus, in the medical field, not only an MRI apparatus with a high magnetic field strength with high resolution is required, but there are still many needs for an MRI apparatus with a low magnetic field intensity. For this reason, it is necessary to change the imaging using different magnetic field strengths depending on the region to be inspected and the inspection situation.
　ところが、従来のＭＲＩ装置においては、使用される磁場強度は固定されているため、各用途に応じて個別のＭＲＩ装置を設けることが必要となる。そのため、複数の設備を設置するための広いスペースが必要となり、さらには設備導入の費用が高額となってしまうという問題がある。 However, in the conventional MRI apparatus, since the magnetic field strength used is fixed, it is necessary to provide an individual MRI apparatus according to each application. Therefore, there is a problem that a large space for installing a plurality of facilities is required, and furthermore, the cost of introducing the facilities becomes high.
　ＭＲＩ装置において磁場強度を変更するには、磁石内の超電導コイルに流れる電流を変化させることで実現することができる。しかしながら、特に液体ヘリウムを用いる低温超電導材料を用いるＭＲＩ装置においては、超電導コイル内を流れる電流を変更する場合には、永久電流スイッチと呼ばれるスイッチを動作させて超電導コイルに外部から電流を供給したり、超電導コイルから電流を取り出すことが必要となる。このとき、永久電流スイッチの動作によって発熱が生じるため、液体ヘリウムが消費（蒸発）されてしまう。そのため、液体ヘリウムの補充等によるコストの増加となってしまったり、さらには急激に温度が上昇すると、爆発的に液体ヘリウムの蒸発が発生するいわゆる「クエンチ」が生じるおそれがある。 In the MRI apparatus, changing the magnetic field strength can be realized by changing the current flowing through the superconducting coil in the magnet. However, especially in an MRI apparatus using a low temperature superconducting material using liquid helium, when changing the current flowing in the superconducting coil, a switch called a permanent current switch is operated to supply current to the superconducting coil from the outside. It is necessary to extract current from the superconducting coil. At this time, since heat is generated by the operation of the permanent current switch, liquid helium is consumed (evaporated). For this reason, if the cost increases due to replenishment of liquid helium, or if the temperature rises rapidly, so-called “quenching” in which evaporation of liquid helium occurs explosively may occur.
　また、低温超電導材料を用いた装置では、装置内の各部品の設計時に液体ヘリウム温度ぎりぎりで動作するように設計されていることが多いため、磁場変更時に大きな温度変動が生じると機器の破損につながる可能性がある。そのため、低温超電導材料を用いた装置における磁場変更はリスクを伴うものであり、専門技術を有する技術者以外では行なうことが困難である。 In addition, devices using low-temperature superconducting materials are often designed to operate just below the liquid helium temperature when designing each component in the device. There is a possibility of connection. Therefore, the magnetic field change in the apparatus using the low-temperature superconducting material involves a risk, and it is difficult to perform it by a person other than an engineer having specialized skills.
　一方で、図１に示した本実施の形態に係るＭＲＩ装置１０のように高温超電導コイル１１５を用いる場合、液体ヘリウムを用いる必要はなくなるが、たとえば超電導材料の接続部分等において超電導状態にできない状況が発生するため、微小の電流損失が生じ得る。そのため、本実施の形態に係るＭＲＩ装置１０では、直流電源装置２００を設けて、高温超電導コイル１１５に対して電源供給を行なうことで、均一な静磁場を安定的に発生させる構成となっている。 On the other hand, when the high-temperature superconducting coil 115 is used as in the MRI apparatus 10 according to the present embodiment shown in FIG. 1, it is not necessary to use liquid helium. However, the superconducting material cannot be brought into a superconducting state, for example, at the connection portion of the superconducting material. As a result, a minute current loss may occur. Therefore, the MRI apparatus 10 according to the present embodiment has a configuration in which a uniform static magnetic field is stably generated by providing the DC power supply device 200 and supplying power to the high-temperature superconducting coil 115. .
　そして、本実施の形態に係るＭＲＩ装置１０においては、入力部３２０からの設定により直流電源装置２００から高温超電導コイル１１５に供給する電流を変化させることによって、１台のＭＲＩ装置１０によって異なる磁場強度を発生させることができる。 In the MRI apparatus 10 according to the present embodiment, the current supplied from the DC power supply apparatus 200 to the high-temperature superconducting coil 115 is changed by the setting from the input unit 320, thereby changing the magnetic field intensity that varies depending on one MRI apparatus 10. Can be generated.
　発生する磁場強度が変化すると、それに応じてＲＦ送信コイル１３０によって送信するＲＦパルス信号の周波数を変更させる必要がある。本実施の形態に係るＭＲＩ装置１０においては、ＲＦ送信コイル１３０は、使用される磁場強度に対応した複数の周波数で共振するように構成されている。たとえば、使用される磁場強度が１．５Ｔと３Ｔである場合には、ＲＦ送信コイル１３０は、１．５Ｔの磁場強度に対応した周波数である６４ＭＨｚと、３Ｔの磁場強度に対応した周波数である１２８ＭＨｚで共振するように構成される。このように、複数の周波数で共振するようにＲＦ送信コイル１３０を設計することによって、磁場強度の変更に対応してＲＦ送信コイルの交換を抑制することができる。なお、３つ以上の磁場強度が切換可能とされる場合には、それらの磁場強度に対応した周波数で共振するようにＲＦ共振コイルを設計することが好ましい。 When the intensity of the generated magnetic field changes, it is necessary to change the frequency of the RF pulse signal transmitted by the RF transmission coil 130 accordingly. In the MRI apparatus 10 according to the present embodiment, the RF transmission coil 130 is configured to resonate at a plurality of frequencies corresponding to the magnetic field strength used. For example, when the magnetic field strengths used are 1.5T and 3T, the RF transmission coil 130 has a frequency corresponding to a magnetic field strength of 1.5T and a frequency corresponding to a magnetic field strength of 3T. It is configured to resonate at 128 MHz. In this way, by designing the RF transmission coil 130 so as to resonate at a plurality of frequencies, it is possible to suppress replacement of the RF transmission coil in response to a change in magnetic field strength. When three or more magnetic field strengths can be switched, it is preferable to design the RF resonance coil so as to resonate at a frequency corresponding to the magnetic field strengths.
　また、超電導マグネット１１０で生成される静磁場の磁場強度が変化すると、当該静磁場に対する周囲からの影響（歪み）も変化し得る。そのため、本実施の形態に係るＭＲＩ装置１０においては、各磁場強度に適合されたシムコイル１５０が設けられる。なお、シムコイル１５０は複数のシムコイルから構成されており、当該複数のシムコイルの組合せによって静磁場を補正するように構成される。そのため、一部のシムコイルについては、異なる磁場強度においても共通に使用される場合がある。 Also, when the magnetic field strength of the static magnetic field generated by the superconducting magnet 110 changes, the influence (distortion) from the surroundings on the static magnetic field can also change. Therefore, in the MRI apparatus 10 according to the present embodiment, a shim coil 150 adapted to each magnetic field strength is provided. The shim coil 150 includes a plurality of shim coils, and is configured to correct a static magnetic field by a combination of the plurality of shim coils. Therefore, some shim coils may be used in common at different magnetic field strengths.
　以上のように、本実施の形態に係るＭＲＩ装置１０においては、超電導マグネット１１０に高温超電導コイル１１５を用いることにより液体ヘリウムを使用しないようにするとともに、静磁場を保持するための直流電源装置２００の出力電流を変化させることによって、１台のＭＲＩ装置で異なる磁場強度を発生させることができるように構成されている。これによって、設備の小型化および省スペース化を実現することができるとともに、ＭＲＩ装置の台数を削減できるためトータルコストを削減することが可能となる。 As described above, in the MRI apparatus 10 according to the present embodiment, the high-temperature superconducting coil 115 is used in the superconducting magnet 110 so that liquid helium is not used, and a DC power supply apparatus 200 for maintaining a static magnetic field. By changing the output current, different magnetic field strengths can be generated by one MRI apparatus. As a result, downsizing and space saving of the equipment can be realized, and the total number of MRI apparatuses can be reduced, so that the total cost can be reduced.
　また、ＲＦ送信コイルを、使用する複数の磁場強度に適合したそれぞれの周波数で共振するように設計することで、磁場強度の変更を容易にすることが可能となる。 In addition, it is possible to easily change the magnetic field strength by designing the RF transmission coil so as to resonate at respective frequencies suitable for a plurality of magnetic field strengths to be used.
　さらに、高温超電導材料で形成された超電導コイルを用い、直流電源装置の電流を変更することによって磁場強度を変更させることができるので、一般的な超電導マグネットの場合と比較して、磁場強度の変更に伴うリスクを低減でき、磁場強度の変更を容易に行なうことが可能になる。そのため、各医療機関において、たとえば曜日ごとに磁場強度の変更を行なうなど、頻繁に磁場強度を変更することが可能となる。 Furthermore, the magnetic field strength can be changed by using a superconducting coil made of high-temperature superconducting material, and changing the current of the DC power supply, so that the magnetic field strength can be changed compared to the case of a general superconducting magnet. Can be reduced, and the magnetic field strength can be easily changed. Therefore, in each medical institution, for example, the magnetic field strength can be frequently changed, for example, the magnetic field strength is changed every day of the week.
　（シムコイルの設計方法）
　本実施の形態に係るＭＲＩ装置においては、１台のＭＲＩ装置で複数の磁場強度を生成することができるが、磁場強度が異なると不均一磁場の程度や分布が変化するため、磁場強度に対応したシムコイルの設定が必要となる。 (Sim coil design method)
In the MRI apparatus according to the present embodiment, a single MRI apparatus can generate a plurality of magnetic field strengths. However, if the magnetic field strength differs, the degree and distribution of the inhomogeneous magnetic field changes. It is necessary to set the shim coil.
　ここで、本実施の形態のように磁場強度が可変である場合には、鉄シムのような固定的な補正磁場では十分に補正することができず、シムコイルを用いて補正磁場も可変とすることが必要となる。この場合、大きな磁場補正のために、シムコイル自体の抵抗値の上昇やシムコイルに流れる電流量が増加してしまい、結果としてシムコイルによる発熱量が増加してしまう。 Here, when the magnetic field intensity is variable as in the present embodiment, it cannot be sufficiently corrected with a fixed correction magnetic field such as an iron shim, and the correction magnetic field is also variable using a shim coil. It will be necessary. In this case, due to the large magnetic field correction, the resistance value of the shim coil itself increases and the amount of current flowing through the shim coil increases, resulting in an increase in the amount of heat generated by the shim coil.
　発生した熱は冷却装置の冷媒により冷却されるが、冷却配管の設置および冷却効率の観点からは、発熱箇所を局所的に集中させることが好ましい。 The generated heat is cooled by the refrigerant of the cooling device, but it is preferable to concentrate the heat generation portion locally from the viewpoint of the installation of the cooling pipe and the cooling efficiency.
　そこで、本実施の形態においては、使用するシムコイルを、強い補正磁場を生成するための「高発熱系シムコイル」と、微調整用の弱い補正磁場を生成するための「低発熱系シムコイル」とに分けることにより、シムコイルの冷却効率を向上させる。 Therefore, in this embodiment, the shim coil to be used is a “high heat generation shim coil” for generating a strong correction magnetic field and a “low heat generation shim coil” for generating a weak correction magnetic field for fine adjustment. By dividing, the cooling efficiency of the shim coil is improved.
　以下、上記の「高発熱系シムコイル」および「低発熱系シムコイル」の設計手法について図３のフローチャートを用いて説明する。 Hereinafter, the design method of the “high heat generation shim coil” and the “low heat generation shim coil” will be described with reference to the flowchart of FIG.
　ＭＲＩ装置は、設置場所ごとに外部からの磁場への影響が異なるため、シムコイルの設計にあたっては、所望の設置場所へのＭＲＩ装置の設置が完了した状態で図３のフローチャートに従って設計を行なう場合がある。 Since the influence on the magnetic field from the outside differs depending on the installation location of the MRI apparatus, when designing the shim coil, the design may be performed according to the flowchart of FIG. 3 in a state where the installation of the MRI apparatus at the desired installation location is completed. is there.
　図３を参照して、ＭＲＩ装置で設定可能な所定の各磁場強度ｂ（たとえば、１．５Ｔおよび３Ｔ）を発生させている状態において、検査領域内の複数の測定箇所の磁場強度を複数回計測する（ステップＳ１００）。この時の測定された磁場強度をＢb,nと表す。ここで、ｎは測定回数（１≦ｎ≦Ｎ）を表しＮは総測定回数を示す。また、ＢはＭ次元ベクトルであり、Ｍは総測定点数を表している。 Referring to FIG. 3, in a state where predetermined magnetic field strengths b (for example, 1.5T and 3T) that can be set by the MRI apparatus are generated, the magnetic field strengths at a plurality of measurement locations in the examination region are measured a plurality of times. Measurement is performed (step S100). The measured magnetic field strength at this time is represented as B b, n . Here, n represents the number of measurements (1 ≦ n ≦ N), and N represents the total number of measurements. B is an M-dimensional vector, and M represents the total number of measurement points.
　次に計測したＢb,nとその平均磁場からの磁場強度変化ΔＢb,nを以下の式（１）のように演算する（ステップＳ１１０）。 Next, the measured B b, n and the magnetic field strength change ΔB b, n from the average magnetic field are calculated as in the following equation (1) (step S110).
　　ΔＢb,n＝Ｂb,n－（ΣbＢb,n）／Ｎ　　…　（１）
　そして、得られた磁場強度変化ΔＢb,nをＭ次元空間上のＮ点の集合と考え、主成分分析を行なう（ステップＳ１２０）。 ΔB b, n = B b, n − (Σ b B b, n ) / N (1)
Then, the obtained magnetic field strength change ΔB b, n is considered as a set of N points on the M-dimensional space, and principal component analysis is performed (step S120).
　次に、主成分分析によって得られた第Ｋ主成分の累積寄与率が、所定のしきい値Ｔ１以下であるか否かを判定する（ステップＳ１３０）。第Ｋ主成分の累積寄与率がしきい値Ｔ１以下の場合（ステップＳ１３０にてＹＥＳ）は、当該主成分が表す磁場分布を形成するために比較的大きな電流をシムコイルに流すことが必要になるため、当該シムコイルを高発熱系シムコイルとして設計する（ステップＳ１４０）。 Next, it is determined whether or not the cumulative contribution ratio of the Kth principal component obtained by the principal component analysis is equal to or less than a predetermined threshold value T1 (step S130). When the cumulative contribution ratio of the K-th principal component is equal to or less than threshold value T1 (YES in step S130), it is necessary to flow a relatively large current through the shim coil in order to form the magnetic field distribution represented by the principal component. Therefore, the shim coil is designed as a high heat generation shim coil (step S140).
　一方、第Ｋ主成分の累積寄与率がしきい値Ｔ１より大きい場合（ステップＳ１３０にてＮＯ）は、次に第Ｋ主成分の累積寄与率が、所定のしきい値Ｔ２以下（Ｔ１＜Ｔ２）であるか否かを判定する（ステップＳ１５０）。 On the other hand, when the cumulative contribution ratio of the K-th principal component is greater than threshold value T1 (NO in step S130), the cumulative contribution ratio of the K-th principal component is next equal to or less than predetermined threshold value T2 (T1 <T2). ) Is determined (step S150).
　累積寄与率がしきい値Ｔ２以下の場合（ステップＳ１５０にてＹＥＳ）は、当該主成分が表す磁場分布を形成するためには比較的低い電流をシムコイルに流せばよいため、当該シムコイルを低発熱系のシムコイルとして設計する。一方、累積寄与率がしきい値Ｔ２よりも大きい場合（ステップＳ１５０にてＮＯ）は、当該主成分が表す磁場分布については補正の必要がないものと判断することができる。 If the cumulative contribution rate is equal to or less than threshold value T2 (YES in step S150), a relatively low current may be passed through the shim coil in order to form the magnetic field distribution represented by the main component. Designed as a system shim coil. On the other hand, if the cumulative contribution rate is greater than threshold value T2 (NO in step S150), it can be determined that there is no need to correct the magnetic field distribution represented by the principal component.
　ステップＳ１２０の主成分分析の結果から得られる各主成分に対して、ステップＳ１３０～Ｓ１６０までの手順を行ない、各主成分が表す磁場分布を形成するためのシムコイルを、高発熱系シムコイルまたは低発熱系シムコイルに分類して設計する。なお、得られた主成分に基づいたコイルの設計手法については、公知の手法を用いることができる。 For each principal component obtained from the result of the principal component analysis in step S120, the procedure from step S130 to S160 is performed, and a shim coil for forming a magnetic field distribution represented by each principal component is a high heat generation shim coil or a low heat generation. Design and classify into system shim coils. As a coil design method based on the obtained main component, a known method can be used.
　このような手順に従ってシムコイルを設計することによって、必要となるシムコイルの数を低減でき、かつ、発熱箇所を局所的に集中させることができる。そのため、シムコイルを冷却するための冷却系をシンプル化し、冷却効率を向上させることができる。 By designing the shim coils according to such a procedure, the number of shim coils required can be reduced and the heat generation points can be concentrated locally. Therefore, the cooling system for cooling the shim coil can be simplified and the cooling efficiency can be improved.
　（シムコイルへの電流供給制御）
　次に、上記のように設計されたシムコイルに供給する電流を決定する手法について説明する。 (Current supply control to shim coil)
Next, a method for determining the current supplied to the shim coil designed as described above will be described.
　可変磁場を形成するための超電導マグネットで生成される磁場の不均一性および不安定性（すなわち、磁場の時空間変化）は、静磁場強度変更に起因する磁場強度変化ΔＢｆ、超電導マグネットの駆動電流の不安定性に起因する磁場強度変化ΔＢｃ、磁場変更直後の長周期リップルノイズに起因する磁場強度変化ΔＢｒ、および遮蔽電流に起因する磁場強度変化ΔＢｅの、４つの磁場強度変化の時空間の関数の加算によりモデル化することができる。 The inhomogeneity and instability of the magnetic field generated by the superconducting magnet for forming the variable magnetic field (that is, the spatiotemporal change of the magnetic field) is caused by the change in the magnetic field strength ΔB f resulting from the change in the static magnetic field strength, Space-time of four magnetic field strength changes: magnetic field strength change ΔB c due to instability of magnetic field, magnetic field strength change ΔB r due to long-period ripple noise immediately after magnetic field change, and magnetic field strength change ΔB e due to shielding current It can be modeled by adding the functions.
　静磁場強度変更に起因する磁場強度変化ΔＢｆは、磁場強度を変更した場合に、マグネットの製作誤差や電磁力による変形などのマグネット内部の構造や周囲環境に起因して生じる磁場強度変化である。そのため、磁場強度を固定している間の時間変動は発生しない。 The change in magnetic field strength ΔB f resulting from the change in the static magnetic field strength is a change in the magnetic field strength caused by the internal structure of the magnet and the surrounding environment such as a magnet manufacturing error and deformation due to electromagnetic force when the magnetic field strength is changed. . For this reason, no time fluctuation occurs while the magnetic field strength is fixed.
　駆動電流の不安定性に起因する磁場強度変化ΔＢｃは、マグネットの駆動電流を供給する直流電源装置の不安定性によって駆動電流が変動することに起因して生じる磁場強度変化である。当該磁場強度変化ΔＢｃの空間的なパターンは基本的には一定であるが、その大きさや向き（符号）については、駆動電流の変動に相関している。 The change in magnetic field strength ΔB c caused by the instability of the drive current is a change in magnetic field strength caused by the fluctuation of the drive current due to the instability of the DC power supply device that supplies the drive current of the magnet. Although spatial pattern of the magnetic field strength change .DELTA.B c is constant in principle, for the size and orientation (sign) is correlated to the variation of the drive current.
　長周期リップルノイズに起因する磁場強度変化ΔＢｒは、磁場変更プロセスを終了した直後から始まり、マグネットの特性から定まる特定の周期で振動しながら単調に減衰する磁場強度変化である。この磁場強度変化ΔＢｒは、時間と空間の関数であり、時間的および空間的位置に応じて変動する。 Field strength change .DELTA.B r due to the long cycle ripple noise begins immediately after the completion of the magnetic field change process, a magnetic field intensity change monotonously attenuated while oscillating in a specific period determined from the characteristics of the magnet. This magnetic field strength change ΔB r is a function of time and space, and varies according to temporal and spatial positions.
　遮蔽電流に起因する磁場強度変化ΔＢｅは、生成される静磁場を抑制する方向に生じる磁場強度変化であり、磁場変更プロセスの開始直後から非常に遅い速度で単調に減少する特性を有している。この磁場強度変化ΔＢｅも時間と空間の関数であり、時間的および空間的位置に応じて変動する。 The magnetic field intensity change ΔB e caused by the shielding current is a magnetic field intensity change that occurs in a direction to suppress the generated static magnetic field, and has a characteristic that decreases monotonously at a very slow rate immediately after the start of the magnetic field change process. Yes. This magnetic field strength change ΔB e is also a function of time and space, and fluctuates according to temporal and spatial positions.
　ここで、磁場強度変化ΔＢｒおよび磁場強度変化ΔＢｅについては、直流電源装置のインピーダンスを調整してマグネットに適合させたり、磁場変更時の電流調整を適切に行なうことによって、ある程度は低減することが可能である。 Here, the magnetic field intensity change ΔB r and the magnetic field intensity change ΔB e are reduced to some extent by adjusting the impedance of the DC power supply device to be adapted to the magnet or by appropriately adjusting the current when changing the magnetic field. Is possible.
　また、磁場強度変化ΔＢｆ、磁場強度変化ΔＢｒおよび磁場強度変化ΔＢｅは、比較的再現性が高いため、事前に測定した磁場の時空間変化のデータを解析することによって推定することが可能である。磁場強度変化ΔＢｆについては、磁場変更プロセスの開始から十分に時間が経過し、磁場強度変化ΔＢｒおよび磁場強度変化ΔＢｅが減衰した後に、十分に多くの測定から算出した磁場強度を平均化することによって磁場強度変化ΔＢｃの影響を排除することによって推定することができる。 In addition, the magnetic field strength change ΔB f , the magnetic field strength change ΔB r, and the magnetic field strength change ΔB e are relatively highly reproducible, and therefore can be estimated by analyzing data of temporal and spatial changes of the magnetic field measured in advance. It is. The magnetic field strength changes .DELTA.B f, elapsed sufficient time from the start of the magnetic field change process, after the magnetic field intensity changes .DELTA.B r and the magnetic field intensity changes .DELTA.B e is attenuated, averaging the magnetic field strength calculated from sufficiently many measurements Thus, it can be estimated by eliminating the influence of the magnetic field strength change ΔB c .
　磁場強度変化ΔＢｒについては、リップルの振動周波数を特定することによって推定することができる。さらに、磁場強度変化ΔＢｅは、磁場強度変化ΔＢｒに比べて減衰速度が十分に遅くかつ変化が大きいため、磁場測定を長時間行なうことによって推定することができる。 The magnetic field strength changes .DELTA.B r, it can be estimated by identifying the oscillation frequency of the ripple. Furthermore, the magnetic field strength change ΔB e can be estimated by performing the magnetic field measurement for a long time because the decay rate is sufficiently slow and the change is large compared to the magnetic field strength change ΔB r .
　なお、磁場強度変化ΔＢｃについては、駆動電流の不規則な変動によって生じるため再現性が低くランダムに変動する。しかしながら、上記のように空間的なパターンが一定であるため、磁場強度変化ΔＢｒを測定する際に空間パターンを予め推定しておくことで、リアルタイムで測定した駆動電流から推定することが可能である。 Note that the magnetic field strength changes .DELTA.B c, reproducible randomly fluctuating low order caused by random variations of the drive current. However, because of the spatial pattern as above is constant, by previously estimating the spatial pattern in measuring the magnetic field intensity change .DELTA.B r, it can be estimated from the drive current measured in real time is there.
　以上のように４つの磁場強度変化が推定されると、これらの磁場強度変化を、上述のシムコイル設計手法の際に示した主成分で分解することによって、複数のシムコイルによって４つの磁場強度変化を補正するのに必要な電流値Ｉｆ，Ｉｃ，Ｉｒ（ｔ），Ｉｅ（ｔ）の組合せを演算により求める。ここで電流値Ｉｆ，Ｉｃ，Ｉｒ（ｔ），Ｉｅ（ｔ）の各々は、シムコイルと同じ数の要素を含むベクトルである。たとえば、ｋ番目のシムコイルに１Ａの電流を流した時に発生する磁場分布をｂｋとした場合には、磁場強度変化ΔＢは、ΔＢ＝Ｉｋｂｋと表すことができる。 When the four magnetic field strength changes are estimated as described above, these magnetic field strength changes are decomposed by the main components shown in the above-described shim coil design method, so that the four magnetic field strength changes are generated by a plurality of shim coils. A combination of current values I f , I c , I r (t), I e (t) necessary for correction is obtained by calculation. Here, each of the current values I f , I c , I r (t), and I e (t) is a vector including the same number of elements as the shim coil. For example, when the magnetic field distribution generated when a current of 1 A is passed through the kth shim coil is b k , the magnetic field strength change ΔB can be expressed as ΔB = I k b k .
　上述のように、磁場強度変化ΔＢｆ、磁場強度変化ΔＢｒおよび磁場強度変化ΔＢｅについては、事前の測定に基づいて予め推定可能であるため、電流値Ｉｆ，Ｉｒ（ｔ），Ｉｅ（ｔ）についても予め算出することが可能である。電流値Ｉｃについては、電流センサによって計測した駆動電流から算出することができるので、事前に求めた電流値Ｉｆ，Ｉｒ（ｔ），Ｉｅ（ｔ）と、駆動電流から算出した電流値Ｉｃを合算することによって、各シムコイルに必要とされる電流を決定することができる。 As described above, since the magnetic field strength change ΔB f , the magnetic field strength change ΔB r, and the magnetic field strength change ΔB e can be estimated in advance based on prior measurements, the current values I f , I r (t), I It is possible to calculate e (t) in advance. Since the current value I c can be calculated from the drive current measured by the current sensor, the current values I f , I r (t), I e (t) obtained in advance and the current calculated from the drive current by summing the values I c, can be determined current required for each shim coil.
　図４は、上記の手順を表したフローチャートである。図４において、ステップＳ２００～Ｓ２４０については、設備導入の際の調整時等のタイミングで、予めオフラインで処理が行なわれ、得られた電流値Ｉｆ，Ｉｃ，Ｉｒ（ｔ），Ｉｅ（ｔ）は制御装置に含まれる記憶部に記憶される。ステップＳ２５０については、記憶された電流値Ｉｆ，Ｉｒ（ｔ），Ｉｅ（ｔ）と電流センサで計測される駆動電流から定まる電流値Ｉｃとに基づいて、オンラインにて処理される。 FIG. 4 is a flowchart showing the above procedure. In FIG. 4, steps S200 to S240 are previously processed off-line in advance at the time of adjustment at the time of equipment introduction, and the obtained current values I f , I c , I r (t), I e (T) is stored in a storage unit included in the control device. Step S250 is processed online based on the stored current values I f , I r (t), I e (t) and the current value I c determined from the drive current measured by the current sensor. .
　ステップＳ２００において、磁場強度ｂを変更したときの磁場強度の時間的変化Ｂｂ（ｔ）および駆動電流の時間的変化Ｉｃ（ｔ）を測定し、その測定データを制御装置内の記憶部に記憶する。 In step S200, the temporal change B b (t) of the magnetic field strength and the temporal change I c (t) of the drive current when the magnetic field strength b is changed are measured, and the measurement data is stored in the storage unit in the control device. Remember.
　次に、記憶された磁場強度の時間的変化Ｂｂ（ｔ）のデータをフーリエ変換することによって、リップルノイズに起因する磁場強度変化ΔＢｒを推定する（ステップＳ２１０）。 Next, the stored magnetic field strength temporal change B b (t) data is Fourier-transformed to estimate the magnetic field strength change ΔB r caused by ripple noise (step S210).
　ステップＳ２２０においては、記憶された磁場強度の時間的変化Ｂｂ（ｔ）のデータからステップＳ２１０で推定した磁場強度変化ΔＢｒ（ｔ）を分離したデータに平滑化処理を施し、収束後の磁場分布を磁場強度変化ΔＢｆと推定するとともに、収束までの変動を磁場強度変化ΔＢe（ｔ）として推定する。 In step S220, the data obtained by separating the magnetic field intensity change ΔB r (t) estimated in step S210 from the stored data of the temporal change B b (t) of the magnetic field intensity is subjected to smoothing processing, and the magnetic field after convergence is obtained. The distribution is estimated as the magnetic field strength change ΔB f and the fluctuation until convergence is estimated as the magnetic field strength change ΔB e (t).
　そして、記憶された磁場強度の時間的変化Ｂｂ（ｔ）から、ステップＳ２１０，Ｓ２２０で推定された磁場強度変化ΔＢｒ（ｔ），ΔＢｆ，ΔＢe（ｔ）を分離したデータから、記憶された駆動電流の時間的変化Ｉｃ（ｔ）に対応する変動を磁場強度変化ΔＢｃとして分離する（ステップＳ２３０）。 Then, from the data obtained by separating the magnetic field intensity changes ΔB r (t), ΔB f , ΔB e (t) estimated in steps S210 and S220 from the temporal change B b (t) of the stored magnetic field intensity, The fluctuation corresponding to the temporal change I c (t) of the drive current is separated as the magnetic field strength change ΔB c (step S230).
　ステップＳ２４０にて、磁場強度変化ΔＢｒ（ｔ），ΔＢｆ，ΔＢe（ｔ），ΔＢｃの各々を、シムコイル設計時の主成分で分解することで、電流値Ｉｒ（ｔ），Ｉｆ，Ｉｅ（ｔ），Ｉｃを算出し、得られた電流値を記憶部に記憶する。ここまでのステップは、オフラインで実施される。 In step S240, each of the magnetic field strength changes ΔB r (t), ΔB f , ΔB e (t), and ΔB c is decomposed with the main components at the time of designing the shim coil, whereby current values I r (t), I f 1 , I e (t), and I c are calculated, and the obtained current value is stored in the storage unit. The steps so far are performed offline.
　そして、実際にＭＲＩ装置を運転する場合には、記憶された電流値Ｉｒ（ｔ），Ｉｆ，Ｉｅ（ｔ）と、電流センサで計測された実際に超電導マグネットに供給される駆動電流Ｉｃとを合算することによって、各シムコイルに対して供給すべきトータル電流を算出し、算出した電流を各シムコイルに供給する。 When the MRI apparatus is actually operated, the stored current values I r (t), I f , I e (t) and the drive current actually supplied to the superconducting magnet measured by the current sensor. by summing the I c, to calculate the total current to be supplied to each shim, and supplies the calculated current to each shim coil.
　以上のような処理を行なうことによって、設計したシムコイルに対して適切な電流を供給することが可能となる。 By performing the above processing, it is possible to supply an appropriate current to the designed shim coil.
　今回開示された実施の形態はすべての点で例示であって制限的なものではないと考えられるべきである。本発明の範囲は上記した説明ではなくて請求の範囲によって示され、請求の範囲と均等の意味および範囲内でのすべての変更が含まれることが意図される。 The embodiment disclosed this time should be considered as illustrative in all points and not restrictive. The scope of the present invention is defined by the terms of the claims, rather than the description above, and is intended to include any modifications within the scope and meaning equivalent to the terms of the claims.
　１０　ＭＲＩ装置、１００　本体装置、１１０　超電導マグネット、１１５　高温超電導コイル、１２０　傾斜磁場コイル、１３０　ＲＦ送信コイル、１４０　ＲＦ受信コイル、１５０　シムコイル、１６０　検査用テーブル、１７０　被検者、２００　直流電源装置、２１０　電流センサ、３００　制御装置、３１０　表示部、３２０　入力部。 10 MRI apparatus, 100 main unit, 110 superconducting magnet, 115 high temperature superconducting coil, 120 gradient magnetic field coil, 130 RF transmitting coil, 140 RF receiving coil, 150 shim coil, 160 testing table, 170 subject, 200 DC power supply, 210 current sensor, 300 control device, 310 display unit, 320 input unit.

Claims (5)

  1. 　複数の磁場強度を設定することが可能な磁気共鳴イメージング装置であって、
    　高温超電導材を含み、被検体が曝露する静磁場を発生するための超電導コイルと、
    　前記超電導コイルに直流電流を供給する直流電源装置と、
    　前記超電導コイルにより生成された磁場分布を補正可能なシムコイルとを備え、
    　前記直流電源装置は、前記超電導コイルへ供給する出力電流の大きさを調整することによって、複数の磁場強度を設定することが可能に構成され、
    　前記磁場分布は、前記複数の磁場強度の各々に対応した電流を前記シムコイルに供給することによって補正される、磁気共鳴イメージング装置。     A magnetic resonance imaging apparatus capable of setting a plurality of magnetic field strengths,
    A superconducting coil containing a high temperature superconducting material for generating a static magnetic field to which the subject is exposed;
    A DC power supply for supplying a DC current to the superconducting coil;
    A shim coil capable of correcting the magnetic field distribution generated by the superconducting coil,
    The DC power supply device is configured to be able to set a plurality of magnetic field strengths by adjusting the magnitude of the output current supplied to the superconducting coil,
    The magnetic resonance imaging apparatus, wherein the magnetic field distribution is corrected by supplying a current corresponding to each of the plurality of magnetic field strengths to the shim coil.
  2. 　所定周波数の励磁パルスを発生させるための送信コイルと、
    　被検体からの磁気共鳴信号を受信するための受信コイルをさらに備え、
    　前記複数の磁場強度は、第１の磁場強度および第２の磁場強度を含み、
    　前記送信コイルは、前記第１の磁場強度に対応する第１の周波数、および前記第２の磁場強度に対応する第２の周波数で共振するように構成される、請求項１に記載の磁気共鳴イメージング装置。     A transmission coil for generating an excitation pulse of a predetermined frequency;
    A receiving coil for receiving a magnetic resonance signal from the subject;
    The plurality of magnetic field strengths include a first magnetic field strength and a second magnetic field strength,
    The magnetic resonance of claim 1, wherein the transmit coil is configured to resonate at a first frequency corresponding to the first magnetic field strength and a second frequency corresponding to the second magnetic field strength. Imaging device.
  3. 　前記第１および第２の磁場強度は、それぞれ１．５Ｔおよび３Ｔであり、
    　前記第１の周波数は６４ＭＨｚ、前記第２の周波数は１２８ＭＨｚである、請求項２に記載の磁気共鳴イメージング装置。     The first and second magnetic field strengths are 1.5T and 3T, respectively.
    The magnetic resonance imaging apparatus according to claim 2, wherein the first frequency is 64 MHz and the second frequency is 128 MHz.
  4. 　前記シムコイルに供給される電流は、前記複数の磁場強度の各々における磁場強度変化の主成分分析に基づいて決定される、請求項１～３のいずれか１項に記載の磁気共鳴イメージング装置。 The magnetic resonance imaging apparatus according to any one of claims 1 to 3, wherein the current supplied to the shim coil is determined based on a principal component analysis of a magnetic field strength change in each of the plurality of magnetic field strengths.
  5. 　前記超電導コイルに供給される駆動電流を計測するための電流計測装置をさらに備え、
    　前記直流電源装置は、前記電流計測装置によって計測される駆動電流の変動に基づいて、前記出力電流を調整する、請求項１～４のいずれか１項に記載の磁気共鳴イメージング装置。     A current measuring device for measuring a drive current supplied to the superconducting coil;
    The magnetic resonance imaging apparatus according to any one of claims 1 to 4, wherein the DC power supply device adjusts the output current based on a variation in driving current measured by the current measuring device.
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Publication number Priority date Publication date Assignee Title
JPH01236043A (en) * 1988-03-17 1989-09-20 Hitachi Medical Corp Nuclear magnetic resonance imaging device
JPH07148142A (en) * 1993-09-30 1995-06-13 Siemens Ag Seaming method for magnetic field in space to be inspected of nuclear spin resonance device
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