KR20010081364A - Receive coil and magnetic resonance imaging method and apparatus - Google Patents

Abstract

PURPOSE: A receive coil, a magnetic resonance imaging method and an apparatus are provided to prevent an electromagnet coupling between loops due to the high impedance generated by parallel resonance in LC circuits. CONSTITUTION: A DC magnetic field generation unit(2) generates the same DC magnet field at its inner space. A gradient coil units(4,4') and a transmitting coil units(6,6') are arranged to face each other in vertical direction with an interval inside of the DC magnetic field generation unit(2). Receive coil units(102,104,106) surrounding each part of patient are attached to an image pick-up table(10). The gradient coil units(4.4') are connected with a gradient driving unit(16). The gradient driving unit(16) feeds a driving signal to the gradient coil units(4,4') so as to generate a gradient magnet field. The transmitting coil units(6,6') are connected with a transmitter(18). The transmitter(18) feeds a driving signal to the transmitting coil unit(6,6') so as to generate an RF magnet field, therefore exciting spins in patient. The receive coil unit receive a magnetic resonance signal generated by spins exited in patient and the received signal are input to a receiver(20). The output of the receiver is input to an A-D converter(22) and the output of the A-D converter are input to the computer(24). The computer(24) stores the digital signal received from the A-D converter(22) at the memory. The computer is connected with a controller(30) and the controller(30) controls the gradient driving(16), the transmitter(18), the receiver(20) and the A-D converter(22) so as to perform a magnetic resonance imaging.

Description

Receiver coil and magnetic resonance imaging method and apparatus thereof {RECEIVE COIL AND MAGNETIC RESONANCE IMAGING METHOD AND APPARATUS}

The present invention relates to a receiving coil and a magnetic resonance imaging method and apparatus, and more particularly, to a receiving coil composed of a plurality of conductive loops having a series capacitor, and a magnetic resonance imaging method and apparatus using the receiving coil.

In a magnetic resonance imaging apparatus, magnetic resonance signals are measured where possible as close to the imaging region, and a receiving coil is placed closely around the patient to improve the signal-to-noise ratio (SNR) of the signal.

Conventional receiving coils include solenoid receiving coils used for spine imaging. Since the spinal column is generally divided into three parts, namely C (neck)-spine, T (chest)-spine, and L (waist)-spine, three types of receiving coils are provided for proper use depending on the part of the spine to be imaged. .

When using a plurality of receiving coils simultaneously, electromagnetic coupling occurs, which changes the resonant frequency of the coil. Therefore, when imaging two or more spinal column portions for the same patient, the respective dedicated receiving coils are properly used and exchanged according to use so that the coupling is prevented.

It is cumbersome to replace the receiving coils when imaging the plurality of spinal column portions as described above, for example, the demand on the patient to change his posture burdens the patient.

On the other hand, the plurality of chucks can be imaged simultaneously if the axial length of the solenoid coil is extended to widen the receiving sensing range in the body axial direction. However, such extended length coils reduce the SNR of the received signal.

It is an object of the present invention to provide a receiving coil composed of a plurality of coils which do not form any coupling, and a magnetic resonance imaging method and apparatus using such a receiving coil.

According to a first aspect of the invention, there is provided a receiving coil consisting of a plurality of conductive loops having series capacitors, each of which comprises a low input-impedance amplifier, The input circuit of is connected in parallel across the capacitor via an inductor.

According to a second aspect of the present invention, a magnetic resonance imaging method is provided, which method includes generating a static magnetic field in a space in which a patient is placed, generating a gradient magnetic field in the space, and generating a high frequency magnetic field in the space. And measuring the magnetic resonance signal from the space, and generating an image based on the measured magnetic resonance signal, wherein measuring the magnetic resonance signal includes a plurality of series capacitors. Is performed using a receiving coil consisting of a conductive loop of s, each of the conductive loops comprising a low input impedance amplifier, the input circuit of the amplifier being connected in parallel across the capacitor via an inductor.

According to a third aspect of the present invention, there is provided a magnetic resonance imaging apparatus, comprising: gradient magnetic field generating means for generating a static magnetic field in a space in which a patient is placed, gradient magnetic field generating means for generating a gradient magnetic field in the space; High frequency magnetic field generating means for generating a high frequency magnetic field in the space, measuring means for measuring a magnetic resonance signal from the space, and image generating means for generating an image based on the magnetic resonance signal measured by the measuring means Wherein the measuring means comprises a receiving coil consisting of a plurality of conductive loops having a series capacitor, each of the conductive loops comprising a low input impedance amplifier, the input circuit of the amplifier being connected across the capacitor via an inductor Is connected in parallel.

According to a fourth aspect of the invention there is provided a self-resonant imaging device as described with respect to the third aspect, wherein the measuring means comprises a second receiving coil consisting of one conductive loop having a series capacitor, and And a frequency selection signal block circuit for connecting the ground side of at least one output signal line of the low input impedance amplifiers to a common ground, wherein the second receiving coil has a sensitivity in a direction perpendicular to the sensitivity direction of the receiving coil. And the loop has a low input impedance amplifier, the input circuit of the low input impedance amplifier being connected in parallel across the capacitor via an inductor.

In any one of the first to fourth aspects of the invention, it is desirable for the loop to have a disable means controlled by an external signal, whereby decoupling performance can be further improved. have.

In addition, in a second or third aspect of the present invention, it is desirable to provide each low input impedance amplifier with a frequency selective signal block circuit connecting the ground side of the output signal line of each low input impedance amplifier to a common ground. In doing so, a received signal having a good SNR is obtained.

Further, in the third or fourth aspect of the present invention, it is preferable to implement the frequency selective signal block circuit as a trap circuit, in which the signal blocking is effectively obtained.

Further, in the fourth aspect of the present invention, it is preferable to implement the second receiving coil in a saddle-type coil, whereby quadrature reception is effectively obtained.

According to the invention, the inductor is connected in parallel to the capacitor via the low input impedance of the amplifier, thereby forming an LC parallel circuit. The high impedance generated by the parallel resonance blocks the coupling between the loops.

Accordingly, the present invention can provide a receiving coil composed of a plurality of coils which does not cause coupling, and a self resonance imaging method and apparatus using such a receiving coil.

Other objects and advantages of the present invention will become apparent from the following detailed description of the preferred embodiments of the present invention with reference to the accompanying drawings.

1 is a block diagram of an apparatus according to an embodiment of the invention,

2 is a view schematically illustrating a configuration of a receiving coil unit in the apparatus shown in FIG. 1;

3 is an electrical circuit diagram of a receiving coil in the apparatus shown in FIG. 1;

4 is an electrical circuit diagram of receiving coils in the apparatus shown in FIG.

5 is an electrical circuit diagram of receiving coils in the apparatus shown in FIG.

6 is an electrical circuit diagram of receiving coils in the apparatus shown in FIG.

7 is an electrical circuit diagram of a receiving coil portion in the apparatus shown in FIG. 1;

8 is an enlarged view of a receiving coil unit in the apparatus shown in FIG. 1;

FIG. 9 is a partial broken line diagram illustrating partial division of a receiving coil portion in the apparatus shown in FIG. 1; FIG.

FIG. 10 illustrates the function of the shape defining member of the receiving coil in the apparatus shown in FIG. 1; FIG.

11 illustrates the function of the shape defining member of the receiving coil portion in the apparatus shown in FIG. 1;

12 is a schematic diagram showing an example of the function of a pulse sequence performed by the apparatus shown in FIG. 1;

13 is a block diagram of an apparatus according to another embodiment of the present invention;

14 is a schematic illustration of the configuration of a receiving coil in the apparatus shown in FIG. 13;

FIG. 15 illustrates electrical connection of receiving coils in the apparatus shown in FIG. 13; FIG.

16 illustrates the electrical connection of the receiving coils in the apparatus shown in FIG. 13;

Explanation of symbols for the main parts of the drawings

4,4 ': gradient coil section 6,6'1: transmission coil section

10: imaging table 20: receiver

22: A-D converter section 24: computer

32: display unit 102, 104, 106: receiving coil unit

362: shape definition member

1 is a block diagram of a magnetic resonance imaging apparatus according to an embodiment of the present invention. The configuration of the device represents one embodiment of the device according to the invention, and the operation of the device represents one embodiment of the method according to the invention.

As shown in Fig. 1, the apparatus includes a static magnetic field generator 2 for generating a homogeneous static magnetic field in its internal space. The static magnetic field generators 2 constitute a pair of magnetic field generators (not shown), such as permanent magnets, which face each other in the vertical direction while maintaining a constant distance, and generate a static magnetic field (vertical magnetic field) of the space therebetween. do.

Similarly, in the internal space of the static magnetic field generating unit 2, inclined coil portions 4, 4 'and transmitting coil portions 6, 6' facing each other in the vertical direction are arranged while maintaining respective distances.

The patient 8 is placed on the imaging table 10 and transferred to a space of the transmission coil portions 6, 6 'facing each other by a conveying means (not shown). The body axis of the patient 8 is orthogonal to the static magnetic field direction. The imaging table 10 is attached with receiving coil portions 102, 104, 106 surrounding each part of the patient 8 to be imaged. Receiving coil parts 102, 104 and 106 represent one embodiment of the receiving coil of the present invention.

The receiving coil section 102 captures the C-spine and is attached to surround the neck of the patient 8. The receiving coil 104 captures the T-spine and is attached to surround the chest of the patient 8. Receiving coil 104 consists of, for example, two coils in pairs. It should be noted that the number of coils in pairs (or groups) is not limited to two but can be any appropriate number. Receiving coil section 106 captures the L-spine and is attached to surround the hip of patient 8. The receiving coil section 106 also consists of two coils, for example in pairs. Again, the number of coils in pairs (or groups) is not limited to two but may be any suitable number. The receiving coil units 102, 104, and 106 will be described in more detail below.

The gradient coil parts 4 and 4 'are connected to the gradient driver 16. The inclined drive unit 16 supplies a drive signal to the inclined coil units 4 and 4 'to generate a gradient magnetic field. The inclined portion 16 together with the inclined coil portion 4.4 'represents one embodiment of the inclined magnetic field generating means of the present invention. The gradient magnetic field to be generated is as follows. That is, a slice gradient magnetic field, a readout gradient magnetic field, and a phase encoding magnetic field.

The transmitting coil parts 6, 6 ′ are connected to the transmitter part 18. The transmitter unit 18 supplies a drive signal to the transmission coil unit 6.6 'to generate an RF magnetic field, thereby exciting the spins in the patient 8.

The receiving coils 102, 104, 106 receive the magnetic resonance signals generated by the spin excited in the patient 8. The receiving coils 102, 104 and 106 receive the magnetic resonance signals at the C-, T- and L- spines, respectively. The receiving coil parts 102, 104, 106 are connected to the inputs of the receiver part 20 and input the received signals to the receiver part 20, respectively.

The receiver unit 20 receives the received signal from one of the receiving coil units 102, 104, 106 via a switch (not shown). Alternatively, three receiver systems corresponding to receiving coils 102, 104, 106 may be provided to receive respective signals. It should be noted that the receiver unit 20 may include a plurality of receiving circuits (not shown) corresponding to the plurality of coils in pairs (or groups).

The output of the receiver section 20 is connected to the input of the A-D converter section 22. The A-D converter section 22 converts the output from the receiver section 20 into a digital signal. The A-D converter section 22 includes an A-D converter circuit corresponding to the receiver circuit. The A-D converter unit 22 may include an A-D conversion system corresponding to three receiving circuits in the receiver unit 20.

The output of the A-D converter section 22 is connected to the computer 24.

The computer 24 receives the digital signal from the A-D converter section 22 and stores it in a memory (not shown). Thus, data space is formed in the memory. When three A-D conversion systems are provided, data space is provided by each system. Each data space constitutes a two-dimensional Fourier space. The computer 24 performs a two-dimensional inverse Fourier transform on the data in the two-dimensional Fourier space to reconstruct the image of the patient 8.

The computer 24 is connected to the control unit 30, and the control unit 30 controls the tilt driver 16, the transmitter unit 18, the receiver unit 20, and the AD converter unit 22 to perform magnetic resonance imaging. .

The computer 24 is operated by a display unit 32 for displaying the reconstructed image and various information output from the computer 24 and a human operator for inputting various commands and information to the computer 24. It is connected to the operation unit 34.

Fig. 2 schematically shows an example of the configuration of the receiving coils 102, 104, and 106 together with the patient 8 and the imaging table 10, where (a) is a plan view and (b) is a cross-sectional view taken along the line AA. . In each figure, three mutually orthogonal directions are indicated by x, y and z. The x-direction is defined as the horizontal direction of the receiving coil parts 102, 104 and 106, the y-direction is defined as the vertical direction of the receiving coil parts 102, 104 and 106 and the z-direction is defined as the axial direction of the receiving coil parts 102, 104 and 106. These definitions apply to the other figures as well.

As shown in Fig. 2, the two partial coils forming the receiving coil portion 106 are semicircular around the hips of the patient 8, and at both ends are connected to the imaging table 10 through their connectors 108. To a corresponding connector (not shown). The lower surface of the imaging table 10 is provided with an electrical circuit 110 which allows each partial coil to be connected with the connector. Thus, two coils are formed, each having a closed loop of one or another suitable number of turns. The receiving coils 104, 102 surround the chest and neck of the patient 8, respectively, and are attached to the imaging table 10 in a configuration similar to that of the receiving coils 106 ,. Each of the receiving coils 104, 102 forms a closed loop.

3 shows an example electrical circuit of a closed loop. The closed loop constitutes the receiving coil. As shown, a closed loop is formed by connecting capacitors 302 and conductors 304 in series. Reference numerals for the capacitors and conductors are indicated for only one position in the figure. Closed loops represent one embodiment for the loop of the present invention, and capacitor 302 represents one embodiment for the capacitor of the present invention. The number of capacitors 302 is not limited to four as shown and may be any suitable number. Conductor 304 represents one embodiment for the conductor of the present invention.

Opposing ends of a capacitor 302 are connected to a preamplifier 306 via an inductor 308, which amplifies the magnetic resonance signal received by the closed loop. Inductor 308 represents one embodiment for an inductor of the present invention. As the preamplifier 306, an amplifier whose input circuit has a sufficiently low impedance, that is, a low input impedance amplifier is used. Preamplifier 306 represents an embodiment of a low input impedance amplifier of the present invention.

Since the preamplifier 306 is a low input impedance amplifier, an LC parallel circuit is formed in substantially such a closed loop by the capacitor 302 and the inductor 308. The resonant frequency of the LC parallel circuit is selected to match the magnetic resonance signal. Thus, when a magnetic resonance signal is received, the LC parallel circuit resonates such that the associated high impedance substantially leaves the closed loop open.

Receiving coil portions 102, 104 and 106 composed of such closed loops are arranged on the imaging table 10 at a predetermined distance from their loop surfaces facing each other. 4 shows the electrical circuits of the receiving coil parts 102, 104 and 106 arranged in the imaging table 10. As shown in FIG. 4 is shown at an angle at which loop surfaces of the coil are seen obliquely. As shown, the receiving coil portion 102 is composed of a single closed loop, and each of the receiving coil portions 104, 106 is composed of two closed loops. It will be readily appreciated that the receiving coil portion 102 may also consist of two or more closed loops, if desired.

In the receiving coil section 102, both ends of the capacitor 422 are connected to the input circuit of the preamplifier 426 through the inductor 428. The capacitor 422 and the inductor 428 constitute a parallel resonance circuit which resonates at the frequency of the magnetic resonance signal.

In the receiving coil unit 104, both ends of the capacitor 422 are connected to the preamplifier 446 through the inductor 448 in the case of one closed loop, and both ends of the capacitor 442 'are connected to the inductor 448 in the other closed loop. ') Is connected to the input circuit of the preamplifier 446'. In these closed loops, the capacitor 442 and the inductor 448, and the capacitor 442 'and the inductor 448' each constitute a parallel resonant circuit which resonates at the frequency of the magnetic resonance signal.

In receive coil section 106, both ends of capacitor 462 in one closed loop are connected to the input circuit of preamplifier 466 via inductor 468, and both ends of capacitor 442 'in the other closed loop. Is connected to the input circuit of preamplifier 466 'via an inductor 468'. In these closed loops, the capacitor 462, the inductor 468, and the capacitor 462 'and the inductor 468' each constitute a parallel resonant circuit that resonates at the frequency of the magnetic resonance signal.

Thus, all of the closed loops are substantially open due to the high impedance generated by the parallel resonance of the LC circuit upon receipt of the magnetic resonance signal. All of the closed loops in the receiving coils 102, 104 and 106 are substantially open when receiving the magnetic resonance signal so that no coupling is generated between the closed loops. Because of such decoupling, each of these closed loops exists as if they existed separately, allowing individual magnetic resonance signals to be received without being affected by neighboring closed loops. That is, the receiving coil parts 102, 104, and 106 together constitute a so-called phased-array coil. It is desirable to integrate a portion consisting of the LC circuit and the preamplifier in each receiving coil portion with the electric circuit 110 under the surface of the imaging table 10, thereby preventing work such as mounting the preamplifier and winding joints. It becomes easy.

Receiving coil sections 102, 104, 106 are provided with controllable disable circuits, respectively, to further ensure decoupling between them. 5 and 6 show examples of the electric circuits of the receiving coil parts 102, 104 and 106 having such disable circuits. The same components as those described with reference to FIG. 4 in FIGS. 5 and 6 are denoted by the same reference numerals, and description thereof is omitted.

As shown in Fig. 5, in the receiving coil section 102, a series circuit composed of an inductor 524 and a diode 526 is connected in parallel with the capacitor 522 in a closed loop. Opposite or reverse bias is selectively applied across the diode 526 from a bias circuit (not shown). The application of the bias is controlled by the controller 30. The capacitor 522, the inductor 524, the diode 526 and the bias circuit constitute the disable circuit.

When forward bias is applied, diode 526 is turned on and the other LC parallel circuit is closed loop. The resonant frequency of the LC parallel circuit is chosen to be equal to the frequency of the self resonant signal, and this closed loop is opened by the additional high impedance generated by the parallel resonance. When reverse bias is applied, diode 526 is turned off so that no LC parallel circuit is formed.

The same disable circuit is also provided in the receiving coils 104 and 106. In particular, the capacitor 542, the inductor 544, the diode 546 and the bias circuit together form a disable circuit for one loop of the receiving coil section 104, and the capacitor 542, the inductor 544 and the diode ( 546 ′ configures a disable circuit for the other loop of the receiving coil unit 104. Similarly, capacitor 562, inductor 564, diode 566 and bias circuit together form a disable circuit for one loop of receive coil portion 106, capacitor 562 ', inductor 564'. The diode 566 ′ and the bias circuit together form a disable circuit for the other loop of the receive coil section 106.

In Fig. 6, a diode 626 is connected in series in a closed loop of the receiving coil portion 102. In FIG. Forward or reverse bias is selectively supplied to the diode 626 from a bias circuit (not shown). The diode 626 and the bias circuit make up the disable circuit. The bias circuit is controlled by the controller 30. By turning off the diode 626 with reverse bias, conduction through the closed loop is interrupted and the closed loop is disabled. By turning on diode 626 with forward bias, conduction through the closed loop is enabled so that the closed loop acts as a receiving coil.

The receiving coils 104, 106 are likewise provided with a disable circuit. In particular, the diode 646 and the bias circuit constitute a disable circuit for one loop of the receive coil portion 104, and the diode 646 ′ and the bias circuit disable the other loop of the receive coil portion 104. Configure the circuit. Similarly, the diode 666 and the bias circuit constitute a disable circuit for one loop of the receive coil portion 106, and the diode 666 ′ and the bias circuit disable the other loop of the receive coil portion 106. Configure the circuit.

In addition, a circuit as shown in FIG. 7 is used to prevent a decrease in the Q value of the closed loop due to the ON resistance of the diode, where a plurality of circuits are connected in parallel to each other, each of which is a capacitor And the diodes are connected in series with the anode and cathode of one diode, so that the polarities of the diodes are inverted alternately and all diodes are connected in series with a bias circuit (not shown). Thus the on resistance of the diodes at both ends of the parallel circuit can be reduced in proportion to the number of parallel lines of the circuit, thereby reducing the effect of the on resistance of the diode in the closed loop.

In such receiving coil portions 102, 104, 106, decoupling with the receiving coil portion (s) used to receive the magnetic resonance signal by operating the disable circuit (s) of the receiving coil portion (s) not used to receive the magnetic resonance signals. This can be achieved perfectly. These disable circuits are used for decoupling with the transmitting coil portion 6.6 'in RF excitation by the transmitting coil portion 6.6'.

Therefore, the receiving coil parts 102, 104 and 106 are decoupled from each other, so that the same effects as in the case where the receiving coil parts are present individually can be obtained. Therefore, the receiving coil parts 102, 104 and 106 can receive respective magnetic resonance signals from the C-, T- and L- spins without being affected by the neighboring receiving coils and reducing the SNR, respectively. . Moreover, magnetic resonance signals from the vertebral columns at a plurality of locations can be received simultaneously with good SNR.

It is desirable to provide a plurality of types of receiving coil parts 102, 104, and 106 having different sizes so as to be selected according to the size of the patient, whereby the measurement of the magnetic resonance signal can be appropriately performed. Moreover, it is preferable to arrange the connectors on the imaging table in a plurality of positions corresponding to the respective connectors so as to be suitable for the difference in the body size of the patient 8.

According to the receiving coils 102, 104 and 106, since the received signals from the closed loops are separated from each other, individual images may be reconstructed from the received signals. In particular, images of C-spindle, T-spindle and L-spindle can be obtained separately. Since the received signals have good SNR, the quality of reconstructed images is increased.

In addition, the individual images are used to obtain combined images of the C- and T-spindles, combined images of the T- and L-spindles, or combined images of all of the C-, T- and L-spindles. These combined images also have high quality.

Thus, since the vertebrae at a plurality of positions can be imaged individually or simultaneously, there is no need to replace the receiving coils each time the imaging area changes, as in conventional systems. Therefore, workability is remarkably improved, and the patient does not feel any burden.

The phase-array coil may consist of an appropriate number of closed loops. Thus, a receiving coil having a good SNR capable of imaging not only the spinal column of the patient 8 but also all body parts can be provided.

8 shows a partial coil of the receiving coil portion 106 in a sheet shape. The internal structure of the receiving coil section 106 in this state is shown as a partially broken view in FIG. The receiving coils 104 and 102 have a similar structure. In FIG. 9, portions in the vertical direction are enlarged for convenience of description.

As shown in FIG. 9, the receiving coil section 106 is composed of a flexible substrate 360. As shown in FIG. The flexible substrate 360 is provided with, for example, an electric path pattern by a printed circuit. A pair of shape defining members 362 on the length of the flexible substrate 360 are provided at the periphery in the longitudinal direction of the upper surface of the flexible substrate 360. ) Is placed. The upper surface of the flexible substrate 360 corresponds to its inner side when a semicircular cylinder is formed. The shape defining member 362 is formed of, for example, plastic.

The shape defining member 362 has a predetermined thickness in the y-direction such that warpage is substantially prevented. The shape defining member 362 includes a plurality of U-shaped notches 364. The notch 364 is cut in the z-direction and is open upwards. The notch 364 has a depth that defines the shape defining member 362. The thickness of the bottom of the notch 364 is extremely reduced, so that sufficient flexibility is obtained, or alternatively, the thickness of the bottom may be zero (0).

Such formation defining member 362 allows only the flexible portion thereof (i.e., the bottom of the notch) to bend when the flexible substrate 360 is bent in the direction in which the cylinder is to be formed, the amount of bending as shown schematically in FIG. Likewise, it is limited only to the extent of closing the opening of the notch 364. The allowable bending amount is determined by the width of the notch. The wider the notch, the larger the bending allowance.

The width and spacing in the x-direction of the notch 364 is determined by the amount of bending in all portions of the flexible substrate 360 when forming the cylinder. Thus, as illustrated by way of example and schematically in FIG. 11, bending of the flexible substrate is to be provided when the cylinder is formed. Although only the left side is shown in FIG. 11, the right side may also be formed symmetrically thereto. Such bending defines the shape of the bending of the cylinder, i. By defining the bent shape, the electromagnetic field conditions of the receiving coil section 106 are fixed, thereby enabling stable imaging.

On the shape defining member 362 and the flexible substrate 360, an impact absorbing member 366, such as a sponge, is provided. Similar shock absorbing members 366 are also provided on the bottom surface of the flexible substrate 360. All of these structures are encased in an envelope 368 to be bound to a connector 108 at both ends of the receiving coil portion 106.

The operation of the device according to the invention will now be described. The patient 8 is placed on the imaging table 10 in the side space (not shown), and at the same time, the receiving coils 102, 104 and 106 are attached to the connector. In this way, three coils of the receiving coil portions 102, 104 and 106 are formed around the neck, chest and hip of the patient. It will be appreciated that these three coils need not be attached, and that the appropriate coil (s) can be selected depending on the purpose of the imaging.

Next, the imaging table 10 is transferred to the internal space of the static magnetic field generating unit 2, and imaging is started. This imaging is performed under the control of the control unit 30. The following description relates to imaging by a spin-echo method as a special example of self-resonant imaging. The spin-echo method uses a pulse sequence as shown in FIG.

12 is a schematic diagram of a pulse sequence for obtaining a magnetic resonance signal (spin-echo signal) for one view.

Such a pulse sequence is repeated 256 times to obtain spin-echo for 256 views, for example.

The process of performing the pulse sequence and obtaining the spin-echo signal is controlled by the controller 30. It will be appreciated that this self-resonant imaging is not limited to only the spin-echo method, but can also be performed by various other techniques such as a gradient echo method.

As shown in Fig. 12 (6), the pulse sequence is divided into four periods (a)-(d) along the time axis. First, RF excitation is obtained by a 90 ° pulse P90 in the period a as shown in Fig. 12 (1). RF excitation is performed by the transmitting coil parts 6, 6 'which are driven by the transmitter part 18.

In accordance with the RF excitation, a fractional gradient magnetic field Gs is applied as shown in Fig. 12 (2). The application of the fractional gradient magnetic field Gs is performed by the gradient coil portions 4 and 14 'which are driven by the gradient drive unit 16. Thus, the spindle at a predetermined fragment within the patient 8 is excited (selective excitation). If the receiving coil portions 102, 104 and 106 include a disable circuit, the operation of the disable circuit is enabled in this period. Thus, decoupling is obtained between the transmitting coil section 6.6 'and the receiving coil sections 102, 104 and 106.

Next, the phase-encoding gradient magnetic field Gp is applied in the period b as shown in Fig. 12 (3). The application of the phase-encoding gradient magnetic field Gp is also performed by the gradient coil portions 4, 4 'which are driven by the gradient driver 16. Thus, phase encoding for spin is achieved.

Also, in the phase-encoding period, rehasing for the spindle is achieved by the fractional gradient magnetic field Gs as shown in Fig. 12 (2). In addition, a read gradient magnetic field Gr is applied as shown in Fig. 12 (4) to dephasing the spindle. The application of the read slope magnetic field Gr is also performed by the gradient coil sections 4, 4 'driven by the gradient drive section 16. FIG.

Then, 180 ° pulse P180 is applied in period C as shown in Fig. 12 (1), causing the spindle to reverse. The inversion of these spins is performed by the transmitting coil parts 6, 6 'driven by the transmitter part 18 at high frequency RF. If the receiving coil portions 102, 104, 106 include a disable circuit, the operation of the disable circuit is enabled in this period, so that decoupling between the transmitting coil portions 6, 6 'and the receiving coil portions 102, 104, 106 is disabled. Will be done.

Next, as shown in Fig. 12 (4), the read gradient magnetic field Gr is applied in the period d. In this way, as shown in Fig. 12 (5), the spin-echo signal MR is generated. This spin-echo signal MR is received by the receiving coil parts 102, 104 and 106. Since the receiving coil portions 102, 104, 106 constitute phased array coils, they have good SNR and C-spin without being affected by neighboring coils. It is possible to receive spin-echo signals of T- and L- spins. If the receiving coil portions 102, 104, 106 include a disable circuit, the unused receiving coil portion (s) are disabled so that decoupling is obtained completely.

The received signal is input to the computer 24 through the receiver unit 20 and the A-D converter unit 22. The computer 24 stores this input signal in memory as measurement data. Thus, spin-echo data for one view is collected in memory.

The operation is repeated for example 256 times in a predetermined cycle. The phase-encoding gradient magnetic field Gp changes with each iteration, resulting in a phase encoding change every time. This is represented by a plurality of broken lines in the waveform of Fig. 12 (3).

Computer 24 performs image reconstruction based on spin-echo data for all views collected in memory, and generates images of C- and T- and L-spints.

The images of the C-spindles T-spindles and L-spindles thus generated are displayed on the display unit 32 as visual images. Alternatively, composite images of C-spindle T-spindles, composite images of T- and L-spindles, or composite images of all these spines may be displayed. Since the received signals have good SNR, the quality of these images is improved.

Up to now, although described with respect to the receiving coil for spinal column imaging, not only the receiving coil for spinal column imaging, but also the receiving coils for other body parts can be formed as phase-array coils having the same configuration, and the same effect is also obtained. You can get it.

Fig. 13 shows a block diagram of a magnetic resonance imaging apparatus according to another embodiment of the present invention. As shown, the apparatus includes a receiving coil unit 120. Except for the receiving coil unit 120, the same components as shown in FIG. 1 are denoted by the same reference numerals, and description thereof will be omitted.

Fig. 14 schematically shows the configuration of the main part of the receiving coil part 120. As shown, the receiving coil part 120 includes a saddle-shaped coil 112 and solenoid coils 114 and 116, which are z- Along the axis are arranged concentrically on their axis. The saddle coil 112 represents one embodiment of the second receiving coil and represents one embodiment of the saddle coil of the present invention.

If the three mutually orthogonal directions are defined as x, y, z and the direction of the static magnetic field is aligned with the y-direction, the saddle coil 112 is a linear path 212, 214, 212 ', 214'. Having closed loops, these paths extending in the z-direction parallel to each other and arcuate paths 232, 234, 232 ′, 234 ′ link the linear paths.

Linear paths 212, 214 lie in the x-z plane, and linear paths 212 ', 214' lie in another x-z plane spaced apart from the plane in the y-direction. Arched paths 232, 234 lie in the x-y plane, linking the first ends of linear paths 212, 212 ′ and linking the first ends of linear paths 214, 214 ′, respectively. The arcuate paths 232 ′, 234 ′ lie in another xy plane spaced apart from the xy plane in the z-axis direction, linking the second ends of the linear paths 212, 214 ′, and the linear paths 214, 212 ′).

The saddle coil 112 forms a generally cylindrical space defined by linear paths 212, 214, 212 ', 214' and arcuate paths 232, 234, 232 ', 234' and the patient is in the cylindrical space. Maintaining the imaging portion of the device receives a magnetic resonance signal generated therefrom, having a sensitivity direction aligned with the x-direction. The received signal is picked up from both ends of a capacitor (not shown) placed in the proper position in the loop, via a low input impedance preamplifier connected through an inductor in the manner described above with respect to FIG.

The solenoid coils 114 and 116 are identical to the receiving coil portions 104 and 106 used in the apparatus shown in Fig. 1, one of which consists of two coil loops. These coil loops are configured to enclose a cylindrical space supporting the patient and receive a magnetic resonance signal having a sensitivity direction aligned with the z-direction.

The received signal of each of the solenoid coils 114, 116 is the same as the receive coils 104, 106 used in the apparatus shown in Fig. 1, one of which consists of two coil loops. These coil loops are configured to enclose a cylindrical space in which the patient is placed and receive a magnetic resonance signal having a sensitivity direction aligned with the z-direction.

The received signal of each of the solenoid coils 114, 116 is connected to a low input impedance pre-connected across an end of a capacitor (not shown) disposed in a suitable position in the loop via an inductor in the same manner as in the receive coils 104, 106. Picked up via amplifier.

Since the saddle coil 112 and the solenoid coils 114 and 116 have respective reception sensitivity in mutually orthogonal directions, the SNR can be improved by the quadrature method by adding the received signals to the saddle coil and the solenoid coil. It becomes possible.

Such saddle coils 112 and solenoid coils 114, 116 are connected to the receiver portion 20 via concentric coils. 15 is a schematic diagram illustrating an example of these connections. As shown, the output signal from the preamplifier 720 picking up the received signal of the saddle coil 112 is transmitted to the receiver section 20 through the concentric cable 712 as described above. As shown, output signals from the preamplifiers 704, 704 'that pick up the received signals of the two loops in the solenoid coil 114 are fed to the receiver section 20 via concentric cables 714, 714'. Is sent. Output signals from preamplifiers 706, 706 ′ which pick up the received signals of the two loops in solenoid coil 116 are transmitted to receiver 20 via concentric cable 716 716 ′.

The outer conductors of the concentric cables 712-716 ′ are connected to a common ground at the side of the receiver portion 20. At this time, the external conductor of the concentric cable 712 is connected to the common ground via a frequency selection signal block circuit, ie, a trap circuit 722 consisting of a parallel circuit of an inductor and a capacitor. The outer conductor of the concentric cable represents an embodiment on the ground side of the output signal line of the present invention.

Since the parallel resonant frequency of the trap circuit 722 becomes equal to the frequency of the received signal, that is, the self resonant signal, the outer conductor of the concentric cable 712 is connected to the ground with high impedance with respect to the received signal. This high impedance prevents coupling between concentric cables 712 and concentric cables 714 and 716 'through a common ground. Therefore, if not, the high frequency current flowing electrically between the concentric cable 712 and the concentric cables 714-716 'is blocked by the coupling through the common ground portion, thereby improving the SNR of the received signal. .

As illustrated in FIG. 16, similar trap circuits 724-726 ′ can be provided for the concentric cables 712-716 as well as the concentric cables 714-716 ′. This cuts off the circuit current more completely, thereby further improving the SNR. It will be readily appreciated that the trap circuit can be used in the magnetic resonance imaging apparatus shown in FIG.

As a result, according to the present invention, the high impedance generated by the parallel resonance of the LC circuit effectively prevents coupling between loops, and also greatly improves the SNR, thereby improving the quality of the side surfaces.

Many other different embodiments of the invention may be applied without departing from the spirit and scope of the invention. It should be understood that the invention is not limited to the specific embodiments described in the detailed description of the present application, except as defined in the appended claims.

Claims (8)

A receiving coil comprising a plurality of conductive loops having series capacitors, each of the loops comprising a low input impedance amplifier, the input circuit of the low input impedance amplifier being connected across the capacitor via an inductor. . Generating a static magnetic field in the space in which the patient is placed, Generating a gradient magnetic field in the space; Generating a high frequency magnetic field in the space; Measuring a magnetic resonance signal from the space, and Generating an image based on the measured magnetic resonance signal; Measuring the self resonant signal is performed using a receiving coil consisting of a plurality of conductive loops having series capacitors, each of the conductive loops comprising a low input impedance amplifier, the input circuit of the amplifier being an inductor. Self-resonance imaging method characterized in that connected to both ends of the capacitor in parallel. Gradient magnetic field generating means for generating a static magnetic field in the space in which the patient is placed; Gradient magnetic field generating means for generating a gradient magnetic field in the space; High frequency magnetic field generating means for generating a high frequency magnetic field in the space; Measuring means for measuring a magnetic resonance signal from the space, and Image generating means for generating an image based on the magnetic resonance signal measured by the measuring means, The measuring means comprises a receiving coil consisting of a plurality of conductive loops having a series capacitor, each of the conductive loops comprising a low input impedance amplifier, the input circuit of the amplifier being parallel across the capacitor via an inductor Self-resonance imaging device, characterized in that connected. The method of claim 3, wherein The measuring means, A second receiving coil consisting of one conductive loop having a series capacitor, A frequency selection signal block circuit for connecting the ground side of at least one output signal line of the low input impedance amplifiers to a common ground, The second receiving coil has sensitivity in a direction perpendicular to the sensitivity direction of the receiving coil and the loop has a low input impedance amplifier, and the input circuit of the low input impedance amplifier is connected in parallel across the capacitor through an inductor. Magnetic resonance imaging device, characterized in that. The method of claim 1, Receiving coil, characterized in that the loop has disable means controlled by an external signal. The method of claim 3, wherein Receiving coil, characterized in that the loop has disable means controlled by an external signal. The method of claim 3, wherein And the frequency selection signal block circuit is a trap circuit. The method of claim 4, wherein And the second receiving coil is a saddle-shaped coil.