JPH07171132A - High frequency coil for magnetic resonance imaging device - Google Patents

Abstract

PURPOSE:To moderate the influence, in a high frequency coil for magnetic resonance imaging device, when a testee is inserted into the coil, to prevent the deterioration of S/N, and reduce the slippage of resonance frequency by the change in size of the subject. CONSTITUTION:Coil conductors placed in a transmitting system and a receiving system, respectively, to form loops larger than a testee are divided at least in one position on the circumferences to form circular first coil conductors 32a, 32b. One or two circular second coil conductors 38a, 38b close to the first coil conductors in non-contact are formed on the inside of the first coil conductors, and the first coil conductors 32a, 32b and the second coil conductors 38a, 38b are connected in high frequency state by a capacitance formed between the both to constitute one conductive loop 31a. Thus, the influence when the testee is inserted into the coil is moderated to prevent the deterioration of S/N, and the slippage of resonance frequency by the change in size of the testee can be reduced.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention has a magnet for generating a static magnetic field in a direction perpendicular to the axial direction of a subject and utilizes the phenomenon of nuclear magnetic resonance (hereinafter abbreviated as "NMR") The present invention relates to a high-frequency coil used in a transmission system and a reception system of a nuclear magnetic resonance imaging apparatus for obtaining a tomographic image of a site, and in particular, S / N by relaxing the influence when a subject is inserted in the coil.
The present invention relates to a high-frequency coil of a magnetic resonance imaging apparatus capable of preventing the deterioration of the magnetic field and reducing the deviation of the resonance frequency due to the change in the size of the subject.

[0002]

2. Description of the Related Art A magnetic resonance imaging apparatus includes a magnetic field generating means for applying a static magnetic field and a gradient magnetic field in a direction perpendicular to a body axis direction of a subject, and a nuclear magnetic resonance for atomic nuclei of atoms constituting the biological tissue of the subject. To irradiate a high-frequency signal in order to wake up, a receiving system to detect the high-frequency signal emitted by the above-mentioned nuclear magnetic resonance, and a signal to perform image reconstruction operation using the high-frequency signal detected by this receiving system And a processing system. Then, while applying a uniform static magnetic field to the subject by the static magnetic field generating means, a high-frequency signal of a frequency that excites nuclear magnetic resonance is applied by a high-frequency coil of the transmission system, whereby the nuclear magnetic resonance signal emitted from the subject. Is detected by the high frequency coil of the receiving system.
At this time, in order to specify the emission position of the nuclear magnetic resonance signal from the subject, imaging is performed by further applying a gradient magnetic field by the gradient magnetic field generating means.

The conventional high-frequency coil in such a magnetic resonance imaging apparatus is constructed as shown in FIGS. 4 and 5. This high-frequency coil is formed in a solenoid type, and such a solenoid coil has a number of loops (number of turns), a connection method of each loop (series or parallel), and a coil conductor of one conductive loop divided by a capacitor. Although there are various types depending on the number of contacts, FIG. 5 shows a type in which two turns are connected in series and the coil conductor of one conductive loop is divided into two by a capacitor.

One of the conductive loops 21a has a capacitor 2 for dividing the coil conductors 22a and 22b divided into two.
3a is connected to form one turn. The other conductive loop 21b has a coil conductor 2 divided into two.
Connect 4a and 24b with a dividing capacitor 23b to
It is composed of turns. Then, each coil conductor 2 described above
The base end portions of 2a, 22b, 24a, and 24b are connected to the tuning substrate 25, and at this portion, the coil conductor 22b of the one conductive loop 21a and the coil conductor 24a of the other conductive loop 21b are divided. 22a → 23a → 22 by connecting with the capacitor 23c
A 2-turn solenoid coil is formed by a connection path of b → 23c → 24a → 23b → 24b. Further, reference characters A and B indicate feeding points, and the feeding point A is the coil conductor 2.
2a and the other feeding point B is connected to the coil conductor 24b. Further, reference numerals 26a and 26b denote tuning capacitors provided at the ends of the coil conductors 22a or 24b, and reference numeral 27 denotes the feeding point A,
The varicap diode provided between B is shown.

FIG. 4 shows a circuit diagram of a high-frequency coil configured as shown in FIG. Since the high frequency coil utilizes LC parallel resonance, it is necessary to tune the resonance frequency of the coil to the magnetic resonance frequency. That is, in FIG. 4, each coil conductor 22a, 2
The resonance frequency of the high frequency coil is determined by the inductance L of 2b, 24a and 24b and the capacitance C of each of capacitors 23a, 23b, 23c, 26a and 26b. Also, the varicap diode 27 is
This is for finely adjusting the change of the resonance frequency due to the change of the size of the subject, and by changing the capacitance of the varicap diode 27 by changing the capacitance of the varicap diode 27 by supplying a direct current to the feeding points A and B. There is. At this time, the DC value was determined by measuring the value at which the magnetic resonance signal becomes maximum for each subject.

[0006]

However, in such a conventional high frequency coil, as shown in FIG. 5, the coil conductors 22a and 22b are inserted into the conductive loops 21a and 21b with respect to the subject. , 24a, 24b were directly adjacent to each other, so that they were easily affected when the subject was inserted into the coil. That is, it is generally known that the Q value of the coil is increased in order to increase the S / N of the obtained tomographic image. However, when f ₀ is the resonance frequency of the coil, the Q value is Q = 2πf ₀ It is expressed as L / R. When the subject is inserted into the coil in such a state, a capacitance is formed between the subject and the coil conductor at high frequencies, and the capacitance C is added as shown by the broken line in FIG. further,
Since the subject is a kind of conductor, the resistance R is added via the capacitance C as shown by the broken line. Therefore, as is clear from the above equation, the addition of this resistance R component causes the Q value to decrease and the S / N to decrease. As a countermeasure against this, conventionally, as shown in FIG. 5, the conductive loops 21a and 21b are divided in the middle and connected by capacitors 23a and 23b to lower the operating voltage of the coil, and the coil conductors 22a, 22b and 2 are connected.
Although the capacitive coupling between 4a and 24b and the analyte was alleviated, that was not enough.

Further, as described above, the capacitance C is formed between the subject and the coil conductor.
As shown in FIG. 3, the resonance frequency of the high-frequency coil using LC parallel resonance may be displaced, and further the deviation of the resonance frequency may be changed due to the change in the size of the subject. For example, the varicap diode 27 causes a predetermined difference for each subject. It was necessary to tune so that the resonance frequency of Therefore, this adjustment may take time, and the imaging time and the subject restraint time may increase.

Therefore, the present invention addresses such a problem, mitigates the influence when the subject is inserted into the coil to prevent S / N deterioration, and reduces the size of the subject. It is an object of the present invention to provide a high frequency coil of a magnetic resonance imaging apparatus capable of reducing the shift of the resonance frequency due to the change of

[0009]

In order to achieve the above object, a high frequency coil of a magnetic resonance imaging apparatus according to the present invention comprises a magnetic field generating means for applying a static magnetic field and a gradient magnetic field to a subject, and a living tissue of the subject. A transmission system that irradiates a high-frequency signal to cause nuclear magnetic resonance in the nuclei of the atoms that make up the, a reception system that detects the high-frequency signal emitted by the above-mentioned nuclear magnetic resonance, and a high-frequency signal detected by this reception system A high-frequency signal having coil conductors that are respectively formed in the transmission system and the reception system of the nuclear magnetic resonance imaging apparatus including a signal processing system that performs image reconstruction calculation using In the high-frequency coil for irradiating and detecting the above, the coil conductor is divided at at least one position on its circumference to form an arc-shaped first coil conductor, One or two arc-shaped second coil conductors that are close to each other in a non-contact manner with the first coil conductor are provided inside the one coil conductor, and the first coil conductor and the second coil conductor are both The capacitors formed between the two are connected at high frequencies to form one conductive loop.

Further, it is preferable to connect a portion divided at least one place on the circumference of the first coil conductor with a capacitor.

Further, it is effective to connect to the first coil conductor a tuning circuit for matching the resonance frequency of the conductive loop formed by the first coil conductor with the magnetic resonance frequency.

Further, a spacer made of a material having a small high frequency loss may be interposed between the first coil conductor and the second coil conductor to keep the distance between the two coil conductors. .

[0013]

The high-frequency coil of the magnetic resonance imaging apparatus configured as described above is connected in a high-frequency manner with the capacitance formed by the arc-shaped first coil conductor and the arc-shaped second coil conductor therebetween. By configuring one conductive loop, the operating voltage of the second coil conductor that is directly adjacent to the subject inside the first coil conductor is lowered, and the second coil conductor and the subject are inserted by inserting the subject. Operates to relax capacitive coupling between. This allows
It is possible to prevent the S / N from deteriorating when the subject is inserted into the coil and reduce the deviation of the resonance frequency due to the change in the size of the subject.

[0014]

Embodiments of the present invention will now be described in detail with reference to the accompanying drawings. 1 is a perspective explanatory view showing an embodiment of a high frequency coil of a magnetic resonance imaging apparatus according to the present invention, and FIG. 2 is a block diagram showing an overall configuration of a magnetic resonance imaging apparatus to which the above high frequency coil is applied.

The magnetic resonance imaging apparatus obtains a tomographic image of a subject by utilizing a nuclear magnetic resonance (NMR) phenomenon. As shown in FIG. 2, a static magnetic field generating magnet 2 and a magnetic field gradient generating system 3 are provided. , Transmission system 4, reception system 5, and signal processing system 6
A sequencer 7 and a central processing unit (CPU) 8.

The static magnetic field generating magnet 2 is for generating a uniform static magnetic field around the subject 1 in a direction orthogonal to the body axis direction thereof. Permanent magnet type, normal conducting type or superconducting type magnetic field generating means is arranged. Magnetic field gradient generation system 3
Is a gradient magnetic field coil 9 wound in three axial directions of X, Y and Z
And a gradient magnetic field power supply 10 for driving each coil, and by driving the gradient magnetic field power supply 10 for each coil in accordance with an instruction from the sequencer 7,
Gradient magnetic fields Gx, Gy, Gz in the triaxial directions of X, Y, Z are applied to the subject 1. The slice plane for the subject 1 can be set by the method of applying the gradient magnetic field.

The transmission system 4 irradiates a high-frequency signal (electromagnetic wave) to cause nuclear magnetic resonance in atomic nuclei of the atoms forming the biological tissue of the subject 1, and includes a high-frequency oscillator 11, a modulator 12 and a high-frequency amplifier. 13 and a high-frequency coil 14a on the transmission side, and the high-frequency pulse output from the high-frequency oscillator 11 is modulated by the modulator 1 in accordance with a command from the sequencer 7.
2, the amplitude-modulated high-frequency pulse is amplified by the high-frequency amplifier 13 and then supplied to the high-frequency coil 14a arranged close to the subject 1, whereby the subject 1 is irradiated with electromagnetic waves. It is like this.

The receiving system 5 detects a high-frequency signal (NMR signal) emitted by nuclear magnetic resonance of atomic nuclei of the living tissue of the subject 1, and includes a high-frequency coil 14b on the receiving side, an amplifier 15 and a quadrature phase detector. 16 and an A / D converter 17, and the high-frequency signal (N of the response of the subject 1 due to the electromagnetic wave emitted from the high-frequency coil 14a on the transmission side) (N
The MR signal) is detected by the high-frequency coil 14b on the receiving side arranged in the vicinity of the subject 1, is input to the A / D converter 17 via the amplifier 15 and the quadrature phase detector 16, and is converted into a digital amount. Further, the quadrature detector 16 samples the two-series collected data at the timing according to the instruction from the sequencer 7, and the signal is sent to the signal processing system 6.

The signal processing system 6 includes a CPU 8, a recording device such as a magnetic disk 18 and a magnetic tape 19, and a CRT.
And the like, and the CPU 8 performs processing such as Fourier transform, correction coefficient calculation, and image reconstruction,
The signal intensity distribution of an arbitrary section or a distribution obtained by performing an appropriate calculation on a plurality of signals is imaged and displayed on the display 20.
It is designed to be displayed as a tomographic image.

The sequencer 7 operates under the control of the CPU 8 and sends various commands necessary for data acquisition of a tomographic image of the subject 1 to the transmission system 4, the magnetic field gradient generation system 3 and the reception system 5, and the above-mentioned NMR. It is a means for generating a sequence for measuring a signal. In FIG. 2, the high frequency coil 14a of the transmission system, the high frequency coil 14b of the reception system, and the gradient magnetic field coils 9, 9 are arranged in the magnetic field space of the static magnetic field generating magnet 2 arranged in the space around the subject 1. Has been done.

Here, in the present invention, the transmission system 4 is used.
And a high-frequency coil 14 provided in the receiving system 5, respectively.
a and 14b are configured as shown in FIG. Figure 1
The embodiment shown in (1) shows an example of a one-turn, two-parallel solenoid coil, and has one conductive loop 31a forming a loop larger than the subject and the other conductive loop 31b. One conductive loop 31a
Is a first coil conductor 32, which is obtained by dividing one coil conductor made of, for example, an arc-shaped copper plate into two on its circumference.
a and 32b, and these coil conductors 32a and 32b are connected by a dividing capacitor 33a to form one turn. Further, the other conductive loop 31b is obtained by dividing one coil conductor, which is also made of an arcuate copper plate, into two coil conductors on the circumference thereof as first coil conductors 34a and 34b, and these coil conductors 34a and 34b are connected to each other. One turn is formed by connecting with a dividing capacitor 33b.

The first coil conductors 32a, 34a, 3
The base parts of 2b and 34b are tuning boards 35, respectively.
Are connected in parallel to each other, and as a whole, this constitutes a one-turn, two-parallel solenoid coil.
Then, the tuning board 35 is provided with feeding points A and B.
Are connected, and the feeding point A is connected to the coil conductors 32a and 34a.
a, and the other feed point B is connected to the coil conductors 32b and 34b.
connected to b. Further, reference numerals 36a, 36b, 3
6c and 36d are the coil conductors 32a, 32b and 3 described above.
A tuning capacitor provided at the ends of 4a and 34b is shown, and a reference numeral 37 is a varicap diode provided between the feeding points A and B. The tuning board 35, the tuning capacitors 36a to 36d, and the varicap diode 37 constitute a tuning circuit for matching the resonance frequency of the conductive loop with the magnetic resonance frequency. Note that, in FIG. 1, a bobbin for holding each coil conductor in a loop shape is omitted for the sake of clarity.

In such a state, two second loops are provided inside the first coil conductors 32a and 32b of the one conductive loop 31a and the first coil conductors 34a and 34b of the other conductive loop 31b. Coil conductors 38a, 38
b, 39a, 39b are provided. Hereinafter, the one conductive loop 31a side will be described as a representative. The second coil conductors 38a, 38b are for reducing the operating voltage of the coil conductors located on the side of the subject 1, and are made of, for example, a semi-arcuate copper plate, and include the dividing capacitor 33a and the tuning. It is provided so as to cover at least both ends of the first coil conductors 32a and 32b divided at the position of the substrate 35. The second coil conductors 38a and 38b are the first coil conductor 32.
The first coil conductors 32a and 32b and the second coil conductor 38 are provided inside the a and 32b in close proximity to each other without contact.
a and 38b are connected in high frequency by the capacitance formed between the two, forming one conductive loop 31a. That is, the second coil conductors 38a and 38b also operate as coils.

Further, as shown in FIG. 3, the first coil conductors 32a and 32b and the second coil conductors 38a and 38b.
A spacer 40 is interposed between the spacer 40 and b. The spacer 40 holds the space between the first coil conductors 32a and 32b and the second coil conductors 38a and 38b, and is made of a material such as Teflon having a small high frequency loss, and the first coil conductor 32a. 32b, which is equal to or slightly wider than the width of 32b. The first coil conductors 32a and 32b, the spacer 40, and the second coil conductors 38a and 38b are adhered to each other with an adhesive agent that causes less high frequency loss. Note that FIG.
In the figure, reference numeral 41 indicates a bobbin.

In the high-frequency coil constructed as described above, for example, the first coil conductors 32a and 32b and the second coil conductors 38a and 38 in the one conductive loop 31a.
If the capacitance formed with b is Cx,
In addition to the capacitance of the capacitor of the conventional LC parallel resonance circuit as shown in FIG. 4, the above capacitance Cx is added to form the resonance circuit. In this case, the capacitance Cx is the capacitance of the capacitor 3 shown in FIG.
Since the capacitance is smaller than that of 3a, 36a, 36b, the operating voltage of the second coil conductors 38a, 38b is lower than the operating voltage of the first coil conductors 32a, 32b, and the subject 1 is inserted into the coil. The effect of doing this can be mitigated. The capacitors 36a and 36b have the same capacitance, and the dividing capacitor 33
It is desirable that the capacity is approximately twice as large as a. If such a capacitance ratio is selected, the upper and lower center lines of the conductive loop 31a will be near the ground potential. Further, the presence of the second coil conductors 38a and 38b reduces the influence of the insertion of the subject, and the change in the resonance frequency of the coil due to the change in the size of the subject 1 to be inserted is reduced. The varicap diode 37 may be omitted. The configuration and operation described above are the same as those of the other conductive loop 31b shown in FIG.
Is exactly the same.

Although FIG. 1 shows an example in which the solenoid coil has one turn and two parallels, the present invention is not limited to this, and a solenoid coil having only one turn may be used. It may be configured as a turn-series solenoid coil. Further, the example in which two second coil conductors are provided inside the first coil conductor is shown, but only one second coil conductor may be provided with an appropriate size. In this case, it is desirable to cover as wide an area as possible with the second coil conductor inside the first coil conductor.

[0027]

Since the present invention is constructed as described above,
The arc-shaped first coil conductor and the arc-shaped second coil conductor are connected in high frequency by the capacitance formed between them to form one conductive loop, so that the inside of the first coil conductor is As a result, the operating voltage of the second coil conductor directly adjacent to the subject is lowered, and the capacitive coupling between the second coil conductor and the subject due to insertion of the subject can be alleviated. As a result, it is possible to prevent the S / N from deteriorating when the subject is inserted into the coil and reduce the deviation of the resonance frequency due to the change in the size of the subject. Therefore, tuning for each subject is not necessary, and the imaging time and the subject restraint time can be shortened.

Further, in the case where a spacer made of a material having a small high frequency loss is interposed between the first coil conductor and the second coil conductor to keep the distance between the two coil conductors, As a result, it is possible to prevent a change in the interval between the coil conductors with time, and a change in the resonance frequency of the entire coil by preventing a change in the capacitance formed at that interval.

[Brief description of drawings]

FIG. 1 is a perspective explanatory view showing an embodiment of a high frequency coil of a magnetic resonance imaging apparatus according to the present invention.

FIG. 2 is a block diagram showing an overall configuration of a magnetic resonance imaging apparatus to which the high frequency coil is applied.

FIG. 3 is an enlarged explanatory view of a main part showing the structure of the high-frequency coil.

FIG. 4 is a circuit diagram showing a conventional high-frequency coil of this type.

FIG. 5 is a perspective explanatory view showing the conventional high frequency coil.

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 ... Subject 2 ... Static magnetic field generation magnet 3 ... Magnetic field gradient generation system 4 ... Transmission system 5 ... Reception system 6 ... Signal processing system 7 ... Sequencer 8 ... CPU 14a ... Transmission side high frequency coil 14b ... Reception side high frequency coil 31a , 31b ... Conductive loop 32a, 32b, 34a, 34b ... First coil conductor 38a, 38b, 39a, 39b ... Second coil conductor 33a, 33b, 36a to 36d ... Capacitor 37 ... Varicap diode 40 ... Spacer

Claims (4)

[Claims] 1. A magnetic field generating means for applying a static magnetic field and a gradient magnetic field to a subject, and a transmission system for irradiating a high frequency signal in order to cause nuclear magnetic resonance in atomic nuclei of atoms constituting the biological tissue of the subject, A nuclear magnetic resonance imaging apparatus comprising a receiving system for detecting a high frequency signal emitted by the nuclear magnetic resonance, and a signal processing system for performing an image reconstruction operation using the high frequency signal detected by the receiving system. In a high-frequency coil for irradiating and detecting a high-frequency signal, which has coil conductors that form a loop larger than the subject and is provided in each of the transmission system and the reception system, divide the coil conductors in at least one place on its circumference This is an arc-shaped first coil conductor, and one or two arc-shaped second coils which are in contact with the first coil conductor inside the first coil conductor in a non-contact manner. A magnetic resonance imaging apparatus characterized in that a conductor is provided, and the first coil conductor and the second coil conductor are connected in high frequency by a capacitance formed between them to form one conductive loop. High frequency coil. 2. The high frequency coil for use in a magnetic resonance imaging apparatus according to claim 1, wherein at least one portion on the circumference of the first coil conductor is connected with a capacitor. 3. A tuning circuit for connecting the resonance frequency of the conductive loop formed by the first coil conductor to the magnetic resonance frequency is connected to the first coil conductor. A high frequency coil of the magnetic resonance imaging apparatus described. 4. A spacer formed of a material having a small high frequency loss is interposed between the first coil conductor and the second coil conductor to maintain a distance between the two coil conductors. The high frequency coil of the magnetic resonance imaging apparatus according to claim 1, 2.