JP6887346B2 - High frequency coil unit and magnetic resonance imaging device - Google Patents

Description

The present invention relates to a high-frequency coil unit and a magnetic resonance imaging device, and more particularly to a technique for irradiating an electromagnetic wave and detecting a magnetic resonance signal in the high-frequency coil unit.

In a magnetic resonance imaging device (hereinafter referred to as "MRI device"), a high-frequency signal (hereinafter referred to as "RF signal") which is an electromagnetic wave is transmitted to a subject placed in a uniform static magnetic field generated by a static magnetic field magnet. ) Is irradiated to excite the nuclear spins in the subject, and an NMR signal, which is an electromagnetic wave generated by the nuclear spins, is received and signal-processed to acquire a magnetic resonance image of the subject.

In this way, the MRI apparatus irradiates the subject with an RF signal. Here, the irradiation of the RF signal and the reception of the NMR signal are performed by an antenna device such as an RF antenna or an RF coil that transmits or receives an electromagnetic wave having a radio frequency. As such an antenna device, a high frequency coil unit called a birdcage type (birdcage type) coil is known (for example, Patent Document 1 and Patent Document 2).

Normally, as shown in FIG. 13, a bird cage type coil has two cylindrical ring conductors 203, a plurality of linear rung (crossbar) conductors 204, a capacitor, a diode, a power feeding cable (not shown), and the like. Then, the ends of the rung conductors arranged at equal intervals on the cylindrical surface and the ring conductor are connected, each capacitor is inserted into the gap 201 provided at equal intervals in the ring conductor 203, and the diode is inserted into the rung conductor 204. It is configured by being inserted into the gap.

The cylindrical bird cage coil 200 of FIG. 13 is surrounded by a cylindrical conductor called an RF shield (not shown). The capacitor provided in the bird cage coil 200 is adjusted by the RF shield, the rung conductor, and the ring conductor so as to resonate at a specific frequency in the MRI apparatus. Normally, the RF shield is further installed on the inner cylindrical surface of the cylindrical gradient magnetic field coil existing outside.

The bird cage type coil is characterized in that the spread of the uniform space of the RF magnetic field created by the irradiated RF signal is higher than that of a simple loop coil or saddle type coil. Due to this feature, the bird cage coil is now the standard transmission coil for tunnel horizontal magnetic field MRI equipment.

In the bird cage type coil, the degree of spread of the uniform space in the cross-sectional direction of the cylinder depends on the number of rungs arranged at equal intervals on the cylinder surface if the cylinder has the same diameter. In other words, in the case of a bird cage type coil used for a 1.5 Tesla transmission coil, the number of rung conductors is usually 16 or 24, and the larger the number, the wider the uniform space in the cross-sectional direction of the cylinder. large.

Normally, the uniformity of the RF magnetic field is high at the center of the cylinder and decreases as it approaches the periphery. In the bird cage type coil, when the number of rung conductors is small, the difference in magnetic field between the space closer to the rung conductor and the space away from the rung conductor between the rung conductor and the rung conductor is large in the vicinity of the rung conductor. Therefore, the uniformity of the magnetic field is reduced.

U.S. Pat. No. 7,680,070 International Publication 2012/059574 Gazette

By the way, in general, when a person enters a cylindrical irradiation coil while lying down, the shoulders, hips, abdomen, arms, etc. of the person spread laterally, so that the shoulders, hips, arms, etc. of a person with a large physique In order to take a clear image of the abdomen with an MRI apparatus, it is desirable that all of the static magnetic field, the gradient magnetic field, and the RF magnetic field can be uniformly generated in a spherical space having a diameter of about 50 cm from the center of the magnet.

On the other hand, for example, a superconducting magnet that generates a strong static magnetic field of 1.5 tesla has a cylindrical shape called a tunnel type, and it is preferable to increase the imaging space, which is the internal space of the tunnel type. It is costly to increase the imaging space by increasing the size. Therefore, when considering creating a wide RF magnetic field space without changing the size of the magnet, it is desirable to generate a uniform RF magnetic field as wide as possible inside without changing the size of the bird cage type coil. For this purpose, it is effective to increase the number of rung conductors as described above, but increasing the number of rung conductors complicates the shape of the coil and increases the number of circuit elements such as capacitors. Therefore, the cost increases.

The present invention has been made in view of the above circumstances, and the imaging space is expanded by enlarging the region where the RF magnetic field is uniformly irradiated without increasing the number of parts and the equipment cost, and a wide range and high image quality are obtained. The purpose is to acquire an MRI image of.

One aspect of the present invention is applied to an MRI apparatus and has a tubular shape that transmits at least one of a high-frequency signal to a static magnetic field space and a nuclear magnetic resonance signal generated from a subject placed in a static magnetic field. The RF coil is provided with a shield made of a sheet-shaped conductor and formed in a tubular shape so as to surround the RF coil, and the RF coil is formed in an annular shape, and the tubular shape is formed at both ends of the tubular shape. Two ring conductors arranged so that the central axis of the ring and the center of the annular shape coincide with each other, and the two ring conductors are arranged along the outer peripheral surface of the cylindrical shape at a predetermined distance from each other in parallel with the central axis of the tubular shape. The ring conductor has a plurality of rungs that electrically connect the two ring conductors and a plurality of capacitors that are inserted between the adjacent rung conductors in the ring conductor, and the rung conductors are arranged in the circumferential direction. Provided is a high frequency coil unit characterized by having N (N is an integer of 2 or more) linear portions arranged substantially parallel to each other.

According to the present invention, it is possible to expand the imaging space by enlarging the region uniformly irradiated with the RF magnetic field without increasing the number of parts and the cost of the apparatus, and to acquire a wide range and high-quality MRI image. ..

It is a schematic block diagram of the MRI apparatus to which the RF coil which concerns on embodiment of this invention is applied. It is a perspective view of the RF coil which concerns on embodiment of this invention. The map of the RF magnetic field generated by the conventional RF coil is shown, and (A) is a map showing the magnetic field generated when about 12 watts are input to each of the two input terminals, for a total of about 24 watts, in gray scale. Yes, (B) is a diagram showing the map shown in (A) with 30 contour lines at equal intervals of 0 to 2 [A / m]. The map of the RF magnetic field generated by the RF coil according to the embodiment of the present invention is shown, and (A) grayscales the magnetic field generated when about 12 watts are input to each of the two input terminals, for a total of about 24 watts. (B) is a diagram showing the map shown in (A) with the intensity of 0 to 2 [A / m] indicated by 30 contour lines. It is the figure which plotted the magnetic field strength on the vertical axis, (A) is the plot of the calculation point of the Z = 0 cross section of the conventional RF coil, (B) is the Z = of the RF coil which concerns on this embodiment. It is a plot of the calculation points of the 0 cross section. It is a figure which shows the diode circuit provided in the rung conductor, (A) is the figure which shows the diode circuit applied to the conventional RF coil, and (B) is the figure which shows the diode circuit applied to the RF coil which concerns on this Embodiment. It is a figure which shows. It is a perspective view of the RF coil which concerns on the modification 1 of this embodiment. It is a perspective view of the RF coil which concerns on the modification 2 of this embodiment. It is a perspective view of the RF coil which concerns on the modification 3 of this embodiment. It is sectional drawing explaining the distance between an RF coil and a shield. It is sectional drawing explaining the distance between an RF coil and a shield. It is a perspective view of the RF coil which concerns on the modification 4 of this embodiment. It is a perspective view of the conventional RF coil.

Hereinafter, the MRI apparatus to which the RF coil unit according to the embodiment of the present invention is applied will be described with reference to the drawings.

[Overall configuration of MRI equipment]
As shown in FIG. 1, it is a schematic block diagram of the MRI apparatus 100 according to this embodiment. The MRI apparatus 100 transmits a magnet 101 that forms a static magnetic field in the measurement space in which the subject 112 is arranged, a gradient magnetic field coil 102 that gives the static magnetic field a magnetic field gradient in a predetermined direction, and a high-frequency signal (RF signal). An RF antenna 103 that transmits to 112 and receives a nuclear magnetic resonance signal (NMR signal) generated from a subject 112, a pulse waveform of an RF signal (RF wave) is generated and transmitted to the RF antenna 103, and the RF antenna While controlling the drive of the transmitter / receiver 104 that processes the NMR signal received by the 103, the gradient magnetic field power supply 109 that supplies the current to the gradient magnetic field coil 102, the transmitter / receiver 104, and the gradient magnetic field power supply 109, and various information. It includes a data processing unit 105 that accepts processing and operations by an operator, a display device 108 for displaying the processing result of the data processing unit 105, and a bed 111 on which the subject 112 is placed.

The gradient magnetic field power supply 109 and the gradient magnetic field coil 102 are connected by a gradient magnetic field control cable 107. Further, the RF antenna 103 and the transceiver 104 are connected by a transmission / reception cable 106 that transmits / receives a signal between the RF antenna 103 and the transceiver 104. The transceiver 104 includes a synthesizer, a power amplifier, a receiving mixer, an analog-digital converter, a transmission / reception changeover switch, and the like (none of which are shown).

The RF antenna 103 includes a multi-channel transmitting or transmitting / receiving antenna that resonates at a predetermined frequency and has two or more channels.
In the example shown in FIG. 1, a single RF antenna is shown as the RF antenna 103 that transmits the RF signal and receives the NMR signal, but the present invention is not limited to this. For example, an RF antenna composed of a plurality of antennas may be used as the RF antenna 103, for example, by combining an RF antenna for wide-range imaging and an RF antenna for local use.

In particular, when photographing each part of the human body in detail, it is often the case that different antennas are used for transmission and reception. A large irradiation antenna installed inside the gradient magnetic field coil that covers the entire body is used for transmission, and a local antenna installed near the surface of the human body is often used for reception. In this case, the local antenna is mostly receive-only. In some cases, a local transmission / reception antenna that is locally installed near the human body and performs both transmission and reception is used. The local transmission / reception antenna in this case is also often composed of a plurality of channels.

The MRI apparatus 100 is classified into a horizontal magnetic field method and a vertical magnetic field method according to the direction of the static magnetic field formed by the magnet 101. In the case of the horizontal magnetic field method to which the RF coil unit according to the present embodiment is applied, the magnet 101 generally has a cylindrical bore (central space) and generates a static magnetic field in the left-right direction in FIG. It is called a tunnel type MRI device. On the other hand, in the case of the vertical magnetic field method, a pair of magnets are arranged vertically with the subject 112 interposed therebetween to generate a static magnetic field in the vertical direction in FIG.

The data processing unit 105 controls the transmitter / receiver 104 and the gradient magnetic field power supply 109, and intermittently irradiates the subject 112 arranged in the static magnetic field with an RF signal from the RF antenna 103 and the gradient magnetic field coil 102. A gradient magnetic field is applied. Further, the RF antenna 103 receives the NMR signal emitted from the subject 112 in resonance with the RF signal, performs signal processing, and reconstructs the image. The subject 112 is, for example, a predetermined part of the human body.

[Structure of RF transmission / reception system]
In the present embodiment, the bird cage type coil (hereinafter, simply referred to as “RF coil”) 210 shown in FIG. 2 is applied as the RF antenna 103. Although not shown here, the RF coil 210 is surrounded by a shield in which a sheet-shaped conductor is formed in a tubular shape. As shown in FIG. 2, the RF coil 210 has two annular ring conductors 203, a plurality of rung conductors 214, a capacitor, a diode, a power feeding cable (not shown), and the like.

More specifically, the ring conductor 203 is formed in an annular shape and is arranged at both ends of the cylindrical RF coil so as to form an end ring so that the central axis of the cylinder and the center of the annular shape coincide with each other. Further, the rung conductors 214 are arranged at equal intervals along the outer peripheral surface of the cylindrical shape, and the end portions of the rung conductors 214 and the ring conductors 203 are connected to each other.

Further, in the RF coil 210, the capacitors are inserted into the gaps 201 provided in the ring conductor 203 at equal intervals, and the diode circuit is inserted into the gap 212 of the rung conductor 214. The capacitor is adjusted to resonate at the frequency of a high frequency signal or a nuclear magnetic resonance signal by a ring conductor 203, a rung conductor 214, and a shield (not shown).

In the RF coil 210 shown in FIG. 2, the shape of the rung conductor 214 is different from that of the conventional bird cage type coil 200 shown in FIG. That is, since the rung conductor 214 has two straight portions 218 that are branched into two at both ends and are located substantially parallel to each other, the two conductors are parallel to each other and the end portion is formed at the axial center portion of the cylindrical axis. It is a connected ring.

By forming the rung conductor 214 in such a shape, the number of rung conductors can be substantially increased without increasing the number of capacitors inserted into the ring conductor, so that a plurality of straight lines constituting the rung conductor can be substantially increased. The distance between the parts can be increased. As a result, the uniform space of the irradiation RF magnetic field generated inside the cylinder of the RF coil 210 can be expanded to the peripheral portion of the cylinder.

It is conceivable that one rung conductor has a plurality of straight portions (two in the present embodiment (FIG. 2)), so that a current loop is formed between the plurality of straight portions. However, by adjusting the capacitor inserted in the ring conductor, current will flow in a plurality of straight lines in substantially parallel. Further, regarding the gradient magnetic fields in the X and Y directions where the generation of overcurrent may be a problem, the loops formed by the rung conductor are arranged symmetrically so as to cancel the magnetic fields from the gradient magnetic field coil, which adversely affects the image quality. It can be said that the overcurrent caused by the gradient magnetic field does not occur.

Here, the magnetic field distribution generated by the RF coil will be described. With the current calculation technology, the magnetic field distribution of the RF coil used for MRI can be calculated by simulation in almost the same manner as in reality. The magnetic field distribution calculated by simulation is shown below.

FIG. 3 shows a map of the RF magnetic field (B1 + circularly polarized light) generated by the conventional RF coil 200 shown in FIG. 13 in a Z = 0 cross section when the cylindrical axis is the Z axis and the magnetic field center is Z = 0. Is shown.
FIG. 3 (A) is a grayscale map showing the magnetic field generated when approximately 12 watts are input to each of the two input terminals, for a total of approximately 24 watts, and FIG. 3 (B) is FIG. 3 (A). ) Is a map showing the intensity of 0 to 2 [A / m] with 30 contour lines at equal intervals. At the center of the magnetic field, a cylindrical container having a diameter of 30 cm and a length of 50 cm containing a saline solution having an electrical conductivity of 0.613 [S / m] is placed instead of a human load. The map shows the magnetic field generated when a total of about 24 watts is input to the two input terminals of the antenna, about 12 watts each, in units of A / m.

Here, the conventional RF coil 200 of FIG. 13 is designed on the assumption that it is used in a tunnel type MRI apparatus of 1.5 Tesla. The RF coil 200 is a bird cage type coil having a diameter of about 610 mm and a length of a rung conductor of about 530 mm, and a capacitor is installed only in a ring conductor portion, which is called a high-pass type coil.

The number of rung conductors is 16, and a diode is installed in the notch in the center of each rung conductor. During the time when RF is not irradiated, the diode is reverse biased to prevent coupling with the receiving coil. Avoid causing coil resonance.
The value of the capacitors installed at 16 points on each of the ring conductors arranged at both ends of the cylinder is about 100 pF, and resonates at 63.8 MHz, which is the RF resonance frequency of the 1.5 Tesla MRI apparatus.

On the other hand, as described above, in the RF coil 210 of FIG. 2, each rung conductor 214 is branched at the end to form an annular shape in which both ends are connected. Therefore, in the RF coil 210, when the cylindrical axis is the Z axis and the magnetic field center is Z = 0, 16 rung conductors substantially correspond to 32 rung conductors in the cross section of Z = 0. It has become. The width of the branched rung conductor 214 is half the width of the rung conductor 202 of the RF coil 200, and the distance between the branched conductors is also arranged at equal intervals in the circumferential direction as exactly 32 rungs with a cross section of Z = 0. It has become so.

FIG. 4 shows a map of the RF magnetic field (B1 + circularly polarized waves) generated by the RF coil 210 according to the present embodiment shown in FIG. 2 in a Z = 0 cross section. FIG. 4A is a grayscale representation of the magnetic field generated when approximately 12 watts are input to each of the two input terminals of the RF coil 210, for a total of approximately 24 watts, and FIG. 4B is FIG. The map shown in (A) shows the intensity of 0 to 2 [A / m] with 30 contour lines. A similar saline phantom is installed at the center of the magnetic field.

As can be seen from FIGS. 3 (A) and 4 (A), in the slightly dark donut-shaped part near the center of the map, the magnetic field in the inner part about 4 cm from the surface of the phantom cross section having a diameter of 30 cm is the electrical conductivity of the saline solution. It is a part that is weakened due to the relative permittivity. In addition, in the part where there is a rung conductor with a diameter of about 61 cm, a black part with a weak magnetic field and a white part with a strong magnetic field appear alternately, and in the part with a rung conductor, the magnetic field is strong, and the space between the rung conductor and the rung conductor is It can be seen that the magnetic field is weakened. Comparing FIG. 3 (A) and FIG. 4 (A), there is a remarkable difference in the magnetic field in the space between the rung conductors.

In FIGS. 3 (B) and 4 (B), the portion where the contour lines are widely spaced (sparse portion) is a space where the magnetic field is uniform. Comparing FIG. 3 (B) and FIG. 4 (B), in the map of FIG. 4 (B), the portion where the contour lines are widely spaced extends closer to the rung than the map of FIG. 3 (A). That is, it can be seen that the uniform space region of the magnetic field is wide. Note that there are no contour lines in the part where the magnetic field strength on the outermost side exceeds 2 [A / m].

FIG. 5 shows the results of confirming the above comparison results more quantitatively.
In the figures shown in FIGS. 5 (A) and 5 (B), the electromagnetic field is calculated with Z = 0 cross sections at intervals of about 5 mm, and each calculation point is defined by the distance from the origin of X = Y = 0 as the horizontal axis. , The magnetic field strength is plotted on the vertical axis. FIG. 5 (A) plots the calculation points of the Z = 0 cross section of the conventional RF coil 200, and FIG. 5 (B) shows the calculation points of the Z = 0 cross section of the RF coil 210 according to the present embodiment. It is a plot. The point where the distance from the origin on the horizontal axis is 0 is the magnetic field strength of the origin.

The magnetic field strength at the origin in FIG. 5 (A) is 1.16 [A / m], and the magnetic field strength at the origin in FIG. 5 (B) is 1.13 [A / m]. The magnetic field strength of is obtained. From this, it can be seen that the irradiation efficiency of the cylindrical center portion of the RF coil 210 according to the present embodiment is substantially the same as that of the conventional RF coil 200.

In the plots of FIGS. 5A and 5B, the magnetic field B1 + largely shifts up and down as the distance from the origin increases. FIG. 5A shows the minimum value 502 and the maximum value 503 of the magnetic field of the conventional RF coil 200 at a distance of 0.25 [m]. Further, FIG. 5B shows the minimum value 512 and the maximum value 513 of the magnetic field at a distance of 0.25 [m] of the RF coil 210 according to the present embodiment.

In the conventional RF coil 200 shown in FIG. 5 (A), the magnetic field strength varies from 0.8 to 1.4 [A / m] in a portion having a diameter of 50 cm, but the present embodiment shown in FIG. 5 (B). It can be seen that the RF coil 210 of the above causes only a variation of about 0.9 to 1.2 [A / m]. This means that even when the shoulder of the human body is arranged in the portion having a diameter of 50 cm, the variation of the RF magnetic field in the portion having a diameter of 50 cm, that is, the peripheral portion of the cylinder is larger in the RF coil 210 according to the present embodiment than in the conventional RF coil 200. It means that it has been improved less. If the variation of the RF magnetic field is small, a uniform image can be acquired, so that the image quality can be improved.

FIG. 6 shows the details of the diode circuit provided on the rung conductor.
FIG. 6A shows the diode circuit 601 of the conventional RF coil 200, and FIG. 6B shows the diode circuit 602 of the RF coil 210 according to the present embodiment.
The diode circuit 601 shown in FIG. 6A is inserted into the gap 202 of the rung conductor 204 shown in FIG. 13 and includes a diode 613 and an inductor 612. In order to pass a DC current through the diode 613, a branch of a conducting wire is provided at both ends of the diode 613, and an inductor 612 is provided so that an RF current does not flow in the path through which the DC current flows.

The diode circuit 601 of FIG. 6A merely shows an example of a diode circuit provided in one rung conductor 204, but the diode circuit 601 provided in each rung conductor 204 is the circumference of a cylinder. It is arranged continuously in the direction so that a DC current can flow through the diode 613 of each rung conductor 204.
At the time of RF irradiation, a DC current is passed to bring the RF coil 200 into a resonance state. At the time of reception, no DC current is applied to bring the RF coil 200 into a non-resonant state. By doing so, the coupling with the receiving antenna is prevented, and the deterioration of the receiving sensitivity is prevented.

The diode circuit 602 shown in FIG. 6B is inserted into the gap 212 of the rung conductor 214 and includes a diode 613 and an inductor 612. The diode 613 is provided in the gap 212 of each straight portion 218 branched at the end of the rung conductor 214, respectively. The inductor 612 is also provided between the diodes 613 inserted into the adjacent straight line portions 218.

Therefore, the number of diodes 13 and inductors 612 provided on the rung conductor 214 is increased by one as compared with the diodes 613 and inductors 612 provided on the rung conductor 204 of the conventional RF coil 200. The diode circuits 602 provided in each rung conductor 214 are connected in series with each other via an inductor 612. By configuring the diode circuit 602, the diode 613 can be inserted into the rung conductor even if the rung conductor 214 branches into two straight portions 218.

As described above, according to the present embodiment, the rung conductor is branched at both ends to include a plurality of straight portions, so that the number of capacitors inserted into the gap of the ring conductor is not increased. Since the number of conductors can be substantially increased, the distance between the rung conductors can be narrowed without increasing the number of parts.

As a result, the uniform space of the irradiation RF magnetic field generated inside the cylinder of the RF coil can be expanded to the peripheral portion of the cylinder. That is, according to the RF coil according to the present embodiment, the region where the RF magnetic field is uniformly irradiated can be expanded to expand the imaging space without increasing the number of parts and the device cost, and this can be used as an MRI device. When applied to, a wide range of high-quality MRI images can be obtained.

(Modification example 1)
In the above-described embodiment, the RF coil 210 including the rung conductor 214 branched at both ends and having two straight portions has been described. The RF coil is not limited to such a shape, and for example, the rung conductor may have a shape having three straight portions.

FIG. 7 shows the RF coil 310 according to this modification. The RF coil 310 includes a rung conductor 314 that is branched at both ends and has three straight portions. In the rung conductor 314 of this modification, since the number of the rung conductors 314 itself is not changed, it is not necessary to increase the number of capacitors inserted in the ring conductor 303. In this example, 16 capacitors are inserted in one ring conductor 303, the number of rung conductors 314 is 16, and each rung conductor 314 has three linear portions 318, which is substantially the same. The same effect as the existence of 48 linear rung conductors can be obtained in, and the uniform space of the magnetic field generated by the RF coil is widened.

(Modification 2)
In the above-described embodiment and the first modification thereof, one rung conductor is branched into a plurality of straight portions at both ends thereof, and the ring conductor is provided with one rung conductor and one capacitor alternately. In this modification, an example will be described in which two independent linear portions that are not branched at both ends and are not connected at both ends form a rung conductor, and the rung conductor and one capacitor are alternately provided.

FIG. 8 shows an example of an RF coil in which the number of capacitors is reduced and a rung conductor 814 having two linear portions 818 and one capacitor are alternately provided. In the example of FIG. 8, the RF coil 810 includes an annular ring conductor 803 and a rung conductor 814, and the rung conductor 814 has two straight portions 818. In FIG. 8, the rung conductor 814 is connected to the ring conductor 803 so that the straight portions 818 are arranged at equal intervals.

The ring conductor 803 inserts the rung conductor 814 and the capacitor alternately, and as a result, one capacitor is inserted every time two linear portions 818 are provided. That is, the ring conductor 803 is provided with two straight portions 818 and a capacitor alternately, and at the connection point between the two straight portions 818 and the straight portion 818 in the ring conductor, the two straight portions 818 are paired with each other. There is no gap between them to install a capacitor.

This is because the number of capacitors is reduced in the ring conductor 803 as compared with the case where the linear rung conductors and the capacitors are alternately arranged one by one. Normally, in a high-pass type bird cage coil having 32 linear rung conductors, 32 capacitors are installed in the ring conductor, but in the RF coil 801 of FIG. 8, the number is reduced to 16 which is half of the number. It can be said. When the number of capacitors is reduced, the number of parts is reduced, so that the cost required for the parts and the manufacturing cost are reduced, and the equipment cost is reduced. Further, since the number of parts is small, there is an advantage that the risk of failure is reduced.

(Modification example 3)
In the above-described embodiment and each modification, an RF coil in which a plurality of rung conductors branched at both ends and having a plurality of straight portions are provided between two annular ring conductors has been described. The RF coil is not limited to such a shape, and the ring conductor may be an annular shape on an ellipse.

FIG. 9 shows an example in which the rung conductor 214 branched into two straight portions 218 at both ends is applied to the elliptical bird cage type coil 910. In the elliptical bird cage coil 910 of FIG. 9, rung conductors 214 are installed at equal intervals in the elliptical circumferential direction. As described above, the present embodiment can be applied even to the elliptical bird cage type coil, and in this case as well, the uniformity of the irradiation magnetic field in the peripheral portion of the rung can be improved. In the elliptical bird cage type coil, the rung conductors 214 do not necessarily have to be installed at equal intervals as in the example shown in FIG. 9, and the rung conductor installation intervals can be changed as appropriate.

As shown in FIG. 10, it is conceivable that the distance (gap) between the rung conductor of the normally elliptical bird cage coil 910 and the shield 950 is equal. However, as shown in FIG. 11, by narrowing the left-right spacing and widening the top-bottom spacing, a wider patient space can be secured in the left-right direction. Further, at this time, the uniformity of the irradiation magnetic field is maintained or improved by narrowing the distance between the rung conductors in the vertical direction and widening in the horizontal direction. The distance between the shield and the RF coil needs to be a predetermined distance (usually about 20 to 30 mm) in order to increase the power efficiency of irradiation.

(Modification example 4)
Further, as shown in FIG. 12, even in the case of an elliptical bird cage type coil, a rung conductor composed of two independent linear rung portions and one capacitor are alternately provided without branching at both ends. be able to. The elliptical bird cage coil 1101 shown in FIG. 12 is also provided with a rung conductor 814 having two straight portions 818 and a capacitor alternately provided. Therefore, at the connection point of the ring conductor with the straight line portion 818, a gap is not provided between the straight line portions 818 which are paired with each other because a capacitor is installed.

In the ring conductor 1103, the number of capacitors is reduced as compared with the conventional elliptical bird cage type coil in which linear rung conductors and capacitors are alternately arranged one by one, so that the cost is reduced and a failure occurs. There is a merit that the risk of As described above, even in the elliptical bird cage type coil 1101, by applying the rung conductor according to the present embodiment, it is possible to obtain an RF coil in which the uniformity of the irradiation magnetic field in the peripheral portion of the rung conductor is improved.

Various additions and changes can be made to the embodiments of the present invention and modifications thereof without departing from the spirit of the invention.

100 ... MRI device, 101 ... magnet, 102 ... gradient magnetic field coil, 103 ... RF antenna, 104 ... transmitter / receiver, 105 ... data processing unit, 106 ... transmit / receive cable, 107 ... gradient magnetic field control cable, 108 ... display device, 109 ... gradient magnetic field power supply, 111 ... bed, 112 ... subject, 200 ... conventional RF coil, 201 ... Gap, 202 ... Gap, 203 ... Ring conductor, 204 ... Lang conductor, 210 ... RF coil, 212 ... Gap, 214 ... Lang conductor, 218 ... Straight part, 310 ... Bird cage type RF coil, 303 ... Ring conductor, 314 ... Lang conductor, 318 ... Straight part, 601 ... Diode circuit, 602 ... Diode circuit, 612 ... inductor, 613 ... Diode, 803 ... Ring conductor, 810 ... Bird cage type RF coil, 814 ... Lang conductor, 818 ... Straight part, 910 ... Elliptical bird cage type RF coil, 950 ...・ ・ Shield 1101 ・ ・ ・ Elliptical bird cage type RF coil 1103 ・ ・ ・ Ring conductor

Claims (8)

It is applied to MRI equipment and has a tubular RF coil that transmits at least one of the high frequency signal to the static magnetic field space and the reception of the nuclear magnetic resonance signal generated from the subject placed in the static magnetic field.
The RF coil
Two ring conductors formed in an annular shape and arranged at both ends of the tubular shape so that the central axis of the tubular shape and the center of the annular shape coincide with each other.
A plurality of rung conductors arranged along the outer peripheral surface of the tubular shape at predetermined intervals parallel to the central axis of the tubular shape and connecting the two ring conductors.
The ring conductor has a plurality of capacitors inserted between the adjacent rung conductors, and has a plurality of capacitors.
Said rung conductor, peripheral N (N is an integer of 2 or more) in the direction have a straight portion which is arranged substantially parallel to each other, the linear portion of the N book that are arranged at equal intervals in the circumferential direction High frequency coil unit featuring. The high-frequency coil unit according to claim 1, wherein the rung conductor has straight portions that are arranged substantially parallel to each other by branching into N lines in the circumferential direction at both ends. The high-frequency coil unit according to claim 1, wherein both ends of each of the straight portions constituting the rung conductor are directly connected to the ring conductor and arranged substantially parallel to each other. The high-frequency coil unit according to any one of claims 1 to 3, wherein the capacitors are arranged alternately with the rung conductor in the ring conductor. It is composed of a sheet-shaped conductor, and further includes a shield formed in a tubular shape so as to surround the RF coil.
The RF coil and the shield are formed in a cylindrical shape having a substantially circular cross section perpendicular to the central axis.
The high-frequency coil unit according to any one of claims 1 to 4, wherein the central axis of the RF coil and the central axis of the shield are arranged so as to coincide with each other. It is composed of a sheet-shaped conductor, and further includes a shield formed in a tubular shape so as to surround the RF coil.
The RF coil and the shield are formed in a tubular shape having a substantially elliptical cross section perpendicular to the central axis.
The high-frequency coil unit according to any one of claims 1 to 4, wherein the central axis of the RF coil and the central axis of the shield are arranged so as to coincide with each other. The high-frequency coil unit according to claim 5 or 6 , wherein the capacitor is adjusted by the ring conductor, the rung conductor, and the shield to resonate at the frequency of the high-frequency signal or the nuclear magnetic resonance signal. A magnetic resonance imaging apparatus comprising the high-frequency coil unit according to any one of claims 1 to 5.

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