JP2012183211A - Rf coil apparatus and magnetic resonance imaging apparatus - Google Patents

Abstract

PROBLEM TO BE SOLVED: To prevent inductive currents from flowing within an RF coil for local transmission even if an RF signal is transmitted from an RF coil for the whole body housed in a magnetic resonance imaging (MRI) apparatus while a connector of the RF coil for local transmission set on a gantry is left unconnected to the MRI apparatus.SOLUTION: In one embodiment, the RF coil apparatus is detachably connected to the magnetic resonance imaging apparatus to transmit an RF signal to a subject as an RF signal current fed from the magnetic resonance imaging apparatus flows through a coil element. The RF coil apparatus has a trap circuit inserted into a pathway of the RF signal current in the coil element. The trap circuit selectively blocks a frequency band including a Lamor frequency according to the input voltage.

Description

  Embodiments described herein relate generally to an RF coil device (radio frequency coil device) and a magnetic resonance imaging device.

  MRI is an imaging method in which a nuclear spin of a subject placed in a static magnetic field is magnetically excited with an RF signal having a Larmor frequency, and an image is reconstructed from an MR signal generated by the excitation. The MRI means magnetic resonance imaging, the RF signal means a radio frequency signal, and the MR signal means a nuclear magnetic resonance signal. .

  In MRI, an RF coil that transmits an RF signal to a subject and detects an MR signal from the subject is used (see, for example, Patent Document 1). There are two types of RF coils, one dedicated to transmission for transmitting RF signals, one dedicated to reception for detecting MR signals, and one for both transmission and reception for transmitting RF signals and detecting MR signals.

  The transmission of the RF signal may be performed, for example, from a whole body RF coil built in the MRI apparatus body or from a movable local transmission RF coil.

JP 2009-142646 A

  In the control unit in the MRI apparatus, the presence of the local reception RF coil and the local transmission RF coil that are separate from the MRI apparatus is recognized by the connector connection. That is, when the connector of the local transmission RF coil is connected to the MRI apparatus, the presence of the local transmission RF coil is normally recognized by the control unit of the MRI apparatus.

  In this state, when an RF signal is transmitted from the whole-body RF coil built in the MRI apparatus, the control unit of the MRI apparatus functions a coupling prevention circuit in the whole-body RF coil. In this case, the function of the coupling prevention circuit in the whole body RF coil prevents a large induced current from flowing in the local transmission RF coil, and there is no possibility of damage to the local transmission RF coil.

  However, in the state in which the connector of the local transmission RF coil installed on the pedestal is forgotten to be connected to the MRI apparatus, the presence of the local transmission RF coil is not recognized by the control unit in the MRI apparatus. In this state, when an RF signal is transmitted from the whole body RF coil side built in the MRI apparatus, the control unit of the MRI apparatus determines that there is no other RF coil apparatus to be protected, Do not allow the anti-coupling circuit to function. In this case, an induction current flows in the local transmission RF coil due to the mutual inductance due to the high-frequency magnetic field, and the local transmission RF coil may be damaged.

  For this reason, even if an RF signal is transmitted from the whole body RF coil built in the MRI apparatus in the state where the connector of the local transmission RF coil installed on the gantry is forgotten to be connected to the MRI apparatus, the local transmission RF coil There has been a demand for a technique for preventing an induced current from flowing in the inside.

  In one embodiment, the RF coil device is detachably connected to the magnetic resonance imaging apparatus, and an RF signal current supplied from the magnetic resonance imaging apparatus flows through the coil element to transmit an RF signal to the subject. And has a trap circuit. This trap circuit is inserted in the path of the RF signal current in the coil element, and selectively cuts off the frequency band including the Larmor frequency according to the input voltage.

  In one embodiment, the MRI apparatus applies a static magnetic field and a gradient magnetic field to the imaging space in which the subject is placed, transmits an RF signal for causing nuclear magnetic resonance, and receives an echo signal emitted by the nuclear magnetic resonance. Thus, the image data of the subject is generated based on the echo signal, and the above-described RF coil device is responsible for transmitting the RF signal.

1 is a block diagram showing a circuit configuration of an entire configuration of an RF coil device according to an embodiment. The circuit diagram which shows the detail of the part containing a protection circuit in the RF coil apparatus of FIG. FIG. 2 is an equivalent circuit diagram of a portion including a protection circuit when the input connector is not connected in the RF coil device of FIG. 1. FIG. 2 is an equivalent circuit diagram of a portion including a protection circuit in a state where the input connector is connected to the MRI apparatus in the RF coil apparatus of FIG. 1. The block diagram which shows the structure which provided the protection circuit in all the rung parts as a modification of the RF coil apparatus of FIG. FIG. 6 is a block diagram showing a configuration in which the wiring of the input power source is shared between the protection circuits as a modification of the RF coil device of FIG. 5. FIG. 7 is a block diagram showing a configuration in which a parallel resonance circuit is further added between the protection circuits as a modification of the RF coil device of FIG. 6. The block diagram which shows the structure which provided the protection circuit in the ring part as a modification of the RF coil apparatus of FIG. 1 is a block diagram showing an overall configuration of an MRI apparatus according to an embodiment.

  Hereinafter, embodiments of the present invention will be described with reference to the accompanying drawings. In addition, in each figure, the same code | symbol is attached | subjected to the same element and the overlapping description is abbreviate | omitted.

(Configuration example of RF coil device)
FIG. 1 is a block diagram showing the overall configuration of the local transmission RF coil apparatus according to the present embodiment in a circuit form. Here, as an example, a case where a birdcage type coil is used will be described.

  As shown in FIG. 1, the RF coil 100A includes a coil support portion 104 indicated by a dotted line, two annular ring portions 106 and 108 indicated by thick lines, and four rung portions 112a, 112b, 112c, and 112d. A coil element is comprised by the ring parts 106 and 108 and the rung parts 112a-112d. Here, for the sake of simplification, the number of rung portions is four, but this is only an example, and the number of rung portions may be four or more, for example.

  The coil support portion 104 is a support member formed in a cylindrical shape, and supports the ring portion 108, the rung portions 112a to 112d, and the like so as to be maintained at predetermined positions.

  Each of the ring portions 106 and 108 is a conductive member formed in a ring shape using, for example, copper foil. Each of the ring portions 106 and 108 is provided so that the diameter and the central axis of the ring coincide with each other and are separated by a predetermined distance.

  Each of the rung portions 112a to 112d is a conductive member formed in a rectangular shape using, for example, copper foil. Each of the rung portions 112a to 112d is provided at a predetermined interval so as to connect the two ring portions 106 and 108, respectively.

  Four capacitors C 1, C 2, C 3, and C 4 are inserted (connected) in series in the ring unit 106. That is, the capacitor C1 is between the rungs 112a and 112d, the capacitor C2 is between the rungs 112d and 112c, the capacitor C3 is between the rungs 112c and 112b, and the capacitor C4 is between the rungs 112b and 112a. Inserted in series.

  Here, as an example, RF coil device 100A includes an input connector (not shown) connected to a connection node between capacitors C1-C2 and a connection node between capacitors C2-C3. As shown by the one-dot chain line in FIG. 1, the connection node between the capacitors C1 and C2 is supplied with an RF signal current from an RF transmitter 46 of an MRI apparatus, which will be described later, via this input connector. The connection node between C3 is set to the ground potential, for example. The above input voltage method is merely an example, and other forms may be used.

  As a feature of the present embodiment, the RF coil device 100A includes a protection circuit 120 inserted in series with the rung portion 112a.

  FIG. 2 is a circuit diagram showing details of a portion including the protection circuit 120 in the RF coil device 100A. As shown in FIG. 2, the protection circuit 120 includes parallel resonant circuits 124 and 126, a trap circuit 128, a first input terminal 130, and a second input terminal 132.

  The parallel resonant circuit 124 includes a coil L1 and a capacitor C9. The inductance value of the coil L1 and the capacitance value of the capacitor C9 are selected so that the resonance frequency of the parallel resonance circuit 124 becomes the Larmor frequency, that is, the impedance of the parallel resonance circuit 124 becomes maximum at the Larmor frequency. . The Larmor frequency is determined in consideration of the static magnetic field strength of the MRI apparatus to which the RF coil apparatus 100A is connected. This prevents a current in a frequency band including the Larmor frequency from flowing into the power supply side to which the first input terminal 130 is connected. That is, the power supply is protected.

  The parallel resonant circuit 126 includes a coil L2 and a capacitor C10. The inductance value of the coil L2 and the capacitance value of the capacitor C10 are also selected so that the impedance of the parallel resonant circuit 124 is maximized at the Larmor frequency. This prevents a current in a frequency band including the Larmor frequency from flowing into the connection destination of the second input terminal 132.

  The trap circuit 128 is a circuit that selectively cuts off a current in a predetermined frequency band. The trap circuit 128 has a configuration in which a coil L3, a capacitor C11, a PIN (P-Intrinsic-N) diode D1, and a PIN diode D2 are connected in parallel, and capacitors C12 and C13 are inserted at predetermined positions.

  Specifically, a DC cut capacitor C13 is inserted in series between the anode of the PIN diode D1 and the cathode of the PIN diode D2. A DC cut capacitor C12 is inserted in series between the cathode diode C11 of the PIN diode D1.

  The trap circuit 128 of the protection circuit 120 is wired so as to be inserted in series in the rung portion 112a, and the PIN diodes D1 and D2 form a cross diode with respect to the rung portion 112a. That is, the node N1 connecting the PIN diodes D1-D2 is connected to one ring unit 106, and the node N2 connecting the PIN diode D1-capacitor C13 is connected to the other ring unit 108.

  When the input connector of the RF coil device 100A is connected to the MRI apparatus, as shown in FIG. 1, the two portions of the ring portion 106 are connected to the RF transmitter 46 and the predetermined voltage line, and the first and second Each of the input terminals 130 and 132 is connected to a predetermined voltage line via the MRI apparatus. Specifically, for example, a constant voltage (for example, ground voltage GND) lower than that of the first input terminal 130 is supplied to the second input terminal 132, and the PIN diodes D1 and D2 are turned on (conductive) to the first input terminal 130. ) Is supplied with a constant positive voltage VCC.

  FIG. 3 is an equivalent circuit diagram of a portion including the protection circuit 120 when the input connector of the RF coil device 100A is not connected. In a state where the input connector is not connected to the MRI apparatus, the voltages of the first input terminal 130 and the second input terminal 132 are released. In this case, since the PIN diodes D1 and D2 are in the off (non-conducting) state, they can be divided into a capacitance Cd1 and a resistance component Rd1, and a capacitance Cd2 and a resistance component Rd2, respectively.

  Here, the capacitance values of the capacitors C11, C12, and C13, the capacitance values (capacitance values of Cd1 and Cd2) and the resistance values (resistance values of Rd1 and Rd2) when reverse voltages are applied to the PIN diodes D1 and D2, and The inductance value of the coil L3 is selected as follows.

  That is, when the node N1 in FIGS. 2 and 3 is an input terminal and the node N2 is an output terminal, the impedance of the parallel resonance circuit 124 is maximized at the Larmor frequency (the resonance frequency of the trap circuit matches the Larmor frequency). To be selected). Thereby, when the PIN diodes D1 and D2 are in the OFF state, the trap circuit 128 selectively cuts off the current in the frequency band including the Larmor frequency.

  The cutoff frequency of the trap circuit 128 is selectively cut off not only in the Larmor frequency under a certain condition in the operating state of the MRI apparatus but also in a predetermined frequency band including the Larmor frequency under the certain condition. It is desirable that the belt has a certain width. This is because, in the operating state of the MRI apparatus, for example, the transmission frequency of the RF signal may be changed as the temperature of the iron shim increases with time, so that it is desirable to include a slight change in the transmission frequency. Therefore, the resonance frequency of the trap circuit 128 does not have to be completely matched with the Larmor frequency, and may be close to the Larmor frequency to such an extent that a current in a predetermined frequency band including the Larmor frequency can be selectively cut off.

  Here, a case is considered where an RF signal having a Larmor frequency is transmitted from a whole-body RF coil built in the MRI apparatus while the input connector of the RF coil apparatus 100A placed on the gantry is not connected. In this case, an induced current having a Larmor frequency tends to flow, for example, along a path indicated by a thick arrow in FIG. 3 due to mutual inductance. However, when the input connector is not connected, that is, when the first and second input terminals 130 and 132 are in the released state, the trap circuit 128 has a maximum impedance for the frequency band including the Larmor frequency as described above. (Resonance state). For this reason, an induced current is prevented from flowing in the RF coil device 100A, and there is no possibility that the RF coil device 100A is damaged.

  FIG. 4 is an equivalent circuit diagram of a portion including the protection circuit 120 in a state where the input connector of the RF coil device 100A is connected to the MRI apparatus. In this case, for example, a positive constant voltage VCC is supplied to the first input terminal 130, and the ground voltage GND is supplied to the second input terminal 132. As a result, the PIN diodes D1 and D2 are both in a state where a forward bias is applied, and thus are turned on (conducted). Therefore, a direct current flows through the path of the first input terminal 130, the coil L1, the PIN diode D1, the PIN diode D2, the coil L2, and the second input terminal 132.

  Then, the values of the capacitors Cd1 and Cd2 and the resistance components Rd1 and Rd2 of the PIN diodes D1 and D2 are different from those shown in FIG. 3 (when the first input terminal 130 and the second input terminal 132 are released). It becomes. Thereby, the frequency band and the resonance frequency at which the impedance of the trap circuit 128 is maximum are shifted from the frequency band including the Larmor frequency.

  That is, when the PIN diodes D1 and D2 are in the ON state, the trap circuit 128 passes a current in a frequency band including the Larmor frequency. Therefore, in this case, if the RF signal current of the Larmor frequency is supplied from the RF transmitter 46, the RF signal current flows through the path indicated by the thick line arrow in FIG. 4, so that the RF coil device 100A as a whole is a high-frequency magnetic field. And function as an RF coil for transmission.

  The cross diode used for the trap circuit 128 may be a normal PN junction diode, but a PIN diode is more desirable as described above. This is because a PIN diode has a higher switching speed than a normal PN junction diode (because the switching characteristics of a high-frequency signal are good).

  As described above, in this embodiment, when the input connector of the RF coil device 100A is forgotten to be connected to the MRI apparatus main body, the first and second input terminals 130 and 132 are released, and the trap circuit 128 has the Larmor frequency. Is selectively cut off. Since the trap circuit 128 is inserted in the path of the RF signal current, no induced current flows in the RF coil device 100A even if a magnetic field of the RF signal from the whole body coil is generated around the trap circuit 128. That is, there is no risk of damage to the RF coil device 100A.

  On the other hand, when the input connector of the RF coil device 100A is connected to the MRI device, the PIN diodes D1 and D2 are turned on, and the resonance frequency of the trap circuit 128 is shifted. In this case, the trap circuit 128 passes an RF signal current having a Larmor frequency, and the RF coil device 100A functions normally as an RF coil for transmission.

  That is, according to the present embodiment, an RF signal is transmitted from the whole body RF coil built in the MRI apparatus in a state where the connector of the transmission RF coil (100A) installed on the gantry is forgotten to be connected to the MRI apparatus. However, it is possible to prevent the induced current from flowing in the transmitting RF coil (100A). In the prior art, none of the transmission RF coil devices that are not built in the MRI apparatus (removably connected to the MRI apparatus) have a trap circuit.

(Modification of RF coil device)
FIG. 5 is a block diagram showing an RF coil device 100B having a configuration in which protection circuits 120 are inserted in series in all the rung portions 112a to 112d as a modification of the RF coil device 100A. In this case, as shown in FIG. 2, wiring is performed so that the nodes N1 and N2 of the trap circuit 128 in each protection circuit 120 are directly connected to the conductors of the rung portions 112a to 112d, respectively.

The first and second input terminals 130 and 132 of each protection circuit 120 may be individually connected to the power supply voltage line VCC and the ground voltage line GND as in the RF coil device 100B, but are common as shown in FIG. It may be. FIG. 6 is a block diagram showing an RF coil device 100C in which the wiring of the input power supply to the first and second input terminals 130 and 132 is common between the protection circuits as a modification of the RF coil device 100B.
In this case, since forward current flows through a total of eight PIN diodes D1 and D2 in the four protection circuits 120, the supply voltage VCC2 of the rightmost positive voltage source in FIG. 6 is the case of the individual connection shown in FIG. For example, 4 times or more.

  Here, it is desirable to reduce the interval (wiring length) between the connection points of the parallel resonant circuit (124 or 126) and the trap circuit 128. This is to prevent the generation of the induced current in the Larmor frequency band as much as possible. Therefore, when the wiring length between the first input terminal 130 of one protection circuit 120 and the second input terminal 132 of the other protection circuit 120 connected thereto is long, as shown in FIG. Parallel resonant circuits 141, 142, 143, 144, and 145 may be further provided between the circuits 120 and between the protection circuit 120 and the power supply line.

  FIG. 7 is a block diagram showing an RF coil device 100D to which a parallel resonant circuit is added as a modification of the RF coil device 100C of FIG. The circuit configuration of the parallel resonance circuits 141 to 145 shown in FIG. 7 is the same as that of the parallel resonance circuits 124 and 126. The capacitance values of the capacitors in the parallel resonant circuits 141 to 145 and the inductance values of the coils are such that the wiring length between the protection circuits 120 and the protection circuit are set so as to cut off the generation of the induced current in the Larmor frequency band as much as possible. Selection may be made according to the wiring length between 120 and the power supply line.

  Note that the protection circuit 120 may be inserted in series with respect to two or three of the rung portions 112a to 112d without having to insert the protection circuit 120 in series with each of the rung portions 112a to 112d. Not shown).

  FIG. 8 is an equivalent circuit diagram showing a configuration of an RF coil device 100E in which protective circuits 120 are inserted in series in the ring portions 106 and 108 in place of the rung portions 112a to 112d as modifications of the RF coil devices 100A to 100D. It is. In this case, as shown in FIG. 8, wiring is performed so that the nodes N1 and N2 in each protection circuit 120 are directly connected to the conductors of the ring portions 106 and 108, respectively.

  Note that the protection circuit 120 may be inserted in series only for one of the ring portions (106 or 108) (not shown). Or you may insert the protection circuit 120 in both the rung parts 112a-112d and the ring parts 106 and 108, respectively. Even in the RF coil devices of the modified examples described above including the RF coil devices 100B to 100E, the same effects as those of the RF coil device 100A can be obtained.

(Configuration example of MRI system)
FIG. 9 is a block diagram showing the overall configuration of the MRI apparatus 20 according to an embodiment. As shown in FIG. 9, the MRI apparatus 20 includes a cylindrical static magnetic field magnet 22 for forming a static magnetic field, a cylindrical shim coil 24 provided with the same axis inside the static magnetic field magnet 22, A gradient coil 26, an RF coil 28, a control system 30, and a gantry (bed) 32 on which the subject P is placed are provided.

  Here, as an example, the X-axis, Y-axis, and Z-axis that are orthogonal to each other in the apparatus coordinate system are defined as follows. First, the static magnetic field magnet 22 and the shim coil 24 are arranged so that their axial directions are orthogonal to the vertical direction, and the axial direction of the static magnetic field magnet 22 and the shim coil 24 is the Z-axis direction. Further, it is assumed that the vertical direction is the Y-axis direction, and the gantry 32 is arranged such that the normal direction of the surface for placing the top plate is the Y-axis direction.

  The control system 30 includes a static magnetic field power supply 40, a shim coil power supply 42, a gradient magnetic field power supply 44, an RF transmitter 46, an RF receiver 48, a bed driving device 50, a sequence controller 56, and a computer 58. .

  The gradient magnetic field power supply 44 includes an X-axis gradient magnetic field power supply 44x, a Y-axis gradient magnetic field power supply 44y, and a Z-axis gradient magnetic field power supply 44z. The computer 58 includes an arithmetic device 60, an input device 62, a display device 64, and a storage device 66.

  The static magnetic field magnet 22 is connected to the static magnetic field power supply 40 and forms a static magnetic field in the imaging space by a current supplied from the static magnetic field power supply 40. The shim coil 24 is connected to a shim coil power source 42 and makes the static magnetic field uniform by a current supplied from the shim coil power source 42. The static magnetic field magnet 22 is often composed of a superconducting coil, and is connected to the static magnetic field power source 40 and supplied with current when excited, but after being excited, it is disconnected. Is common. The static magnetic field magnet 22 may be formed of a permanent magnet without providing the static magnetic field power supply 40.

  The gradient magnetic field coil 26 includes an X-axis gradient magnetic field coil 26 x, a Y-axis gradient magnetic field coil 26 y, and a Z-axis gradient magnetic field coil 26 z, and is formed in a cylindrical shape inside the static magnetic field magnet 22. The X-axis gradient magnetic field coil 26x, the Y-axis gradient magnetic field coil 26y, and the Z-axis gradient magnetic field coil 26z are connected to the X-axis gradient magnetic field power source 44x, the Y-axis gradient magnetic field power source 44y, and the Z-axis gradient magnetic field power source 44z, respectively.

  The X-axis gradient magnetic field power supply 44x, the Y-axis gradient magnetic field power supply 44y, and the Z-axis gradient magnetic field power supply 44z respectively supply the X-axis gradient magnetic field coil 26x, the Y-axis gradient magnetic field coil 26y, and the Z-axis gradient magnetic field coil 26z A gradient magnetic field Gx in the axial direction, a gradient magnetic field Gy in the Y-axis direction, and a gradient magnetic field Gz in the Z-axis direction are formed in the imaging space.

  That is, the gradient magnetic fields Gx, Gy, and Gz in the three-axis directions of the apparatus coordinate system are synthesized, and the slice selection direction gradient magnetic field Gss, the phase encoding direction gradient magnetic field Gpe, and the readout direction (frequency encoding direction) gradient as the logical axes. Each direction of the magnetic field Gro can be arbitrarily set. The gradient magnetic fields in the slice selection direction, the phase encoding direction, and the readout direction are superimposed on the static magnetic field.

  The RF transmitter 46 generates Larmor frequency RF pulses (RF current pulses) for causing nuclear magnetic resonance based on the control information input from the sequence controller 56. The RF transmitter 46 transmits the generated RF pulse to the transmission RF coil 28 or another transmission RF coil device connected to the MRI apparatus 20.

  The RF coil 28 includes a whole-body coil 28a (indicated by a broken line in FIG. 9) for transmitting and receiving an RF pulse built in the gantry, and an RF pulse receiver provided near the gantry 32 or the subject P. There are local coils. In the present embodiment, the MRI apparatus 20 further includes RF coil apparatuses 100A to 100E for local transmission.

  The transmission RF coil 28 and the RF coil devices 100A to 100E receive an RF pulse from the RF transmitter 46 and transmit it to the subject P as an RF signal. The receiving RF coil 28 receives an MR signal (high-frequency echo signal) generated when the nuclear spin inside the subject P is excited by the RF signal, and this MR signal is detected by the RF receiver 48. Is done.

The RF receiver 48 performs various signal processing such as pre-amplification, intermediate frequency conversion, phase detection, low-frequency amplification, and filtering on the detected MR signal, and then performs A / D (analog to digital) conversion. Then, raw data (raw data) which is digitized complex data is generated. The RF receiver 48 inputs the generated raw data of the MR signal to the sequence controller 56.
The arithmetic device 60 performs system control of the entire MRI apparatus 20.

  The sequence controller 56 stores control information necessary for driving the gradient magnetic field power supply 44, the RF transmitter 46, and the RF receiver 48 in accordance with a command from the arithmetic device 60. The control information here is, for example, sequence information describing operation control information such as the intensity, application time, and application timing of the pulse current to be applied to the gradient magnetic field power supply 44.

  The sequence controller 56 drives the gradient magnetic field power supply 44, the RF transmitter 46, and the RF receiver 48 in accordance with the stored predetermined sequence, whereby the X-axis gradient magnetic field Gx, the Y-axis gradient magnetic field Gy, the Z-axis gradient magnetic field Gz, and RF Generate a pulse. Further, the sequence controller 56 receives the raw data (raw data) of the MR signal input from the RF receiver 48 and inputs this to the arithmetic device 60.

  The couch driving device 50 is connected to the arithmetic device 60 via the sequence controller 56. The sequence controller 56 moves the top plate of the gantry 32 by controlling the bed driving device 50 in accordance with a command from the arithmetic device 60, and thereby performs imaging by, for example, a moving table method or a stepping-table method.

  Hereinafter, an example of the operation when imaging is performed using any of the RF coil devices 100A to 100E of the MRI apparatus 20 will be described.

  Any of the local transmission RF coil devices 100A to 100E is connected to the MRI apparatus via a connector (not shown), and is set on the subject P by the gantry 32. Thereby, the arithmetic unit 60 of the MRI apparatus 20 recognizes the presence of the RF coil apparatus (any one of 100A to 100E). Next, a static magnetic field is formed in the imaging space by the static magnetic field magnet 22 excited by the static magnetic field power supply 40. Further, a current is supplied from the shim coil power source 42 to the shim coil 24, and the static magnetic field formed in the imaging space is made uniform.

  When an imaging start instruction is input from the input device 62 to the arithmetic device 60, the arithmetic device 60 inputs an imaging condition including a pulse sequence to the sequence controller 56. The sequence controller 56 drives the gradient magnetic field power supply 44 according to the input pulse sequence, thereby forming a gradient magnetic field in the imaging space and supplying an RF pulse from the RF transmitter 46 to the RF coil devices 100A to 100E. Thus, an RF signal is transmitted to the subject P.

  Therefore, the MR signal (echo signal) generated by the nuclear magnetic resonance inside the subject P is received by the RF coil 28 and detected by the RF receiver 48. The RF receiver 48 performs predetermined signal processing on the detected MR signal, and A / D converts this to generate raw data that is a digitized MR signal. The RF receiver 48 inputs the generated raw data to the sequence controller 56.

  The sequence controller 56 inputs the raw data to the arithmetic device 60, and the arithmetic device 60 arranges the raw data as k-space data in its internal k-space database. The arithmetic device 60 reconstructs image data by performing image reconstruction processing including Fourier transform on the k-space data, temporarily stores this, and then performs predetermined image processing on this to generate display image data. Then, the display image data is stored in the storage device 66.

  Thus, in the MRI apparatus 20 of the present embodiment, any one of the RF coil apparatuses 100A to 100E is used for local transmission. For this reason, if the RF coil device (any one of 100A to 100E) installed on the gantry 32 is forgotten to be connected to the MRI device 20, an RF signal is transmitted from the whole body coil 28a for transmission and reception. However, there is no possibility that the RF coil device (any one of 100A to 100E) is damaged.

(Supplementary items of the embodiment)
[1] In the above embodiment, the RF coil devices 100 </ b> A to 100 </ b> E have been described as an example of a birdcage type. The embodiment of the present invention is not limited to such an aspect. The technical idea of the above embodiment for preventing the induced current when the connector is not connected by the trap circuit is applicable to other transmitting RF coil devices such as a saddle coil, a solenoid coil, and a surface QD (Quadrature Detection) coil. .

  [2] Although the RF coil devices 100A to 100E have been described as those for local transmission, the embodiment of the present invention is not limited to such a mode. For example, when the subject is small, such as an infant, the RF coil devices 100A to 100E can also be used as whole body transmission RF coils. That is, the embodiment of the present invention is not limited to an RF coil device for local transmission, but can be applied to a transmission RF coil device for whole body.

  Further, the technical idea of the above-described embodiment for preventing the induced current when the connector is not connected by the trap circuit is not limited to the RF coil device dedicated to transmission, but can also be applied to the RF coil device for both transmission and reception. .

  [3] Although several embodiments of the present invention have been described, these embodiments are presented as examples and are not intended to limit the scope of the invention. These embodiments can be implemented in various other forms, and various omissions, replacements, and changes can be made without departing from the spirit of the invention. These embodiments and their modifications are included in the scope and gist of the invention, and are also included in the invention described in the claims and the equivalents thereof.

20 MRI apparatus 22 Magnet for static magnetic field 24 Shim coil 26 Gradient magnetic field coil 26 x X-axis gradient magnetic field coil 26 y Y-axis gradient magnetic field coil 26 z Z-axis gradient magnetic field coil 28 RF coil 30 Control system 32 Base 40 Static magnetic field power supply 42 Shim coil power supply 44 Gradient magnetic field 44 Power supply 44x X-axis gradient magnetic field power supply 44y Y-axis gradient magnetic field power supply 44z Z-axis gradient magnetic field power supply 46 RF transmitter 48 RF receiver 50 Bed driving device 56 Sequence controller 58 Computer 60 Computing device 62 Input device 64 Display device 66 Storage device 100A- 100E RF coil device 104 Coil support part 106, 108 Ring part 112a-112d Lang part 120 Protection circuit 124, 126, 141-145 Parallel resonance circuit 128 Trap circuit 130 1st input terminal 132 2nd input terminal C1-C13 Condensate D1, D2 PIN diode L1~L3 coil P subject

Claims (5)

An RF coil apparatus that is detachably connected to a magnetic resonance imaging apparatus, and that transmits an RF signal to a subject when an RF signal current supplied from the magnetic resonance imaging apparatus flows through a coil element,
An RF coil device having a trap circuit that is inserted into a path of the RF signal current in the coil element and selectively cuts off a frequency band including a Larmor frequency according to an input voltage. The RF coil device according to claim 1, wherein
The RF coil device, wherein the trap circuit selectively cuts off a frequency band including the Larmor frequency when the input voltage to the trap circuit is in an open state. The RF coil device according to claim 1, wherein
The trap circuit has a cross diode that shifts a resonance frequency of the trap circuit in a direction away from the Larmor frequency by switching from an off state to an on state in accordance with the input voltage to the trap circuit. Coil device. The RF coil device according to claim 3, wherein
The RF coil device, wherein the cross diode is composed of two PIN diodes. Applying a static magnetic field and a gradient magnetic field to the imaging space in which the subject is placed and transmitting an RF signal for causing nuclear magnetic resonance, receiving an echo signal emitted by nuclear magnetic resonance, and based on the echo signal A magnetic resonance imaging apparatus for generating image data of a subject,
5. A magnetic resonance imaging apparatus comprising the RF coil device according to claim 1, which is responsible for transmitting the RF signal.

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