JPH0454939A - High frequency coil for magnetic resonance imaging apparatus - Google Patents

Abstract

PURPOSE:To enable reduction is coupling between two conductive loops by making the two conductive loops composing a high frequency coil narrower in conductor width thereof at a crossing part therebetween to reduce an area opposed thereto. CONSTITUTION:For example, in the case of a vertical magnetic field type QD coil, a solenoid coil 22 is wound on an outer circumferential surface of a cylindrical resin bobbin 21 in a circumferential direction as one conductive loop while a saddle-shaped coil 23 is arranged as the other conductive loop with a direction of receiving orthogonal to a direction of receiving of the solenoid coil 22, and the coils 22 and 23 are divided with a capacitor 24 to lower an operation voltage. The solenoid coil 22 and the saddle-shaped coil 23 are made narrower in conductor width at a crossed part 25 therebetween to reduce an area opposed thereto. Here, with a reduction in area S of planar conductive plates A1 and A2, an electric capacitance C between the both also becomes smaller. This lessens a floating capacitance as formed the two coils 22 and 23 thereby enabling the lowering a coupling by a capacitive coupling between both the coils.

Description

DETAILED DESCRIPTION OF THE INVENTION [Field of Industrial Application] The present invention relates to a magnetic resonance imaging apparatus that obtains a tomographic image of a desired part of a subject (human body) by utilizing the nuclear magnetic resonance (hereinafter abbreviated as rNMRJ) phenomenon. Regarding a high-frequency coil used in a transmission system or a reception system, which is formed into a set of two conductive loops with their sensitivity directions orthogonal to each other, it is possible to reduce coupling between the two conductive loops in particular. The present invention relates to a high-frequency coil for a magnetic resonance imaging device.

[Prior Art] A magnetic resonance imaging apparatus includes a magnetic field generating means that applies a static magnetic field and a gradient magnetic field in a direction perpendicular to the body axis direction of 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; The image forming apparatus is configured to include a signal processing system that performs image reconstruction calculations using high-frequency signals detected by the system. Then, while applying a uniform static magnetic field to the subject using the static magnetic field generating means, a high-frequency signal with a frequency that excites nuclear magnetic resonance is applied by a high-frequency coil of the transmission system, and thereby a nuclear magnetic resonance signal is emitted from the subject. was detected by a high-frequency coil in the receiving system. At this time, in order to specify the emission position of the nuclear magnetic resonance signal from the subject, imaging was performed by further applying a gradient magnetic field using a gradient magnetic field generating means.

Conventionally, a high-frequency coil in such a magnetic resonance imaging apparatus uses one conductive loop, such as a solenoid coil or a saddle-shaped coil, to receive nuclear magnetic resonance signals in one direction. On the other hand, S
In order to improve the /N ratio, there is a method in which two conductive loops are formed as a set with their sensitivity directions orthogonal to each other, and receive nuclear magnetic resonance signals in two directions. A high-frequency coil made by combining the latter two conductive loops is called a quadrature receiving coil (Qua
rdrature Detection Co11s:
Hereinafter referred to as "rQl) coil", the conventional
As the D coil, a combination of a saddle-shaped coil and a saddle-shaped coil has been proposed, for example, as a horizontal magnetic field type. However, when a combination of a saddle-shaped coil and a saddle-shaped coil is used, especially in a vertical magnetic field type magnetic resonance imaging apparatus, the direction of the static magnetic field coincides with the receiving direction, making it impossible to receive with good sensitivity. Therefore, recently, for example, a combination of a solenoid coil and a saddle-shaped coil has been proposed as a vertical magnetic field type QD coil.

[Problem to be solved by the invention]

However, among such conventional high-frequency coils, especially vertical magnetic field type QD coils, because they are a combination of different coils, such as a solenoid coil and a saddle-shaped coil, coupling between each coil is not likely to occur. there were. Here, coupling refers to the fact that when a high-frequency current is passed through one coil, the high-frequency current leaks to the other coil. When such coupling occurs, each coil This acts as a loss in each coil, reducing the sensitivity of the high-frequency coil as a whole. Therefore, the S/N ratio of the obtained image may deteriorate.

Here, the cause of coupling in the above-mentioned high-frequency coil is that the distance between the intersections of the two coils is close to each other, so a stray capacitance is formed between them, causing capacitive coupling that leaks to the other side, or one of the coils. Inductive coupling can be considered, where the magnetic flux causes an imbalance with the magnetic flux of the other coil. Coupling due to inductive coupling can be reduced by arranging a good conductor, such as a copper plate, near the coil to adjust the unbalance of the magnetic flux, and is not a problem. On the other hand, coupling due to capacitive coupling can be reduced by increasing the distance between the coils at the intersection of the two coils to reduce the stray capacitance formed between the coils. That is, as shown in FIG. 5, two flat conductor plates A1. A1.A are arranged close to each other in parallel (corresponding to the intersection of two coil conductors), when both A1. The distance between A2 is d, and each plane conductor plate A
□.

If the area of A2 is S and the dielectric constant between them is ε, then the above two planar conductor plates A1. The electric capacitance C between A2 is expressed by the following formula.

εSC=□...(1)
As is clear from this equation (1), two flat conductor plates A1. By increasing the distance d between A2, the electric capacitance C between them becomes smaller.

Therefore, in the past, the distance between the intersections of two coils was increased to reduce the stray capacitance formed between each coil, thereby reducing the coupling due to capacitive coupling.

However, in this case, as the nuclear magnetic resonance frequency becomes higher, the interval between each coil must be increased, which increases the size of the entire high-frequency coil.

Furthermore, as the distance of at least one of the coils from the subject increases, the sensitivity further decreases and the S/N ratio deteriorates.

SUMMARY OF THE INVENTION An object of the present invention is to provide a high-frequency coil for a magnetic resonance imaging apparatus that can solve these problems and reduce the coupling between two conductive loops (coils).

[Means to solve the problem]

In order to achieve the above object, the high-frequency coil of the magnetic resonance imaging apparatus according to the present invention includes a magnetic field generating means that applies a static magnetic field and a gradient magnetic field to a subject, and a nuclear magnetic field that applies a nuclear magnetic field to the nuclei of atoms constituting the living tissue of the subject. A transmitting system that irradiates high-frequency signals to cause resonance, a receiving system that detects the high-frequency signals emitted by the above-mentioned nuclear magnetic resonance, and image reconstruction calculations are performed using the high-frequency signals detected by this receiving system. A signal processing system is provided in the transmission system or reception system of the magnetic resonance imaging apparatus, and the two conductive loops are formed as a set with their sensitivity directions orthogonal to each other, and the two conductive loops are configured to transmit high-frequency signals to the subject. In a high-frequency coil whose sensitivity direction is orthogonal to the static magnetic field and detects high-frequency signals emitted by nuclear magnetic resonance,
The width of the conductor at the intersection of the two conductive loops is narrowed to reduce the opposing area.

Further, the conductor width of either one of the intersections of the two conductive loops may be formed narrow.

Furthermore, it is effective to use a solenoid coil and a saddle-shaped coil as the two conductive loops.

[For production]

The high-frequency coil of the present invention configured in this manner reduces the stray capacitance formed at the intersection by narrowing the conductor width and reducing the opposing area at the intersection of the two conductive loops that make up the coil. It works like that. Thereby, the capacitive coupling can be relaxed and the coupling between the two conductive loops can be reduced without increasing the interval between the intersections of the two conductive loops.

〔Example〕

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

FIG. 1 is a perspective explanatory diagram showing an embodiment of the high-frequency coil of the magnetic resonance imaging apparatus according to the present invention, FIG. 2 is a circuit diagram showing the principle and connection of the high-frequency coil, and FIG. 3 is a diagram showing the principle and connection of the high-frequency coil. 1 is a block diagram showing the overall configuration of a magnetic resonance imaging apparatus to which this is applied.

The above-mentioned magnetic resonance imaging apparatus includes nuclear magnetic resonance (NMR).
) phenomenon to obtain a tomographic image of a subject, and as shown in FIG. 3, a static magnetic field generating magnet 2.

A magnetic field gradient generation system 3, a transmission system 4, a reception system 5, a signal processing system 6, a sequencer 7, and a central processing unit (CPU) 8
It consists of:

The static magnetic field generating magnet 2 generates a uniform static magnetic field around the subject l in the body axis direction or in a direction perpendicular to the body axis, and is used to generate a uniform static magnetic field in a certain expanse of space around the subject l. A magnetic field generating means of a permanent magnet type, a normal conduction type, or a superconducting type is arranged. The magnetic field gradient generation system 3 consists of gradient magnetic field coils 9 wound in the three axes directions of x, y, and z, and a gradient magnetic field power supply 10 that drives each coil. By driving the gradient magnetic field power supply 10 of x,
Gradient magnetic field G x in the three wheel directions of y and z.

Gy and Gz are applied to the subject 1. Depending on how this gradient magnetic field is applied, a slice plane for the subject 1 can be set.

The transmission system 4 irradiates high frequency signals (electromagnetic waves) to cause nuclear magnetic resonance in the nuclei of atoms constituting the living tissue of the subject 1, and includes a high frequency oscillator 11, a modulator 12, a high frequency amplifier 13, and a transmitter. The high-frequency pulse output from the high-frequency oscillator 11 is amplitude-modulated by the modulator 12 according to the command from the sequencer 7, and the amplitude-modulated high-frequency pulse is sent to the high-frequency amplifier 1.
The subject 1 is irradiated with the electromagnetic waves by supplying the electromagnetic waves to the high-frequency coil 14a placed close to the subject 1 after being amplified in step 3.

The receiving system 5 detects a high frequency signal (NMR signal) emitted by nuclear magnetic resonance of the atomic nucleus of the living tissue of the subject 1, and includes a high frequency coil 14b, an amplifier 15, a quadrature phase detector 16, and a receiving side high frequency signal (NMR signal). /D conversion@17,
The high frequency signal (NMR signal) of the response of the subject 1 due to the electromagnetic wave irradiated from the transmitting side high frequency coil 14a is detected by the high frequency coil 14b placed close to the subject 1, and is detected by the amplifier 15 and the quadrature phase detector. A/ through 16
The signal is input to the D converter 17 and converted into a digital quantity.

Furthermore, the collected data of two bases is sampled by the quadrature phase detector 16 at the timing according to the command from the sequencer 7, and the signal is sent to the signal processing system 6.

This signal processing system 6 consists of a CPU 8, a recording device such as a magnetic disk 18 and a magnetic tape 19, and a display 20 such as a CRT, and the CPU 8 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 20. In addition, the sequencer 7
Means that operates under the control of U8 and sends various commands necessary for data collection of tomographic images of the subject 1 to the transmission system 4, magnetic field gradient generation system 3, and reception system 5, and generates a sequence for measuring the NMR signal. This is the result. In addition, in FIG. 3, the high-frequency coil 14a on the transmitting side, the high-frequency coil 14b on the receiving side, and the gradient magnetic field coils 9, 9 are connected to the subject 1.
It is arranged in the magnetic field space of the static magnetic field generating magnet 2 arranged in the space around.

Here, in the present invention, for example, the high-frequency coil 14b on the receiving side is formed as a set of two conductive loops with their sensitivity directions orthogonal to each other, and the high-frequency coil 14b is formed as a set with two conductive loops whose sensitivity directions are orthogonal to each other. The receiving direction for detecting the high frequency signal is arranged perpendicular to the static magnetic field generated by the static magnetic field generating magnet 2, and the conductor width at the intersection of the two conductive loops is narrowed to reduce the opposing area.

That is, for example, in the case of a vertical magnetic field type QD coil, as shown in FIG. As a conductive loop, the saddle-shaped coil 23 directs its reception direction to the solenoid coil 22.
Each coil 22, 23 is arranged orthogonal to the receiving direction of
is divided by a capacitor 24 to lower the operating voltage. The purpose of the saddle-shaped coil 23 is to improve the reception sensitivity of a portion corresponding to the top of the head of the subject 1 inserted into the high-frequency coil 14b.

Coil member 2 located near one end of resin bobbin 21
3a and 23b are deformed and protrude outward in the longitudinal direction of the resin bobbin 21. Further, the intersection portion 25 between the solenoid coil 22 and the saddle-shaped coil 23 is designed to reduce the capacitive coupling between the two coils, for example, by approximately 6.
■There is an interval of approximately However, even with such a spacing, the capacitive coupling between the two coils cannot be reduced to a practical level. Therefore, as shown in FIG. 1, the width of the conductor at the intersection 25 of the solenoid coil 22 and the saddle-shaped coil 23 is narrowed to reduce the opposing area. In this case, the area S of the planar conductor plates A□ and A2 in FIG.
becomes smaller. Therefore, the stray capacitance formed between the two coils 22 and 23 is reduced, and coupling due to capacitive coupling between the two can be reduced.

In addition, according to what the present inventors confirmed experimentally,
The opposing area of the intersection 25 of the above-mentioned coils 22 and 23 is:
For example, it is desirable to set it as 100-400 m2. If it is smaller than this value, the sensitivity will be lowered (decreased S/N ratio), and if it is larger, capacitive coupling will increase, making it impractical. However, the above values vary depending on the resonance frequency of the magnetic resonance imaging apparatus used, and are not limited to the above values. In addition, in the example of FIG. 1, the solenoid coil 22 and the saddle-shaped coil 23
Although the conductor width is shown as narrowing continuously from one intersection 25 to the next intersection 25, the present invention is not limited to this, and the above-mentioned coils 22 and 23 actually intersect. Only a portion may be narrowed. Furthermore, the first
In the figure, the conductor width of both the solenoid coil 22 and the saddle-shaped coil 23 is shown to be narrow at the intersection portion 25, but the conductor width of only one of the coils may be narrowed. In this case, the conductor width of one of the coils to be narrowed needs to be narrower than the conductor widths of both coils.

Furthermore, in FIG. 1, the solenoid coil 22
Although both coils 22 and 23 are divided by the capacitor 24, the coils 22 and 23 do not necessarily need to be divided by the capacitor 24. In addition, in FIG. 1, the saddle-shaped coil 23 to be combined with the solenoid coil 22 is a coil member 23a located on the negative end thereof.
, 23b are deformed by protruding outward, but the present invention is not limited to this, and a normal shape may be used.

FIG. 2 is a circuit diagram showing the principle and connections of the high frequency coil 14b constructed as described above. In the figure, the coil tuning circuit and the like are omitted for simplification of explanation. In the figure, the direction of the static magnetic field is indicated by arrow S, and the magnetization vector rotating in one plane is 90 degrees
induces the same signal with a phase difference of degrees. Here, since the solenoid coil 22 and the saddle-shaped coil 23 are arranged with their axial directions perpendicular to each other, a high frequency signal (NMR signal) is detected accompanied by mutually independent random noise. Possible sources of this noise are the resistance of each coil 22.23 and the equivalent resistance from the subject 1 due to the magnetic coupling and electrical coupling of these coils 22.23.

The signals from the solenoid coil 22 and the saddle-shaped coil 23 are amplified by the first amplifier 15a or the second amplifier 15b in the amplifier 15, respectively, and then input to the shifter 26. This shifter 26 is a phase shifter 27
, an attenuator 28 and an adder 29.

Then, the phase of the signal from the solenoid coil 22 is shifted by 90 degrees by the phase shifter 27 to match the phase with the signal from the saddle-shaped coil 23. On the other hand, the saddle-shaped coil 23 and the solenoid coil 22 have unequal sensitivities; for example, when the sensitivity of the former is 1", that of the latter is 1", 4".
It becomes. Therefore, in this case, a high S/N ratio cannot be obtained unless the signal addition ratio in the adder 29 is changed. The optimal addition ratio at this time is 12÷1.4
2=0.51. Therefore, an attenuator 28 is inserted in the signal path of the saddle coil 23 so that when the signal from the solenoid coil 22 is set to 1'', the signal from the saddle coil 23 becomes '0.51''. is being adjusted. In this way, after the signal intensities from both coils 22 and 23 are combined, the adder 29 adds the two signals, and the signal is output from the shifter 26. The output signal from the shifter 26 is then sent to the quadrature phase detector 16 shown in FIG.

In this way, when the phases of the signals from both coils 22 and 23 are matched by the phase shifter 27 and added by the adder 29, the noise becomes a little large, but the detected signal becomes considerably large, and as a result, the S/N ratio becomes large. Become. For example, if the dimensions and shapes of one coil 22 and the other coil 23 are the same, and the equivalent resistance from the object 1 described above is also the same, the detection signal will be doubled and the noise will be /7 times, resulting in As a result, the S/N ratio is improved by a factor of /7.

In the above description, the solenoid coil 22 and the saddle-shaped coil 23 are combined as a vertical magnetic field QD coil, but the present invention is not limited to this, and the present invention is not limited to this. coil 23
and other saddle-shaped coils 23, or combinations of various other types of coils, by narrowing the conductor width at the intersection of the two coils and reducing their facing area, the Capacitive coupling between the two coils can be reduced. The connection of the high-frequency coil 14b when the saddle-shaped coils 23, 23 having the same shape as described above are combined is as shown in FIG. At this time, since the two coils constituting the high-frequency coil 14b are a combination of saddle-shaped coils 23 and 23 with equal sensitivity, the addition ratio of the signals in the adder 29 is 1.
:1 may be used. Therefore, inside the shifter 26',
There is no need to insert an attenuator 28 to change the addition ratio of the signals from the two coils as shown in FIG.

Furthermore, in the above explanation, an example was described in which the present invention was applied to the high-frequency coil 14b on the receiving side in FIG.
The present invention is not limited to this, and may also be applied to the high-frequency coil 14a on the transmitting side.

〔Effect of the invention〕

Since the present invention is configured as described above, the high frequency coil 1
Two conductive loops (22,
23) By narrowing the conductor width of the crossing portion 25 and reducing the opposing area thereof, the stray capacitance formed at the crossing portion 25 can be reduced. Thereby, the capacitive coupling can be relaxed and the coupling between the two conductive loops (22, 23) can be reduced without increasing the interval between the intersections 25 of the two conductive loops (22, 23). Therefore, the overall sensitivity of the high-frequency coils 14a, 14b is improved, and it is possible to prevent the S/N ratio of the obtained image from decreasing. Furthermore, since the interval between the intersections 25 of the two conductive loops (22, 23) does not need to be increased or can be made smaller than in the past, the high-frequency coils 14a, 1
The whole of 4b can be made into lJ1 type.

[Brief explanation of the drawing]

FIG. 1 is a perspective explanatory diagram showing an embodiment of a high-frequency coil of a magnetic resonance imaging apparatus according to the present invention, FIG. 2 is a circuit diagram showing the principle and connection of the high-frequency coil, and FIG. 3 is a diagram to which the above-mentioned high-frequency coil is applied. FIG. 4 is a block diagram showing the overall configuration of the magnetic resonance imaging apparatus. FIG. 4 is a circuit diagram showing the connection of high-frequency coils according to another embodiment. FIG. FIG. 3 is an explanatory diagram for explaining capacity. 1... Subject, 2... Static magnetic field generating magnet, 3...
・Magnetic field gradient generation system, 4... Transmission system, 5... Receiving system. 6...Signal processing system, 7...Sequencer, 8...
・CPU, 14a... High frequency coil on the transmitting side,
14b... Receiving side high frequency coil, 15...
Amplifier, 22... Solenoid coil, 23...
・Saddle-shaped coil, 24... Capacitor, 25... Crossing part, 26... Shifter Figure 4

Claims (3)

[Claims] (1) A magnetic field generation means for applying a static magnetic field and a gradient magnetic field to the subject; a transmission system for irradiating high-frequency signals to cause nuclear magnetic resonance to the nuclei of atoms constituting the living tissue of the subject; The above-mentioned transmission system of a magnetic resonance imaging apparatus comprising a receiving system that detects a high-frequency signal emitted by nuclear magnetic resonance, and a signal processing system that performs image reconstruction calculation using the high-frequency signal detected by the receiving system; Sensitivity for detecting a high-frequency signal emitted by irradiating a high-frequency signal onto the subject or by nuclear magnetic resonance, which is provided in the receiving system and is formed into a set of two conductive loops with their sensitivity directions orthogonal to each other. A high-frequency coil for a magnetic resonance imaging apparatus, characterized in that the high-frequency coil is disposed with a direction perpendicular to a static magnetic field, and the conductor width at the intersection of the two conductive loops is narrowed to reduce the facing area thereof. (2) The high-frequency coil for a magnetic resonance imaging apparatus according to claim 1, wherein the conductor width of one of the intersections of the two conductor loops is formed narrow. (3) The high-frequency coil for a magnetic resonance imaging apparatus according to claim 1 or 2, wherein a solenoid coil and a saddle-shaped coil are used as the two conductive loops.