WO2015072301A1

WO2015072301A1 - Magnetic resonance imaging apparatus

Info

Publication number: WO2015072301A1
Application number: PCT/JP2014/078184
Authority: WO
Inventors: 津田　宗孝
Original assignee: 株式会社日立メディコ
Priority date: 2013-11-12
Filing date: 2014-10-23
Publication date: 2015-05-21
Also published as: JPWO2015072301A1

Abstract

The purpose of the present invention is to stabilize the magnetic field strength of the superconductive magnet in driven mode operation and achieve improvement of MRI image quality. A magnetism locking sample (122) is disposed at a position where a magnetic field generated by the superconductive coil (105) is applied and which does not physically interfere with the subject (101). A magnetic field locking unit (123) irradiates a high frequency magnetic field on the magnetic field locking sample (122), detects the NMR signal (B) generated by the magnetic field locking sample (122), and adjusts the strength of the static magnetic field according to the detection results to make the static magnetic field constant. Preferably the magnetic field locking sample comprises an atomic nuclide that generates a nuclear magnetic resonance signal (B) of a magnetic resonance frequency differing from the magnetic resonance frequency of the nuclear magnetic resonance signal (A) detected by the measurement unit.

Description

Magnetic resonance imaging system

The present invention relates to a magnetic resonance imaging apparatus (Magnetic-Resonance-Imaging apparatus, hereinafter referred to as an MRI apparatus), and in particular, an MRI apparatus using a driven-mode superconducting magnet that is operated by connecting a power source for applying a current to a superconducting coil. Relates to stabilization of magnetic field strength.

There are two types of MRI devices in medical facilities: a device with a magnetic field strength of 0.5 Tesla or less using a permanent magnet and a device with a magnetic field strength of 1.0 Tesla or more using a superconducting magnet. Permanent magnet MRI systems have the advantages of low installation costs and low operating costs. On the other hand, the superconducting magnet cooling the superconducting wire, such as NbTi and Nb ₃ Sn below the critical temperature, etc. have to be installed in the service area of the liquid helium to be cryogen, but many restrictions on its installation, the magnetic field strength Since the detection sensitivity of the nuclear magnetic resonance (hereinafter referred to as NMR) signal increases in proportion, it has a high diagnostic ability.

As shown in Patent Document 1, a superconducting magnet is generally operated in a mode in which a superconducting coil is connected in a loop and a permanent current flows. However, as shown in Patent Document 2, in order to obtain a higher magnetic field strength and to use a high-temperature superconducting wire that does not use liquid helium, the superconductivity of a driven mode NMR apparatus or MRI apparatus in which a current is constantly supplied from a power source. Magnets have also been developed.

However, as shown in Patent Document 2, the stability of the magnetic field strength of the driven mode superconducting magnet is about 5 ppm per hour due to the temperature drift of the power source.
Therefore, in the technique disclosed in Patent Document 2, the magnetic field measuring instrument probe is arranged in the magnetic field space, and the main magnetic field is induced so as to induce a magnetic field in the opposite direction to the fluctuation of the magnetic field of the main coil measured by the magnetic field measuring instrument probe. The coil current is adjusted to achieve a stable magnetic field strength of 0.01 ppm per hour.

JP 2005-124721 JP JP 2008-020266 A

In the MRI system, the subject is placed in a uniform magnetic field generated by a superconducting magnet. In addition, due to restrictions on the size of the superconducting magnet and the manufacturing cost, the uniform magnetic field space is manufactured so as to substantially match the size of the subject. In other words, the uniform magnetic field space of 3 ppm or less required for MRI examination is a spherical space with a diameter of about 40 centimeters. Outside this space, the magnetic field uniformity rapidly deteriorates, that is, has a magnetic field strength gradient in which the magnetic flux density changes.

In addition, in the MRI apparatus, in order to select the examination site of the subject and to give position information to the NMR signal, the magnetic field strength is adjusted along the three axes orthogonal to each other, that is, the x, y, and z axes of the magnetic field space. A gradient magnetic field is applied so as to intentionally generate a gradient. That is, since a gradient magnetic field is applied during the MRI examination, the magnetic field strength of the uniform magnetic field space generated by the superconducting magnet changes along the x, y, and z axes.

For this reason, in order to stabilize the magnetic field strength of the superconducting magnet in the driven mode operation, the technique of Patent Document 2 that compensates for the drift of the power source with the signal measured by the magnetic field measuring instrument probe is adopted for the superconducting magnet of the MRI apparatus. In doing so, the following problems arise.

(1) It is difficult to place the subject and the magnetic field measuring probe simultaneously in the uniform magnetic field space where the subject is placed. For this reason, it is not easy to measure the magnetic field intensity with the magnetic field measurement probe while detecting the NMR signal of the subject.

(2) Since the magnetic field in the uniform magnetic field space is changed by the gradient magnetic field operated during the MRI examination, it is difficult to accurately measure the uniform magnetic field strength with the magnetic field measurement probe.

(3) The high-frequency signal detected by the magnetic field measurement probe interferes with the high-frequency signal detected from the subject of the MRI examination.

(4) It is difficult for the operator to grasp whether the magnetic field compensation operates and the magnetic field strength is accurate.

In Patent Document 2, there is no description regarding the positional relationship between the subject and the magnetic field measuring probe, the interference of the detection signal, and the like, and it is assumed that the magnetic field is measured in a state where the subject is not arranged. For this reason, the above problem cannot be solved by the technique of Patent Document 2.

The present invention has been made in view of the above problems, and an object thereof is to stabilize the magnetic field strength of the superconducting magnet in the driven mode operation and achieve the improvement of the image quality of the MRI image.

In order to solve the above problems, the MRI apparatus of the present invention includes a superconducting coil that forms a static magnetic field in an imaging space, a magnet power supply that continuously supplies current to the superconducting coil while the superconducting coil forms a static magnetic field, A gradient magnetic field generator for generating a gradient magnetic field in the imaging space; a high-frequency magnetic field generator for irradiating a subject placed in the imaging space with a high-frequency magnetic field; a measurement unit for detecting a nuclear magnetic resonance signal (A) of the subject; A sample for magnetic lock and a magnetic lock part are provided. The sample for magnetic lock is disposed at a position where the magnetic field generated by the superconducting coil 105 is applied and does not physically interfere with the subject. The magnetic field locking unit irradiates the magnetic field locking sample with a high frequency magnetic field, detects the nuclear magnetic resonance signal (B) generated by the magnetic field locking sample, adjusts the strength of the static magnetic field according to the detection result, Keep it constant.

According to the present invention, it is possible to stabilize the magnetic field strength of the driven mode superconducting magnet and to improve the image quality of the MRI image.

The block diagram which shows the whole structure of the MRI apparatus of this embodiment 1 is a block diagram showing the circuit configuration of the measurement system high-frequency unit 112 and the magnetic field lock system high-frequency unit 123 of the apparatus of FIG. Sectional drawing of the superconducting magnet which comprises the MRI apparatus of this embodiment Sectional drawing which shows the structure which has arrange | positioned the sample for a magnetic field lock in several places in a top plate Sectional drawing which shows the structure which has arrange | positioned the sample for magnetic field lock | rock in the holding | maintenance part of a high frequency receiver coil Side view of sample for magnetic field lock 2A is an explanatory diagram showing input / output signals, and FIG. 2B is an explanatory diagram showing waveforms of input / output signals. Timing chart showing operation of imaging pulse sequence and magnetic field lock of embodiment Flow chart showing the operation of the MRI apparatus of the first embodiment Flow chart showing the operation of the MRI apparatus of the second embodiment

The MRI apparatus of the present invention, as shown in FIGS. 1 to 3, continues the current to the superconducting coil 105 while the superconducting coil 105 forms a static magnetic field in the imaging space 102 and the superconducting coil 105 forms a static magnetic field. Magnet power supply 106, a gradient magnetic field generator (109, 110) that generates a gradient magnetic field in the imaging space 102, and a high-frequency magnetic field generator (111, 110) that irradiates the subject 101 disposed in the imaging space 102 with a high-frequency magnetic field 112) and measuring units (112, 113) for detecting a nuclear magnetic resonance (NMR) signal (A) of the subject 101. Furthermore, the MRI apparatus of the present invention includes a magnetic lock sample 122 and a magnetic lock portion (123).

The magnetic locking sample 122 is disposed at a position where the magnetic field generated by the superconducting coil 105 is applied and does not physically interfere with the subject 101. The magnetic field locking unit (123) irradiates the magnetic field locking sample 122 with a high frequency magnetic field, detects the NMR signal (B) generated by the magnetic field locking sample 122, adjusts the strength of the static magnetic field according to the detection result, Make the static magnetic field constant.

As described above, in the present invention, the magnetic field locking sample 122 is arranged at a position not interfering with the subject 101, and the magnetic field locking unit (123) adjusts the static magnetic field intensity based on the NMR signal (B) of the magnetic field locking sample 122. By doing so, it is possible to adjust the static magnetic field uniformity while measuring the NMR signal (A) of the subject 101. As a result, the magnetic field strength in the imaging space of the superconducting magnet in the driven mode operation where the superconducting coil is always supplied with current from the magnet power source during the generation of the magnetic field can be maintained in a uniform state with high accuracy. Can be improved.

As the magnetic field locking sample 122, it is desirable to use a sample containing an atomic species that generates an NMR signal (B) having a magnetic resonance frequency different from the magnetic resonance frequency of the NMR signal (A) detected by the measurement unit (112). Thereby, electrical mutual interference between the NMR signal (A) and the NMR signal (B) can be prevented and simultaneously detected. For example, when the NMR signal (A) detected by the measurement unit (112) is an NMR signal of a hydrogen nucleus (proton) ¹ H, the magnetic field locking sample 122 includes a fluorine nucleus ¹⁹ F, a deuterium nucleus ² H, and A material containing any of the phosphorus nuclei ¹⁵ P can be used.

The magnetic field lock unit (123) temporarily stops adjusting the strength of the static magnetic field according to the detection result of the NMR signal (B) while the gradient magnetic field generation unit (109, 110) is generating the gradient magnetic field. It is desirable to do. This prevents the magnetic field lock unit (123) from detecting the static magnetic field in which the gradient magnetic field is superimposed and adjusting the static magnetic field, thereby preventing the static magnetic field from being balanced without being affected by the gradient magnetic field. Can be improved once. For example, the magnetic field lock unit (123) maintains the adjustment amount immediately before the stop while the gradient magnetic field generation unit (109, 110) generates the gradient magnetic field.

The high-frequency magnetic field generator (111, 112) includes, for example, a high-frequency transmitter coil 111 and a high-frequency signal generator (112) that supplies a high-frequency signal for causing the high-frequency transmitter coil 111 to generate a high-frequency magnetic field. The magnetic field lock unit (123) supplies a high-frequency signal for generating a high-frequency magnetic field to the high-frequency transmitter coil (703) for magnetic field locking that irradiates the magnetic field locking sample 122 with high frequency, and the high-frequency transmitter coil (703) for magnetic field locking. And a magnetic field locking high-frequency signal generator (123). It is desirable that a reference frequency generator 201 for supplying a common reference frequency signal is connected to the high-frequency signal generation unit (112) and the magnetic field locking high-frequency signal generation unit (123).

As a result, even when the frequency of the reference frequency signal fluctuates, the correlation between the high-frequency magnetic field of the high-frequency transmitter coil 111 and the high-frequency magnetic field of the high-frequency transmitter coil for magnetic field locking (703) can be maintained. (123) can accurately detect the fluctuation of the static magnetic field.

For example, as shown in FIG. 3, the magnetic field locking sample 122 is always arranged outside the imaging space 102 and at a position where it does not physically interfere with the measurement units (112, 113) and the table 118 on which the subject 101 is mounted. Can be configured. Alternatively, as shown in FIG. 4, the magnetic field locking sample 122 is arranged at one or more places in the top plate 119 on which the subject 101 of the table 118 is mounted, and is inserted into the imaging space 102 together with the subject 101. It may be.

As shown in FIG. 4, when the magnetic field locking samples 122 are arranged at a plurality of locations on the top plate 119 of the table 118, the magnetic field locking unit (123) selects any one of the plurality of magnetic field locking samples 122. , And can be used for detection of NMR signal (B).

Further, as shown in FIG. 5, the magnetic field locking sample 122 can be arranged in the holding unit (801) that holds the high-frequency receiver coil 113 of the measurement unit (112, 113).

The magnetic field locking sample 122 is disposed in a container (701) (see, for example, FIG. 6) having a size that is not affected by the intensity gradient of the magnetic field generated by the superconducting coil 105 at the position where the magnetic field locking sample 122 is disposed. desirable.

The magnetic field lock unit (123), for example, as shown in FIGS. 2 and 7, uses the signal having the same frequency as the magnetic resonance frequency of the magnetic field locking sample 122 as a reference signal 302, and outputs the NMR signal (B) (301). A configuration including a phase detection unit (213) for phase detection may be employed.

In this case, the magnetic field lock unit (123) adjusts the magnetic field strength of the imaging space 102 so that the phase detection output (303) becomes zero (phase lock state). In this case, the magnetic field lock unit (123) preferably has a display unit (217) that notifies the operator that the output of the phase detection unit (213) has become zero (phase lock state). The measurement unit (112) preferably measures the NMR signal (A) of the subject 101 in a state where the output of the phase detection unit (213) becomes zero (phase lock state).

The magnetic field lock unit (123) can be configured to adjust the strength of the static magnetic field by changing the output current of the magnet power source 106 in accordance with the detection result of the NMR signal (B).

In addition, the magnetic field lock unit (123) adjusts the strength of the static magnetic field by changing the current supplied to the correction coil 215 that generates the correction magnetic field in the imaging space 102 according to the detection result of the NMR signal (B). May be.

Hereinafter, embodiments of the present invention will be specifically described with reference to the accompanying drawings. Note that components having the same function are denoted by the same reference symbols throughout the drawings for describing the embodiments of the invention, and the repetitive description thereof is omitted.

(First embodiment)
<Overall configuration of MRI system>
FIG. 1 is a block diagram showing the overall configuration of the MRI apparatus of the present embodiment, FIG. 2 is a block diagram showing the circuit configuration, and FIG. 3 is a cross-sectional view of the superconducting magnet 104. FIGS. 1 to 3 all illustrate an MRI apparatus installed in a medical facility and taking a medical diagnostic image of a patient's head, which is a subject 101.

A superconducting magnet 103 that generates a uniform static magnetic field in the imaging space 102 includes an iron yoke 104 having two magnetic poles serving as NS poles, a pair of superconducting coils 105, and a magnet power source 106. Further, a correction coil 215 that generates a correction magnetic field that corrects the static magnetic field generated by the superconducting magnet 103 is disposed between the superconducting magnet 103 and the imaging space 102. The examination region of the subject 101 is disposed at the center of the imaging space 102 where a uniform static magnetic field is generated.

As shown in FIG. 3, the superconducting magnet 103 is further provided with a vacuum vessel 107 containing the superconducting coil 105 and a refrigerator 108 for keeping the superconducting coil 105 at a low temperature, in addition to the iron yoke 104. The iron yoke 104 weighs 14 tons and has a C-shaped cross-sectional shape with a part being an opening. For example, the shape is determined so as to secure a magnetic flux density that generates a magnetic field strength of 0.5 Tesla at an opening of 55 centimeters and to minimize the magnetic flux leaking out of the iron yoke 104.

Also, the opening of the iron yoke 104 has a pair of magnetic poles 601 processed into a concave surface in order to generate a uniform magnetic field. Around the magnetic pole 601, a doughnut-shaped vacuum vessel 107 in which a pair of superconducting coils 105 is incorporated is incorporated. The iron yoke 104 has a function of supporting the vacuum vessel 107.

With such a magnet configuration, the front (y-axis) and the left and right sides (x-axis) of the imaging space 102 have nothing to obstruct the field of view and can provide an open inspection environment.

The superconducting coil 105 is thermally connected to the refrigerator 108 by a heat conducting member and is cooled to a temperature of 20 Kelvin to maintain a stable superconducting state. For example, a current of 160 amperes is applied to the superconducting coil 105 from the magnet power source 106, and the imaging space 102 has, for example, a z-axis (for example, 0.5 z Tesla intensity). Has become a common practice). The imaging space 102 is a spherical space with a diameter of 40 cm, for example, and is also called FOV (FieldFOof View). The magnetic field uniformity of the imaging space is 3 ppm or less when the magnetic field lock unit (123) is not operating.

The magnet power supply 106 also incorporates a compressor that supplies high-pressure helium gas to the refrigerator 108 and a sensor circuit that monitors the operating state of the superconducting magnet 103.

The gradient magnetic field coil assembly 109 is attached to the two magnetic poles 601, and a gradient magnetic field having a gradient in magnetic field strength is generated in three axial directions orthogonal to each other in the imaging space 102. Although not distinguished in FIGS. 1 and 3, the gradient coil assembly 109 has three types of coils, x, y, and z, laminated thereon.

For example, when a positive current flows through the z gradient magnetic field coil, the z gradient magnetic field coil attached to the upper magnetic pole generates a magnetic flux in the same + z axis direction as the magnetic flux generated by the superconducting coil 105, and superimposes its density. Increase. On the other hand, the z gradient magnetic field coil attached to the lower magnetic pole generates a magnetic flux along the −z axis in the direction opposite to the magnetic flux generated by the superconducting coil 105, and reduces its density. As a result, a gradient magnetic field in which the magnetic flux density increases from bottom to top along the z-axis of the imaging space 102 can be created. The x gradient magnetic field coil changes the magnetic flux density generated by the superconducting coil 105 along the x axis of the imaging space 102, and the y gradient magnetic field coil changes along the y axis of the imaging space 102. Gradient magnetic field power supply 110 that operates independently is connected to each of the gradient magnetic field coils of x, y, and z. For example, by supplying a current of 500 amperes to each, the magnetic field strength of 25 millitesla changes in 1 meter. A gradient magnetic field of 25 mT / m can be generated.

Furthermore, a pair of high-frequency transmitter coils 111 are incorporated on the imaging space 102 side of the gradient coil assembly 109. The high-frequency transmitter coil 111 has a flat plate structure so as not to hinder an open inspection environment, and a coil conductor is printed and wired so that a magnetic flux parallel to the xy plane of the imaging space 102 is generated. A plurality of capacitive elements are incorporated in the high-frequency transmitter coil 111 (not shown in the figure), and here is a 21.28 MHz LC resonance circuit. A high frequency magnetic field is generated in the imaging space 102 by flowing a high frequency current of 21.28 MHz from a high frequency power source incorporated in the measurement system high frequency unit 112.

By combining the static magnetic field, the gradient magnetic field, and the high-frequency magnetic field described above, an NMR phenomenon occurs in the hydrogen nucleus (proton) at a specific part of the subject 101, and x, y in the process of Larmor precession of the hydrogen nucleus (proton) The spatial information is given by the z gradient magnetic field pulse.

Thus, in order to detect the Larmor precession of the hydrogen nucleus (proton) to which the spatial information is given as the NMR signal (A), a high-frequency receiver coil 113 is attached to the examination site of the subject 101. Like the high-frequency transmitter coil 111, the high-frequency receiver coil 113 is an LC resonance circuit in which a capacitive element is incorporated (not shown in the figure) and resonates at 21.28 MHz. The difference from the high-frequency transmitter coil 111 is that it has a shape that fits the body shape of the examination site so as to detect the Larmor precession of the nuclear spin as an electrical signal by electromagnetic induction with high efficiency. FIGS. 1 to 3 show a high-frequency receiver coil 113 that detects the head of the subject 101.

The NMR signal (A) detected by the high-frequency receiver coil 113 is amplified, detected, and converted from analog to digital by an amplifier in the measurement system high-frequency unit 112, and passes through an interface circuit called a sequencer 114 to the computer. Recorded at 115. The high-frequency receiver coil 113 and the measurement system high-frequency unit 112 constitute a measurement unit.

In the computer 115, the NMR signal is subjected to arithmetic processing such as Fourier transform, and processed into a tomographic image and a spectrum distribution map effective for medical diagnosis. These data are stored in a storage device (not shown) of the computer 115 and displayed on the display 116. On the other hand, the computer 115 is operated via the sequencer 114 to operate the gradient magnetic field power source 110 and the measurement system high-frequency unit 112 in accordance with various pulse sequences so that a tomographic image or the like for diagnosis can be obtained from the examination site of the subject 101. Control function. For this reason, an input device 117 for selecting the type of pulse sequence and the like by the operator of the MRI apparatus is connected to the computer 115.

In addition, a subject 101 is mounted in front of the superconducting magnet 103, and a table 119 for loading and unloading the inspection site to the center of the imaging space 102, and a table for moving the table 119 in the loading and unloading direction, etc. 118 is arranged. The superconducting magnet 103 and the table 118 are installed in an examination room 120 that is shielded from electromagnetic waves.

Outside the examination room 120, there are a gradient magnetic field power supply 110, a measurement system high frequency unit 112, a magnetic field lock system high frequency unit 123, a magnet power supply 106, a sequencer 114, a computer 115, an input device 117, and a display device 116. Is arranged. The gradient magnetic field power source 110, the measurement system high frequency unit 112, the magnetic field lock system high frequency unit 123, and the magnet power source 106 are connected to the gradient magnetic field coil assembly 109, the high frequency transmitter coil 111 / It is connected to the receiver coil 113, the magnetic field locking high-frequency transmitter coil (703), and the superconducting coil 105. This prevents electromagnetic waves generated by the computer 115 and other devices from entering the high frequency receiver coil 113 as noise.

<Magnetic field lock sample>
In this embodiment, the magnetic field locking sample 122 is disposed at a position where the magnetic field generated by the superconducting coil 105 is applied and does not physically interfere with the subject 101, and the NMR signal (B) is converted into the magnetic field. It is detected by the lock unit (123).

Here, a fluorine compound (for example, CFCl ₃ solution or liquefied CFC) containing fluorine nuclei ¹⁹ F is used as the magnetic field locking sample 122. As shown in FIG. 6, the magnetic field locking sample 122 is enclosed in a spherical microcapillary 701 having an inner diameter of 5 mm, for example, and the microcapillary 701 is sealed. The size of the microcapillary 701 is designed such that the magnetic field locking sample 122 is not affected by the strength gradient of the magnetic field generated by the superconducting coil 105 (the magnetic field strength distribution can be ignored).

For example, as shown in FIG. 3, a magnetic lock sample 122 is disposed on the outer peripheral side surface of the high-frequency transmitter coil 111 under the top plate 119. In this case, the microcapillary 701 is about 5 mm in diameter. With this size, the magnetic field strength distribution of the superconducting coil 105 in the space below the top plate 119 can be ignored.

On the surface of the microcapillary 701, a solenoid coil 703 that generates a high-frequency magnetic field that excites the NMR phenomenon of the magnetic field locking sample and detects the generated NMR signal (B) is wound. That is, the solenoid coil 703 serves both as a magnetic field locking high-frequency transmitter coil and a magnetic field locking high-frequency receiver coil. The shaft of the solenoid coil 703 is incorporated and held in a holder 704 formed of Teflon (registered trademark) resin so that the axis of the solenoid coil 703 is orthogonal to the z-axis of the imaging space 102. The holder 704 is fixed to the lower part of the top plate 119 as shown in FIG. 3, for example.

The position of the magnetic field locking sample 122 shown in FIG. 3 shows a magnetic field strength higher than the magnetic field strength of the imaging space 102 due to the influence of the magnetic pole 601. For example, when the magnetic field strength of the imaging space 102 is 0.5 Tesla, the position of the magnetic field locking sample 122 under the top plate 119 is 0.51 Tesla. A high frequency magnetic field of 20.42 MHz, which is the magnetic resonance frequency of the magnetic field locking sample (fluorine nucleus ¹⁹ F) 122 to which a static magnetic field of 0.51 Tesla is applied, is irradiated from the solenoid coil 703. That is, a solenoid coil 703 having a resonance frequency of 20.42 MHz is used, a high-frequency current having a frequency of 20.42 MHz is supplied to the solenoid coil 703 from a magnetic field lock transmitter 209 described later, and a high-frequency magnetic field having a magnetic resonance frequency of the magnetic-field locking sample 122 is solenoidally supplied. Irradiate from coil 703. Since the static magnetic field strength of the superconducting coil 105 differs depending on the position where the magnetic field locking sample 122 is disposed, the high frequency magnetic field having a magnetic resonance frequency corresponding to the position is irradiated from the solenoid coil 703 to the magnetic field locking sample 122.

<Circuit configuration of measurement unit and magnetic field lock unit>
The MRI apparatus of the present embodiment includes a measurement unit (measurement system high frequency unit) 112 for generating a high frequency magnetic field from the high frequency transmitter coil 111 toward the subject 101 and measuring the NMR signal (A), and a solenoid coil 703. A magnetic field lock unit (magnetic field lock system high frequency unit) 123 for generating a high frequency magnetic field from the magnetic field to the magnetic field locking sample 122, detecting the NMR signal (B), and adjusting the static magnetic field is provided.

These circuit configurations and functions will be described in more detail with reference to FIG.

As shown in FIG. 2, the measurement system high-frequency unit 112 includes a measurement system transmitter 202, a high-frequency power amplifier 203 with a gate function, a high-frequency amplifier 204, a detector 205, an audio amplifier 206, and an A / D conversion circuit 207. Is included. On the other hand, the magnetic field lock system high frequency unit 123 includes a magnetic field lock system transmitter 209, a high frequency switch 210, a high frequency power amplifier 211, a preamplifier 212, a phase detector (PSD) 213, and an integral type DC amplifier 214. And a voltage comparison amplifier 216.

A reference frequency generator 201 is commonly connected to the measurement system transmitter 202 of the measurement system high frequency unit 112 and the magnetic field lock system transmitter 209 of the magnetic field lock system high frequency unit 123. The reference frequency transmitter 201 generates a 10 MHz reference frequency signal by a crystal resonator. The crystal resonator used is subjected to a long-term aging treatment, and the change with time can be ignored. It is housed in a temperature-compensated container and has a stability of the order of ^10-8 .

The operation of the measurement system high frequency unit 112 will be described. The measurement system transmitter 202 generates a frequency of 21.28 MHz at which a hydrogen nucleus (proton) ¹ H causes an NMR phenomenon with a magnetic field intensity of 0.5 Tesla from a reference frequency signal of 10 megahertz. The high frequency signal of 21.28 MHz is amplified by a high frequency power amplifier 203 with a gate function, modulated into a high frequency pulse, and applied to the high frequency transmitter coil 111. The high-frequency transmitter coil 111 irradiates the subject 101 with a high-frequency magnetic field having a frequency of 21.28 MHz. Thereby, an NMR signal (A) of the hydrogen nucleus (proton) of the subject 101 is generated.

The NMR signal (A) is detected by the high frequency receiver coil 113 and amplified to about 60 dB by the high frequency amplifier 204. The amplified NMR signal is detected by the detector 205 using the 21.28 MHz high frequency signal synthesized by the measurement system transmitter 202 as a reference signal, and changed to an audible frequency band. The NMR signal (A) converted to the audible frequency is amplified by the audio amplifier 206, converted into a digital signal by the A / D conversion circuit 207, and delivered to the computer 115 via the sequencer 114.

The sequencer 114 controls the irradiation timing of the high-frequency magnetic field from the high-frequency transmitter coil 111 and the measurement timing and intensity of the NMR signal (A) by the high-frequency receiver coil 113 according to the imaging pulse sequence selected by the operator.

Further, the sequencer 114 inputs the gradient signal of the analog signal to the gradient magnetic field power supply 110 of x, y, z via the D / A conversion circuit 208 in accordance with the imaging pulse sequence, so that the gradient coil assembly 109 Gradient magnetic fields in the x, y, and z directions are applied to the subject 101 at a predetermined timing and intensity. As a result, the sequencer 114 executes a desired imaging pulse sequence, and the obtained NMR signal (A) is reconstructed by the computer 115 to generate a tomographic image or the like.

Next, the operation of the magnetic field lock unit (magnetic field lock type high frequency unit) 123 will be described. The magnetic field lock system transmitter 209 generates a high frequency signal of 20.42 MHz in which the fluorine nucleus ¹⁹ F undergoes nuclear magnetic resonance with a magnetic field of 0.51 Tesla from the 10 MHz reference frequency signal of the reference frequency generator 201. Thus, by generating the high frequency signal 21.28 MHz of the hydrogen nucleus (proton) ¹ H and the high frequency signal 20.42 MHz of the fluorine nucleus ¹⁹ F from the same reference frequency signal, the reference frequency signal drifts due to temperature or the like. Even in this case, a complete correlation is maintained.

The high frequency signal of 20.42 MHz generated by the magnetic field lock system transmitter 209 is amplified by the high frequency power amplifier 211 via the high frequency switch 210 (the function of the circuit 210 of this high frequency switch will be described later), and the top plate 119 This is applied to the solenoid coil 703 wound around the magnetic field locking sample 122 arranged below. The solenoid coil 703 irradiates the magnetic field locking sample (fluorine nucleus ¹⁹ F) 122 with a high frequency magnetic field of 20.42 MHz. The magnetic field locking sample 122 generates a 20.42 MHz NMR signal (B) when the magnetic field is 0.51 Tesla.

The NMR signal (B) generated by the magnetic field locking sample 122 is received by the solenoid coil 703 and amplified by the preamplifier 212. In this way, a method in which the solenoid coil 703 that generates a high-frequency magnetic field that excites nuclear spins is also used as a solenoid coil that detects an NMR signal is called a single coil method. This method is a widely known technique (see, for example, Pulse and Fourier Transform NMR published by Academic Press by Farrar & Becker).

The NMR signal (B) amplified by the preamplifier 212 is phase-detected by a phase detector (PSD) 213 using a 20.42 MHz high-frequency signal generated by the magnetic field lock system transmitter 209 as a reference signal 302.

Here, the phase detector 213 will be described in more detail with reference to FIGS. 7 (a) and 7 (b). As shown in FIG. 7 (a), when the NMR signal (B) 301 and the reference signal 302 are input to the phase detector 213 and these phases are the same (referred to as being synchronized), FIG. ), The output signal 303 becomes zero. On the other hand, when the phase of the input signal 301 advances with respect to the reference signal 302, as shown in FIG. 7B, the output signal 303 swings to the plus side, and when the phase difference is 180 °, the output signal 303 reaches the maximum value. Show. Conversely, when the phase is delayed, the output signal 303 swings to the negative side, and when the phase is delayed by 180 °, the output signal 303 shows a negative maximum value.

As described above, when the static magnetic field strength of the imaging space 102 is a desired 0.5 Tesla, there is a relationship of a static magnetic field strength of 0.51 Tesla at a position below the top plate 119 where the magnetic field locking sample 122 is disposed. When the magnetic field intensity is 0.51 Tesla, the magnetic field locking sample (fluorine nucleus ¹⁹ F) 122 outputs an NMR signal (B) having a resonance frequency of 20.42 MHz, and therefore is synchronized with the frequency of the reference signal 302, and the phase detector 213 The output will be zero.

On the other hand, because the current supplied to the superconducting coil 105 drifts, the static magnetic field formed by the superconducting coil 105 in the imaging space 102 is greater than 0.5 Tesla, and the static magnetic field strength at the position where the magnetic field locking sample 122 is disposed. However, when it becomes larger than 0.51 Tesla, the frequency of the NMR signal (B) of the magnetic field locking sample (fluorine nucleus ¹⁹ F) 122 also increases. As a result, the phase of the NMR signal (B) 301 advances from the reference signal 302, and the output signal 303 of the phase detector (PSD) 213 generates a positive value.

The positive output signal 303 of the phase detector (PSD) 213 is integrated with the polarity inverted by the integrating DC amplifier 214, amplified to a desired value, and applied to the magnetic field correction coil 215. As a result, the correction coil 215 generates a magnetic field in a direction that reduces the static magnetic field strength generated by the superconducting coil 105. Therefore, the magnetic field intensity in the imaging space 102 returns to the value of 0.5 Tesla again after the magnetic field generated by the magnetic field correction coil 215 is reduced. As a result, the magnetic field intensity at the position where the magnetic field locking sample 122 is disposed also returns to 0.51 Tesla again, so that the NMR signal (B) 301 and the reference signal 301 are synchronized again, and the output signal of the phase detector 213 is zero. The output value of the integral type DC amplifier 214 does not change.

Conversely, when the static magnetic field formed by the superconducting coil 105 in the imaging space 102 is smaller than 0.5 Tesla and the static magnetic field strength at the position where the magnetic field locking sample 122 is disposed is smaller than 0.51 Tesla, the magnetic field locking sample ( The frequency of the NMR signal (B) of the fluorine nucleus ¹⁹ F) 122 is also lowered. For this reason, the phase of the NMR signal (B) 301 is delayed from the reference signal 302, and the output signal 303 of the phase detector 213 generates a negative value. The negative output signal 303 of the phase detector 213 is integrated with the polarity inverted by the integrating DC amplifier 214, amplified to a desired value, and applied to the magnetic field correction coil 215. The correction coil 215 generates a magnetic field in a direction that increases the static magnetic field strength generated by the superconducting coil 105. As a result, the magnetic field intensity of the imaging space 102 returns to the value of 0.5 Tesla again, and the magnetic field intensity at the position where the magnetic field locking sample 122 is arranged also returns to 0.51 Tesla again, so the NMR signal (B) 301 and the reference signal 301 is synchronized again, and the output signal of the phase detection circuit 213 becomes zero.

Thus, according to the present embodiment, the magnetic field lock unit (magnetic field lock system high frequency unit) 123 determines the correction magnetic field generated by the correction coil 215 in accordance with the frequency of the NMR signal (B) of the magnetic field lock sample 122. Since feedback control is performed, the magnetic field strength of the imaging space 105 can be stabilized to an order (0.01 ppm) equivalent to the stability 10 ⁻⁸ of the reference frequency of the reference frequency generator 201.

The magnetic field locking sample 122 is disposed at a position where it does not physically interfere with the subject 101, and the NMR signal (B) has a different frequency from the NMR signal (A) of the subject 101 and is electrically interfered. Therefore, even when the imaging pulse sequence is being executed, feedback control can be performed to stabilize the static magnetic field strength.

A state where the magnetic field intensity matches the reference frequency, that is, a state where the output signal of the phase detector 213 is zero is called magnetic field lock-on.

The voltage comparison amplifier 216 receives the output signal of the phase detector 213 and causes the display device 217 to display the magnetic field lock-on state. Display device 217 may be a pilot lamp or an image display device. As a result, the operator can grasp that the magnetic field is in a stable state and can instruct a desired operation (such as starting an imaging pulse sequence).

The voltage comparison amplifier 216 passes a signal 218 indicating the magnetic field lock / on state to the computer 115 via the sequencer 114. As a result, the computer 115 can execute an imaging pulse sequence in a magnetic field lock-on state.

In the present embodiment, the same reference frequency generator 201 is connected to the measurement system transmitter 202 of the measurement system high frequency unit 112 and the magnetic field lock system transmitter 209 of the magnetic field lock system high frequency unit 123, and a common reference frequency signal is transmitted. Supply. This correlates the Larmor frequency of the fluorine nucleus ¹⁹ F in the magnetic field lock system and the Larmor frequency of the hydrogen atom ¹ H in the measurement system. This effect will be described in more detail.

The Larmor precession of nuclear magnetic resonance of nuclear spins is determined by the nuclide-specific gyromagnetic ratio γ and the magnetic field strength B ₀ , and the frequency f is expressed by the following equation (1).

2πf = γB ₀ ... (1)
When the Larmor frequency f _F (= 20.42 MHz) of the fluorine nucleus ¹⁹ F spin and the gyromagnetic ratio γ _F (= 25.18) of the fluorine nucleus ¹⁹ F are substituted into the equation (1), the magnetic field strength B 0 becomes 0.51 Tesla. When the magnetic resonance condition of the hydrogen nucleus (proton) ¹ H used for imaging is verified in this magnetic field lock-on state, that is, the magnetic rotation ratio γ _H (= 26.75) of the hydrogen nucleus (proton) and the magnetic field intensity B ₀ Substituting (= 0.5 Tesla) into equation (1), the Larmor frequency f _H of the hydrogen nucleus (proton) ¹ _H is 21.28 MHz.

When the frequency of the reference frequency generator 201 is slightly displaced due to temperature drift or the like, the Larmor frequency of the fluorine nucleus ¹⁹ F of the magnetic field locking sample 122 is displaced, and the output of the phase detector 213 is not zero, and the magnetic field strength feedback loop As a result of this control, the magnetic field strength B _{0 in the} imaging space 102 is slightly displaced. In this case, the Larmor frequency of the hydrogen nucleus (proton) ¹ H slightly changes with a slight displacement of the magnetic field strength B ₀ , but the high-frequency signal generated by the measurement system transmitter 202 and the reference input to the detector 205 frequency of the signal is also because it is generated from the output signal of the same reference frequency generator 201, in correlation with the displacement of the magnetic field strength B _0, by the same percentage slightly displaced. Therefore, the resonance condition (1) of the hydrogen nucleus (proton) ¹ H is maintained corresponding to the change in the magnetic field strength B ₀ , and the detection accuracy of the NMR signal (A) does not change.

As described above, in the present embodiment, the magnetic field strength of the superconducting magnet 103 can be set to 10 ⁻⁸ order (0.01 ppm) equivalent to the stability of the reference frequency generator 201 by the magnetic field lock-on, and the measurement system high frequency By connecting the same reference frequency generator 201 to the unit 112 and the magnetic field lock system high frequency unit 123 to obtain the correlation of the frequencies, the resonance condition is maintained even when the reference frequency fluctuates due to temperature drift or the like. The NMR signal (A) can be detected with high accuracy and stability.

<Gradient magnetic field operation and magnetic field lock operation>
Since the NMR signal (A) and the NMR signal (B) do not interfere with each other, it has been explained that the magnetic field can be locked even during the imaging sequence, but in the actual imaging pulse sequence, a gradient magnetic field pulse is applied, A gradient magnetic field is superimposed on the static magnetic field. For this reason, the magnetic field lock unit (magnetic field lock system high frequency unit) 123 stabilizes the original static magnetic field with high accuracy when feedback control is performed on the static magnetic field on which the gradient magnetic field pulse is superimposed by the NMR signal (B). I can't.

Therefore, the operation of the magnetic field lock high-frequency unit 121 is stopped during the period when the gradient magnetic field is applied. It is desirable that the correction state (feedback control amount) immediately before application of the gradient magnetic field is maintained. With this configuration, it is possible to stabilize the magnetic field strength of the imaging space 102 by compensating only for the magnetic field strength generated by the superconducting coil 105 and other magnetic field changes due to other disturbances without being affected by the magnetic field variation due to the gradient magnetic field. it can. While the application time of the gradient magnetic field is usually a pulse operation of several milliseconds, the magnetic field change due to the temperature drift of the magnet power supply 106 or disturbance magnetic field disturbance is a change over a long time of several seconds or more, so the gradient magnetic field is applied. Even if the magnetic field feedback control (magnetic field lock operation) is stopped for a certain period, the magnetic field strength of the imaging space 102 can be stabilized with high accuracy.

The process for stopping the magnetic field locking operation during the period when the gradient magnetic field is applied will be described with reference to FIG. FIG. 8 shows an imaging pulse sequence (timing magnetic field pulse application timing of the gradient magnetic field power supply 110, high frequency magnetic field irradiation timing and NMR signal (A) detection timing of the measurement system high frequency unit 112), and magnetic field lock high frequency unit 123. It is the figure which showed the relationship with this operation | movement as a timing chart.

The imaging pulse sequence in FIG. 8 is a sequence for imaging the cross-section of the patient's head (the x-z plane perpendicular to the y-axis that is the body axis) by the spin echo method.

Process (1): First, the gradient magnetic field power supply 110 supplies a drive current to the gradient coil assembly 109, and a y gradient magnetic field pulse (slice selection gradient magnetic field pulse) 401 for determining a cross section to be photographed is an imaging space. Applied to the object 101 of 102. At the same time, a high frequency current of 21.28 MHz is supplied from the high frequency power amplifier 203 to the high frequency transmitter coil 111, and the high frequency magnetic field 402 is irradiated. The hydrogen nuclei (protons) ¹ H of the selected cross section transition from the thermal equilibrium state to the resonance excited state.

Process (2): Next, the resonance-excited hydrogen nucleus (proton) ¹ H undergoes Larmor precession and returns to its original thermal equilibrium state according to the relaxation time constant (referred to as the longitudinal relaxation time T1). Return. In this relaxation process, z gradient magnetic field pulse (phase encoding gradient magnetic field pulse) 403-1 and x gradient magnetic field pulse (frequency encoded gradient magnetic field pulse) 404 are applied to add the position information of the xz plane to the Larmor precession. . During application of the x gradient magnetic field pulse 404, the NMR signal (A) 405 is detected by the high-frequency receiver coil 113.

Process (3): Wait for a waiting time according to the time parameter selected in the spin echo method.

If the waiting time has elapsed, repeat steps (1), (2) and (3). The number of repetitions is, for example, 256 times and matches the number of pixels of the MRI image. By repeating this, the phase encoding gradient magnetic field is applied by changing the z gradient magnetic field pulses 403-1, 403-2,... 403-256 and their intensities in 256 steps. The 256 repetition periods from the process (1) to the process (3) become an imaging examination period 406.

During this period 406, the magnetic field lock unit (magnetic field lock system high-frequency unit) 123 applies the gradient

magnetic field pulses

401, 403, and 404 for the time of the process (1) and the process (2). A rest period 407 in which 122 NMR signals (B) are not detected is set.

And, the magnetic field correction current which is the output of the integral type DC amplifier 214 holds the value. The time of the process (3) is an operation period 408 in which the NMR signal (B) of the magnetic field locking sample 122 is detected and the magnetic field correction current is updated.

The magnetic field lock is also in the operation period 408 during the imaging examination pause period 409, which is the patient loading / unloading time and waiting time before and after the imaging examination period 406.

Switching between the operation period 408 and the rest period 409 of the magnetic field lock unit (magnetic field lock type high frequency unit) 123 is performed by the high frequency switch 210 of FIG. That is, the high frequency switch 210 is controlled to be turned on / off by the control signal 219 from the sequencer 114. The control signal 219 is linked to the control signal 220 of the D / A conversion circuit 208 that is an input signal of the gradient magnetic field power supply 110.

During the rest period 407 in which the gradient magnetic field is applied, the high frequency switch 210 is turned off by the control signal 219, and the 20.02 MHz high frequency signal output from the magnetic field lock system transmitter 209 is input to both the high frequency power amplifier 211 and the phase detector 213. It will not be done.

Therefore, not only the fluorine nucleus 19F is not resonantly excited, but also the phase detector 213 has no reference signal 302, so its output signal becomes zero. Since the input signal of the integrating DC amplifier 214 is zero, the output of the integrating DC amplifier 214 is maintained at the previous value. Since the applied current of the magnetic field correction coil 215 does not change, the value of the correction magnetic field so far is held. On the other hand, in the operation period 408 in which no gradient magnetic field is applied, the high frequency switch 210 is turned on, and the magnetic field lock unit (magnetic field lock system high frequency unit) 123 performs feedback control of the magnetic field correction coil 215 according to the detected NMR signal (B). And stabilize the static magnetic field.

<Flow of operation of superconducting magnet of MRI system>
Next, the overall operation of the superconducting magnet of the MRI apparatus will be described with reference to FIG. FIG. 9 is a flowchart showing daily operations of the MRI apparatus of this embodiment.

Prior to the inspection on that day, a static current of 160 amperes is supplied from the magnet power source 106 to the superconducting coil 105 of the superconducting magnet 103 to generate a static magnetic field (step 501). This operation may be performed by an operation of the input device 117 by the operator, or may be performed using an automatic startup function programmed in advance in the computer 115.

When the excitation work of the superconducting magnet 103 is completed, the magnetic field lock system high-frequency unit 123 operates, and the magnetic signal from the integrating DC amplifier 214 is obtained using the NMR signal (B) of the fluorine nucleus ¹⁹ F obtained from the magnetic field locking sample 122. The current supplied to the correction coil 215 is feedback-controlled. Thereby, the magnetic field strength of the imaging space 102 becomes, for example, 0.5 Tesla (magnetic field locked / on state) and is maintained (step 502). An output signal of the voltage comparison amplifier 216 is input to the computer 115. At the same time, the display device 217 displays that the magnetic field is locked / on.

The operator confirms that the magnetic field is locked on by the display device 217, and then carries the first subject 101 into the imaging space 102 (step 503). Although the normal human body does not affect the magnetic field, it may have magnetism depending on the wearing of the magnetic body or some medical implants. Although the distribution is once changed, the magnetic field intensity of the imaging space 102 is maintained in the magnetic field lock-on state of 0.5 Tesla by the feedback control of the magnetic field lock system high-frequency unit 123. Normally, the change in magnetic field due to the influence of the subject 101 is slight, and the change speed (operation of loading and unloading the subject 101) is sufficiently covered by the response speed of the magnetic field lock high-frequency unit 123. The magnetic field lock will never be locked out.

Next, an imaging examination (imaging) of the examination region of the subject 101 is performed by executing an imaging pulse sequence suitable for the examination purpose. At this time, during a period when the gradient magnetic field is not operating, the magnetic field lock high-frequency unit 123 detects the NMR signal of the magnetic field locking sample 122 and feedback-controls the magnetic field strength of the imaging space 102. During the operation of the gradient magnetic field, the feedback control is stopped and the previous correction amount is maintained. Thus, the magnetic field strength of the imaging space 102 can be maintained at 0.5 Tesla even during execution of the imaging pulse sequence (step 504).

When the first imaging test of the subject 101 is completed, it is carried out from the imaging space 102. During this carry-out operation, the magnetic field strength is maintained in a lock-on state of 0.5 Tesla (step 505).

Then, it is determined whether or not the next subject 101 is inspected (step 506). If the next subject 101 is inspected, the process returns to step 503 and steps 503 to 505 are repeated. During this time, the magnetic field lock-on function is maintained, and the operator only has to confirm the display device 217 for magnetic field lock-on. Of course, since the magnetic field lock-on signal 218 is constantly input to the computer 115, the imaging inspection in the magnetic field lock-out state is not performed.

On the other hand, if there is no examination of the next subject 101 in step 506, the MRI apparatus determines whether to proceed to the end operation or to wait for the subject 101 that is not reserved, such as an urgent patient, based on a predetermined criterion. (Step 507). For example, when instructed to proceed to the end operation by the operator's instruction, or when the end time of the medical facility has elapsed, it is decided to shift to demagnetization work of the superconducting magnet 103, and otherwise wait Can do. In the case of standby, the process returns to step 506.

If completed, demagnetization work of superconducting magnet 103 is performed (step 508). The degaussing operation is performed by executing an operation instructed by the operator via the input device 117, or by performing a predetermined automatic degaussing operation by the computer 115.

When entering the degaussing operation, the magnetic field lock type high frequency unit 123 finishes its operation and enters a magnetic field lockout state (step 508).

The flow of operation of the MRI apparatus for one day has been described above, but the magnetic field lock-on state is maintained during the above steps 503 to 507. Accordingly, the magnetic field attenuation due to the slight resistance component of the superconducting coil 105, the change in the output current due to the temperature drift of the magnet power source 106, the magnetic field fluctuation due to the disturbance due to the movement of the magnetic material, etc. are compensated, and 0.5 Tesla is stably maintained.

In the above-described embodiment, the example in which the correction of the static magnetic field in the imaging space 102 is performed by the correction coil 215 has been described. However, the present invention is not limited to this method. It is also possible to directly correct the static magnetic field generated by the superconducting coil 105 by feedback-controlling the current supplied from the magnet power source 106 to the superconducting coil 105 according to the output signal 303 of the phase detector 213.

The inductance of the correction coil 215 can be several hundred millihenries, which is three orders of magnitude smaller than that of a general superconducting coil 105. Therefore, the correction coil 215 can be corrected more accurately than when the superconducting coil 105 is feedback controlled. It is feasible. Therefore, it is possible to adopt a hybrid method in which both corrections are combined. That is, a method in which the magnetic field intensity is roughly adjusted by feedback control of the magnet power supply 106 and then finely adjusted by the correction magnetic field coil 215 can be used. Alternatively, the magnetic field strength can be roughly adjusted by the magnet power source 106 only when the correction magnetic field coil 215 is usually used for adjustment and the magnetic field strength greatly changes.

In the above-described embodiment, the fluorine nucleus ¹⁹ F is used as the magnetic field locking sample, but deuterium nucleus ² H heavy water (D ₂ O, ² H ₂ O, or deuterated chloroform CDCl ₃ ) is used. Also good. In this case, the frequency synthesized by the magnetic field lock system transmitter 209 is 3.333 MHz. Since both are different from the magnetic resonance frequency of hydrogen nucleus (proton) ¹ H, a high-frequency signal of the magnetic field lock system is not induced in the high-frequency receiver coil 113, and a stable MRI image of the subject 101 is obtained. can get.

(Second embodiment)
In the second embodiment, an example in which the magnetic field locking sample 122 is disposed in the holding unit (801) that holds the high-frequency receiver coil 113 will be described.

The high-frequency receiver coil 113 has a structure in which a coil conductor is covered with a holding portion 801 having a desired shape made of an insulating material. The coil conductor and the holding unit 801 have a shape corresponding to the examination site. For example, as shown in FIG. 5, the high-frequency receiver coil 113 for the human head has a cylindrical holding unit 801 and includes a head receiving base. ing. In the example of FIG. 5, an example in which the magnetic field locking sample 122 is disposed inside the head receiving base of the holding unit 8 is shown.

Thus, when the magnetic field locking sample 122 is arranged inside the high frequency receiver coil 113, the magnetic field locking sample 122 is located in the imaging space 102 (imaging region (FOV)), but the magnetic field locking sample 122 Does not appear in the imaging image because it does not contain hydrogen nuclei (protons) ¹ H in the imaging examination.

In the case of the second embodiment, the magnetic field locking sample 122 can be arranged in the imaging space 102 and closest to the center thereof. As a result, the fluorine nucleus ¹⁹ F of the magnetic field locking sample 122 exists in a static magnetic field with the same intensity and the same intensity as the hydrogen nucleus (proton) ¹ H of the imaging examination, so the NMR signal (B) is converted into an NMR signal ( It can be detected with higher accuracy with higher correlation with A). Further, the magnetic field locking sample 122 can be arranged at an optimum position and size for each high-frequency receiver coil 113. Therefore, it is possible to stabilize the magnetic field in the imaging space 102 with high accuracy.

In the second embodiment, since the magnetic field locking sample 122 is disposed in the high-frequency receiver coil 113, the start timing of the magnetic field locking operation is different from that in the first embodiment. The operation flow of the superconducting magnet 103 of the second embodiment will be described with reference to FIG.

First, a predetermined current of 160 amperes is supplied to the superconducting magnet 103 prior to the inspection on the day, and a static magnetic field is generated in the superconducting coil 105 (step 501). This operation is the same as step 501 in FIG. When the excitation operation of the superconducting magnet 103 is completed, the operator attaches the high-frequency receiver coil 113 suitable for the purpose of the imaging examination to the first subject 101 and carries it into the imaging space 102 (step 1001). By this step 1001, since the magnetic field locking sample 122 provided in the high frequency receiver coil 113 is arranged in the imaging space 102, the magnetic field locking system high frequency unit 123 outputs the NMR signal (B) of the magnetic field locking sample 122. It becomes possible to detect, feedback control of the static magnetic field is disclosed, and the magnetic field lock-on state is entered (step 1002). The output signal 218 of the voltage comparison amplifier 216 is input to the computer 115, and at the same time, the magnetic field lock-on state is displayed on the display device 217.

Next, step 504 is performed in the same manner as in the first embodiment, and an imaging test of the test site of the subject 101 is performed.

When the first MRI examination of the subject 101 is completed and taken out from the imaging space 102, the magnetic field locking sample 122 is also carried out from the magnetic field formed by the superconducting coil 105, so the magnetic field locking function is released (step 1003). .

After step 506 for determining whether or not the next subject 101 is inspected, if there is the next subject 101, the process returns to step 1001 to carry the subject 101 into the imaging space 102. Here again, the magnetic field lock-on state is established (step 1002).

In step 506, if there is no reservation for the next subject 101, the MRI apparatus proceeds to step 507 and determines whether to proceed to step 1004 of demagnetization of the superconducting magnet or return to step 506 and wait. In step 1004, the demagnetization work of the superconducting magnet 103 is performed as in step 508 of the first embodiment. Since the magnetic field lock has already been completed in step 1003, only demagnetization is performed in step 1004.

Other configurations and operations are the same as those in the first embodiment, and thus description thereof is omitted.

(Third embodiment)
As a third embodiment, an embodiment in which magnetic field locking samples 122 are arranged at a plurality of locations in a top board 119 where the subject 101 is carried in and out will be described with reference to FIG.

As shown in FIG. 4, magnetic field locking samples 122-1, 122-2, and 122-3 are respectively placed in different positions (three locations in FIG. 4) along the body axis direction of the subject 101 in the top plate 119. It is arranged. In the case of this configuration, the position of the top 119 with respect to the imaging space 102 changes according to the examination site of the subject 101. For example, in the case of the examination of the head of the subject 101, the magnetic field locking sample 122-1 is disposed closest to the center of the imaging space 102. Therefore, the magnetic field lock high-frequency unit 123 receives the position information of the top plate 119 from the computer 115 via the sequencer 114, and accordingly, the magnetic lock closest to the imaging space 102 among the plurality of magnetic field locking samples 122 is received. Sample 122-1 is selected, and feedback control of the static magnetic field is performed using the NMR signal (B) of the sample 122-1 for magnetic lock.

In the third embodiment, the start and end timing of the magnetic field lock is the timing when the subject is carried in / out, and is the same as the flow of FIG. 10 of the second embodiment. Other configurations and operations are the same as those in the first embodiment, and thus description thereof is omitted.

101 subject, 102 imaging space, 103 superconducting magnet, 105 superconducting coil, 106 magnet power supply, 109 gradient magnetic field coil assembly, 111 high frequency transmitter coil, 112 measurement system high frequency unit, 113 high frequency receiver coil, 114 sequencer, 115 computer, 119 Top plate, 122 Magnetic field lock sample, 123 Magnetic field lock system high frequency unit, 201 Reference frequency generator, 209 Magnetic field lock system transmitter, 210 High frequency switch, 211 High frequency power amplifier, 213 Phase detector (PSD), 214 Integral DC amplifier 215 Magnetic field correction coil 217 Lock-on display device 701 Microcapillary 702 Fluorine compound

Claims

A superconducting coil that forms a static magnetic field in the imaging space, a magnet power supply that continuously supplies current to the superconducting coil while the superconducting coil forms the static magnetic field, and a gradient magnetic field that generates a gradient magnetic field in the imaging space A generator, a high-frequency magnetic field generator that irradiates a subject placed in the imaging space with a high-frequency magnetic field, a measurement unit that detects a nuclear magnetic resonance signal (A) of the subject,
A magnetic field locking sample disposed at a position where a magnetic field generated by the superconducting coil is applied and does not physically interfere with the subject;
Irradiating the magnetic field locking sample with a high frequency magnetic field, detecting a nuclear magnetic resonance signal (B) generated by the magnetic field locking sample, adjusting the intensity of the static magnetic field according to the detection result, and making the static magnetic field constant A magnetic resonance imaging apparatus having a magnetic field lock unit.
The magnetic resonance imaging apparatus according to claim 1, wherein the magnetic field locking sample is a magnetic resonance signal (B) having a magnetic resonance frequency different from a magnetic resonance frequency of the nuclear magnetic resonance signal (A) detected by the measurement unit. A magnetic resonance imaging apparatus characterized by including a nuclear species that generates).
2. The magnetic resonance imaging apparatus according to claim 1, wherein the magnetic field lock unit includes the static magnetic field corresponding to a detection result of the nuclear magnetic resonance signal (B) while the gradient magnetic field generation unit generates the gradient magnetic field. A magnetic resonance imaging apparatus characterized by temporarily stopping the adjustment of the intensity of the magnetic resonance imaging apparatus.
4. The magnetic resonance imaging apparatus according to claim 3, wherein the magnetic field lock unit holds an adjustment amount immediately before the stop while the gradient magnetic field generation unit generates the gradient magnetic field. Resonance imaging device.
The magnetic resonance imaging apparatus according to claim 1, wherein the high-frequency magnetic field generation unit includes a high-frequency transmitter coil and a high-frequency signal generation unit that supplies a high-frequency signal for generating the high-frequency magnetic field to the high-frequency transmitter coil,
The magnetic field lock unit includes a high frequency transmitter coil for magnetic field locking that irradiates the magnetic field locking sample with a high frequency, and a high frequency signal for magnetic field locking that supplies the high frequency signal for generating the high frequency magnetic field to the high frequency magnetic field transmitter coil for magnetic field locking. A reference frequency generator for supplying a common reference frequency signal is connected to the high-frequency signal generation unit and the magnetic field locking high-frequency signal generation unit.
The magnetic resonance imaging apparatus according to claim 1, further comprising a table on which the subject is mounted,
The magnetic resonance imaging apparatus, wherein the magnetic field locking sample is disposed outside the imaging space and at a position that does not physically interfere with the measurement unit and the table.
The magnetic resonance imaging apparatus according to claim 1, further comprising a table for carrying the subject mounted thereon,
The magnetic resonance imaging apparatus, wherein the magnetic field locking sample is disposed at one or more locations in a top plate on which the subject of the table is mounted, and is inserted into the imaging space together with the subject.
The magnetic resonance imaging apparatus according to claim 1, wherein the measurement unit includes a high-frequency receiver coil that receives the nuclear magnetic resonance signal (A), and a holding unit that holds the high-frequency receiver coil.
The magnetic resonance imaging apparatus, wherein the magnetic field locking sample is provided in the holding unit of the high-frequency receiver coil.
2. The magnetic resonance imaging apparatus according to claim 1, wherein the magnetic field locking sample is disposed in a container having a size that is not affected by an intensity gradient of a magnetic field generated by the superconducting coil at a position where the magnetic field locking sample is disposed. A magnetic resonance imaging apparatus.
2. The magnetic resonance imaging apparatus according to claim 1, wherein the magnetic field lock unit performs phase detection of the nuclear magnetic resonance signal (B) using a signal having the same frequency as the magnetic resonance frequency of the magnetic field locking sample as a reference signal. And a magnetic resonance imaging apparatus, wherein the magnetic field intensity of the imaging space is adjusted so that the output of the phase detection becomes zero.
11. The magnetic resonance imaging apparatus according to claim 10, wherein the magnetic field locking unit includes a display unit that notifies an operator that the output of the phase detection unit has become zero.
11. The magnetic resonance imaging apparatus according to claim 10, wherein the measurement unit measures a nuclear magnetic resonance signal (A) of the subject in a state where the output of the phase detection unit is zero. Resonance imaging device.
2. The magnetic resonance imaging apparatus according to claim 1, wherein the magnetic field lock unit adjusts the intensity of the static magnetic field by changing an output current of the magnet power supply according to a detection result of the nuclear magnetic resonance signal (B). A magnetic resonance imaging apparatus.
The magnetic resonance imaging apparatus according to claim 1, further comprising a correction coil that generates a correction magnetic field in the imaging space,
The magnetic field locking unit adjusts the intensity of the static magnetic field by changing a current supplied to the correction coil in accordance with a detection result of the nuclear magnetic resonance signal (B). .
The magnetic resonance imaging apparatus according to claim 7, wherein the magnetic field locking samples are respectively arranged at a plurality of locations on the top plate of the table,
The magnetic resonance imaging apparatus, wherein the magnetic field locking unit selects at least one of the plurality of magnetic field locking samples and detects the nuclear magnetic resonance signal (B).