JP5010623B2 - Coil device for magnetic resonance diagnostic equipment - Google Patents

Description

  The present invention relates to a coil device used in a magnetic resonance diagnostic apparatus.

  In general, in a magnetic resonance diagnostic apparatus such as a magnetic resonance imaging (MRI) apparatus, it is necessary that the magnetic field does not leak outside the imaging (diagnosis) area. For this reason, the present inventor has proposed an active shield gradient coil (ASGC) designed so as not to leak a magnetic field to the outside as a gradient magnetic field coil (Patent Document 1, Patent) Reference 2 and Patent Reference 3).

  When such an active shield type gradient magnetic field coil is used, magnetic coupling with an RF coil for transmitting and receiving a high-frequency magnetic field essential for a magnetic resonance diagnostic apparatus becomes a problem. Therefore, an RF shield for blocking magnetic coupling is disposed between the two. That is, the arrangement of the coil device of the magnetic resonance diagnostic apparatus is, in order from the outside, a static magnetic field magnet, an active shield type gradient magnetic field coil, an RF shield, and an RF coil. A generally known RF shield is made of a conductor or a magnetic material. However, in a magnetic resonance diagnostic apparatus, it is not preferable to use a magnetic material as an RF shield. Therefore, a conductor such as copper or aluminum is usually used. It is done.

  The active shield type gradient magnetic field coil is designed so that the magnetic field does not leak outside, but naturally the gradient magnetic field is generated inside (diagnosis region) where the subject is arranged. As described above, since the RF shield made of a conductor is disposed between the active shield type gradient magnetic field coil and the RF coil, eddy current is generated on the RF shield when the value or polarity of the gradient magnetic field changes. Will occur. An eddy magnetic field is generated by this eddy current, and there is a drawback that the waveform of the pulse-shaped gradient magnetic field is usually distorted.

  FIG. 16 is a diagram for explaining the deterioration of the gradient magnetic field waveform due to the eddy current. Since the eddy magnetic field due to the eddy current generated on the RF shield is generated in a direction that cancels the gradient magnetic field, the original gradient magnetic field waveform becomes dull.

  In recent years, high-speed imaging methods such as an echo planar imaging method (EPI) have become popular in MRI apparatuses, but high-speed imaging methods require a steep rising characteristic of a gradient magnetic field. Therefore, it is desired to suppress the influence of eddy current generated in the RF shield as much as possible.

  Conventionally, in order to reduce distortion of the gradient magnetic field due to eddy current, it has been proposed to make a slit in the RF shield to shorten the time constant of eddy current. As a role of the RF shield, a magnetic field having a relatively low frequency such as a gradient magnetic field (up to about 100 KHz) passes, and a high-frequency magnetic field of several MHz to several tens of MHz such as an excitation pulse is required to be cut off. It depends. That is, putting a slit in the RF shield is to generate C coupling, which is to increase the low frequency impedance and to decrease the high frequency impedance.

  However, when the number of slits is increased and C coupling is increased, the loss (resistance component) included in the C coupling also increases, resulting in a drawback that the high frequency shielding effect of the RF shield is lowered. Therefore, when comparing RF shields having the same thickness, the maximum shielding effect is obtained when no slit is formed. However, as described above, if the slit is not inserted, the gradient magnetic field waveform is distorted.

  Further, the eddy current magnetic field generated on the RF shield leaks to the outside of the gradient coil, that is, to the static magnetic field magnet. Therefore, although it is an active shield type gradient magnetic field coil, it is magnetically coupled to a conductor outside the active shield type gradient magnetic field coil via a conductor inside the active shield type gradient magnetic field coil. This magnetic coupling causes an eddy current to be generated in the conductor outside the active shield type gradient coil, that is, the heat shield body constituting the static magnetic field magnet, and the conductor generates heat, thereby causing an increase in the boil-off amount of helium. .

Patent 62-143012 USP 4,737,716 specification USP 4,733,189 specification

  Thus, in the conventional magnetic resonance diagnostic apparatus using the active shield type gradient magnetic field coil, an RF shield is disposed between the RF coil and the active shield type gradient magnetic field coil in order to block the magnetic coupling between them. An eddy current is generated on the RF shield by the magnetic coupling between the RF shield and the active shield type gradient magnetic field coil, and the eddy magnetic field due to the eddy current distorts the magnetic field in the imaging region. Further, since the magnetic field of the eddy current generated on the RF shield leaks to the outside of the gradient coil, an eddy current is also generated in the conductor outside the active shield type gradient coil, thereby the outside of the gradient coil device. There was a problem that the conductor of the heat generated.

  It is an object of the present invention to provide a shield device for a magnetic resonance diagnosis apparatus in which an RF shield is disposed between the RF coil and an active shield type gradient magnetic field coil to shield the magnetic coupling between the RF coil and the active shield type gradient magnetic field coil. Without degrading the effect, the gradient magnetic field is distorted by the eddy current on the RF shield, and the magnetic field of the eddy current leaks to the outside of the gradient coil and prevents the conductor outside the active shield type gradient coil from generating heat. That is.

  In order to solve the above problems and achieve the object, the present invention uses the following means.

(1) According to the embodiment, in the coil device for a magnetic resonance diagnostic apparatus including the gradient coil and the RF coil provided inside the gradient coil, the gradient coil is a main component that generates a gradient magnetic field in the imaging region. Active shielding type comprising a coil and an active shield coil provided outside the main coil and connected in series to the main coil to generate a magnetic field for preventing the gradient magnetic field generated from the main coil from leaking outside A gradient coil that is connected to the active shield type gradient coil and includes a current waveform shaping circuit that corrects the waveform of the drive current supplied from the gradient magnetic field power source to the main coil or the active shield coil, and is supplied to the main coil The waveform of the drive current supplied to the active shield coil is different from the waveform of the drive current supplied to the active shield coil .

  As described above, according to the present invention, an eddy magnetic field caused by an eddy current generated on the RF shield is suppressed by providing a counter shield made of a conductor at least outside the main coil of the active shield type gradient magnetic field coil. Can do.

  Further, by providing the coil outside at least the main coil of the active shield type gradient magnetic field coil, the eddy magnetic field due to the eddy current generated on the RF shield can be almost completely canceled.

  In addition, by providing a means for overshoot correction of the drive signal supplied to the active shield type gradient magnetic field coil, the eddy magnetic field caused by the eddy current generated on the RF shield can be almost completely eliminated without adding a shield, a coil or the like. Can be countered.

  Further, by providing means for correcting an undershoot of the drive signal supplied to the active shield coil of the active shield type gradient magnetic field coil, an eddy magnetic field due to an eddy current generated on the RF shield without adding a shield, a coil, etc. Can be almost completely counteracted.

  In addition, by providing means for overshoot correction of the drive signal supplied to the main coil of the active shield type gradient magnetic field coil, the eddy magnetic field due to the eddy current generated on the RF shield can be reduced without adding a shield, a coil or the like. Can be almost completely countered.

1 is a block diagram showing a schematic configuration of a magnetic resonance imaging apparatus including a first embodiment of a gradient coil apparatus according to the present invention. The figure which shows the structure of 1st Embodiment of the gradient magnetic field coil assembly of the MRI apparatus by this invention. The figure which shows the structure of the modification of 1st Embodiment. The figure which shows a magnetic field strength waveform in order to demonstrate the principle of 1st Embodiment. The figure which shows a magnetic field strength waveform in order to demonstrate the effect | action of 1st Embodiment. The figure which shows the structure of 2nd Embodiment of the gradient magnetic field coil assembly of the MRI apparatus by this invention. The figure which shows the coil pattern of the active shielding type counter coil of 2nd Embodiment. The figure which shows the power supply of the active shielding type gradient magnetic field coil and active shielding type counter coil of 2nd Embodiment. The figure which shows the structure of 3rd Embodiment of the gradient magnetic field coil assembly of the MRI apparatus by this invention. The figure which shows the power supply of the active shielding type gradient magnetic field coil of 3rd Embodiment. The figure which shows the waveform of a drive signal in order to demonstrate the effect | action of 3rd Embodiment. The figure which shows the power supply of 4th Embodiment of the gradient coil assembly of the MRI apparatus by this invention. The figure which shows the waveform of a drive signal in order to demonstrate the effect | action of 4th Embodiment. The figure which shows the power supply of 5th Embodiment of the gradient magnetic field coil assembly of the MRI apparatus by this invention. The figure which shows the waveform of a drive signal in order to demonstrate the effect | action of 5th Embodiment. The figure explaining the fault of the conventional active shielding type gradient magnetic field coil.

  Hereinafter, embodiments of the present invention will be described in detail with reference to the drawings.

  Embodiments of a coil device for a magnetic resonance diagnostic apparatus according to the present invention will be described below with reference to the drawings.

(First embodiment)
FIG. 1 is a diagram showing a configuration of an MRI apparatus having a gradient magnetic field coil apparatus according to the first embodiment of the present invention. A static magnetic field magnet 1, a gradient magnetic field coil 2, and an RF coil 3 are provided in this order from the outside in a gantry 20 in which a cylindrical imaging space is formed so as to accommodate a subject at the center. The static magnetic field magnet 1, the gradient magnetic field coil 2, and the RF coil 3 are attached to the gantry 20 by a support member (not shown). The static magnetic field magnet 1 generates a static magnetic field Bo in the imaging space along the Z axis that is the body axis direction of the subject, and is configured using, for example, a superconducting coil or a normal conducting coil. The gradient coil 2 has an X-axis gradient magnetic field whose magnetic field strength changes linearly along the X-axis, a Y-axis gradient magnetic field whose magnetic field strength changes linearly along the Y-axis, and a magnetic field strength linearly along the Z-axis. A Z coil set, an X coil set, and a Y coil set are configured so that the changing Z-axis gradient magnetic field can be generated independently. The RF coil 3 generates a high-frequency magnetic field (RF pulse) that acts on the magnetization spin of a nuclide to be diagnosed in the subject, and is used to detect an echo signal generated by magnetic resonance. The subject P on the bed 13 is inserted into an imageable region in the gantry 20 (a spherical region where an imaging magnetic field is formed, and diagnosis is possible only in this region).

  The static magnetic field magnet 1 is driven by a static magnetic field control device 4. The transmitter / receiver coil 3 is driven by the transmitter 5 when magnetic resonance is excited, and is coupled to the receiver 6 when an echo signal is detected. The X coil set, Y coil set, and Z coil set of the gradient magnetic field coil 2 are driven by an X axis gradient magnetic field power source 7, a Y axis gradient magnetic field power source 8, and a Z axis gradient magnetic field power source 9, respectively.

  The X-axis gradient magnetic field power supply 7, the Y-axis gradient magnetic field power supply 8, the Z-axis gradient magnetic field power supply 9, and the transmitter 5 are driven by the sequencer 10 according to a predetermined sequence, and the X-axis gradient magnetic field Gx, Y-axis gradient magnetic field Gy, Z-axis gradient magnetic field The magnetic field Gz and the RF pulse are generated according to a predetermined pulse sequence described later. In this case, the X-axis gradient magnetic field Gx, the Y-axis gradient magnetic field Gy, and the Z-axis gradient magnetic field Gz mainly include, for example, a frequency encode gradient magnetic field, a read gradient magnetic field Gr, a phase encode gradient magnetic field Ge, and a slice gradient magnetic field. Used as Gs, respectively. The computer system 11 drives and controls the sequencer 10, captures an echo signal as an echo signal received by the receiver 6, and performs predetermined signal processing to generate a tomographic image of the subject. indicate.

  FIG. 2 is a longitudinal sectional view of the gradient coil assembly of the MRI apparatus according to the first embodiment of the present invention. The coil assembly is housed in a cylindrical housing 1, and a cylindrical opening 15 into which a bed (not shown) on which a subject is placed is inserted is provided at the center of the housing 1. In the housing 1, an active shield type gradient magnetic field coil 3, an RF shield 9, and an RF coil 11 are provided in order from the outside. Although not shown, a static magnetic field magnet is disposed outside the housing 1. There are permanent magnets, normal conducting magnets, and superconducting magnets as static magnetic field magnets, but superconducting magnets are often used because of the ease of generating a high magnetic field and the stability of the magnetic field.

  The active shield type gradient magnetic field coil 3 is provided outside the main coil 7 and the main coil 7 that generates a gradient magnetic field having a linear gradient of magnetic field intensity in a predetermined imaging region 17 in the opening 15. The active shield coil 5 generates a magnetic field in the opposite direction to the gradient magnetic field for preventing the gradient magnetic field generated by the above from leaking outside the active shield type gradient magnetic field coil 3. The active shield type gradient magnetic field coil 3 includes three coils of an X channel, a Y channel, and a Z channel. In general, the active shield coil 5 has the same coil pattern as the main coil 7 and is connected in series with the main coil 7. Further, a passive counter shield 13 unique to the present invention is provided outside the active shield type gradient magnetic field coil 3. A gradient magnetic field power source for applying a pulse current is also connected to the main coil 7 and the active shield coil 5.

  The RF shield 9 is made of a conductor such as copper or aluminum as in the conventional example, but is not provided with a slit (cut).

  The passive counter shield 13 is made of a metal having good conductivity such as copper or aluminum, and is provided so as to be magnetically coupled to the RF shield 9. The shape may be a cylindrical shape, a flat plate shape, or a curved shape, and may be provided with a slit. For this reason, the eddy current generated on the RF shield 9 is suppressed. Moreover, the leakage magnetic field to outer conductors, such as a heat shield body of a static magnetic field magnet, can be reduced. The counter shield 13 has a higher suppression effect when the conductivity is higher, and even with the same material, the greater the thickness, the higher the suppression effect.

  The arrangement of the passive counter shield 13 is not limited to the outside of the active shield type gradient coil 3 but may be provided between the main coil 7 and the active shield coil 5 of the active shield type gradient coil 3 as shown in FIG. . In short, it may be outside the main coil 7. Further, the passive counter shield 13 may be provided integrally with the active shield type gradient magnetic field coil 3 instead of being provided separately from the active shield type gradient magnetic field coil 3.

  Next, the operation of the first embodiment will be described.

  An eddy current (hereinafter referred to as a primary eddy current) for canceling the gradient magnetic field generated by the active shield type gradient magnetic field coil 3 is generated in the RF shield 9 disposed inside the active shield type gradient coil 3. This primary eddy current generates an eddy magnetic field (hereinafter referred to as a primary eddy magnetic field). Since the active shield type gradient magnetic field coil 3 does not leak the magnetic field generated from the main coil 7 to the outside, this primary eddy magnetic field creates an eddy magnetic field outside the active shield type gradient magnetic field coil 3. For this reason, when the inside of the coil device is viewed from the passive counter shield 13, this is equivalent to the presence of a non-shielded gradient magnetic field coil that responds in time as shown in FIG.

  Therefore, if a passive counter shield 13 made of a conductor is disposed near the outside of the active shield type gradient magnetic field coil 3 or between the main coil 7 and the active shield coil 5, an undesired shield generated in the RF shield 9 is obtained. Strongly coupled with the primary eddy magnetic field to be. Therefore, the primary eddy magnetic field leaked to the outside from the RF shield 9 is eddy current (hereinafter referred to as secondary eddy current) on the passive counter shield 13 which is a conductor provided outside the active shield gradient magnetic field coil 3. Is generated). Due to the secondary eddy current, an eddy magnetic field having a waveform as shown in FIG. 4B (hereinafter referred to as a secondary eddy magnetic field) is generated.

  Just as the primary eddy field cancels the gradient field, the secondary eddy field cancels the primary eddy field. As a result, a synthetic magnetic field as shown in FIG. 4C, which is a combination of the primary eddy magnetic field and the secondary eddy magnetic field, is generated inside the passive counter shield 13. That is, the primary eddy magnetic field generated by the RF shield 9 is suppressed by its own eddy magnetic field (secondary eddy magnetic field) generated by the passive counter shield 13. The maximum value of the primary eddy magnetic field (FIG. 4A) by the RF shield 9 is −0.387 (T). On the other hand, when the passive counter shield 13 according to the present invention is used, the primary vortex The maximum value of the intensity of the combined eddy magnetic field (FIG. 4C) of the magnetic field and the secondary eddy magnetic field is reduced to −0.150 (T).

  FIG. 5A shows an ideal rise time waveform of the gradient magnetic field (broken line) and an actual rise time waveform of the gradient magnetic field affected by the primary eddy magnetic field suppressed by the passive counter shield 13 ( (Solid line). According to the effect of suppressing the primary eddy magnetic field, the distortion of the gradient magnetic field in the imaging region 17 is reduced. In addition, since the RF shield 9 is not provided with a slit, the shielding effect on the RF coil 11 is not deteriorated at all.

  However, regarding the magnetic field response by the passive counter shield 13, the time constant of the magnetic field response by the passive counter shield 13 is usually generated on the RF shield 9 in order to increase the shielding effect and reduce the resistivity of the conductor. Larger than the vortex time constant. Therefore, although the maximum value of the eddy magnetic field strength can be reduced, the influence has a long time effect. However, this can be easily corrected by correcting the current waveform supplied to the active shield type gradient magnetic field coil 3 to an appropriate waveform.

  One of the major advantages of this method is that this correction current can be reduced.

  Furthermore, by using a metal having a sufficiently small resistivity as the passive counter shield 13, the actual rising waveform (solid line) can be brought closer to the ideal waveform (broken line) as shown in FIG. 5B.

  As described above, according to the present embodiment, by disposing the passive counter shield 13 made of a conductor and magnetically coupled to the RF shield outside the main coil 7 of the active shield gradient magnetic field coil 3, The eddy current generated on the RF shield 9 is suppressed, the distortion of the pulsed gradient magnetic field in the imaging region is reduced by the eddy magnetic field due to the eddy current, and the leakage magnetic field to the outer conductor such as the heat shield body of the static magnetic field magnet is also reduced. Can be reduced.

  Hereinafter, other embodiments of the coil device according to the present invention will be described. In the description of the other embodiments, the same parts as those in the first embodiment are denoted by the same reference numerals, and the detailed description thereof is omitted.

(Second Embodiment)
In the first embodiment, a passive counter shield 13 made of a conductor is provided outside the active shield gradient magnetic field coil 3 to suppress the eddy magnetic field generated by the RF shield. However, in the second embodiment, the active shield gradient magnetic field coil 3 is suppressed. An active shield type counter coil having the same principle as that of the active shield type gradient magnetic field coil 3 is provided outside the magnetic field coil 3, and a drive pulse is applied to the coil, so that the eddy magnetic field generated by the RF shield is almost completely eliminated. It generates a magnetic field that cancels out.

  FIG. 6 is a longitudinal sectional view of the gradient coil assembly of the MRI apparatus according to the second embodiment of the present invention. The first embodiment is different from the first embodiment in that an active shield counter coil (ASCC) 21 is provided outside the active shield gradient coil 3 in place of the passive counter shield 13. Is the same. One passive counter shield 13 acts on the active shielding gradient coil 3 of three channels of X, Y, and Z, while the active shielding counter coil 21 acts on each channel. Therefore, it is preferable to provide three counter coils 21 for the X channel, Y channel, and Z channel, but it is also possible to provide a counter shield only for the active shield type gradient magnetic field coil of at least one channel. Since the active shield type gradient coil 3 of each channel is normally composed of four saddle coils, the active shield type counter coil 21 of each channel is also composed of four saddle coils. The active shield counter coil 21 is formed in a cylindrical shape, a plate shape, or a curved shape.

  The active shield type gradient magnetic field coil 3 is provided with an active shield coil having a coil pattern that does not leak the magnetic field generated by the main coil 7 to the outside, but the coil pattern of the active shield type counter coil 21 is the same as this. Designed. That is, the active shield counter coil 21 has a coil pattern that generates a magnetic field that cancels the eddy magnetic field generated in the RF shield 9.

  FIG. 7 shows a coil pattern of one of the four saddle coils constituting the active shield counter coil 21 of the X channel. The horizontal axis is the axial position (Z position), and the vertical axis is the azimuth angle (circumferential position).

  FIG. 8 shows the configuration of the power supply of the active shield type gradient coil (ASGC) 3 and the active shield type counter coil (ASCC) 21. A drive signal having a waveform according to a predetermined pulse sequence is input from the sequence controller 19 to the gradient magnetic field amplifier (G amplifier) 23. A drive current corresponding to this drive signal is supplied from the gradient magnetic field amplifier 23 to the active shield type gradient magnetic field coil 3. The drive signal supplied from the sequence controller 19 to the gradient magnetic field amplifier 23 is also supplied to the shielding magnetic field amplifier (C amplifier) 27 via the waveform shaping circuit 25. The waveform shaping circuit 25 corrects the drive signal waveform according to the vortex response function, and outputs the corrected signal to the shielding magnetic field amplifier (C amplifier) 27. In addition, you may perform correction | amendment of this drive signal not according to an over-response function but according to the result of having actually measured the leakage magnetic field which leaks outside from the active shielding type gradient magnetic field coil 3. FIG.

  A drive current corresponding to the corrected drive signal is supplied from the shield amplifier 27 to the active shield counter coil 21. Therefore, in synchronization with the timing at which the gradient magnetic field is generated from the active shield type gradient magnetic field coil 3, a magnetic field opposite to the eddy magnetic field due to the eddy current generated on the RF shield 9 is generated from the active shield type counter coil 21. To do. Thereby, the eddy magnetic field from the RF shield 9 is completely canceled, and the gradient magnetic field distortion is eliminated.

  Similar to the passive counter shield of the first embodiment, the arrangement of the active shield counter coil 21 is not limited to the outside of the active shield gradient coil 3 and is active with the main coil 7 of the active shield gradient coil 3. It may be provided between the shield coils 5. In short, it may be outside the main coil 7. The active shield counter coil 21 may be provided integrally with the active shield gradient magnetic field coil 3 instead of being provided separately from the active shield gradient magnetic field coil 3.

  As described above, according to the second embodiment, by providing the active shield counter coil 21 that generates a magnetic field that cancels the eddy magnetic field generated by the RF shield, the magnetism between the RF shield and the outside world is provided. It is possible to remove the magnetic coupling almost completely and to correct the magnetic field in the imaging region.

(Third embodiment)
In the first and second embodiments, a counter shield or counter coil is provided at least outside the main coil 7 of the active shield type gradient magnetic field coil 3, and the eddy magnetic field generated by the RF shield 9 is suppressed or canceled by these magnetic fields. It is like that. In the third embodiment, the active shield type gradient magnetic field coil 3 does not shield the eddy magnetic field generated on the RF shield 9. Since the active shield type gradient magnetic field coil 3 is a shield type gradient magnetic field coil, eddy current correction was not required unlike the non-shield type gradient magnetic field coil, but it was used to correct the eddy current on the RF shield. Is a feature of the third embodiment.

  FIG. 9 is a longitudinal sectional view of the gradient coil assembly of the MRI apparatus according to the third embodiment. That is, in the third embodiment, the passive counter shield 13 is omitted from the first embodiment shown in FIG. 2, or the active shield counter coil 21 is omitted from the second embodiment shown in FIG.

  FIG. 10 shows the configuration of the power source of the active shield type gradient magnetic field coil 3. A drive signal having a waveform according to a predetermined pulse sequence is input from the sequence controller 19 to the gradient magnetic field amplifier (G amplifier) 31 via the vortex correction circuit 29. The eddy correction circuit 29 performs overshoot correction on the drive signal of the active shield type gradient magnetic field coil 3 so as to compensate in advance for the dullness of the gradient magnetic field waveform caused by the eddy magnetic field. A drive current corresponding to this drive signal is supplied from the gradient magnetic field amplifier 31 to the active shield type gradient magnetic field coil 3.

  11A to 11C show the drive signal waveform input to the vortex correction circuit 29, the correction signal waveform (overshoot) for vortex correction added to the drive signal, the input drive signal and the correction signal. An output waveform from the vortex correction circuit 29, which is a combined signal, is shown. The correction waveform shown in FIG. 11B corresponds to a waveform of a magnetic field opposite to the eddy magnetic field generated from the RF shield 9. This correction waveform is obtained from the response function of eddy current. Similar to the second embodiment, the waveform may be corrected according to the result of actual measurement of the leakage magnetic field leaking from the active shield type gradient magnetic field coil 3 to the outside. A drive current having a waveform corresponding to the combined waveform shown in FIG. 11C is supplied from the gradient magnetic field amplifier 31 to the active shield type gradient magnetic field coil 3. For this reason, the active shield type gradient magnetic field coil 3 generates an ideal rectangular waveform gradient magnetic field and a magnetic field opposite to the eddy magnetic field. The opposite magnetic field and the eddy magnetic field generated by the RF shield 9 cancel each other. Therefore, only a gradient magnetic field having an ideal rectangular waveform remains, distortion of the gradient magnetic field in the imaging region is eliminated, and a leakage magnetic field to an outer conductor such as a heat shield body of a static magnetic field magnet can be reduced.

  The active shield type gradient magnetic field coil 3 includes three coils of X channel, Y channel, and Z channel. However, if the eddy correction circuit 29 of this embodiment is provided for at least one channel of the active shield type gradient magnetic field coil. Good.

(Fourth embodiment)
In the fourth embodiment, focusing on the fact that the main coil 7 and the active shield coil 5 of the active shield type gradient magnetic field coil 3 originally generate opposite magnetic fields, at least one of the main coil 7 and the active shield coil 5 is used. The waveform of the drive pulse supplied to one of the coils is corrected, the magnetic field generated by either one is adjusted, and the eddy magnetic field of the RF shield 9 is canceled. As modes for correcting the waveform of the drive pulse, (i) a drive signal having a dull waveform (undershoot waveform) is input to the active shield coil 5 of the active shield type gradient magnetic field coil 3, and distortion due to the eddy magnetic field of the RF shield 9 is corrected. There are two possible cases: (ii) a drive signal having an overshoot waveform is input to the main coil 7 of the active shield type gradient magnetic field coil 3 to compensate for distortion caused by the eddy magnetic field of the RF shield 9.

  The fourth embodiment employs the aspect (i). Therefore, the longitudinal sectional view of the gradient coil assembly is the same as that of the third embodiment shown in FIG. The fourth embodiment is based on the fact that the active shield coil 5 of the active shield type gradient magnetic field coil 3 has a coil pattern that almost completely cancels the magnetic field generated by the main coil 7. Therefore, the active shield type gradient magnetic field coil 3 fulfills its original function only when the same drive current flows through the active shield coil 5 and the main coil 7. By the way, since the RF shield 9 is in contact with the inside of the active shield type gradient magnetic field coil 3, an eddy current (current distribution identical to the coil pattern of the main coil 7) is generated so as to cancel the magnetic field generated by the main coil 7. . Therefore, by applying a correction current for correcting the eddy magnetic field to the active shield coil 5, the eddy magnetic field generated in the RF shield 9 can be canceled. On the other hand, the magnetic field by the same drive current flowing through the main coil 7 and the active shield coil 5 is shielded.

  FIG. 12 shows the configuration of the power source of the active shield type gradient magnetic field coil 3. A drive signal having a waveform according to a predetermined pulse sequence is supplied from a sequence controller (not shown) to the main coil 7 via a gradient magnetic field amplifier (G amplifier) 31. The main coil 7 is connected to the active shield coil 5 in series. The drive current flowing through the main coil 7 is divided into two, one is supplied to the active shield coil 5 as it is, and the other is a current waveform shaping circuit 33 and a shunt amplifier 35. To the series circuit.

For this reason, a gradient magnetic field corresponding to the waveform of the drive signal output from the sequence controller is generated from the main coil 7. A series circuit including a current waveform shaping circuit 33 and a shunt amplifier (S amplifier) 35 functions as a correction current supply source that generates a correction current. The active shield coil 5 is supplied with a current obtained by combining the correction current (undershoot current) generated by the current waveform shaping circuit 33 and the drive current flowing through the main coil 7 . That is, by connecting the correction current supply source to the active shield coil 5, the drive current supplied to the main coil 7 and the shield coil 5 can have different waveforms.

  FIG. 13 shows the principle of this correction. By generating a correction magnetic field (differential current between the main coil 7 and the shield coil 5) from the current waveform shaping circuit 33 so as to cancel the eddy magnetic field generated on the RF shield 9, the active shield type gradient magnetic field coil 3 is RF shielded. The magnetic field which cancels the eddy magnetic field which generate | occur | produced on 9 can be generated.

  If the correction current supply source of the fourth embodiment is connected to the active shield coil 5 of the active shield type gradient magnetic field coil of at least one of the X channel, the Y channel, and the Z channel, as in the third embodiment. Good.

(Fifth embodiment)
The fifth embodiment employs the above aspect (ii). Therefore, the longitudinal sectional view of the gradient coil assembly is the same as that of the third embodiment shown in FIG. Similarly to the fourth embodiment, the fifth embodiment is based on the fact that the active shield coil 5 of the active shield type gradient magnetic field coil 3 has a coil pattern that almost completely cancels the magnetic field generated by the main coil 7. Therefore, the active shield type gradient magnetic field coil 3 fulfills its original function only when the same drive current flows through the active shield coil 5 and the main coil 7. By the way, since the RF shield 9 is in contact with the inside of the active shield type gradient magnetic field coil 3, an eddy current (current distribution identical to the coil pattern of the main coil 7) is generated so as to cancel the magnetic field generated by the main coil 7. . Therefore, by applying a correction current for correcting the eddy magnetic field to the main coil 7, the eddy magnetic field generated in the RF shield 9 can be canceled. On the other hand, the magnetic field by the same drive current flowing through the main coil 7 and the active shield coil 5 is shielded.

  FIG. 14 shows the configuration of the power source of the active shield type gradient magnetic field coil 3. A drive signal having a waveform according to a predetermined pulse sequence is supplied from an unshown sequence controller to the active shield coil 5 via a gradient magnetic field amplifier (G amplifier) 31. The active shield coil 5 is connected in series with the main coil 7, and the drive current flowing through the active shield coil 5 is supplied to the main coil 7. A correction current supply source composed of a series circuit of a current waveform shaping circuit 39 and a shunt amplifier 37 is connected in parallel to the main coil 7, and the overshoot current of the drive current flowing through the main coil 7 is the main coil 5 and the active shield coil 5. To the connection point.

  For this reason, a correction drive current in which an overshoot waveform is added to the drive signal waveform output from the sequence controller flows through the main coil 7.

  FIG. 15 shows the principle of this correction. By generating a correction magnetic field that cancels the eddy magnetic field generated on the RF shield 9 from the current waveform shaping circuit 33, the active shielding gradient coil 3 generates a magnetic field that cancels the eddy magnetic field generated on the RF shield 9. be able to.

  If the correction current supply source of the fifth embodiment is also connected to the active shield coil 5 of the active shield type gradient magnetic field coil of at least one of the X channel, the Y channel, and the Z channel, as in the third embodiment. Good.

  The present invention is not limited to the above-described embodiment, and can be implemented with various modifications. For example, the first or second embodiment may be combined with the third, fourth, or fifth embodiment, or the fourth or fifth embodiment may be combined. Further, although the MRI apparatus has been described as the magnetic resonance diagnostic apparatus, the present invention is not limited to this, and can be applied to all diagnostic apparatuses applying other magnetic resonance phenomena such as an NMR spectrum analyzer.

  DESCRIPTION OF SYMBOLS 3 ... Active shield type gradient magnetic field coil, 5 ... Active shield coil, 7 ... Main coil, 9 ... RF shield, 11 ... RF coil, 13 ... Passive counter shield, 21 ... Active shield counter coil

Claims (7)

In a coil device for a magnetic resonance diagnostic apparatus having a gradient magnetic field coil and an RF coil provided inside the gradient magnetic field coil,
The gradient coil is provided on the outside of the main coil for generating a gradient magnetic field in the imaging region and is connected in series with the main coil, and the gradient magnetic field generated from the main coil leaks to the outside. An active shield type gradient magnetic field coil comprising an active shield coil for generating a magnetic field to prevent this,
A current waveform shaping circuit that is connected to the active shield type gradient magnetic field coil and corrects the waveform of the drive current supplied from the gradient magnetic field power source to the main coil or the active shield coil;
A coil device for a magnetic resonance diagnostic apparatus , wherein the drive current supplied to the main coil and the drive current supplied to the active shield coil have different waveforms . The coil device according to claim 1, wherein the correction unit corrects the waveform of the drive current in accordance with a response function of eddy current generated in the RF shield. 3. The coil device according to claim 1, wherein the correction unit is provided for an active shield type gradient magnetic field coil of at least one of X, Y, and Z channels. 4. The coil device according to claim 1, wherein the correction unit corrects undershoot correction of a waveform of a drive current supplied to the main coil or the active shield coil. 5. 4. The coil device according to claim 1, wherein the correction unit performs overshoot correction on a waveform of a drive current supplied to the main coil or the active shield coil. 5. The correction means includes a current waveform circuit to which a drive current flowing through the main coil is supplied,
  The coil according to any one of claims 1 to 5, wherein a combined current of a drive current flowing through the main coil and a current generated from the current waveform circuit is supplied to the active shield coil. apparatus. The correction means includes a current waveform circuit to which a drive current flowing through the shield coil is supplied,
  The coil device according to any one of claims 1 to 5, wherein a combined current of a drive current flowing through the shield coil and a current generated from the current waveform circuit is supplied to the main coil. .

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