JP3382865B2 - RF magnetic field forming apparatus and magnetic resonance imaging apparatus - Google Patents

Description

Detailed Description of the Invention

[0001]

TECHNICAL FIELD The present invention relates to an RF coil (ra).
BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a radio frequency coil, an RF magnetic field forming apparatus, and a magnetic resonance imaging method and apparatus, and in particular, an RF having a plurality of electric paths through which a current for forming an RF magnetic field flows and an annular electric path connecting the ends of these electric paths. Coil and R using such an RF coil
The present invention relates to an F magnetic field forming apparatus, and a magnetic resonance imaging method and apparatus using such an RF magnetic field forming apparatus.

[0002]

2. Description of the Related Art The direction of a static magnetic field is perpendicular to the body axis of a subject,
In a so-called vertical magnetic field type magnetic resonance imaging apparatus, since an open static magnetic field space is formed, an RF coil for generating an RF magnetic field has an RF having a loop surface parallel to a magnetic pole surface of the static magnetic field generator. A coil is used. In this type of RF coil, R is applied in the direction parallel to the loop surface of the coil.
An F magnetic field is generated to form a magnetic field for RF excitation perpendicular to the static magnetic field direction.

As an RF coil for generating such an RF magnetic field, for example, one as shown in FIG. 17 is used. As shown in the figure, when the three directions perpendicular to each other in the three-dimensional space are x, y, and z, and the z direction is the static magnetic field direction, the RF coil has two xy spaced apart in the z direction.
It has two loops formed on each side. The two loops are connected in series to form a one-stroke electric path.

The two loops are a pair of main paths 202 and 204 extending in the x direction at the center of each of them, and a main path 2
It has a pair of 02 ', 204', a connecting path 206, 208, 206 ', 208' on both sides of the main path, and a connecting path 210 in the z direction. By connecting the main paths 202 to 204 ′ with the communication paths 206 to 210 in a single stroke, a current flows in the main paths 202 and 204 in the same direction,
Currents in the same direction flow in the opposite directions to the main paths 202 'and 204'.

In this relationship, the currents flowing through the main paths 202, 204 and 202 ', 204' respectively cause
Magnetic fields as shown in FIG. 18 are generated, respectively, and a magnetic field resulting from the combination thereof is generated in the y direction, that is, the direction parallel to the coil loop plane.

Using two RF coils having such a structure,
As shown by a solid line and a broken line in FIG. 19, respectively, when the main paths are combined in a relationship in which they intersect each other perpendicularly in the xy plane, and two RF coils are supplied with RF signals having phases different from each other by 90 degrees, the two RF coils are RF generated in each
By synthesizing the vectors of the magnetic fields, it is possible to obtain an RF magnetic field rotating in a plane parallel to the xy plane. An RF coil composed of a combination of such two RF coils is a quadrature type R coil.
It is called an F coil.

[0007]

In the RF coil as described above, an RF magnetic field is also generated by the current in the communication path, and this magnetic field attenuates the magnetic field for RF excitation by the main path, so that the efficiency of RF magnetic field generation is poor, and it is desirable. There is a problem that a large excitation power is required to obtain the magnetic field strength of.

Therefore, the current flowing through the main path and the communication path becomes large, and the high frequency power absorption of the object, that is, the hall SAR (whole specific a).
There has been a problem that the bsorption rate) becomes large.

Further, the RF magnetic field generated by the communication path is
In addition to attenuating the RF excitation, the subject's local high frequency power absorption or local SAR (local s)
There was a problem of increasing the specific absorption rate.

The present invention has been made to solve the above-mentioned problems, and an RF coil and an RF magnetic field forming apparatus for efficiently generating an RF magnetic field, and a magnetic resonance imaging method using such an RF magnetic field forming apparatus. And to realize the device.

Another object of the present invention is to realize an RF coil and an RF magnetic field forming apparatus having a small SAR of a subject, and a magnetic resonance imaging method and apparatus using such an RF magnetic field forming apparatus.

[0012]

(1) A first aspect of the present invention for solving the above-mentioned problems is to provide a radial electric path through which a current for forming an RF magnetic field flows in a space accommodating a subject, and the radial electric path. And an annular electric path connecting ends of the radial electric path at a position farther from the space than the electric path.

(2) A second aspect of the present invention which solves the above-mentioned problems is that a plurality of electric paths are parallel to each other through which an electric current for forming an RF magnetic field flows in a space accommodating a subject, and the plurality of electric paths. And a ring-shaped electric path that connects the ends of the plurality of electric paths at positions farther from the space than the plurality of electric paths, respectively.

(3) According to a third aspect of the present invention for solving the above-mentioned problems, a radial electric path through which a current for forming an RF magnetic field flows in a space accommodating a subject, and the radial electric path is more preferable than the radial electric path. The phase is 90 degrees at two points of the annular electric path connecting the ends of the radial electric path at a position far from the space and the annular electric path whose directions viewed from the center of the annular electric path differ from each other by 90 degrees. And an RF signal supply means for supplying different RF signals, respectively.

(4) According to a fourth aspect of the present invention which solves the above-mentioned problems, a plurality of electric paths parallel to each other through which an electric current for forming an RF magnetic field flows in a space accommodating a subject, and the plurality of electric paths. At both ends of the electric path, a circular electric path connecting the ends of the plural electric paths at positions farther from the space than the plural electric paths and a direction viewed from the center of the circular electric path are different from each other by 90 degrees. An RF magnetic field forming device, comprising: RF signal supplying means for supplying RF signals having phases different from each other by 90 degrees to two points of the annular electric path.

(5) According to a fifth aspect of the present invention for solving the above-mentioned problems, a static magnetic field is formed in a space accommodating a subject, a gradient magnetic field is formed in the space, and an RF magnetic field is formed in the space. A magnetic resonance imaging method for measuring a magnetic resonance signal from the space and generating an image based on the measured magnetic resonance signal, wherein the formation of the RF magnetic field is performed by the RF magnetic field forming device according to (3) and ( 4) A magnetic resonance imaging method characterized in that the magnetic resonance imaging method is performed by any one of the RF magnetic field forming devices described in 4).

(6) A sixth aspect of the present invention for solving the above problems is a static magnetic field forming means for forming a static magnetic field in a space accommodating a subject, and a gradient magnetic field forming means for forming a gradient magnetic field in the space. An RF magnetic field forming means for forming an RF magnetic field in the space; a measuring means for measuring a magnetic resonance signal from the space; and an image generating means for generating an image based on the magnetic resonance signal measured by the measuring means. It is a magnetic resonance imaging device which has, Comprising: The said RF magnetic field formation means is either the RF magnetic field formation apparatus of (3), or the RF magnetic field formation apparatus of (4), It is characterized by the above-mentioned. It is a magnetic resonance imaging apparatus.

In any one of the first to sixth inventions, the RF close to the annular electric path is provided.
Providing a shield further increases the RF excitation efficiency,
Moreover, it is preferable in that the SAR is further reduced.

Further, in the first invention or the third invention, it is preferable that the radial electric paths are opposed to each other across the space in order to improve the uniformity of the RF magnetic field strength distribution.

(Operation) In the present invention, the effect of reducing the RF excitation magnetic field by the annular electric path is reduced by increasing the distance between the annular electric path and the subject. Also, the SAR of the subject is reduced.

[0021]

BEST MODE FOR CARRYING OUT THE INVENTION Hereinafter, embodiments of the present invention will be described in detail with reference to the drawings. The present invention is not limited to the embodiment. FIG. 1 shows a block diagram of the magnetic resonance imaging apparatus. This device is an example of an embodiment of the present invention. The configuration of this device shows an example of an embodiment relating to the device of the present invention. The operation of the apparatus shows an example of an embodiment relating to the method of the present invention.

As shown in FIG. 1, in this apparatus, the static magnetic field generator 2 forms a uniform static magnetic field in its internal space. The static magnetic field generator 2 is an example of an embodiment of the static magnetic field forming means in the present invention. The static magnetic field generator 2 includes a pair of magnetic generators (not shown) such as permanent magnets, which are opposed to each other in the vertical direction in the drawing with a space therebetween, and a static magnetic field (vertical magnetic field) is formed in the opposed space. is doing. The magnetic generator is not limited to a permanent magnet, but may be a superconducting electromagnet, a normal conducting magnet, or the like.

Gradient coil units 4, 4'and transmission coil units 6, 6'are provided in the internal space of the static magnetic field generating unit 2 and similarly face each other in the vertical direction with a space therebetween. An RF shield (not shown) is provided between the gradient coil units 4 and 4'and the transmission coil units 6 and 6 '. The RF shield is made of a conductor such as copper foil. The transmission coil units 6 and 6'are examples of embodiments of the RF coil of the present invention. The transmission coil units 6 and 6'will be described later.

In the space where the transmission coil units 6 and 6'face each other,
The subject 8 is mounted on the imaging table 10 and carried in and carried out by a carrying unit (not shown). The body axis of the subject 8 is orthogonal to the direction of the static magnetic field. The receiving coil unit 106 is attached to the imaging table 10 so as to surround the imaging region of the subject 8. The receiving coil unit 106 is, for example, for L spine (lumbar spine) imaging,
It is attached so as to surround the waist of the subject 8. The receiving coil unit 106 is not limited to the L spine,
It can be installed at a position corresponding to a desired imaging region.

A gradient drive unit 16 is provided in the gradient coil units 4 and 4 '.
Are connected. The gradient driving unit 16 includes the gradient coil unit 4,
A drive signal is given to 4'to generate a gradient magnetic field. Gradient coil section 4, 4'and gradient drive section 16
The part consisting of is an example of an embodiment of the gradient magnetic field forming means in the present invention. There are three types of gradient magnetic fields that are generated: a slice gradient magnetic field, a read gradient magnetic field and a phase encode gradient magnetic field.

A transmitting unit 18 is connected to the transmitting coil units 6 and 6 '. The transmitter 18 applies a drive signal to the transmitter coils 6 and 6 ′ to generate an RF magnetic field, and thereby excites a spin in the body of the subject 8. The portion including the transmission coil units 6 and 6 ′ and the transmission unit 18 is an example of the embodiment of the RF magnetic field forming apparatus of the present invention. It is also an example of an embodiment of the RF magnetic field forming means in the present invention.

The receiving coil section 106 receives a magnetic resonance signal generated by the excited spin. Receiver coil unit 10
Reference numeral 6 is connected to the input side of the receiving unit 20 and inputs the received signal to the receiving unit 20. The output side of the receiver 20 is connected to the input side of an analog-to-digital converter 22. The analog / digital converter 22 converts the output signal of the receiver 20 into a digital signal. The output side of the analog / digital converter 22 is connected to a computer 24.

The computer 24 inputs the digital signal from the analog / digital converter 22 and stores it in a memory (not shown). Receiver coil unit 10
6, a portion including the receiving unit 20, the analog / digital converting unit 22, and the computer 24 is an example of the embodiment of the measuring means in the present invention.

A data space is formed in the memory. The data space constitutes a Fourier space. The computer 24 performs an inverse Fourier transform on the data in the Fourier space to reconstruct the image. The computer 24 is an example of the embodiment of the image generating means in the present invention.

The computer 24 is connected to the control unit 30. The controller 30 is connected to the gradient driver 16, the transmitter 18, the receiver 20, and the analog / digital converter 22. The control unit 30 controls the gradient drive unit 16, the transmission unit 18, the reception unit 20, and the analog / digital conversion unit 22 based on a command given from the computer 24, and executes magnetic resonance imaging.

A display unit 32 and an operation unit 34 are connected to the computer 24. The display unit 32 is the computer 2
The reconstructed image and various information output from No. 4 are displayed. The operation unit 34 is operated by an operator and inputs various commands and information to the computer 24.

FIG. 2 shows a schematic configuration of the transmission coil units 6 and 6 '. In the figure, the three-dimensional structure of the RF coil which is the main part of the transmission coil units 6 and 6'is shown. Three mutually perpendicular directions in the three-dimensional space are defined as x, y, and z. The z direction is the direction of the static magnetic field. As shown in the figure, the transmission coil unit 6 has an electric path formed on the xy plane, and the transmission coil unit 6'has an electric path formed on another xy plane separated in the z direction.

The transmission coil section 6 has an annular electric path 60.
Have two. The electric path 602 is an example of the embodiment of the annular electric path according to the present invention. Inside the annular electric path 602, a plurality of electric paths 604 extending radially from the central portion are provided. The electric path 604 is an example of the embodiment of the radial electric path according to the present invention. The number of radial electric paths 604 is eight, for example.
The number is not limited to 8 and may be an integral multiple of 4. Note that the electrical path 604 is represented by a single point. The electric path 604 has the electric path 60 at one end thereof.
2 and the other ends thereof are connected to the central circular electric path 606, respectively.

On the electric path 602, the RF signal is supplied from the transmitting unit 18 to two locations whose directions viewed from the center inside are different from each other by 90 degrees. The phases of the RF signals supplied to these two points are different from each other by 90 degrees. The transmitter 18 is an example of an embodiment of the RF signal supply means in the present invention.

The transmitting coil section 6'is provided with an annular electric path 6
02 '. The electric path 602 ′ is an example of the embodiment of the annular electric path according to the present invention. Inside the annular electric path 602 ′, a plurality of electric paths 604 ′ radially extending from the central portion are provided. Electrical pathway 6
04 'is an example of embodiment of the radial electric path in this invention. The number of radial electric paths 604 ′ is eight, for example. The number is not limited to 8 and may be an integral multiple of 4. It should be noted that the electrical path 604 'is represented by a single point. The electric paths 604 'are connected at one end to the electric path 602' and at the other end to the central circular electric path 606 '.

On the electric path 602 ', the RF signal is supplied from the transmitting unit 18 to two locations whose directions viewed from the center inside are different from each other by 90 degrees. The phases of the RF signals supplied to these two points are different from each other by 90 degrees. The transmitter 18 is an example of an embodiment of the RF signal supply means in the present invention.

The transmitting coil portions 6 and 6'as described above face each other in a mirror image relationship, and form the RF coil of the present apparatus. The RF signals are supplied to the transmission coil units 6 and 6 ′, for example, as shown in FIG. 3, two RF signals output from the transmission unit 18 and having phases different from each other by 90 degrees are respectively supplied to a power splitter 182. ，
This is done by dividing the power into equal power and giving the same at 184.
Supplying the RF signal in this manner is preferable in that the transmitter coil units 6 and 6'are RF-driven exactly under the same conditions. An RF power amplifier may be provided between the splitters 182 and 184 and the transmission coil units 6 and 6 ', if necessary.

Since the RF signals having the same power and the same phase are simultaneously supplied from the common signal source to the two coil loops facing each other, the electromagnetic coupling between the two coils does not pose a problem. Further, since the distance between both coils is sufficiently large, electrostatic coupling does not pose a problem. As a result, an RF coil having excellent frequency characteristics can be obtained.

The signal line for supplying the RF signal to both coils is
For example, an appropriate signal line such as a coaxial cable may be used, and the wiring can be freely routed.
Therefore, it is easy to process the signal line so as not to hinder the carry-in of the subject 8 and it does not hinder the releasability of the magnetic field space.

The electric circuit of the transmission coil section 6 is shown in FIG.
As shown in the figure, the annular electric path 602 has capacitors 608 in series in each section divided by the radial electric path 604. The capacitor 6
The coding of 08 is represented at one place. A variable capacitor for adjusting the capacitance is connected in parallel to some of the capacitors. These variable capacitors are used to precisely adjust the quadrature of the 0 and 90 degree phases.

The RF signals of 0-degree phase and 90-degree phase from the transmitter 18 are applied to both ends of the capacitor 608 in two sections whose directions from the center are different from each other by 90 degrees.

Focusing on the RF signal of 0 degree phase, the ratio of the current flowing through each electric path is as shown in FIG. That is, in the outer annular electric path 602, the current ratio is 0 in the section where the RF signal is fed and the section on the opposite side, and the current ratio is 1 in the two sections whose directions are different by 90 degrees, and the others. The current ratio is 0.7 in the section. Also in the inner annular electric path 606, the same current ratio is obtained in each section corresponding to each section of the outer electric path 602.

In the radial electric path 604, the current ratios of the four electric paths located on both sides of the two sections where the electric current ratio is 0 in the annular electric path 602 are all 0.
7, the current ratios of the four electric paths located on both sides of the two sections having the current ratio of 1 are 0.3.

With one polarity of the RF signal, the direction of the current in each electrical path is as indicated by the arrows. That is, in the outer annular electric path 602, the current ratio flows in the counterclockwise and clockwise directions on the right and left sides of the section where the current ratio is 0, and the inner annular electric path 606 opposes them. It flows around. In the radial electric path 604, the currents branched from the above currents flow such that the currents that are symmetrical with respect to the center have directions opposite to each other.

When such a current flows, the inner ring-shaped electric path 606 is thereby provided on the back side in the drawing, that is, on the side facing the transmission coil section 6 ', as shown in FIG.
As indicated by the dotted line arrow, a diametrical magnetic field perpendicular to the electric path having a current ratio of 1 is generated.

For other polarities of the RF signal, the direction of the current is all opposite to the above and the direction of the magnetic field is also reversed. Therefore, the transmission coil unit 6 produces an RF magnetic field corresponding to the RF signal.

Focusing on the 90-degree phase current, the current ratio of each electric path for one polarity of the RF signal is as shown in FIG. Since the supply location of the RF signal is spatially different from the 0 degree phase signal by 90 degrees, FIG. 7 corresponds to FIG. 5 rotated 90 degrees counterclockwise. Therefore, the direction of the RF magnetic field is also the direction obtained by rotating the direction shown in FIG. 6 by 90 degrees counterclockwise, as shown by the arrow indicated by the one-dot chain line in FIG.

The RF magnetic field generated by the transmitting coil section 6 is 2 above.
Vector synthesis of two RF magnetic fields. Since the two RF magnetic fields have a phase difference of 90 degrees, the combined RF magnetic field is a rotating magnetic field that rotates at the frequency of the RF signal. That is, the transmission coil unit 6 is a quadrature type RF coil.

The transmission coil section 6'has the same circuit configuration and produces a similar rotating magnetic field. However, in the transmission coil unit 6 ′, the directions of the electric currents in the respective electric paths are all opposite to the electric paths in the transmission coil unit 6 which are in a mirror image relationship. Such a reversal of the direction of the current can be easily performed by, for example, reversing the connection of the signal lines connected to both ends of the capacitor 608. In this case, the RF magnetic field generated immediately above the inner annular electric path 606 ′, that is, on the side facing the transmission coil unit 6, is
An RF magnetic field is formed in the same direction as the above-mentioned magnetic field by the transmission coil unit 6 and the sum of the two.

As a result, an RF magnetic field rotating in a plane perpendicular to the z direction is generated in the space between the transmitting coil portions 6 and 6 '. The intensity profile of this RF magnetic field in the z direction (p
profile) is the center of the static magnetic field (Z / 2), as shown in FIG. 9, for example, by adding the RF magnetic fields of the two coils.
It has uniformity over a wide range including.

The electric paths of the transmission coil portions 6 and 6'are made of, for example, a conductor foil. FIG. 10 shows an example of a circuit pattern of the transmission coil unit 6 made of a conductor foil. As shown in the figure, the electric paths 602, 604 and 606 are all made of a conductor such as copper foil. The thickness of the copper foil is, for example, about several tens of μm. The width of the electric paths 602 and 604 is, for example, about several cm to several tens of cm. The circuit pattern is formed on a support member (not shown).

The annular electric path 602 has slits 610 for each section.
A capacitor 608 connecting the electric path 602 in series is provided between the capacitors 10. The illustration of the capacitor 608 is omitted. In addition, the numbering of the slit 610 is represented by one place.

Since the radial electric path 604 is a wide current path, the ratio of the portion where no current flows between the adjacent electric paths 604 is determined by the electric path 604.
It can be made much smaller than if it were composed of electric wires. As a result, it is possible to mitigate the decrease in magnetic field strength in the middle of the adjacent electric paths 604 and mitigate the non-uniformity of the magnetic field strength distribution.

The electric path 606 is not a circular ring as described above, but a disk shape with its central portion closed. As a result, the current flows in a distributed manner throughout the disk, so that the magnetic field does not drop in the central part due to the absence of current in the central part, as in the case of a ring, and magnetic field formation can be optimized. .

A step is provided between the annular electric path 602 and the radial electric path 604. That is, A-A
As shown in the cross section in FIG. 11, a step is provided so that the annular electric path 602 is farther from the center of the static magnetic field than the radial electric path 604. This increases the distance between the subject 8 carried into the static magnetic field space and the annular electric path 602, and the attenuation effect of the RF excitation magnetic field by the annular electric path 602 is weakened.

Therefore, the excitation power for obtaining a predetermined RF magnetic field strength is reduced, and efficient RF excitation can be performed. Further, since the RF current flowing in the electric path is reduced, the hole SAR of the subject 8 is reduced.
Further, since the distance between the subject 8 and the circular electric path 602 is large, the local SAR of the subject 8 becomes small in combination with the decrease in the current.

The circuit pattern of the transmission coil section 6'is configured to have a mirror image relationship with the transmission coil section 6. For this reason, the distance between the subject 8 and the annular electric path 602 ′ is increased, and the attenuation effect of the RF excitation magnetic field by the annular electric path 602 ′ is weakened. Therefore, similarly to the above, efficient RF excitation can be performed, and the Hall SAR and the local SAR can be reduced.

By providing the step, for example, as shown in FIG. 12, the distance between the RF shield 700 provided between the gradient coil 4 and the electric path 602 is shortened. Therefore, the mirror (mir) of the RF current flowing in the electric path 602 is
ror) effect becomes large. The Miller effect is equivalent to a current (mirror current) having a polarity opposite to that of the RF current flowing through the electric path 602 at a position symmetrical to the electric path 602 with respect to the RF shield 700. As this effect is enhanced, the canceling action due to the mirror current is increased, and the RF magnetic field due to the current flowing in the electric path 602 is reduced. Therefore, the efficiency of RF excitation is further improved, and the hole / local SA of the subject 8 is also increased.
R is further reduced.

FIG. 13 shows another example of the magnetic resonance imaging apparatus. This device is an example of an embodiment of the present invention. The configuration of this device shows an example of an embodiment relating to the device of the present invention. The operation of the apparatus shows an example of an embodiment relating to the method of the present invention.

In FIG. 13, the same parts as those shown in FIG. 1 are designated by the same reference numerals and the description thereof will be omitted. FIG.
As shown in FIG.
Have. The static magnetic field generator 21 is an example of an embodiment of the static magnetic field forming means in the present invention. Static magnetic field generator 21
Is composed of, for example, a superconducting electromagnet. The static magnetic field generator 21 generates a static magnetic field in its internal space. The direction of the static magnetic field is parallel to the body axis of the subject 8 and is a so-called horizontal magnetic field.

A substantially cylindrical gradient coil section 41 is provided in the internal space of the static magnetic field generating section 21, and a body coil section 61 is provided inside the gradient coil section 41. The body coil portion 61 is an example of the embodiment of the RF coil of the present invention. Gradient coil section 41 and body coil section 61
An RF shield (not shown) is provided between them. The subject 8 is housed in the internal space of the body coil unit 61.

The body coil section 61 is driven by the transmitting section 18 to generate an RF magnetic field. Further, the magnetic resonance signal generated from the subject 8 is detected and input to the receiving unit 20. That is, the body coil unit 61 is used for transmitting and receiving RF signals.

FIG. 14 shows a schematic structure of the body coil portion 61. As shown in the figure, the body coil portion 61 is
It has a plurality of electrical paths 612 that are parallel to each other. It should be noted that the coding of the electrical path is represented at one place. Electric path 61
2 is an example of an embodiment of a plurality of parallel electric paths in the present invention. Both ends of the plurality of electric paths 612 are connected by annular electric paths 614 and 614 ′, respectively. The annular electric paths 614 and 614 ′ are an example of an embodiment of the annular electric path according to the present invention.

In the annular electric paths 614 and 614 ', capacitors (not shown) are provided in each section constituted by the connection points of the plurality of electric paths 612. In either one of the annular electric paths 614, 614 ',
An RF signal feeding section (not shown) is provided at each of two locations that differ in direction from the center by 90 degrees. These power supply units are also used as RF signal extraction units.
Such an RF coil is a birdcage coil (bird).
Also called a cage coil).

In the birdcage coil, a plurality of electrical paths 612 contribute to the generation of the RF excitation magnetic field. The annular electric paths 614 and 614 ′ at both ends do not contribute to the generation of the excitation magnetic field and function as a return path of the RF current.

The diameter of the annular electric paths 614 and 614 'is
The diameter is larger than the diameter of the cylindrical space formed by the arrangement of the plurality of electric paths 612, and both ends of the electric path 612 are bent toward the outside of the cylindrical space in order to connect to it. As a result, the annular electric paths 614,
614 ′ is farther from the subject 8 housed in the cylindrical space than the electrical path 612. As a result, the damping effect of the excitation magnetic field by the annular electric paths 614, 614 'is reduced, and the excitation efficiency is improved. In addition, the hole / local SAR of the subject 8 becomes smaller accordingly. These effects are further promoted by an increase in the mirror effect due to the proximity to the RF shield 700 ′ provided between the gradient coil 41 and the gradient coil 41, as shown in FIG.

The operation of this apparatus will be described. The operation of the apparatus shown in FIG. 1 will be described, but the operation of the apparatus shown in FIG. 13 is similar. The operation of this device proceeds under the control of the control unit 30. As one of the specific examples of magnetic resonance imaging,
Gradient echo
Imaging by the method will be described. The magnetic resonance imaging is not limited to the gradient echo method, and for example, a spin echo method or an EPI (echo) method.
Other suitable methods such as planar imaging) may be used.

For imaging by the gradient echo method,
For example, a pulse sequence (pulses) as shown in FIG.
e sequence) is used. The pulse sequence progresses from left to right along the time axis t. The execution of the pulse sequence is controlled by the control unit 30.

As shown in (1) of FIG. 16, RF excitation is performed by α ° pulses. Due to the above-described configuration of the RF coil and the mirror effect, the RF excitation is efficiently performed and the hole / local SAR of the subject 8 is kept low, so that the temperature of the subject 8 as a whole and locally increases. It can be kept low.

RF excitation is performed with good uniformity over a wide range. During RF excitation, a slice gradient magnetic field Gs is applied as shown in (2). As a result, spins at a predetermined portion of the subject 8 are selectively excited. Following this selective excitation, the spin is rephased (r
ephase).

Next, as shown in (4), phase encoding is performed by the phase encoding gradient magnetic field Gp. Next, spin dephase (dep) is performed by the read gradient magnetic field Gr.
is performed, and then the polarity of the gradient magnetic field is reversed to generate a magnetic resonance signal (gradient echo: gradient echo).
(echo) is read.

By repeating such a pulse sequence for a predetermined repetition time TR (repetition time), gradient echoes are collected one by one. The magnitude of the phase encode gradient magnetic field Gp is changed for each TR, and magnetic resonance signals of a plurality of views are collected in the memory. The computer 24 performs a two-dimensional inverse Fourier transform on the data collected in the memory to generate a tomographic image of the imaged site. The tomographic image is displayed on the display unit 32.

[0073]

As described above in detail, according to the present invention, the RF coil and the R coil for efficiently generating the RF magnetic field are provided.
It is possible to realize an F magnetic field forming apparatus, and a magnetic resonance imaging method and apparatus using such an RF magnetic field forming apparatus.

Further, it is possible to realize an RF coil and an RF magnetic field forming apparatus having a small SAR of a subject, and a magnetic resonance imaging method and apparatus using such an RF magnetic field forming apparatus.

[Brief description of drawings]

FIG. 1 is a block diagram of an apparatus according to an example of an embodiment of the present invention.

2 is a diagram showing a schematic configuration of a transmission coil unit in the apparatus of FIG.

3 is a diagram showing a schematic configuration of a transmission coil unit in the apparatus of FIG.

4 is a diagram showing an electric circuit of a transmission coil unit in the apparatus of FIG.

5 is a diagram showing a current distribution in a transmission coil unit in the device of FIG.

6 is an R generated by a transmission coil unit in the apparatus of FIG.
It is a figure which shows F magnetic field.

7 is a diagram showing a current distribution in a transmission coil unit in the device of FIG.

8 is an R generated by a transmission coil unit in the apparatus of FIG.
It is a figure which shows F magnetic field.

9 is an R generated by a transmission coil unit in the apparatus of FIG.
It is a figure which shows the intensity distribution of F magnetic field.

10 is a diagram showing an example of a pattern of an electric path of a transmission coil unit in the apparatus of FIG.

11 is a cross-sectional view taken along the line AA of the electric path pattern shown in FIG.

FIG. 12 is a cross-sectional view taken along the line AA showing the electric path pattern shown in FIG. 10 together with the cross-sectional view of the RF shield.

FIG. 13 is a block diagram of an apparatus according to an example of an embodiment of the present invention.

14 is a diagram showing a schematic configuration of a body coil section in the apparatus of FIG.

FIG. 15 is a schematic view showing a body coil portion of the apparatus shown in FIG.
It is a figure which shows the schematic structure shown with a shield.

FIG. 16 is a schematic diagram showing an example of a pulse sequence for magnetic resonance imaging.

FIG. 17 is a schematic configuration diagram of a conventional example of a transmission coil unit.

FIG. 18 is an explanatory diagram of an RF magnetic field according to a conventional example of a transmission coil unit.

FIG. 19 is a schematic configuration diagram of a conventional example of a transmission coil unit.

[Explanation of symbols]

2 Static magnetic field generator 4,4 ', 41 Gradient coil section 6,6 'Transmission coil unit 61 body coil 8 subject 10 Imaging table 106 receiver coil unit 16 Gradient drive 18 Transmitter 20 Receiver 22 Analog-to-digital converter 24 computer 30 control unit 32 display 34 Operation part 602 to 606, 612, 614 electrical path 182,184 power splitter 608 capacitor

─────────────────────────────────────────────────── --Continued from the front page (56) References JP-A-8-280652 (JP, A) JP-A-7-255699 (JP, A) JP-A-9-117434 (JP, A) JP-A-6- 121779 (JP, A) (58) Fields surveyed (Int.Cl. ⁷ , DB name) A61B 5/055

Claims (4)

(57) [Claims] 1. A radial electric path through which a current for forming an RF magnetic field flows in a space accommodating a subject, and an end of the radial electric path is connected at a position farther from the space than the radial electric path. RF with annular electrical path
An RF coil pair in which two coils are paired and each of the two RF coils is arranged to face each other across the space; RF signal supply means for supplying RF signals having phases different from each other by 90 degrees to two points of the annular electric path having directions different from each other by 90 degrees as viewed from the center, the annular signal being supplied with the RF signals. RF signal supply means in which two positions of the electric path are the same in the RF coil pair, and the phase relationship of the RF signals supplied to the respective RF coils is also the same in the RF coil pair. An RF magnetic field forming apparatus comprising: 2. The RF signal supply means includes a first power splitter for inputting one of two RF signals having phases different from each other by 90 degrees and a second power splitter for inputting the other of the two RF signals. Power splitters, each of the power splitters being
The RF magnetic field forming apparatus according to claim 1, wherein a signal is output to each of the RF coils. 3. A static magnetic field forming means for forming a static magnetic field in a space accommodating a subject, a gradient magnetic field forming means for forming a gradient magnetic field in the space, and an RF magnetic field forming means for forming an RF magnetic field in the space. A magnetic resonance imaging apparatus comprising: a measuring unit that measures a magnetic resonance signal from the space; and an image generating unit that generates an image based on the magnetic resonance signal measured by the measuring unit. Is the RF magnetic field forming device according to claim 1 or 2, wherein the static magnetic field forming means and the gradient magnetic field forming means are arranged to face each other across the space, From the inside of the device where the space is present to the outside thereof, R
A magnetic resonance imaging apparatus characterized in that an F magnetic field forming means, the gradient magnetic field forming means, and the static magnetic field forming means are arranged in this order. 4. The magnetic resonance imaging apparatus according to claim 3, wherein an RF is provided between the RF magnetic field forming means arranged on the inner side and the gradient magnetic field forming means arranged on the outer side.
A magnetic resonance imaging apparatus having a shield.