JPH0371132B2 - - Google Patents

Description

[Detailed Description of the Invention] [Field of Industrial Application] The present invention relates to nuclear magnetic resonance technology that uses nuclear magnetic resonance (hereinafter abbreviated as "NMR") to obtain a tomographic image of a desired part of a subject (human body). Regarding the receiving coil of a resonance imaging device, especially in the receiving coil where the coil diameter is variable according to the subject and the coil conductor length can be adjusted to the optimum state, it is possible to correct changes in the resonance frequency due to changes in the coil diameter. The present invention relates to a receiving coil for a nuclear magnetic resonance imaging apparatus that can widen the adjustment range of conductor length and obtain high-quality images.

[Conventional technology]

A nuclear magnetic resonance imaging apparatus includes a magnetic field generating means that applies a static magnetic field and a gradient magnetic field to a subject, and a transmission system that irradiates high-frequency signals to cause nuclear magnetic resonance to the nuclei of atoms constituting the living tissue of the subject. , a receiving system that detects high-frequency signals emitted by the above-mentioned nuclear magnetic resonance, and a signal processing system that performs image reconstruction calculations using the high-frequency signals detected by this receiving system. Here, a coil is normally used to detect the high frequency signal in the above-mentioned receiving system, and saddle type, solenoid type, and various types of receiving coils are considered. Since the sensitivity of this receiving coil directly affects the S/N ratio of the reconstructed image, many studies have been made to improve it. For example, DIHoult and
According to RERichards' formula expressing the S/N ratio, it is said that the S/N ratio can be improved by reducing the size (coil diameter) of the receiving coil for high-frequency signals, and this has been confirmed through experiments.

For this reason, as shown in FIG. 4, as a receiving coil 1 of a conventional nuclear magnetic resonance imaging apparatus, a main connector 3 is provided at the base end of the coil conductor 2, and a coil for varying the coil diameter is provided at the distal end. By providing a plurality of diameter adjustment connectors 4a, 4b, 4c and connecting the main connector 3 to any of the coil diameter adjustment connectors 4a, 4b, 4c, the coil diameter can be adjusted to suit the subject or its imaging area. A variable coil conductor length has been proposed in which the length of the coil conductor can be adjusted to the optimum state for the subject. The equivalent circuit of this receiving coil 1 is
As shown in FIG. 5, coil diameter adjustment connectors 4a, 4b, 4c are installed at inductances _L1 , _L2 , and _L3 corresponding to the lengths of the coil conductors 2, respectively.
is provided, and the main connector 3 is connected to this by switching appropriately. In addition, the fifth
In the figure, C ₁ is a resonant capacitor, C ₂ is a matching capacitor, and 5 is a variable capacitance element.

[Problem to be solved by the invention]

However, in such a receiving coil 1,
If the coil diameter is changed according to the subject to be examined and its imaging region, and the coil conductor length is adjusted appropriately according to the subject, the resonant frequency will change as the coil diameter changes. It was hot. For example, to minimize the coil diameter,
When the main connector 3 is connected to the first coil diameter adjustment connector 4a in FIG. 4, the length of the coil conductor is from the main connector 3 to the first coil diameter adjustment connector 4a, and as shown in FIG. The inductance of the receiving coil 1 is _L1 .
Then, the inductance _L1 , the resonant capacitance _C1 , and the matching capacitance _C2 tune the signal frequency. Next, in order to slightly increase the coil diameter, if the main connector 3 is connected to the second coil diameter adjustment connector 4b in FIG. This is the part up to the adjustment connector 4b, and the inductance of the receiving coil 1 in FIG. 5 is the sum of _L1 and _L2 . This inductance (L ₁ + L ₂ ), resonance capacitance C ₁ and matching capacitance C ₂ will tune the signal frequency, but the inductance changed this time
It is necessary to correct the tuning shift due to L ₂ by changing the capacitance C ₃ of the variable capacitance element 5. moreover,
In Fig. 4, when the main connector 3 is connected to the third coil diameter adjustment connector 4c to increase the coil diameter, the coil conductor length becomes even larger, and in Fig. 5, the inductance of the receiving coil 1 becomes even larger. The capacitance C ₃ of the variable capacitance element 5 must be changed significantly to correct the tuning shift due to the change in inductance.

However, the variable capacitance element 5 is controlled by voltage, and there is a limit to the amount of correction.
Therefore, if the coil diameter is made too large and the coil conductor length is made too long, it becomes impossible to correct the tuning deviation due to the increase in the inductance of the receiving coil 1 by using the variable capacitance element 5 alone. For this reason, the inductance of the receiving coil 1 cannot be increased too much, and as a result, the range of adjustment of the coil conductor length by changing the coil diameter must be limited to a narrow range. Therefore, it is not possible to set the optimum coil diameter depending on the subject and its imaging region, and the S/N ratio is reduced, making it impossible to obtain high-quality images.

Therefore, an object of the present invention is to provide a receiving coil for a nuclear magnetic resonance imaging apparatus that can solve such problems.

[Means to solve the problem]

In order to achieve the above object, a receiving coil of a nuclear magnetic resonance imaging apparatus according to the present invention includes a magnetic field generating means for applying a static magnetic field and a gradient magnetic field to a subject;
a transmitting system that irradiates high-frequency signals to cause nuclear magnetic resonance to the nuclei of atoms constituting the living tissue of the subject; a receiving system that detects the high-frequency signals emitted by the nuclear magnetic resonance; It is installed in the receiving system of the nuclear magnetic resonance imaging apparatus, which comprises a signal processing system that performs image configuration calculations using high-frequency signals detected by the system, and the coil diameter is variable according to the subject and the coil conductor length is provided. In a receiving coil that can be adjusted to an optimal state, resonance occurs when the coil diameter is changed between the coil conductor and a plurality of coil diameter adjustment connectors provided at a predetermined interval at the tip of the coil conductor. Resonance capacitors are provided to correct frequency changes, and when the coil diameter is changed appropriately by connecting the main connector provided at the base end of the coil conductor to one of the coil diameter adjustment connectors, the above The connection switching of the resonant capacitors is performed at the same time.

[Effect]

In the receiving coil of the nuclear magnetic resonance imaging apparatus configured in this way, in order to change the coil diameter, the main connector provided at the proximal end of the coil conductor is provided at a predetermined interval at the distal end of the coil conductor. By connecting to one of the multiple coil diameter adjustment connectors provided above, the change in resonance frequency when changing the coil diameter provided at each of the above coil diameter adjustment connectors can be corrected according to the change in coil diameter. It operates by simultaneously switching and connecting the resonant capacitors.

〔Example〕

Embodiments of the present invention will be described in detail below with reference to the accompanying drawings.

FIG. 1 is a circuit diagram showing an equivalent circuit of the receiving coil of the nuclear magnetic resonance imaging apparatus according to the present invention, FIG. 2 is an enlarged explanatory diagram showing the main parts of the receiving coil, and FIG. 1 is a block diagram showing the overall configuration of a nuclear magnetic resonance imaging apparatus to which this is applied.

The above-mentioned nuclear magnetic resonance imaging apparatus obtains a tomographic image of a subject by using the nuclear magnetic resonance (NMR) phenomenon, and as shown in FIG. , a transmission system 9, a reception system 10, a signal processing system 11, a sequencer 12,
A central processing unit (CPU) 13 is provided.

The static magnetic field generating magnet 7 generates a strong and uniform static magnetic field around the subject 6 in a direction perpendicular to the body axis direction, and is a permanent magnet placed in a certain expanse of space around the subject 6. A magnetic field generating means of a normal conduction type, a normal conduction type, or a superconductivity type is arranged. The magnetic field gradient generation system 8 consists of gradient magnetic field coils 14 wound in the three axes of X, Y, and Z, and a gradient magnetic field power supply 15 that drives each coil, and each generates a magnetic field according to instructions from the sequencer 12. By driving the gradient magnetic field power supply 15 of the coil, gradient magnetic fields Gx,
Gy and Gz are applied to the subject 6. Depending on how this gradient magnetic field is applied, a slice plane for the subject 6 can be set. The transmission system 9 irradiates high frequency signals to cause nuclear magnetic resonance in the nuclei of atoms constituting the living tissue of the subject 6, and includes a high frequency oscillator 16 and a modulator 17.
It consists of a high frequency amplifier 18 and a transmitting coil 19a, and transmits the high frequency pulses output from the high frequency oscillator 16 to the modulator 1 according to instructions from the sequencer 12.
7, and this amplitude-modulated high-frequency pulse is amplified by a high-frequency amplifier 18 and then supplied to a transmitting coil 19a placed close to the subject 6, whereby the subject 6 is irradiated with electromagnetic waves. It's becoming like that. The receiving system 10 detects a high frequency signal (NMR signal) emitted by nuclear magnetic resonance of the atomic nucleus of the biological tissue of the subject 6, and includes a solenoid-shaped receiving coil 19b, an amplifier 20, a quadrature phase detector 21, and a /D converter 22,
The electromagnetic wave (NMR signal) in response to the electromagnetic wave irradiated from the transmitting coil 19a of the subject 6 is detected by the receiving coil 19b placed close to the subject 6, and is transmitted via the amplifier 20 and quadrature phase detector 21. The signals are input to the A/D converter 22 and converted into digital quantities, and then sampled by the quadrature phase detector 21 at the timing according to the command from the sequencer 12 to obtain two series of collected data, and the signals are sent to the signal processing system 11. It is starting to be sent. This signal processing system 11 consists of a CPU 13, a recording device such as a magnetic disk 23 and a magnetic tape 24, and a display 25 such as a CRT.The CPU 13 performs processing such as Fourier transform, correction coefficient calculation, image reconstruction, etc. The signal intensity distribution of an arbitrary cross section or the distribution obtained by performing appropriate calculations on a plurality of signals is converted into an image and displayed as a tomographic image on the display 25. Further, the sequencer 12 operates under the control of the CPU 13 and sends various commands necessary for data collection of tomographic images of the subject 6 to the transmission system 9, the magnetic field gradient generation system 8, and the reception system 10. In addition, in FIG. 3, the transmitter coil 19a, the receiver coil 19b, and the gradient magnetic field coil 14
is arranged in the magnetic field space of the static magnetic field generating magnet 7 arranged in the space around the subject 6.

Here, in the present invention, the receiving coil 1
Similar to the conventional example shown in FIG. 4, 9b has a main connector 3 provided at the base end of the coil conductor 2, and predetermined coil diameter adjustment connectors 4a, 4b, 4c at the distal end for varying the coil diameter. By providing a plurality of connectors at intervals and connecting the main connector 3 to any of the coil diameter adjustment connectors 4a, 4b, and 4c, the coil diameter can be varied according to the subject and its imaging region, and the coil conductor The length of the coil conductor 2 can be adjusted to the optimum condition for the subject, and as shown in FIG. connectors 4a, 4b, 4c and the coil conductor 2
A capacitor 26 having a predetermined capacitance is provided between each of them as a resonant capacitor for correcting a change in the resonant frequency when the coil diameter is changed.

As shown in FIG. 1, the equivalent circuit of this receiving coil 19b is such that coil diameter adjustment connectors 4a, 4b, 4c are placed at inductances L ₁ , L ₂ , L ₃ corresponding to the lengths of the coil conductors 2, respectively. are provided, and these coil diameter adjustment connectors 4
A, 4b, 4c and the coil conductor 2 are provided with resonant capacitors C ₄ , C ₅ , C ₆ having predetermined capacities for correcting changes in the resonant frequency when the coil diameter is changed, respectively. Diameter adjustment connectors 4a-4
By appropriately switching and connecting the main connector 3 to C, the resonance capacitances C ₄ , C ₅ and C ₆ can be switched at the same time. In FIG. 1, reference numeral _C2 represents a matching capacitance, and reference numeral 5 represents a variable capacitance element.

Next, the receiving coil 19b configured in this way
This section explains the use and operation of . For example, if the coil diameter is to be minimized depending on the subject or its imaging site, the main connector 3 is connected to the first coil diameter adjustment connector 4a in FIG. 4. At this time, the length of the coil conductor of the receiving coil 19b is from the main connector 3 to the first coil diameter adjustment connector 4a, and in FIG. 1, the inductance of the receiving coil 19b is L ₁ , and the resonance capacitance is becomes C ₄ . Then, the signal frequency is tuned by the inductance _L1 , the resonant capacitor _C4 , and the matching capacitor _C2 . Next, in order to slightly increase the coil diameter, if the main connector 3 is connected to the second coil diameter adjustment connector 4b in FIG. This is the part up to the adjustment connector 4b, and is the receiving coil 19 in FIG.
The inductance of b is (L ₁ +L ₂ ), and
The resonant capacitance is switched to _C5 . Then, the signal frequency is tuned by this inductance (L ₁ +L ₂ ), resonance capacitance C ₅ and matching capacitance C ₂ .
At this time, the tuning deviation due to the inductance L ₂ that has changed this time can be corrected by setting the resonant capacitor C ₅ to a value different from the above-mentioned resonant capacitor C ₄ . Furthermore, if the coil diameter is increased by connecting the main connector 3 to the third coil diameter adjustment connector 4c in FIG. 4, the length of the coil conductor becomes even larger, and the inductance of the receiving coil 19b in FIG. ( _L1 + _L2 + _L3 )
As , the resonant capacitance switches to _C6 . Then, the signal frequency is tuned by this inductance (L ₁ +L ₂ +L ₃ ), resonance capacitance C ₆ , and matching capacitance C ₂ . At this time, the correction of the tuning shift due to the inductance (L ₂ + L ₃ ) that has changed this time is:
This can be achieved by setting the resonant capacitance C ₆ to a value different from the resonant capacitances C ₄ and C ₅ described above.

In this way, as the coil conductor length of the receiving coil 19b increases, the inductance becomes L ₁ , (L ₁ +L ₂ ),
(L ₁ +L ₂ +L ₃ ), the resonant capacitance C ₄ ,
Find the values of C ₅ and C ₆ , and
By connecting the capacitor 26 (see FIG. 2) having a predetermined capacity to each of the coil diameter adjustment connectors 4a to 4c, the signal frequency can always be tuned. At this time, the variable capacitance element 5 shown in FIG. 1 only needs to be finely tuned to the signal frequency. Therefore, the amount of correction of the variable capacitance element 5 for changes in the coil conductor length does not need to be very large.

In addition, in FIGS. 1 and 4, the coil diameter adjustment connectors are provided at three locations, and three resonance capacitors are also provided, but the present invention is not limited to this. They may be provided at four or more locations, and four or more resonance capacitors may also be provided.

[Effects of the Invention] Since the present invention is constructed as described above, in order to change the coil diameter, the main connector 3 provided at the proximal end of the coil conductor 2 is connected to the distal end of the coil conductor 2 at a predetermined interval. By connecting to any one of the plurality of coil diameter adjustment connectors 4a to 4c provided at intervals, the coils provided at each of the coil diameter adjustment connectors 4a to 4c can be adjusted according to changes in the coil diameter. Resonant capacitors C ₄ to _{C 6} that correct changes in resonant frequency when changing the diameter can be switched and connected at the same time. Therefore, the resonant frequency of the receiving coil 19b changes as a result of the change in the inductance of the coil conductor 2 due to the change in the coil conductor length due to the change in the coil diameter. Just by connecting to the coil diameter adjustment connectors 4a to 4c, correction can be performed by simultaneously switching and connecting the correction resonance capacitors _C4 to _C6 . Thereby, the resonant frequency of the receiving coil 19b can be kept constant. Therefore, the coil diameter can be changed by increasing the number of coil diameter adjustment connectors to which resonance capacitances are connected without being limited by the variable capacitance element 5, which has a limited correction amount as in the past. The range of adjustment of the coil conductor length can be increased. From this, it is possible to set the optimum coil diameter according to the subject and the region to be imaged, and it is possible to improve the S/N ratio and obtain high-quality images.

[Brief explanation of drawings]

FIG. 1 is a circuit diagram showing an equivalent circuit of a receiving coil of a nuclear magnetic resonance imaging apparatus according to the present invention, and FIG.
FIG. 3 is a block diagram showing the overall configuration of a nuclear magnetic resonance imaging apparatus to which the receiving coil is applied. FIG. 4 is a receiving coil of the present invention and a conventional example. FIG. 5 is a circuit diagram showing an equivalent circuit of a conventional receiving coil. 1, 19b...Reception coil, 2...Coil conductor, 3
... Main connector, 4a to 4c... Coil diameter adjustment connector, 6... Subject, 7... Static magnetic field generation magnet, 8... Magnetic field gradient generation system, 9... Transmission system, 10... Receiving system, 11
...Signal processing system, 26...Capacitor, _C2 ...Matching capacitance, _C4 to _C6 ...Resonance capacitance, _L1 to _L3 ...Inductance.

Claims (1)

[Claims] 1. A magnetic field generating means that applies a static magnetic field and a gradient magnetic field to the subject, a transmission system that irradiates high-frequency signals to cause nuclear magnetic resonance in the nuclei of atoms constituting the living tissue of the subject, and the above-mentioned nuclear magnetic field. In the receiving system of a nuclear magnetic resonance imaging apparatus, the receiving system includes a receiving system that detects high-frequency signals emitted by resonance, and a signal processing system that performs image reconstruction calculations using the high-frequency signals detected by the receiving system. established,
A receiving coil whose coil diameter is variable according to the subject and whose coil conductor length can be adjusted to an optimum state includes a plurality of coil diameter adjustment connectors provided at a predetermined interval at the tip of the coil conductor; Resonance capacitors are provided between the coil conductors to compensate for changes in resonance frequency when the coil diameter is changed, and a main connector provided at the base end of the coil conductor and one of the coil diameter adjustment connectors described above. 1. A receiving coil for a nuclear magnetic resonance imaging apparatus, characterized in that when the coil diameter is appropriately changed by connecting the receiving coil, the connection switching of the resonance capacitance is performed simultaneously.