JP6600601B2 - High frequency array coil and magnetic resonance imaging apparatus - Google Patents

Description

  The present invention relates to a magnetic resonance imaging (MRI) apparatus, and more particularly, to an RF coil (Radio Frequency Coil) that detects irradiation of a high-frequency magnetic field and a nuclear magnetic resonance signal.

  The MRI apparatus is a medical diagnostic imaging apparatus that causes magnetic resonance to occur in a nuclear spin in an arbitrary cross section that crosses an examination target, and obtains a tomographic image in the cross section from a generated nuclear magnetic resonance signal.

  In MRI imaging, a high frequency magnetic field is irradiated by an RF coil while applying a gradient magnetic field to a subject placed in a static magnetic field. When a high-frequency magnetic field is irradiated, a nuclear spin in the object, for example, a nuclear spin of a hydrogen atom is excited, and a circularly polarized magnetic field is generated as a nuclear magnetic resonance signal when the excited nuclear spin returns to an equilibrium state. This signal is detected by an RF coil, and signal processing is performed to image the hydrogen nucleus distribution in the living body.

  The RF coil has a loop portion (coil loop) that performs irradiation and detection. The smaller the coil loop, the narrower the sensitivity region, but the higher the sensitivity. On the other hand, if the coil loop is enlarged, the sensitivity region can be expanded. As described above, in the RF coil, there is a trade-off relationship between the high sensitivity and the wide sensitivity area. Further, since the nuclear magnetic resonance signal is a signal of a rotating magnetic field generated in a direction perpendicular to the static magnetic field, it is preferable that the RF coil be disposed in a direction in which the magnetic field can be irradiated and detected in a direction perpendicular to the static magnetic field.

  As described above, the smaller the RF coil, the higher the sensitivity, but the sensitivity range becomes narrower. As a solution to this problem, there is a multi-channel array coil in which a plurality of RF coils are arranged in an array (see, for example, Non-Patent Document 1). A multi-channel array coil has a high sensitivity and a wide sensitivity region, and therefore has become the mainstream of current reception RF coils. In recent years, high-speed imaging (for example, see Non-Patent Document 2) using a difference in spatial sensitivity of each sub-coil of a multi-channel array coil has become widespread. High-speed imaging can be performed at higher speeds as the number of channels increases. Therefore, the number of multi-channel array coils is further increased, and at present, super multi-channel array coils such as 32 channels and 128 channels are widely used. Hereinafter, in a multi-channel array coil or a super multi-channel array coil (collectively referred to as a multi-channel array coil), each RF coil is referred to as a sub-coil.

  Usually, when RF coils having the same resonance characteristics are arranged close to each other, they interfere by magnetic coupling. Since the performance of the RF coil deteriorates due to the magnetic coupling, it is essential to remove the magnetic coupling between the sub-coils in the multi-channel array coil. In Non-Patent Document 1, magnetic coupling is reduced to the maximum by arranging so that part of coil loops of adjacent subcoils overlap. Here, this is called an overlap method. Further, by using a low-input preamplifier, an inductor, and a capacitor, a part of the coil loop is made high impedance, thereby reducing interference from other than the sub-coil. Here, this is called a preamplifier decoupling method. That is, many conventional multi-channel array coils solve the problem of performance degradation due to magnetic coupling by the overlap method or the preamplifier decoupling method.

  Another problem with multi-channel array coils is that the depth of the entire sensitivity region is reduced. In particular, an ultra-multi-channel array coil developed for high-speed imaging is a collection of sub-coils with a narrow sensitivity region due to the miniaturization of each sub-coil that constitutes the array coil, and has deeper sensitivity than an array coil composed of fewer channels. Lower. Therefore, it is difficult to obtain a beautiful image in the deep part of the subject. With respect to this problem, Patent Document 1 discloses a technique for increasing the depth sensitivity by causing intentional magnetic coupling to a multichannel array coil.

  By the way, the RF coil may have a shape that can be divided in order to improve operability. For example, some limb RF coils can be divided up and down on a plane parallel to the center axis of the cylinder, and can be easily set on the subject. In addition, there is a head / neck RF coil that can be divided between the occipital side and the face side, and can be switched between the head and the neck by changing the attachment on the face side. In these splittable RF coils, the loop portion often straddles the dividing line in order to remove the magnetic coupling. Therefore, connection parts are added so that the divided loops are connected.

International Publication No. 2016/035510 Specification

Roemer PB et al., “The NMR Phased Array”, Journal of Magnetic Resonance, USA, 1990, 16, p. 192-225 Klaas PP et al., “SENSE: Sensitivity Encoding for MRI”, Journal of Magnetic Resonance, USA, 1999, 42, p. 952-962

  In the separable RF coil, as described above, a connecting part for the loop portion to cross the dividing line is added to remove the magnetic coupling. For this reason, there is a problem that the cost increases or the loss due to the connection resistance at the connection point deteriorates the performance of the subcoil. In addition, there is a risk that mechanical deterioration is likely to occur due to poor connection at the connection portion or repeated attachment / detachment.

  If connection parts are not used in the division part, the subcoils arranged adjacent to each other across the division line cannot be arranged so that a part of the coil loop overlaps. That is, magnetic coupling removal by the overlap method cannot be performed. Although magnetic coupling removal by the preamplifier decoupling method can be adopted, this method reduces the magnetic coupling to the maximum and does not completely remove the magnetic coupling. Further, if the distance between the sub-coils arranged adjacent to each other with the dividing line interposed therebetween is reduced so that the magnetic coupling is reduced, the arrangement density of the array coils there is lowered, and the sensitivity in the vicinity thereof is lowered. Therefore, in an array coil in which coil loops are arranged at high density without using connection parts in the division part, there are not a few subcoils arranged adjacent to each other across the division line, and the performance of the array coil deteriorates. To do.

  On the other hand, Patent Document 1 discloses a technique for increasing the depth sensitivity by intentional magnetic coupling, but when the array coil is arranged in a direction in which a magnetic field can be irradiated and detected in a direction perpendicular to the static magnetic field, The arrangement and adjustment conditions of the subcoil for increasing the sensitivity are not disclosed.

  The present invention has been made in view of the above circumstances, and an object of the present invention is to provide a technique for increasing the sensitivity of a deep portion, realizing high speed and high image quality in a multi-channel array coil of an MRI apparatus.

  The high-frequency array coil of the present invention is configured to magnetically couple at least one subcoil constituting the high-frequency array coil with an adjacent subcoil. The single subcoil alone does not resonate at the magnetic resonance frequency, but resonates at the magnetic resonance frequency when magnetically coupled to the adjacent subcoil, and when the current flowing through the one subcoil flows through the adjacent subcoil due to the magnetic coupling, The phase difference of the current is adjusted so as to be within a predetermined range in relation to the angle formed by the central axis of each subcoil.

  That is, the array coil of the present invention includes a first subcoil having a first loop coil portion made of a conductor and a second subcoil having a second loop coil portion made of a conductor, and transmits or receives a high-frequency signal. The first sub-coil is a tuned high-frequency array coil, wherein the first sub-coil has a resonance frequency different from a nuclear magnetic resonance frequency and is magnetically coupled to the second sub-coil. The first loop coil is adjusted to form a current path that circulates in the loop of the one loop coil portion and the loop of the second loop coil portion, and to resonate at a nuclear magnetic resonance frequency. The phase difference between the current flowing through the loop of the first portion and the current flowing through the loop of the second loop coil portion is the center axis of the first loop coil portion and the first Respect of the phase difference angle and the same and opposite sign formed by the central axis of the loop coil portion is adjusted to a range of 90 degrees ±.

  The MRI apparatus of the present invention includes a high-frequency coil that transmits a high-frequency magnetic field or receives a nuclear magnetic resonance signal, and the high-frequency coil is the above-described high-frequency array coil. This high-frequency array coil is preferably arranged such that the central axes of the first and second loop coil portions are substantially perpendicular to the static magnetic field direction of the MRI apparatus.

  ADVANTAGE OF THE INVENTION According to this invention, in the multi-channel array coil of an MRI apparatus, the technique which raises a deep part sensitivity and implement | achieves high speed and high image quality can be provided.

BRIEF DESCRIPTION OF THE DRAWINGS It is an external view of the MRI apparatus with which this invention is applied, (a) is an MRI apparatus of a horizontal magnetic field system, (b) is an MRI apparatus of a vertical magnetic field system. 1 is a block diagram of an MRI apparatus to which the present invention is applied. (A) is explanatory drawing for demonstrating the structure of the birdcage type | mold RF coil used as a transmission RF coil, (b) is a figure for demonstrating the magnetic coupling prevention circuit between transmission / reception by the side of a transmission coil. (A) is a figure for demonstrating the structure of the array coil used as a receiving RF coil, (b) and (c) are the figures for demonstrating the magnetic coupling prevention circuit between transmission / reception by the side of a receiving coil. . It is a figure for demonstrating arrangement | positioning of a transmission RF coil and a reception RF coil. (A) And (b) is a figure for demonstrating arrangement | positioning of the array coil of 1st embodiment. (A)-(e) is a figure for demonstrating operation | movement of the array coil of 1st embodiment. (A)-(c) is a figure for demonstrating operation | movement of the array coil of 1st embodiment. It is a figure for demonstrating operation | movement of the array coil of 1st embodiment. It is a figure for demonstrating the modification of the array coil of 1st embodiment. It is a figure for demonstrating the structure and arrangement | positioning of the array coil of 2nd embodiment. (A) And (b) is a figure for demonstrating the structure of the subcoil contained in the array coil of 2nd embodiment. It is explanatory drawing for demonstrating the structure and arrangement | positioning of the array coil of 3rd embodiment. (A) And (b) is explanatory drawing for demonstrating the structure of the subcoil contained in the array coil of 3rd embodiment. It is a figure explaining the relationship between the magnetic coupling means employ | adopted as the array coil of 3rd embodiment, and an arrangement | positioning angle. It is explanatory drawing for demonstrating the modification of the array coil of 3rd embodiment. (A) is explanatory drawing for demonstrating the transmission array coil of 4th embodiment, (b) is explanatory drawing for demonstrating the magnetic coupling prevention circuit between transmission / reception of 4th embodiment. .

  Hereinafter, embodiments to which the present invention is applied will be described. Hereinafter, in all the drawings for explaining the embodiments of the present invention, those having the same function are denoted by the same reference numerals, and repeated explanation thereof is omitted.

[MRI system configuration]
First, the overall configuration of the MRI apparatus common to the embodiments will be described with reference to FIGS. 1 and 2. FIG. 1 is an external view showing a typical example of an MRI apparatus to which the present invention is applied. In the figure, a coordinate system 090 is used in which the direction opposite to the static magnetic field direction is the z direction, and the two directions perpendicular thereto are the x direction and the y direction, respectively. Hereinafter, the same applies to all drawings in this specification.

  FIG. 1A shows a horizontal magnetic field type MRI apparatus 100 using a tunnel magnet that generates a static magnetic field with a solenoid coil. The inspection object 103 is inserted in a static magnetic field space formed in the bore of the tunnel-type magnet while being placed on the table 102, and is arranged so that the inspection site substantially coincides with the center of the static magnetic field. Is called. FIG. 1B shows a hamburger type (open type) vertical magnetic field type MRI apparatus 101 in which magnets are separated into upper and lower sides in order to enhance the feeling of opening. Also in this MRI apparatus 101, the inspection object 103 is inserted between the upper and lower magnets in a state where it is placed on the table, and imaging is performed.

  Each of the MRI apparatuses 100 and 101 can be applied to each embodiment of the present invention. Further, the present invention is not limited to the form of the MRI apparatus shown in FIG. 1, and can be applied to various known MRI apparatuses. Hereinafter, the MRI apparatus 100 having the horizontal magnetic field type magnet 110 will be described as an example.

  FIG. 2 is a block diagram showing a schematic configuration of the MRI apparatus 100. As shown in the figure, the MRI apparatus 100 includes a horizontal magnetic field type magnet 110, a gradient magnetic field coil 131, a transmission RF coil 151, a reception RF coil 161, a gradient magnetic field power supply 132, a shim coil 121, a shim power supply 122, and a high frequency. A magnetic field generator 152, a receiver 162, a magnetic coupling prevention circuit driving device 180, a computer (PC) 170, a sequencer 140, and a display device 171 are provided. Reference numeral 102 denotes a table on which the inspection object 103 is placed. The magnet 110 constitutes a static magnetic field generation unit that forms a static magnetic field, and the gradient magnetic field coil 131 and the gradient magnetic field power supply 132 constitute a gradient magnetic field generation unit that forms a gradient magnetic field.

  The gradient magnetic field coil 131 is connected to the gradient magnetic field power supply 132 and generates a gradient magnetic field. The shim coil 121 is connected to the shim power source 122 and adjusts the uniformity of the magnetic field. A transmission RF coil 151 (hereinafter referred to as a transmission coil) is connected to the high-frequency magnetic field generator 152 and irradiates (transmits) a high-frequency magnetic field to the inspection target 103. A reception RF coil 161 (hereinafter referred to as a reception coil) is connected to the receiver 162 and receives a nuclear magnetic resonance signal from the inspection target 103. Although not shown, the transmission coil 151 and the reception coil 161 are connected to circuits (magnetic coupling prevention circuits) that prevent magnetic coupling between the transmission coil 151 and the reception coil 161, respectively. The magnetic coupling prevention circuit driving device 180 is connected to these magnetic coupling prevention circuits.

  The sequencer 140 sends commands to the gradient magnetic field power source 132, the high-frequency magnetic field generator 152, and the magnetic coupling prevention circuit driving device 180 to operate them. The command is sent in accordance with an instruction from the computer (PC) 170. The sequencer 140 also sets a magnetic resonance frequency as a reference for detection by the receiver 162 in accordance with an instruction from the computer (PC) 170. For example, according to a command from the sequencer 140, a high-frequency magnetic field is applied to the inspection target 103 through the transmission coil 151. A nuclear magnetic resonance signal generated from the inspection object 103 by irradiating a high-frequency magnetic field is detected by the receiving coil 161 and detected by the receiver 162.

  A computer (PC) 170 controls the operation of the entire MRI apparatus 100 and performs various signal processing. For example, a signal detected by the receiver 162 is received via an A / D conversion circuit, and signal processing such as image reconstruction is performed. Then, the result is displayed on the display device 171. The detected signal and measurement conditions are stored in the storage medium 172 as necessary. In addition, the sequencer 140 is made to send an instruction so that each device operates at a preprogrammed timing and intensity. Further, when it is necessary to adjust the static magnetic field uniformity, the sequencer 140 sends a command to the shim power supply 122 to cause the shim coil 121 to adjust the magnetic field uniformity.

  The MRI apparatus of the present invention is characterized in that a specific high-frequency array coil is adopted as a transmission coil (including a transmission / reception coil) and / or a reception coil. An embodiment in which a high-frequency array coil is used as the transmission coil will be described.

  First, an outline of a transmission coil and a reception coil common to an MRI apparatus using a high-frequency array coil as a reception coil will be described, and then each embodiment of the high-frequency array coil will be described.

[Outline of transmit RF coil and receive RF coil]
As an example, an RF coil (birdcage RF coil) having a birdcage shape is used as the transmission RF coil 151, and an array coil in which two RF coils (surface coils) having a loop shape are arranged as the reception RF coil 161 is used. A case will be described as an example. FIG. 3 shows an example of a birdcage type RF coil 300, and FIG. 4 shows an example of an array coil 400.

  FIG. 3A is a block diagram for explaining the configuration of the birdcage type RF coil 300 of the present embodiment. The birdcage type RF coil 300 according to the present embodiment includes a plurality of linear conductors 301, end conductors 302 that connect the ends of the respective linear conductors 301, and a capacitor 303 that is inserted into the end conductors 302. The resonance frequency of the birdcage type RF coil 300 can be adjusted by adjusting an adjustment element such as the capacitor 303, and is adjusted to the resonance frequency of the element to be excited. In this embodiment, the magnetic resonance frequency of the hydrogen nucleus is adjusted so that the hydrogen nucleus can be excited.

  The birdcage type RF coil 300 of the present embodiment includes two input ports 311 and 312 for supplying power. The first input port 311 and the second input port 312 are configured so that transmission signals having a phase difference of 90 degrees are input and a high-frequency magnetic field is efficiently applied to the subject 103. This high-frequency magnetic field is a rotating magnetic field that rotates in a plane that is substantially perpendicular to the central axis of the birdcage type RF coil 300.

  A circuit for preventing magnetic coupling with the array coil and a transmission / reception magnetic coupling prevention circuit 210 are inserted in series in each linear conductor 301 of the birdcage type RF coil 300. The transmission / reception magnetic coupling prevention circuit 210 is connected to the magnetic coupling prevention circuit driving device 180 of the MRI apparatus 100 as described above.

  FIG. 3B is a view for explaining the configuration of the inter-transmission / reception magnetic coupling prevention circuit 210 inserted into the straight conductor 301 of the birdcage type RF coil 300 and the connection with the magnetic coupling prevention circuit driving device 180. The inter-transmission / reception magnetic coupling prevention circuit 210 includes a PIN diode 211 and a control signal line 212. The PIN diode 211 is inserted in series with the straight conductor 301, and the control signal line 212 is connected to both ends of the PIN diode 211. The control signal line 212 is connected to the magnetic coupling prevention circuit driving device 180. A choke coil is inserted into the control signal line 212 in order to avoid high frequency mixing.

  The PIN diode 211 normally has a high resistance (off), and has a characteristic of being substantially in a conductive state (on) when the value of the direct current flowing in the forward direction of the PIN diode 211 exceeds a certain value. In this embodiment, this characteristic is used, and on / off of the PIN diode 211 is controlled by a direct current output from the magnetic coupling prevention circuit driving device 180. That is, at the time of high-frequency signal transmission, a control current for turning on the PIN diode 211 is supplied, and the birdcage type RF coil 300 is caused to function as the transmission RF coil 151. Further, when receiving a nuclear magnetic resonance signal, the control current is stopped, and the birdcage type RF coil 300 is made to have a high impedance and opened.

  As described above, in the present embodiment, by controlling the direct current (control current) from the magnetic coupling prevention circuit driving device 180, the birdcage type RF coil 300 is made to function as the transmission RF coil 151 at the time of high-frequency signal transmission, and the nuclear magnetism. When receiving the resonance signal, the magnetic coupling with the array coil which is the reception RF coil 161 is removed as an open state.

  On the other hand, the array coil 400 used as the reception RF coil 161 has a structure in which a plurality of subcoils are arranged side by side. FIG. 4A shows an array coil 400 having two subcoils as an example. Here, the two subcoils 410 constituting the array coil 400 are referred to as a first subcoil 410A and a second subcoil 410B, respectively. Each of the first sub-coil 410A and the second sub-coil 410B is a surface coil having a loop configured on a plane. The first subcoil 410A and the second subcoil 410B each receive a nuclear magnetic resonance signal. The received signal is sent to the receiver 162.

Hereinafter, the constituent elements of each subcoil 410 constituting the array coil 400 are omitted in particular when the subcoil 410 does not need to be distinguished.
The first subcoil 410A includes a loop coil unit 420 (first loop coil unit 420A) that receives a nuclear magnetic resonance signal, a low (input) impedance signal processing circuit 430 (first low impedance signal processing circuit 430A), A magnetic coupling adjustment unit 441 (first magnetic coupling adjustment unit 441A) that connects the loop coil unit 420 and the low impedance signal processing circuit 430; The magnetic coupling adjustment unit 441 includes at least one of a capacitor and an inductor. The magnetic coupling adjustment unit 441 includes an adjustment element for adjusting the magnetic coupling with the second subcoil (particularly, the magnitude and phase difference of the current flowing through both coils during magnetic coupling). Details thereof will be described in an embodiment of an array coil described later.

  The loop portion (first loop 421A) of the first loop coil portion 420A is formed of a conductor. The first loop coil section 420A includes a capacitor 424A that is inserted in series with the inductor component of the first loop 421A. The inductor component and the capacitor 424A constitute a parallel resonance circuit. This capacitor 424A is referred to as a first parallel capacitor 424A in order to distinguish it from other capacitors.

  In addition, a capacitor 422A for adjusting the resonance frequency and a transmission / reception magnetic coupling prevention circuit 220 are inserted in series in the first loop 421A. The capacitor 422A is referred to as a first series capacitor 422A in order to distinguish it from other capacitors. Here, a case where two first series capacitors 422A are provided is illustrated, but the number of first series capacitors 422A may be one or more.

  As described above, the first subcoil 410A of the present embodiment is the first magnetic coil adjustment unit 441A and the first loop coil 421A inserted in series with respect to the inductor component of the first loop 421A as an adjustment circuit element. A series capacitor 422A and a first parallel capacitor 424A inserted in series with the inductor component and having the first loop coil section 420A as a parallel resonant circuit are provided. By adjusting the values of these circuit elements, the resonance frequency of the subcoil (single subcoil) and the array coil can be adjusted to a desired resonance frequency.

  One terminal on the loop coil unit 420 side of the low impedance signal processing circuit 430 is connected to one end of the parallel capacitor 424 of the loop coil unit 420 via the magnetic coupling adjustment unit 441. The other terminal on the loop coil unit 420 side of the low impedance signal processing circuit 430 is directly connected to the other end of the parallel capacitor 424 of the loop coil unit 420. The other terminal of the low impedance signal processing circuit 430 that is not on the loop coil section 420 side is connected to the receiver 162.

  The transmission / reception magnetic coupling prevention circuit 220 removes magnetic coupling with the birdcage type RF coil 300 which is the transmission RF coil 151.

  In one embodiment, the second subcoil 410B has the same configuration as the first subcoil 410A. That is, the second subcoil 410B includes a second loop coil unit 420B, which is a parallel resonance circuit, a second low impedance signal processing circuit 430B, a second loop coil unit 420B, and a second low impedance signal processing circuit 430B. And a second magnetic coupling adjustment unit 441B. The second loop coil section 420B includes a loop formed by a conductor (second loop 421B), a second series capacitor 422B inserted in series with the inductor component of the second loop 421B, A second parallel capacitor 424B inserted in series with the inductor component and having the second loop coil section 420B as a parallel resonant circuit.

  Furthermore, the array coil 400 includes a magnetic coupling prevention circuit 220 that prevents magnetic coupling with the birdcage type RF coil 300. This inter-transmission / reception magnetic coupling prevention circuit 220 is inserted in series in the loop 421 of each subcoil constituting the array coil 400, and together with the inter-transmission / reception magnetic coupling prevention circuit 210 inserted in the birdcage type RF coil 300, drives the magnetic coupling prevention circuit. Connected to device 180. FIG. 4B shows a configuration example of the inter-transmission / reception magnetic coupling prevention circuit 220 inserted into the loop 421 and the connection between the inter-transmission / reception magnetic coupling prevention circuit 220 and the magnetic coupling prevention circuit driving device 180.

  In the illustrated example, the transmission / reception magnetic coupling prevention circuit 220 includes a PIN diode 221, an inductor 222, and a control signal line 223. The PIN diode 221 and the inductor 222 are connected in series, and are connected in parallel to the capacitor 423. Note that the capacitor 423 is a capacitor inserted into the loop 421. A control signal line 223 is connected to both ends of the PIN diode 221. The control signal line 223 is connected to the magnetic coupling prevention circuit driving device 180. A choke coil is inserted into the control signal line 223 (not shown) in order to avoid high frequency mixing. Inductor 222 and capacitor 423 are adjusted to resonate in parallel at the frequency of the received nuclear magnetic resonance signal.

  A parallel resonance circuit generally has a characteristic of high impedance (high resistance) at a resonance frequency. Therefore, when a current flows through the PIN diode 221, the PIN diode 221 is turned on, and the capacitor 423 of the loop 421 enters into a high impedance state in parallel with the inductor 222 at the frequency of the received nuclear magnetic resonance signal. Therefore, a part of the loop coil unit 420 becomes high impedance and is opened at the frequency of the nuclear magnetic resonance signal to be received, and the subcoil 410 having the loop coil unit 420 is also opened.

  Thus, when the current flows through the PIN diode 221 and is turned on, the magnetic coupling between each of the subcoils 410A and 410B and the birdcage type RF coil 300 is removed. Therefore, the magnetic coupling between the array coil 400 using each subcoil 410 as a coil element and the birdcage type RF coil 300 is also removed.

  The number of transmission / reception magnetic coupling prevention circuits 220 inserted in the subcoil 410 is not limited to this. Two or more may be inserted into each loop 421. The magnetic coupling can be sufficiently lowered by inserting a plurality.

  The configuration of the transmission / reception magnetic coupling prevention circuit 220 is not limited to the above configuration. For example, a cross diode 221m may be used instead of the PIN diode 221 as in a modification of the inter-transmission / reception magnetic coupling prevention circuit 220m shown in FIG. As a result, when a large signal flows through the conductors constituting the loop 421, the cross diode 221m is turned on, and the capacitor 423 of the loop 421 resonates in parallel with the inductor 222 at the frequency of the received nuclear magnetic resonance signal to generate high impedance. It becomes a state. In this case, the magnetic coupling prevention circuit driving device 180 may not be provided.

  FIG. 5 shows an arrangement example of the above-described birdcage type RF coil 300 and array coil 400 in the MRI apparatus. As shown in the drawing, the birdcage type RF coil 300 is arranged such that its axis is coaxial with the central axis of the magnet 110. That is, it is arranged in the MRI apparatus so that the static magnetic field direction and the magnet axial direction are parallel. Since transmission signals having a phase difference of 90 degrees are input, the high-frequency magnetic field generated from the birdcage type RF coil 300 becomes a rotating magnetic field in a plane substantially orthogonal to the static magnetic field direction. The array coil 400 is disposed in the birdcage type RF coil 300. At this time, the central axis of the loop 421 of the sub-coil 410 constituting the array coil 400 is substantially orthogonal to the static magnetic field direction, and a component orthogonal to the static magnetic field direction of the nuclear magnetic resonance signal generated from the inspection object is detected.

  Based on the configuration of the transmission coil (RF coil 300) 151 and the reception coil (array coil 400) 161 described above, a specific embodiment of the array coil 400 that is a reception coil will be described below.

<< First Embodiment >>
The high-frequency array coil of this embodiment includes two subcoils (a first subcoil and a second subcoil), and includes means for magnetically coupling to the second subcoil when receiving a signal from the first subcoil, and by magnetic coupling. A characteristic is that the phase difference between the currents flowing through both subcoils is adjusted to a specific range in relation to the angle formed by the central axes of the loops of the two subcoils. Further, the magnetic coupling is asymmetric and is not magnetically coupled to the first subcoil when viewed from the second subcoil.

Hereinafter, the details of the high-frequency array coil of this embodiment will be described with reference to FIG. In FIG. 6, the same elements as those in FIG. 4 are denoted by the same reference numerals, and redundant description is omitted.
In FIG. 6, the inter-transmission / reception magnetic coupling prevention circuit 220 shown in FIG. 4 is representatively shown by the capacitor 423 constituting the circuit. A low input impedance signal amplifier 431 is shown as the low impedance signal processing circuit 430. Furthermore, the magnetic coupling adjustment unit 441 may be at least one of a capacitor and an inductor, but here, an inductor is illustrated.

  As shown in FIG. 6A, the first subcoil 410A further includes a first coupled inductor 451A, and the second subcoil 410B further includes a second coupled inductor 451B. The first subcoil 410A is magnetically coupled to the second subcoil 410B with a mutual inductance M between the first coupled inductor 451A and the second coupled inductor 451B. The coupled inductor 451 extends a part of the loop 421 and is disposed at the extended tip. With this configuration, the two subcoils 410A and 410B can be magnetically coupled even when they are disposed at relatively distant positions.

Next, adjustment of the array coil 400 based on the above-described magnetic coupling will be described. As described above, the adjustment is performed mainly using the capacitors 422 and 424 of the loop coil unit 420 and the inductor (or capacitor) of the magnetic coupling adjustment unit 441 as adjustment elements.
First, in the array coil 400 of the present embodiment, the first sub-coil 410A and the second sub-coil 410B are adjusted so that each can receive a nuclear magnetic resonance signal.

  The first subcoil 410A is adjusted so that the resonance frequency of the first subcoil 410A alone is different from the nuclear magnetic resonance frequency that is the frequency of the nuclear magnetic resonance signal to be transmitted and received. The first sub-coil 410A is intentionally magnetically coupled to the second sub-coil 410B, whereby a current path that circulates in each of the first loop 421A and the second loop 421B is formed. Adjustment is made so as to resonate at the magnetic resonance frequency (f0).

  Further, the first subcoil 410A is adjusted so that the series resonance circuit composed of the magnetic coupling adjustment unit 441A and the capacitor 424A resonates at f0. On the other hand, the second subcoil 410B is adjusted so that the series resonance circuit including the magnetic coupling adjustment unit 441B and the capacitor 424B resonates at a frequency different from f0. For example, it is adjusted to be lower than f0.

  In such adjustment, the second subcoil 410B is magnetically coupled via the mutual inductance M when viewed from the first subcoil 410A. This is because the series resonance circuit composed of the magnetic coupling adjustment unit 441B and the capacitor 424B resonates at a frequency different from f0, so that when the signal of the first subcoil 410A is received, both ends of the capacitor 424B of the second subcoil 410B have high resistance. This is because it must not.

  On the other hand, when viewed from the second subcoil 410B, the first subcoil 410A is not magnetically coupled. This is because the series resonance circuit composed of the magnetic coupling adjustment unit 441A and the capacitor 424A resonates at f0, and therefore, when the signal of the second subcoil 410B is received, both ends of the capacitor 424A of the first subcoil 410A have high resistance. .

  In the array coil 400 of this embodiment, in addition to the adjustment of the resonance frequency described above, the phase of the current flowing through the first subcoil 410A and the second subcoil is adjusted when the signal of the first subcoil 410A is received. Specifically, the phase difference of the current flowing through the second subcoil 410B due to magnetic coupling (hereinafter referred to as current phase difference) with respect to the phase of the current flowing through the first subcoil 410A is defined as the first and second subcoils 410A, Adjustment is made so as to obtain an appropriate phase difference in relation to the arrangement angle 410B. As shown in FIG. 6B, the sub-coil arrangement angle is an angle formed by the center axis 11A of the first sub-coil 410A (loop 421A) and the center axis 11B of the second sub-coil 410B (loop 421A). Angle between axes). The current phase difference is adjusted in a range of ± 90 degrees with respect to the angle between the central axes, preferably in a range of ± 60 degrees. When the sign of the angle is also taken into account, it is precisely ± 60 degrees with respect to the phase that is the same as the angle and has the opposite sign.

  The magnitude of the current flowing through the second sub-coil 410B and the current phase difference due to the magnetic coupling include the magnitude of the inductance (capacitance in the case of a capacitor) of the magnetic coupling adjustment unit 441B, the magnitude of the impedance of the low input impedance signal amplifier 431B, It controls by the magnitude | size of 451A and 451B and the magnitude | size of the coupling coefficient.

Hereinafter, the principle of adjusting the current phase difference will be described.
The current flowing through the first subcoil 410A can be expressed by the following equation (1).
[Equation 1]
I ₁₁ = k ₁ sin (2πf ₀ t + θ) (1)
Here, k ₁ represents the magnitude of the current, and θ represents the current phase offset with respect to time t = 0. When this current flows through the first subcoil 410A, a magnetic field that fluctuates in the second subcoil 410B is generated. Therefore, an induced electromotive force V is generated in the second subcoil 410B so as to cancel the magnetic flux generated by the first subcoil 410A across the loop 421B of the second subcoil 410B. Using the proportional coefficient k ₂ , the induced electromotive force V is
[Equation 2]
V = k ₂ (dI ₁₁ / dt)
= 2πf ₀ k ₁ k ₂ cos (2πf ₀ t + θ) (2)
It is expressed.

  The second subcoil 410B magnetically coupled to the first subcoil 410A can be represented by an equivalent circuit 600 shown in FIG. 7A (from Non-Patent Document 1). In this equivalent circuit 600, an inductor 650, a capacitor 622, and a capacitor 624 are inserted in series with respect to the voltage source 690 of the voltage V expressed by Expression (2), and an inductor 641 and a resistor 633 are inserted in parallel. . Here, as viewed from the voltage source 690, the series resonant circuit including the inductor 650, the capacitor 622, and the capacitor 624 is assumed to have a resistance component that is ideally small enough to be ignored, and only the parallel resonant circuit portion viewed from the voltage source 690 is used. When extracted, an equivalent circuit 610 as shown in FIG.

The impedance characteristics of such a parallel resonant circuit 610 are shown in FIGS. 7 (c) and 7 (d). FIG. 7C shows the impedance strength, and FIG. 7D shows the impedance phase. As can be seen from these impedance characteristics, the parallel resonant circuit 610, with respect to the resonance frequency f _R, has an inductive reactance dominant at frequencies lower than f _R, the current phase lags the voltage phase. will dominate the capacitive reactance at a frequency higher than f _R, the current phase is ahead of the voltage phase.

Here, it is assumed that the resonance frequency of the equivalent circuit 600 of the magnetically coupled second subcoil 410B is the frequency (nuclear magnetic resonance frequency) f0 of the signal to be detected. In that case, since the series resonant circuit including the inductor 650, the capacitor 622, and the capacitor 624 is adjusted to resonate at f0, the impedance at f0 is zero. Therefore, as shown by the equivalent circuit 601 in FIG. 7E, only the impedance Z of the parallel resonant circuit composed of the inductor 641 and the capacitor 624 remains. Phase of the impedance Z, as shown in FIG. 7 (d), is 0 in the resonance point f _R. In the preamplifier decoupling method, since f _R = f0 is adjusted, the phase of the current flowing through the second subcoil 410B is not different from the phase of the voltage source 690. Therefore, from the formula (1) and the formula (2), in the preamplifier decoupling method, the phase of the current generated in the second subcoil 410B by magnetic coupling is 90 degrees with respect to the phase of the current flowing in the first subcoil. Cannot be controlled.

In contrast, in the second subcoil 410B of this embodiment, the resonance frequency (f _R ) of the resonance circuit including the magnetic coupling adjustment unit 441B and the capacitor 424B is adjusted to be lower than f0. Therefore, capacitive reactance is dominant at f0, and the current phase precedes the voltage phase. That is, it can be adjusted to be greater than 90 degrees with respect to the phase of the current flowing through the first subcoil. Therefore, the current phase difference can be changed depending on the size of the magnetic coupling adjustment unit 441B (inductance or capacitance).

For the second subcoil 410B, the resonance frequency (f _R ) of the resonance circuit including the magnetic coupling adjustment unit 441B and the capacitor 424B may be adjusted to be higher than f0. In this case, as shown in FIGS. 7B and 7C, inductive reactance is dominant at f0, and the current phase is delayed from the voltage phase. That is, it can be adjusted to be smaller than 90 degrees with respect to the phase of the current flowing through the first subcoil. Therefore, even when the angle between the central axes is smaller than 90 degrees, the current phase difference can be adjusted to a desired value by the size of the magnetic coupling adjustment unit 441B (inductance or capacitance).

  The order of adjusting the resonance frequency and the current phase difference is not limited. For example, first, the mutual inductance M of the coupled inductors 451A and 451B and the resonance frequency of each loop coil unit 420 are adjusted. Next, the current phase difference is adjusted by the magnetic coupling adjustment unit 441. Finally, the capacitor 422 of each loop coil unit 420 is finely adjusted as necessary.

  Next, the principle of detection of a nuclear magnetic resonance signal (rotating magnetic field) in the high-frequency array coil 400 of the present embodiment adjusted as described above will be described.

  First, as a premise, the arrangement of the high-frequency array coil 400 in the MRI apparatus will be described with reference to FIGS. In FIG. 6A, as shown in the coordinate system 090, the vertical direction on the paper surface is the X-axis direction, the horizontal direction is the Z-axis direction (the direction opposite to the static magnetic field direction), and the vertical direction to the paper surface is the Y-axis direction. In FIG. 6B, as shown in the coordinate system 090, the vertical direction on the paper is the Y-axis direction, the horizontal direction is the X-axis direction, and the vertical direction is the Z-axis direction.

  Here, as an example, the loops 421A and 421B of the first loop coil portion 420A and the second loop coil portion 420B are arranged on a cylindrical surface with the Z axis as the central axis. An axis passing through the center of the loop 421 and perpendicular to the curved surface of the loop 421 is taken as a central axis 11. In the example shown in FIG. 6B, the central axis 11A of the first first loop coil portion 420A and the central axis 11B of the second loop coil portion 420B are on substantially the same XY plane perpendicular to the static magnetic field direction. . Here, attention is paid to the detection sensitivity of the rotating magnetic field in the clockwise direction in the XY plane as viewed from the positive side of the Z axis. The arrangement angle between the first sub-coil 410A and the second sub-coil 410B is defined by the angle formed by the respective central axes 11A and 11B. Here, the direction opposite to the direction of the rotating magnetic field in which the angle between the central axes is to be detected is assumed to be positive, and is represented by −180 degrees to +180 degrees. Thus, when the direction opposite to the direction of the rotating magnetic field for which the arrangement angle is to be detected is defined as positive, the current phase difference is adjusted to be the same as the arrangement angle (the difference is 0). In the definition of the arrangement angle, the angle of the central axis 11B with respect to the central axis 11A is “positive”, and is +127 degrees in the illustrated example.

In general, the detection sensitivity of each subcoil 410 is proportional to the rotating magnetic field generated by each subcoil 410 per unit power supply from the reciprocity theorem. The magnetic fields created by the first sub-coil 410A and the second sub-coil 410B in the deep portions are oscillating magnetic fields 510A and 510B as shown in FIG. 6B, respectively. The principle of improving the detection sensitivity of the rotating magnetic field will be described with reference to FIG. FIG. 8A shows the relationship between the oscillating magnetic field and the rotating magnetic field. The oscillating magnetic field 510 can be represented by a combination of a clockwise rotating magnetic field 520 and a counterclockwise rotating magnetic field 530. Here, attention is paid only to the clockwise rotating magnetic field, and the magnitude thereof is defined as sensitivity S. FIG. 8B shows a rotating magnetic field when the current phase difference between the two subcoils is 0 degrees. The clockwise rotating magnetic field 520A generated in the deep part by the first subcoil 410A is directed in the same direction as the central axis 11A at a certain time t, and rotates clockwise therefrom at an angular velocity of 2πf0. Its size is referred to as _{S 11.} On the other hand, the clockwise rotating magnetic field 520B generated by the current flowing through the second subcoil 410B due to the magnetic coupling is directed toward the central axis 11B at the same time because the current phase difference is 0 degree, and from there the angular velocity clockwise Rotates at 2πf0. Its size is referred to as _{S 21.} In this case, the synthetic clockwise rotating field 540 becomes a combination of the clockwise rotating field 520A and clockwise rotating field 520B, a size _{S 1} is has become less than _{S 11.}

FIG. 8C shows a case where the current phase difference is equal to the central axis angle. The clockwise rotating magnetic field 520A is the same as that in FIG. On the other hand, since the clockwise rotating magnetic field 520B is rotated by the current phase difference at the time t, it is directed in the same direction as the clockwise rotating magnetic field 520A. Therefore synthetic clockwise rotating magnetic field 540 is also oriented in the same direction as the clockwise rotating field 520A, its size S ₁ is greater than S _11. When the difference between the current phase difference and the central axis angle is Δφ, the magnitude S ₁ of the synthetic clockwise rotating magnetic field 540 is expressed as follows.
[Equation 3]
S ₁ = S ₁₁ + S ₂₁ cos (Δφ) (3)

  FIG. 9 shows the relationship of the center sensitivity with respect to the difference Δφ between the current phase difference and the center axis angle expressed by the equation (3). As can be seen from FIG. 9, in principle, if the difference between the current phase difference and the angle between the central axes is less than ± 90 degrees, it is strengthened as a rotating magnetic field.

  In order to confirm this, an electromagnetic field simulator was used for calculation. As the array coil, an array coil in which square subcoils 410A and 410B each having a side of 100 mm are arranged on a cylindrical surface having a radius of 122 mm was modeled. The input impedance of the low input impedance signal amplifier 431 was 2Ω. As the center sensitivity, the combined sensitivity of the square sum of squares of the sensitivity of the two subcoils 410A and 410B was calculated.

  When the central axis angle of the loop is set to 127 degrees and the inductance of the magnetic coupling adjustment unit 441B is 55 nH, the current phase difference is 125 degrees (that is, the difference between the current phase difference and the central axis angle is 2 degrees) The center sensitivity was 0.825 A / m / √W. On the other hand, when the loop center axis angle is 233 degrees (-127 degrees) and the inductance of the magnetic coupling adjustment unit 441B is 55 nH, the current phase difference is 125 degrees (that is, the current phase difference and the center axis angle). The central sensitivity was 0.779 A / m / √W. For comparison, in the case of the two subcoils 410 that prevent magnetic coupling by the preamplifier decoupling method, the center sensitivity is 0.791 A / m / √W when the angle between the center axes of the loops is 127 degrees.

  From the results of this simulation, it was confirmed that the sensitivity at the deep portion was improved as compared with the conventional method (preamplifier decoupling method) when the difference between the current phase difference and the central axis angle was less than ± 90 degrees.

On the other hand, as described above, there exists a condition in which the rotating magnetic field weakens due to magnetic coupling. Therefore, as can be seen from the example of the simulation result, when the two subcoils 410A and 410B are magnetically coupled to each other, even if the rotating magnetic field is strengthened on the one hand, it may be weakened on the other hand. Therefore, in the present embodiment, by adjusting the subcoil 410A so as not to be magnetically coupled when viewed from the subcoil 410B, an effect is obtained only when the rotating magnetic field is strengthened.
In the adjustment of the current phase difference described above, the case where the current phase difference is adjusted with respect to the predetermined center axis angle of the subcoil has been described. However, the current phase difference is within an appropriate range in relation to the center axis angle of the subcoil. For example, the case where the relationship between the current phase difference and the center-to-center angle is adjusted by adjusting the arrangement of the subcoils with respect to a predetermined current phase difference is also included in the present embodiment.

  As described above, the high-frequency array coil of the present embodiment includes a first subcoil having a first loop coil portion made of a conductor and a second subcoil having a second loop coil portion made of a conductor, The high-frequency array coil is adjusted as follows. First, itself is adjusted so that a high-frequency signal (a signal having a nuclear magnetic resonance frequency or a nuclear magnetic resonance signal) can be transmitted or received. The first sub-coil has a resonance frequency of the first sub-coil that is different from a nuclear magnetic resonance frequency and is magnetically coupled to the second sub-coil, so that the loop of the first loop coil section and the second sub-coil Each of the loop coil sections is adjusted to form a current path that circulates and resonates at the nuclear magnetic resonance frequency. Further, the phase difference between the current flowing in the loop of the first loop coil portion and the current flowing in the loop of the second loop coil portion is such that the central axis of the first loop coil portion and the second loop The phase difference is the same as the angle formed by the central axis of the coil part and is adjusted to the range of the phase difference of ± 90 degrees (preferably ± 60 degrees) with the opposite sign. On the other hand, the second subcoil is a subcoil in which the resonance frequency of the second subcoil alone is a nuclear magnetic resonance frequency, and itself transmits or receives the high-frequency signal.

  According to the present embodiment, it is possible to realize a high depth sensitivity that exceeds the array coil in which magnetic coupling is prevented by using the preamplifier decoupling method. Therefore, the array coil of this embodiment can realize high image quality. In addition, since the array coil of this embodiment realizes asymmetric magnetic coupling and adjustment of the current phase difference by adjusting the values of the circuit elements of each subcoil 410, the structure is not complicated. Furthermore, the sensitivity distribution in the imaging region of both subcoils is different in the array coil of this embodiment. Therefore, the array coil of this embodiment operates as an array coil applicable to the MRI speed-up technology.

  The array coil according to the present embodiment and the MRI apparatus using the array coil have been described with reference to the drawings. The array coil according to the present embodiment has two subcoils constituting the array coil viewed from one side and the other is magnetically coupled. When viewed from the other side, one is not magnetically coupled, and the phase difference of the current flowing through the subcoil when magnetically coupled is adjusted to a predetermined range in relation to the subcoil arrangement angle. As long as it has the characteristic that it is, it is not limited to the said description and drawing, A various change is possible. For example, although the figure shows the case where the shape and size of the two subcoils are the same, the shape and size may be different. Hereinafter, some modified examples will be described.

<Modification Example 1 of First Embodiment>
In the configuration shown in FIG. 6, the coupled inductor 451 is used as means for magnetically coupling the two subcoils, but the coupled inductor may not be provided. An example in that case (high-frequency array coil 401) is shown in FIG. In this modification, the two arrangements are arranged with the sub-coils 410A and 410B close to each other and magnetically coupled by the magnetic field generated by the loop 421. Thereby, the number of parts can be reduced and the configuration can be simplified. Further, since the loss due to the coupled inductor 451 is eliminated, the sensitivity of the subcoil 410 can be improved.
Furthermore, since the two subcoils 410A and 410B are physically separated, it is possible to configure a split type array coil.

<Modification 2 of the first embodiment>
In the configuration shown in FIG. 6B, the case where the two subcoils are arranged at a specific center axis angle (127 degrees) of more than 90 degrees and less than 180 degrees is shown, but the arrangement of the coils is not limited to this. Absent. In the range of 0 to 180 degrees or 0 to -180 degrees, the current phase difference and the central axis angle can be adjusted to a desired range by the adjustment procedure described above.

Depending on the arrangement angle, for example, the shape of the loop surface may be changed from a flat shape to a curved surface, and a means for magnetically coupling the two subcoils (whether a coupled inductor is used) is appropriately selected. Can do. When a coupled inductor is used as the magnetic coupling means, the coupling coefficient of the coupled inductors 451A and 451B is increased with respect to a value that minimizes the magnetic coupling. This is because the current phase is inverted when the coupling coefficient is large and small with respect to the value that minimizes the magnetic coupling.
According to this modification, the degree of freedom of arrangement of the second subcoil 410B increases.

<Modification 3 of the first embodiment>
In the first embodiment, the current phase difference is adjusted by the magnetic coupling adjustment unit 441B, but the adjustment method is not limited to this. For example, it may be controlled by the magnitude of the impedance of the low input impedance signal amplifier 431B, the magnitude of the coupled inductors 451A and 451B, and the magnitude of the coupling coefficient. By adjusting the magnitude of the impedance of the low input impedance signal amplifier 431B, the magnitude of the current phase difference component of 90 degrees or -90 degrees is changed, thereby changing the combined current phase difference. Further, by adjusting the size of the coupled inductors 451A and 451B and the size of the coupling coefficient, the combined current phase difference changes in order to change the magnitude of the current phase difference component of 180 degrees or -180 degrees. This increases the degree of freedom of adjustment.

<Application example of first embodiment>
The high-frequency array coil of the first embodiment can be used as a single array coil unit, and a plurality of array coil units can be arranged in a predetermined direction, for example, in the circumferential direction or axial direction of a cylinder.

<< Second Embodiment >>
The high-frequency array coil of the present embodiment includes the first subcoil and the second subcoil, and includes a means for magnetically coupling the first subcoil and the second subcoil as in the first embodiment. The array coil of the present embodiment is characterized in that the first subcoil and the second subcoil are coils that operate as one reception channel. The array coil of this embodiment may be combined with other subcoils or similar array coil units, with the first subcoil and the second subcoil as one array coil unit.

  The MRI apparatus of this embodiment is basically the same as that of the first embodiment except that the configuration of the high-frequency array coil used as a receiving coil is different, and the following focuses on the configuration different from the first embodiment. And explain. In this embodiment, the direction of the static magnetic field generated by the horizontal magnetic field type magnet 110 is the Z-axis direction of the coordinate system 090.

  The high-frequency array coil of this embodiment includes one or more array coil units each having two subcoils adjusted to be magnetically coupled. The two subcoils constituting the array coil unit are substantially the same as the array coil of the first embodiment, but the low impedance signal processing circuit (low impedance signal amplifier) is connected to only one subcoil, and one set is one channel. It is different to make a coil.

  Hereinafter, a configuration example of the high-frequency array coil of the present embodiment will be described with reference to FIGS. 11 and 12. Here, as an example, a description will be given of a high-frequency array coil 700 that includes two array coil units 790 and six other subcoils 750 and has a splittable structure.

  FIG. 11 is a view of the high-frequency array coil 700 as viewed from the Z-axis direction, and shows the configuration and arrangement of the high-frequency array coil 700. In FIG. 11, two sets of coils surrounded by a dotted-line square are array coil units 790, and the other six coils are subcoils 750. These subcoils are disposed so as to surround the circumferential direction with respect to the Z axis and to have an overlap region between adjacent subcoils. The area of the overlap region is determined so as not to be magnetically coupled to each other. The inspection object 103 is arranged in a cylindrical space in the high frequency array coil 700. The high-frequency array coil 700 can be physically divided by a line (partition line) 12 indicated by a one-dot chain line in the drawing.

  Since the configuration of the six subcoils 750 is the same, one subcoil will be described with reference to FIG. The subcoil 750 has a series capacitor 722 having a function of preventing magnetic coupling with the transmission coil with respect to the loop 721, a capacitor 723 that adjusts the resonance frequency, a capacitor 724 that adjusts matching, an adjustment inductor 740, a low An input impedance signal amplifier 731 is provided. The series resonance circuit composed of the adjustment inductor 740 and the capacitor 724 is adjusted so that the resonance frequency is f0 of the nuclear magnetic resonance frequency.

  Each array coil unit 790 includes a first subcoil 710A and a second subcoil 710B, as shown in FIG. 12 (b). Unlike the second subcoil 410B of the first embodiment, the second subcoil 710B includes only the second loop coil portion 720B. One or more capacitors for adjusting the resonance frequency and a magnetic coupling adjustment unit 741B are inserted in the second loop coil unit 720B. The magnetic coupling adjustment unit 741B includes at least one of a capacitor and an inductor. Here, an example of a capacitor is shown. Between the first subcoil 710A and the second subcoil 710B, there is a dividing line 12 indicating a boundary dividing the array coil 700. That is, the first subcoil 710A and the second subcoil 710B are physically separated by the dividing line 12, but are disposed at such a distance that the loop coil portions 720 are magnetically coupled with each other with a predetermined mutual inductance M. . The value of the mutual inductance M is adjusted by adjusting the distance. Further, the angle of the central axis 11B of the second subcoil 710B with respect to the central axis 11A of the first subcoil 710A is arranged at 45 degrees in the direction opposite to the direction of the rotating magnetic field (clockwise) of interest (FIG. 11).

  FIG. 12B shows the first subcoil 710A and the second subcoil 710B with the same shape and the same size, but the shape and size may be different depending on the subcoil. For example, in the example shown in FIG. 11, the first subcoil 710A is larger in size than the second subcoil 710B.

Next, adjustment of the array coil unit 790 having the above configuration will be described.
The first subcoil 710A is adjusted so as to receive a nuclear magnetic resonance signal. The first subcoil 710A is adjusted so that the resonance frequency of the first subcoil 710A alone is different from the nuclear magnetic resonance frequency f0 that is the frequency of the nuclear magnetic resonance signal to be transmitted and received. The second subcoil 710B is composed of only the loop coil portion 720B and does not receive a nuclear magnetic resonance signal alone. The resonance frequency of the second subcoil 710B is adjusted to be larger than f0 by the value of the capacitance (inductance in the case of an inductor) of the magnetic coupling adjustment unit 741B.

  By intentionally magnetically coupling the two subcoils 710A and 710B, a circulating current path is formed in the first loop coil portion 720A and the second loop coil portion 720B, respectively, and resonance occurs at f0. Adjusted. Thereby, it can be considered that the first subcoil 710A is combined with the second subcoil 710B arranged in a non-contact manner across the dividing line 12 and operates as one coil. The second subcoil 710B is disposed at a position where it is magnetically coupled to the first subcoil 710A. By adjusting the mutual inductance M adjusted at this position (distance between the coils), the resonance frequency of the first subcoil 710A alone and the resonance frequency of the first subcoil 710A when the second subcoil 710B is arranged are made different. be able to.

  Further, the array coil unit 790 is configured such that the phase difference of the current flowing through the second subcoil 710B is adjusted to 45 ° ± 60 ° with respect to the phase of the current flowing through the first subcoil 710A when the first subcoil 710A receives the signal. Yes. This means that it is ± 60 degrees with respect to the arrangement angle of the central axis 11B with respect to the central axis 11A. The magnitude of the current flowing through the second subcoil 710B by the magnetic coupling and the phase difference of the current are adjusted by the magnitude of the capacitance of the magnetic coupling adjustment unit 741B.

In the array coil unit 790, the principle that the current phase difference can be adjusted by the magnetic coupling adjustment unit 741B is the same as in the first embodiment. That is, as described with reference to FIGS. 7C and 7D, the first subcoil 710A receives a signal by changing the resonance frequency (f _R ) of the second subcoil 710B by the magnetic coupling adjustment unit 741B. The difference in the phase of the current flowing in the second subcoil 710B with respect to the phase of the current flowing in the first subcoil 710A can be changed. In this embodiment, since the resonance frequency of the second subcoil 710B is adjusted to be larger than f0, the second subcoil 710B operates as an inductive reactance (inductor) when viewed from the first subcoil 710A. Therefore, as shown in FIG. 7D, the current phase difference can be adjusted to 0 degree to 90 degrees. By matching the current phase difference with the arrangement angle of the central axis 11B with respect to the central axis 11A in this way, the current flowing through the second subcoil 710B strengthens the rotating magnetic field and improves the depth sensitivity when receiving a signal from the first subcoil 710A.

  In the adjustment of the current phase difference described above, the example in which the magnetic coupling adjustment unit 741B is a capacitor and adjusts the capacitance is shown, but the present invention is not limited thereto. For example, an inductor may be used. In this case as well, the same effect can be obtained by changing the resonance frequency of the second subcoil 710B.

  In the above description, the second subcoil 710B has an angle between the central axes of 45 degrees with respect to the first subcoil 710A, and the resonance frequency of the second subcoil 710B is greater than f0. However, the present invention is not limited to this. . When the center axis angle is 90 degrees, the resonance frequency of the second subcoil 710B is made equal to the magnetic resonance frequency f0, and when the center axis angle is 0 degrees to 90 degrees, the resonance frequency of the second subcoil 710B is set to the magnetic resonance frequency. When it is larger than f0 and the angle between the central axes is 90 degrees to 180 degrees, the resonance frequency of the second subcoil 710B is made smaller than the magnetic resonance frequency f0. Thereby, it can adjust to a desired electric current phase difference with respect to all the angles between central axes of 0 degree-180 degree | times, and the arrangement | positioning freedom degree of 2nd subcoil 710B becomes large.

  The array coil unit 790 adjusted in this way operates as one receiving channel by forming a current path around each loop conductor by magnetically coupling the first subcoil 710A and the second subcoil 710B. The first subcoil 710A and the second subcoil 710B are arranged so as to have a predetermined overlap area and are not magnetically coupled to the subcoils 750 on both sides. In addition, since the two array coil units 790 are disposed at a position of approximately 180 degrees with the inspection object 103 interposed therebetween, high depth sensitivity can be obtained over the entire inspection object 103.

As described above, the high-frequency array coil of the present embodiment adjusts the resonance frequency to the nuclear magnetic resonance frequency in addition to the first subcoil and the second subcoil having the same configuration as the high-frequency array coil of the first embodiment. And one or more subcoils disposed so as not to be magnetically coupled to each other. However, in the high frequency array coil of the present embodiment, the first low impedance signal processing circuit is connected only to the first subcoil among the first subcoil and the second subcoil.
The high-frequency array coil of the present embodiment includes two or more sets of the first subcoil and the second subcoil that are physically separated, the two sets being arranged at axially symmetric positions, and the two sets of the first subcoil and the second subcoil. The first subcoil and the second subcoil are separated from each other as a dividing line.

  According to this embodiment, it is possible to improve the depth sensitivity as in the first embodiment. The array coil unit 790 operates as one reception channel, but has a first subcoil 710A and a second subcoil 710B that are physically separated from each other, so that the array coil unit 790 can be divided without any parts connecting between them. Can do. Therefore, the array coil 700 can be divided and can be easily set to the inspection object. In addition, since the array coil unit 790 operates as one reception channel, the overlapping area can be adjusted so as to minimize the magnetic coupling with the adjacent subcoil 750. Therefore, it can be adjusted so as not to be magnetically coupled to the adjacent subcoil without connecting parts at the division boundary, and the performance of the array coil can be prevented from deteriorating. Further, the connection parts can be reduced or eliminated at the boundary where the division is performed, and the cost due to the addition of the connection parts can be reduced. Furthermore, the performance of the subcoil is prevented from being deteriorated by the influence of the loss due to the connection resistance at the connection point. In addition, the failure risk can be reduced because there is no mechanical deterioration due to poor connection at the connection part or repeated attachment / detachment.

<Modification of Second Embodiment>
The array coil 700 shown in FIG. 12 has a structure in which two array coil units 790 and other subcoils 750 are combined and arranged on the circumference so as to surround the inspection target 103, but the array coil unit 790 and subcoil 750 to be combined are combined. The number and the way of combination are not limited to the structure shown in FIG.

  For example, the array coil unit 790 can be used alone as a receiving coil. It is also possible to replace one of the two array coil units 790 in FIG. 12 with one or more subcoils 750, or conversely replace the upper and lower three subcoils 750 with the array coil units 790, respectively. In the latter case, for example, the second subcoil of the first array coil unit overlaps the first subcoil of the second array coil unit, and the second subcoil of the second array coil unit is third. The array direction of the first subcoil and the second subcoil in each array coil unit 790 is made to be the same so as to overlap the first subcoil of the array coil unit. In an array coil in which a plurality of array coil units 790 are arranged, a coupled inductor can be adopted as means for magnetically coupling the first and second subcoils for the array coil unit 790 that is not divided by the dividing line 12.

Furthermore, it is not essential that the array coil unit 790 and the subcoil 750 surround the entire periphery of the inspection target 103, and a structure that covers a part in the circumferential direction can also be adopted.
As mentioned above, although the modification of 2nd embodiment was illustrated, there exist various modifications, such as combining another type (for example, butterfly type) subcoil and the array coil of 1st embodiment suitably, and making it an array coil. obtain. And the MRI apparatus incorporating the array coils of these modified examples is also included in the MRI apparatus of the present invention.

<< Third embodiment >>
The high-frequency array coil of this embodiment is characterized in that it includes three subcoils (hereinafter referred to as subcoil sets) that have been subjected to specific adjustments. This subcoil set can constitute one array coil alone or in combination with other subcoils. In this embodiment, a high-frequency array coil that is a combination will be described.

  The MRI apparatus according to the present embodiment is basically the same as the first embodiment except that the configuration of the array coil used as the receiving coil is different, and the following focuses on the configuration different from the first embodiment. explain. In this embodiment, the direction of the static magnetic field generated by the horizontal magnetic field type magnet 110 is the Z-axis direction of the coordinate system 090. As in the first embodiment, this embodiment can be applied to both the MRI apparatus 100 including the horizontal magnetic field type magnet 110 and the MRI apparatus 101 including the vertical magnetic field type magnet 111. In the following description, as an example, a case where the MRI apparatus 101 including the vertical magnetic field type magnet 111 is used as an array coil for the head will be described.

  First, the structure of the high-frequency array coil 800 of this embodiment will be described with reference to FIGS. 13 is a view showing the arrangement of the high-frequency array coil 800, FIG. 14A is a development view when the high-frequency array coil 800 is opened on a plane, and FIG. 14B constitutes the high-frequency array coil 800. It is a figure which shows the subcoil 810. FIG.

  As shown in FIG. 13, the array coil 800 includes seven subcoils 810A to 810G. As shown in FIGS. 14A and 14B, each sub-coil 810 includes a capacitor 822 that adjusts the resonance frequency with respect to the loop 821, and a series capacitor 823 that has a function of preventing magnetic coupling with the transmission coil. , A capacitor 824 for adjusting matching, a magnetic coupling adjustment unit 841, and a low input impedance signal amplifier 831. Adjacent subcoils 810G and 810C, 810C and 810E, 810E and 810A, 810A and 810D, 810D and 810B, and 810B and 810F are arranged so that adjacent subcoils have overlap regions. The area of the overlap region is determined so as not to be magnetically coupled to each other. Further, the subcoils 810D to 810G are adjusted so that the resonance frequency of the series resonance circuit including the magnetic coupling adjustment unit 841 and the capacitor 824 becomes the nuclear magnetic resonance frequency f0.

  The subcoils 810D to 810G arranged so as not to be magnetically coupled to each other operate independently and receive nuclear magnetic resonance signals. On the other hand, the sub-coils 810A to 810C are subjected to predetermined adjustment (described later) with respect to the arrangement of the resonance frequency and the strength of magnetic coupling. A combination of these three subcoils is referred to herein as a subcoil set 890. In FIG. 14A, for convenience of drawing, the power supply point (output port) of the subcoil is shown on the left side in 810A to 810C and on the right side in 810D to 810G, but the position of the power supply point is arbitrary.

  Next, the arrangement and adjustment of the array coil 800 will be described with reference to FIG. FIG. 13 is a view of the array coil 800 viewed from the Z axis, and targets a rotating magnetic field in the clockwise direction on the XY plane, as in the first embodiment. The array coil 800 is disposed so as to surround the Z axis that is parallel to the direction of the static magnetic field, and the central axis of each subcoil is in the XY plane that is substantially perpendicular to the Z axis. The central axis 11B of the subcoil 810B is disposed at 45 degrees with respect to the central axis 11A of the subcoil 810A, and the central axis 11C of the subcoil 810C is disposed at −45 degrees with respect to the central axis 11A of the subcoil 810A. Of the three subcoils constituting the subcoil set 890, the subcoil 810A located at the center is the first subcoil, the subcoil 810B arranged in the positive direction of the angle is the second subcoil, and the subcoil 810C arranged in the negative direction of the angle is the first Called three sub-coils.

  Based on such an arrangement, adjustment of the resonance frequency and current phase difference of the subcoil set 890 will be described.

  The first subcoil 810A, like the subcoil 810D other than the subcoil set 890, is adjusted so that the resonance frequency of the series resonance circuit including the magnetic coupling adjustment unit 841A and the capacitor 824A is the nuclear magnetic resonance frequency f0. . On the other hand, in the second subcoil 810B and the third subcoil 810C, the series resonance circuit including the magnetic coupling adjustment unit 841 and the capacitor 824 is adjusted such that the resonance frequency is different from the nuclear magnetic resonance frequency f0. The first subcoil 810A and the second subcoil 810B are magnetically coupled through mutual inductance M1 by adjusting the distance between the two loops, and the first subcoil 810A and the second subcoil 810C are inserted into both the loops. The coupled inductors 851A and 851C are magnetically coupled via the mutual inductance M2.

  This magnetic coupling is asymmetric. That is, when viewed from the first subcoil 810A, the second subcoil 810B is magnetically coupled via the mutual inductance M1. This is because the series resonance circuit composed of the magnetic coupling adjustment unit 841B of the second subcoil 810B and the capacitor 824B resonates at a frequency different from f0, so that when the signal of the first subcoil 810A is received, the capacitor 824B of the second subcoil 810B This is because both ends do not have high resistance. Similarly, since both ends of the capacitor 824C of the third subcoil 810C do not have high resistance, the third subcoil 810C is magnetically coupled through the mutual inductance M2 when viewed from the first subcoil 810A.

  On the other hand, when viewed from the second subcoil 810B, the first subcoil 810A is not magnetically coupled. This is because the series resonance circuit composed of the magnetic coupling adjustment unit 841A of the first subcoil 810A and the capacitor 824A resonates at f0, so that when the signal of the second subcoil 810B is received, both ends of the capacitor 824A of the first subcoil 810A have high resistance. It is because it becomes. Similarly, when viewed from the third subcoil 810C, the first subcoil 810A is not magnetically coupled.

  Due to such magnetic coupling, current flows through the second subcoil 810B and the third subcoil 810C when the signal of the first subcoil 810A is received. Here, the phase difference of the current flowing through the second subcoil 810B is adjusted to 45 degrees ± 60 degrees with respect to the phase of the current flowing through the first subcoil 810A. This means that it is ± 60 degrees with respect to the arrangement angle of the central axis 11B with respect to the central axis 11A. Similarly, when the signal of the first subcoil 810A is received, the phase difference of the current flowing through the third subcoil 810C is adjusted to −45 degrees ± 60 degrees with respect to the phase of the current flowing through the first subcoil 810A. The magnitude of the current flowing through the second subcoil 810B and the third subcoil 810C due to magnetic coupling and the phase difference between the current are the magnitude of the inductance (capacitance in the case of a capacitor) of the magnetic coupling adjustment unit 841, the low input impedance signal amplifier 831 It can be controlled by the size of the impedance, the size of the coupled inductor 851, and the size of the coupling coefficient.

  At this time, a negative current phase difference can be realized by increasing the coupling coefficient of the coupled inductors 851A and 851C with respect to a value that minimizes the magnetic coupling.

  In the example shown in FIG. 14, the magnetic coupling between the first subcoil 810A and the second subcoil 810B is realized by adjusting the distance at which both loops are arranged, and the magnetic coupling between the second subcoil 810A and the third subcoil 810C. The coupling is realized by coupling inductors 851A and 851C inserted in both loops. Which one is adopted as the magnetic coupling means depends on the relationship between the direction of the rotating magnetic field to be detected and the sub-coil arrangement angle. Determined. FIG. 15 shows the magnitude of the mutual inductance caused by the arrangement of the subcoils. The point where the mutual inductance becomes 0 is the arrangement by the overlap method. When the magnetic coupling is realized by the distance between the subcoils, as shown in FIG. 15, the magnitude of the mutual inductance of the subcoils depends on the size of the subcoils but varies depending on the arrangement angle. The hatched portion in FIG. 15 represents the mutual inductance necessary for increasing the depth sensitivity. It is necessary to reverse the sign of the current phase difference depending on whether the arrangement angle is positive or negative. With respect to the region where the graph of FIG. 15 is not included in the shaded area, it is necessary to make the mutual inductance within the shaded area using a coupled inductor as the magnetic coupling means. For example, the arrangement of the subcoil 810C with respect to the subcoil 810A is located in the region (A) in FIG. 15, and a coupled inductor is essential as a magnetic coupling means. On the other hand, the arrangement of the subcoil 810B with respect to the subcoil 810A is located in the region (B) in FIG. 15, and the necessary mutual inductance can be adjusted only by adjusting the distance between the loops. That is, in this range, both the inter-loop distance and the coupling inductor can be employed as the magnetic coupling means.

  As described above, the high-frequency array coil of the present embodiment has a third loop coil portion on the opposite side of the second sub-coil with the same first sub-coil as that of the high-frequency array coil of the first embodiment. And a third subcoil having the third subcoil. The first subcoil is magnetically coupled to the third subcoil to form a current path that circulates in the loop of the first loop coil unit and the loop of the third loop coil unit, respectively, The phase difference between the current flowing in the loop of the first loop coil section and the current flowing in the loop of the third loop coil section is adjusted to resonate at the magnetic resonance frequency, and the first loop coil Is adjusted to the same range as the angle formed by the central axis of the portion and the central axis of the third loop coil portion, and in the range of the phase difference of ± 90 degrees with the opposite sign. Further, the high frequency array coil of the present embodiment has a resonance frequency adjusted to a nuclear magnetic resonance frequency between the first subcoil and the second subcoil and between the first subcoil and the third subcoil, respectively. A fourth subcoil arranged so as not to be magnetically coupled is further provided.

  According to this embodiment, it is possible to improve the depth sensitivity as in the first embodiment. Thereby, high image quality can be realized. In addition, since the sensitivity distribution of each subcoil 810 is different, it functions as a receiving array coil for high-speed imaging, and imaging speed can be increased.

  Although the array coil of the third embodiment has been described above, the array coil of the present embodiment has the configuration shown in FIG. 14 as long as it includes a subcoil set including three subcoils adjusted in a specific relationship. Not limited to this, various modifications and applications are possible. Further, the subcoil set alone may be used for the MRI apparatus. However, as shown in FIG. 13 and FIG. 14, the arrangement density can be increased and the sensitivity can be improved by inserting another subcoil between the three subcoils constituting the subcoil set. As another modification, it is also possible to arrange two sub-coil sets on the top and bottom or on the left and right as in the array coil shown in FIG.

  Further, in the third embodiment, the coupled inductor 851 is inserted only into the subcoils 810A and 810C, but this is not a limitation. For example, a plurality of sub coils may be inserted. Thereby, magnetic coupling with the subcoil farther than the adjacent subcoil can be reduced. This is effective when the magnetic coupling with the subcoil adjacent to each other cannot be removed.

<Modification of Third Embodiment>
In the third embodiment, the array coil 800 having a structure in which subcoils are arranged in a row in the circumferential direction with respect to the Z axis has been described as an example. However, the present invention is not limited to this. For example, a plurality of array coils having the same structure may be arranged in the Z-axis direction. The modification of this embodiment can be used as an abdominal array coil for the MRI apparatus 100 including the horizontal magnetic field type magnet 110, for example. An example of the array coil 801 in this case is shown in FIG. In this modification, an array coil in which subcoils are arranged in a line in the circumferential direction is referred to as an array coil unit.

  The array coil 801 shown in FIG. 16 has a structure in which four array coil units (two 891A and two 891B) are arranged in the Z direction. Of the four array coil units, the array coil unit 891A has the same configuration as the high-frequency coil 800 in the third embodiment described above. On the other hand, array coil unit 891B has subcoils arranged in the same manner as array coil unit 891A, but the resonance frequency of the series resonance circuit of all the subcoils is adjusted to f0, minimizing magnetic coupling between subcoils. It is an array coil unit. These two types of array coil units 891A and 891B are arranged in the Z-axis direction. At that time, the array coil units 891A are arranged so as to be at both ends in the Z-axis direction, and two array coil units 891B are arranged therebetween. The overlap region is adjusted using the overlap method so that the array coil units adjacent in the Z direction are not magnetically coupled to each other.

  In the present modification, adjustment is made so that the arrangement angle of the subcoil 810 that is magnetically coupled in the array coil unit 891A matches the current phase difference, as in the first embodiment. Furthermore, the array coil units arranged in the Z direction are arranged so as not to be magnetically coupled by a known magnetic coupling removal technique, and can be multi-channeled also in the Z direction.

  According to this modification, the same effect as that of the first embodiment, that is, the improvement of the deep sensitivity can be obtained. Furthermore, by increasing the number of channels in the Z direction, it is possible to increase the imaging speed in the Z direction. In addition, the sensitivity of array coils without magnetic coupling tends to decrease at both ends in the Z direction. According to this modification, the sensitivity at both ends in the Z direction can be improved by arranging the array coil units 891A at both ends in the Z direction.

  In this modification, four array coil units are arranged in the Z direction, but the present invention is not limited to this. For example, it may be three or five. Thereby, an array coil having a desired number of channels can be designed.

  In this modification, the array coil units 891A are arranged at both ends in the Z direction, but the arrangement location is not limited to this. For example, the array coil units 891A and 891B may be alternately arranged. Also, all four may be array coil units 891A. Thereby, the sensitivity can be further improved.

  As mentioned above, although each embodiment of the array coil of this invention was described, the array coil of this invention is not limited to these embodiments. For example, among the elements described in the embodiments, some non-essential elements may be added or deleted. Further, these embodiments can be combined as long as there is no technical contradiction.

  The MRI apparatus of the above embodiment includes a static magnetic field generation unit, a gradient magnetic field generation unit that forms a gradient magnetic field in the static magnetic field space formed by the static magnetic field generation unit, and a transmission that irradiates the static magnetic field space with a high-frequency magnetic field pulse. A coil and a receiving coil that receives a nuclear magnetic resonance signal generated from an inspection object arranged in the static magnetic field space, and the high-frequency array coil of each embodiment is used as the receiving coil. The array coil is arranged such that the central axis of the first loop coil portion and the central axis of the second loop coil portion are substantially perpendicular to the direction of the static magnetic field generated by the static magnetic field generating magnet. Yes.

  In the MRI apparatus of the present embodiment, the configuration other than the receiving coil is not limited to the illustrated one, and various changes can be made. For example, although the example using the birdcage type RF coil 300 as the transmission RF coil 151 has been shown, the type of the transmission RF coil 151 is not limited to this. For example, a transmission array coil as shown in the fourth embodiment described below may be used.

<< Fourth Embodiment >>
Next, a fourth embodiment of the present invention will be described. The MRI apparatus of each embodiment described above includes a high frequency array coil that is magnetically coupled to an adjacent subcoil at the time of signal reception and expands a sensitivity region, or a high frequency array coil that includes such a high frequency array coil as a unit. Used for. However, the MRI apparatus of this embodiment uses this array coil as the transmission coil 151.

  The MRI apparatus of the present embodiment has basically the same configuration as the MRI apparatus 100 of the first embodiment. Hereinafter, the present embodiment will be described focusing on the configuration different from the first embodiment.

  As described above, in this embodiment, a high-frequency array coil including a plurality of subcoils 410 is used for the transmission coil 151 instead of the birdcage type RF coil 300.

  FIG. 17A is a diagram for explaining a transmission array coil 900 used as the transmission coil 151 of the present embodiment. Similar to the high frequency array coil 400 of the first embodiment, the transmission array coil 900 of this embodiment includes a plurality of subcoils 910. Here, a case where two subcoils 910A and 910B are provided will be described as an example. The number of subcoils 910 is not limited to this.

  Each subcoil 910 has basically the same configuration as the subcoil 410 of the first embodiment. However, it is not the low input impedance signal processing circuit 430 but the low output impedance signal processing circuit 930 that is connected to the loop coil unit 920 via the magnetic coupling adjustment unit 941. Here, a case where a low output impedance RF amplifier (low output impedance signal amplifier) 931 is provided as the low (output) impedance signal processing circuit 930 will be described as an example.

  The transmission array coil 900 includes a transmission / reception magnetic coupling prevention circuit 210 as a transmission / reception magnetic coupling prevention circuit that prevents magnetic coupling between the transmission coil 151 and the reception coil 161. The transmission / reception magnetic coupling prevention circuit 210 includes a PIN diode 211 and a control signal line 212 connected in series to the loop of the loop coil unit 920, as shown in FIG. Since these functions are the same as those of the inter-transmission / reception magnetic coupling removal circuit 210 shown in FIG. 3B, the redundant description will be omitted, and only the operation during high-frequency magnetic field transmission will be focused on thereafter, and the receiving coil 161 will be opened. This will be described as a state.

  Each circuit element provided in the loop coil portion 920 of each subcoil 910 is basically the same as the array coil 400 of the first embodiment. Adjustment of the sub-coil (adjustment of resonance frequency, adjustment of current magnitude and phase during magnetic coupling) includes a series capacitor 922 inserted in series in the loop 921 of the loop coil unit 920A, and a parallel capacitor 924 inserted in parallel. This is performed by the magnetic coupling adjustment unit 941.

The detailed operation of the transmission array coil 900 of the present embodiment in the above configuration will be described.
At the time of signal transmission, the low output impedance signal amplifier 931 viewed from each of the subcoils 910 constituting the transmission array coil 900 has a low impedance. Therefore, each subcoil 910 of the transmission array coil 900 constitutes the same current distribution as each subcoil 410 of the array coil 400 of the first embodiment. For example, hydrogen nuclear magnetic resonance at a static magnetic field strength of 1.5 T (Tesla). Resonates at a frequency of 64 MHz.

  The MRI apparatus of this embodiment includes a static magnetic field generation unit, a gradient magnetic field generation unit that forms a gradient magnetic field in the static magnetic field space formed by the static magnetic field generation unit, and a transmission coil that irradiates the static magnetic field space with a high-frequency magnetic field pulse And a receiving coil for receiving a nuclear magnetic resonance signal generated from the inspection object arranged in the static magnetic field space, and using the high-frequency array coil of each embodiment as the transmitting coil. The high-frequency array coil is arranged so that the central axis of the first loop coil portion and the central axis of the second loop coil portion are substantially perpendicular to the direction of the static magnetic field generated by the static magnetic field generating magnet. Has been.

  Since each subcoil 910 constituting the transmission array coil 900 of this embodiment resonates at a desired frequency (for example, 64 MHz), RF can be transmitted efficiently. At the same time, the first loop 921A of the first loop coil portion 920A of the first subcoil 910A is coupled to the second loop 921B of the second loop coil portion 920B of the second subcoil 910B by magnetic coupling. Since the difference in phase of the current flowing in the second subcoil 910B due to magnetic coupling with the phase of the current flowing in the first subcoil 910A is adjusted to match the arrangement angle of the two subcoils, Similarly, the depth sensitivity is improved. Therefore, transmission efficiency can be improved and power required for imaging can be reduced.

  Further, the second subcoil 910B is not coupled to the first loop 921A of the first loop coil portion 920A of the first subcoil 910A, and has a sensitivity region different from that of the first subcoil 910A. For this reason, it functions as a multi-channel.

  Therefore, the transmission array coil 900 of this embodiment can realize a coil having a wide RF transmission area while maintaining the number of channels. Further, since this is realized by adjusting the arrangement of the sub-coils 910 and the values of the circuit elements, the structure is not complicated.

  The first implementation of the types of circuit elements used in the low impedance signal processing circuit, the resonance frequency used when adjusting the circuit elements, the location of each subcoil, the presence or absence of a coupled inductor, the loop shape, the size of each subcoil, the magnetic coupling mode, etc. Variations of the embodiment are also applicable to this embodiment. Also in the present embodiment, as in the first embodiment, both the MRI apparatus 100 including the horizontal magnetic field type magnet 110 and the MRI apparatus 101 including the vertical magnetic field type magnet 111 can be applied.

  Although the embodiments of the present invention have been described above, the configurations described in the embodiments and the modifications thereof may be used in combination as appropriate.

  ADVANTAGE OF THE INVENTION According to this invention, in the multi-channel array coil of an MRI apparatus, the technique which raises a deep part sensitivity, and implement | achieves high speed and high image quality can be provided.

11: Loop central axis, 12: Dividing line, 090: Coordinate system, 100: MRI apparatus, 102: Table, 103: Inspection object, 110: Magnet, 111: Magnet, 121: Shim coil, 122: Shim power supply, 131: Inclination Magnetic field coil, 132: Gradient magnetic field power supply, 140: Sequencer, 151: Transmission RF coil, 152: High frequency magnetic field generator, 161: Reception RF coil, 162: Receiver, 170: Computer, 171: Display device, 180: Magnetic coupling Prevention circuit drive device, 210: magnetic coupling prevention circuit between transmission and reception, 211: PIN diode, 212: control signal line, 220: magnetic coupling prevention circuit between transmission and reception, 220m: magnetic coupling prevention circuit between transmission and reception, 221: PIN diode, 221m : Cross diode 222: Inductor 223: Control signal line 300: Bird cage RF coil, 301: straight conductor, 302: end conductor, 303: capacitor, 311: input port, 312: input port, 400: array coil, 410: subcoil, 420: loop coil section, 421: loop, 422: series Capacitor, 423: capacitor, 424: capacitor, 430: low impedance signal processing circuit, 431: low input impedance signal amplifier, 441: magnetic coupling adjustment unit, 451: coupling inductor, 510: oscillating magnetic field, 520: clockwise rotating magnetic field, 530: counterclockwise rotating magnetic field, 540: synthetic clockwise rotating magnetic field, 600: equivalent circuit, 601: equivalent circuit, 622: capacitor, 624: capacitor, 633: resistor, 641: inductor, 650: inductor, 690: voltage source 700: Array coil, 710: Subco 720: loop coil section, 721: loop, 722: series capacitor, 723: capacitor, 724: capacitor, 731: low input impedance signal amplifier, 740: adjustment inductor, 741: magnetic coupling adjustment section, 750: subcoil, 790 : Array coil unit, 800: Array coil, 801: Array coil, 810: Subcoil, 822: Series capacitor, 823: Capacitor, 824: Capacitor, 831: Low input impedance signal amplifier, 841: Magnetic coupling adjustment unit, 851: Coupling Inductor, 890: Subcoil set, 891: Array coil unit, 900: Transmitting array coil, 910: Subcoil, 920: Loop coil section, 921: Loop, 930: Low impedance signal processing circuit, 931: Low output impedance Signal amplifier.

Claims (15)

A first subcoil having a first loop coil portion made of a conductor;
A second sub-coil having a second loop coil portion made of a conductor, and a high-frequency array coil adjusted to transmit or receive a high-frequency signal,
In the first subcoil, the resonance frequency of the first subcoil alone is different from the nuclear magnetic resonance frequency, and the first subcoil is magnetically coupled to the second subcoil. Each of the loops in the loop coil section is adjusted to form a current path that circulates and resonates at the nuclear magnetic resonance frequency.
The phase difference between the current flowing through the loop of the first loop coil portion and the current flowing through the loop of the second loop coil portion is determined by the central axis of the first loop coil portion and the second loop coil portion. Adjusted to be within the same range as the angle formed by the central axis and the phase difference with the opposite sign of ± 90 degrees.
High frequency array coil characterized by The high-frequency array coil according to claim 1,
The phase difference between the current flowing in the loop of the first loop coil portion and the current flowing in the loop of the second loop coil portion is in the range of the same phase difference and a phase difference of ± 60 degrees opposite to the angle;
High frequency array coil characterized by The high-frequency array coil according to claim 1,
The second subcoil is a high-frequency array coil, wherein a resonance frequency of the second subcoil alone is a nuclear magnetic resonance frequency. The high-frequency array coil according to claim 1,
The high-frequency array coil, wherein the second sub-coil is a sub-coil that transmits or receives the high-frequency signal. The high-frequency array coil according to claim 1,
The high-frequency array coil, wherein the first sub-coil and the second sub-coil are disposed at physically separated positions. The high-frequency array coil according to claim 1,
The first loop coil unit and the second loop coil unit each include a coupled inductor, and the first subcoil and the second subcoil are magnetically coupled by the coupled inductor. Array coil. The high-frequency array coil according to claim 1,
At least one of the first subcoil and the second subcoil further includes a magnetic coupling adjustment unit that connects the loop coil unit and a first low impedance signal processing circuit to which the subcoil is connected,
The loop coil portion is
A series capacitor inserted in series with the inductor component of the loop;
A parallel capacitor inserted in series with the inductor component and having the loop coil portion as a parallel resonant circuit;
The high-frequency array coil is adjusted by adjusting values of an adjustment circuit element of the magnetic coupling adjustment unit, the series capacitor, and the parallel capacitor. The high-frequency array coil according to claim 1,
In addition to the first subcoil and the second subcoil, the high frequency array coil further includes one or more subcoils arranged so that a resonance frequency is adjusted to a nuclear magnetic resonance frequency and not magnetically coupled to each other. The high-frequency array coil according to claim 8,
A high frequency array coil, wherein a first low impedance signal processing circuit is connected to only the first subcoil of the first subcoil and the second subcoil. The high-frequency array coil according to claim 8,
Two or more sets of the first sub-coil and the second sub-coil that are physically separated are included, the two sets being arranged at axially symmetrical positions, and between the two sets of the first sub-coil and the second sub-coil. A high-frequency array coil that can be divided using a dividing line. The high-frequency array coil according to claim 1,
A third subcoil having a third loop coil portion on the opposite side of the second subcoil around the first subcoil,
The first subcoil is magnetically coupled to the third subcoil to form a current path that circulates in the loop of the first loop coil unit and the loop of the third loop coil unit, respectively, It is adjusted to resonate at the magnetic resonance frequency,
The phase difference between the current flowing in the loop of the first loop coil portion and the current flowing in the loop of the third loop coil portion is the difference between the central axis of the first loop coil portion and the third loop coil portion. It is adjusted to the same range as the angle formed by the central axis and the phase difference is in the range of ± 90 degrees with the opposite sign.
High-frequency array coil characterized by The high-frequency array coil according to claim 11,
The fourth sub-coil is arranged between the first sub-coil and the second sub-coil and between the first sub-coil and the third sub-coil so that the resonance frequency is adjusted to the nuclear magnetic resonance frequency and is not magnetically coupled to each other. A high-frequency array coil further comprising a subcoil. An array coil composed of a plurality of subcoils arranged along the circumferential direction of the virtual cylindrical body is a high-frequency array coil in which a plurality of array coils are arranged along the axial direction of the virtual cylindrical body,
The high frequency array coil according to claim 8, wherein at least one of the plurality of array coils is the high frequency array coil according to claim 8.   A static magnetic field generator, a gradient magnetic field generator that forms a gradient magnetic field in the static magnetic field space formed by the static magnetic field generator, a transmission coil that irradiates the static magnetic field space with a high-frequency magnetic field pulse, and the static magnetic field space And a receiving coil for receiving a nuclear magnetic resonance signal generated from the inspection object, wherein the high-frequency array coil according to claim 1 is used as the transmitting coil and / or receiving coil. Resonance imaging device. The magnetic resonance imaging apparatus according to claim 14,
The high-frequency array coil is arranged so that the central axis of the first loop coil portion and the central axis of the second loop coil portion are substantially perpendicular to the direction of the static magnetic field generated by the static magnetic field generating magnet. A magnetic resonance imaging apparatus.

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