JPH08257007A - Magnetic resonance imaging device - Google Patents

Abstract

PURPOSE: To improve visibility of a reconstituted image by providing a magnetic resonance imaging device with the second gradient magnetic field applying means which is arranged so as to enclose a gradient magnetic coil and also has another gradient magnetic coil for applying the second gradient magnetic field on an area including a photographing site in magnetostatic field, and with its control means. CONSTITUTION: This magnetic resonance imaging device is provided with a magnet part for generating static magnetic field, a gradient magnetic field part for adding position information to the static magnetic field, a transmitting/receiving part for selective exciting and MR signal receiving, and a control/operating part for system control and reconstituting an image. The magnet part is provided with the magnet 1 and generates the static magnetic field Ho in Z axial direction of a cylindrical diagnostic space into which a testee P placed on a bed 2 is inserted. The gradient magnetic field part is provided with the first and the second gradient magnetic field coils 3, 4 by which two gradient magnetic field of different linear area are generated. Futhermore, the gradient magnetic field part is provided with a gradient magnetic field power source 5 and a gradient magnetic field sequencer 6. The gradient magnetic field sequencer 6 is provided with a computer and receives a command signal from a controller 7.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a magnetic resonance imaging apparatus (hereinafter referred to as MRI apparatus), and more particularly to an MRI apparatus using a small gradient magnetic field coil system.

[0002]

2. Description of the Related Art In an MRI apparatus, for the purpose of knowing the position of each spin, the component of the magnetic field in the Z direction (direction parallel to the static magnetic field) has a spatial primary gradient in three directions of x, y, and z. A gradient magnetic field coil composed of the windings of the system is used.
The gradient magnetic field is generated by, for example, a gradient magnetic field coil including a pair of parallel coils (loop coils) and four saddle (saddle type) coils.

By the way, as problems in the MRI apparatus, speeding up of imaging, high resolution, shortening of echo time, and the like are mentioned. In order to solve this problem, the intensity of the gradient magnetic field is increased, the gradient magnetic field is started up at high speed, or switching (switching of the gradient magnetic field) is performed.
Is considered to be performed at high speed.

One of the means for realizing such a high intensity of the gradient magnetic field, high-speed start-up, high-speed switching, etc. is a small size in which the diameter of the gradient magnetic field coil is made small and its axial length is shortened. An MRI apparatus using the gradient magnetic field coil system (hereinafter, also referred to as small G coil (system)) has been proposed. By using this small G coil system, the above-mentioned high-speed imaging can be realized without reinforcing the gradient magnetic field power supply.

[0005]

However, the small G
In the MRI apparatus using the coil system, the linearity of the gradient magnetic field at the peripheral portion of the imaging region (hereinafter, also referred to as an imaging region) deteriorates. Therefore, in the reconstructed image, a portion near the magnetic field central portion, slice, lead, Distortion has occurred in the phase encoding direction. In other words, the area that can be imaged without distortion is small, so the visibility of the reconstructed image deteriorates. Further, the deterioration of the linearity of the gradient magnetic field causes nonuniformity of the image depending on the position of the slice thickness.

Further, in the outside of the imaging area (hereinafter, the area outside the imaging area is referred to as the non-imaging area), the magnetic field strength is reduced or inverted, so that the signal from the non-imaging area is used to generate an image. It sometimes mixes up in the image as folding back. The influence of this folding back will be described with reference to FIG. According to FIG. 24, the aliasing of the signal from the non-imaging region of the small G coil system 20 is included in the excited spins, so that the obtained reconstructed image has the artifacts as shown in FIG. I got it. This artifact has been a major cause of deteriorating the visibility of the reconstructed image.

The present invention has been made in order to solve all of the above-mentioned problems, and a first object thereof is to prevent the occurrence of artifacts due to a folding signal from outside the imaging area,
This is to improve the visibility of the reconstructed image.

A second object of the present invention is to compensate for the deterioration of the linearity of the gradient magnetic field of the small G coil system, so that the distortion in the slice, the lead, the phase encode direction and the nonuniformity due to the position of the slice thickness. It is to obtain a reconstructed image having no property.

[0009]

In order to achieve the above object, the magnetic resonance imaging apparatus according to claim 1 is provided so as to cover a local imaging region of a subject placed in a static magnetic field, A first gradient magnetic field applying unit having a gradient magnetic field coil for applying a first gradient magnetic field for slicing, phase encoding, and reading to the imaging region;
A high-frequency magnetic field applying unit for applying a high-frequency magnetic field of a required frequency to a region including the imaging region, and magnetic resonance from the imaging region based on the gradient magnetic field and the high-frequency magnetic field applied along the acquisition sequence of image data. A magnetic resonance imaging apparatus configured to collect signals and reconstruct an image based on the magnetic resonance signals, the magnetic resonance imaging apparatus being arranged so as to include the gradient magnetic field coil in the static magnetic field, and including the imaging region. Second gradient magnetic field applying means having another gradient magnetic field coil for applying the second gradient magnetic field to the region, and controlling the application of the first gradient magnetic field, the second gradient magnetic field, and the high frequency magnetic field. And a control means.

Particularly, in the magnetic resonance imaging apparatus according to the second aspect, the control means applies the gradient magnetic field pulse for slicing in the acquisition sequence by the second gradient magnetic field application means and the phase in the acquisition sequence. The gradient magnetic field pulse for encoding and the gradient magnetic field pulse for reading are applied by the first gradient magnetic field applying means.

Further, in particular, in the magnetic resonance imaging apparatus of the third aspect, the acquisition sequence is a spin echo method.

Further, in the magnetic resonance imaging apparatus according to the fourth aspect, the acquisition sequence is a gradient field echo method.

Furthermore, in the magnetic resonance imaging apparatus according to a fifth aspect of the present invention, the control means performs saturation excitation means for performing a first sequence for saturation excitation of a region adjacent to the imaging region, and a second imaging means. And signal collecting means for performing the sequence.

Particularly, in the magnetic resonance imaging apparatus according to the sixth aspect, the control means is adapted to automatically and continuously execute the second sequence after the end of the first sequence.

Further, in particular, in a magnetic resonance imaging apparatus according to a seventh aspect, the saturation excitation means is the second
Means for applying a predetermined gradient magnetic field pulse by the gradient magnetic field applying means, and means for applying a high frequency magnetic field pulse of a frequency corresponding to the adjacent region by the high frequency magnetic field applying means in a state where the gradient magnetic field pulse is applied. I have it.

Further, in the magnetic resonance imaging apparatus described in claim 8, the saturation excitation means applies a predetermined gradient magnetic field pulse with a predetermined polarity by the first gradient magnetic field applying means, and the gradient magnetic field thereof. A means for applying a predetermined gradient magnetic field pulse with a polarity opposite to the polarity by the second gradient magnetic field applying means is provided in synchronism with the pulse application timing.

Further, in order to achieve the above object, in a magnetic resonance imaging apparatus according to a ninth aspect, the magnetic resonance imaging apparatus is provided so as to cover a local imaging region of a subject placed in a static magnetic field, and the imaging region is provided. X, which applies various gradient magnetic fields including slice, phase encode, and read,
The gradient magnetic field applying means having y and z coils, and the high frequency magnetic field applying means for applying a high frequency magnetic field of a required frequency to a region including the imaging region, the gradient magnetic field being applied along an image data acquisition sequence. And a magnetic resonance imaging apparatus that collects a magnetic resonance signal from the imaging region based on a high-frequency magnetic field and reconstructs an image based on the magnetic resonance signal, and executes an imaging sequence in the acquisition sequence. Previously, it is provided with a saturation excitation current supply means for supplying a saturation excitation current for previously saturation-exciting a region adjacent to the imaging region to the x, y and z coils.

Particularly, in the magnetic resonance imaging apparatus according to the tenth aspect, the z-coil is a pair of coils arranged to face each other along the z-direction parallel to the direction of the magnetic flux of the static magnetic field, and the y-coil is: A plurality of pairs of coils arranged to face each other along ay direction orthogonal to the z direction, wherein the x coils are arranged to face each other along an x direction orthogonal to the z direction and the y direction. And a supply means for supplying a current for applying a gradient magnetic field for slicing, phase encoding, and reading to each of the pair coils, while the saturation excitation current supply means includes a z-coil. One of the paired coils forming the y coil, one of the plurality of paired coils forming the y coil, and the plurality of paired coils forming the x coil. Le of at least one orientation of the saturation excitation current supplied to the coil of the one, is set to be in the opposite direction to the direction of current for the gradient magnetic field application flow to one of the Peakoiru.

Further, in order to achieve the above object, claim 1
The magnetic resonance imaging apparatus described in 1 is provided so as to cover a local imaging region of a subject placed in a static magnetic field, and various types including slicing, phase encoding, and reading with respect to the imaging region. And a gradient magnetic field applying means having an imaging x, y, z coil for applying the gradient magnetic field, and a high frequency magnetic field applying means for applying a high frequency magnetic field of a required frequency to a region including the imaging region, and collecting image data. A magnetic resonance imaging apparatus configured to collect magnetic resonance signals from the imaging region based on the gradient magnetic field and the high frequency magnetic field applied along a sequence, and reconstruct an image based on the magnetic resonance signals, Saturation excitation x, y, and z coils for pre-saturating a region adjacent to the imaging region are arranged around the imaging region, and the excitation x, The saturation excitation current to the z coil comprises a saturation excitation current supplying means for supplying before imaging sequence is executed in the acquisition sequence.

Further, in the magnetic resonance imaging apparatus according to the twelfth aspect, the z-coil is a pair of coils arranged opposite to each other along the z-direction parallel to the direction of the magnetic flux of the static magnetic field, and the y-coil is: A plurality of pairs of coils arranged to face each other along ay direction orthogonal to the z direction, wherein the x coils are arranged to face each other along an x direction orthogonal to the z direction and the y direction. And the saturation excitation current supply means includes one of a pair of coils forming the saturation excitation z-coil, one of a plurality of pairs of coils forming the saturation excitation y-coil, and the x-coil coil. The direction of the current supplied to at least one coil of one of the plurality of paired coils constituting the coil corresponds to the execution of the imaging sequence. Image for x, y, is set to be in the opposite direction to the direction of the current flowing in one of Peakoiru constituting the z-coil.

Further, in a magnetic resonance imaging apparatus according to a thirteenth aspect, a first coil group including the imaging x, y and z coils and a second coil group including the saturation excitation x, y and z coils. The coil group and the coil group are arranged so as to be concentric in the xy cross section.

[0022]

In the magnetic resonance imaging apparatus according to any one of claims 1 to 4, there is provided a first gradient magnetic field coil provided so as to cover a local imaging region of a subject placed in a static magnetic field.
A first gradient magnetic field applied by the gradient magnetic field applying means of
The second gradient magnetic field applied by the second gradient magnetic field applying means having another gradient magnetic field coil arranged so as to include the gradient magnetic field coil, and the high frequency magnetic field applying means required for the region including the imaging region. The control unit controls the application (state) including the application timing, the application intensity, the application polarity, and the like of the high-frequency magnetic field applied at the frequency.

For example, under the control of the control means, the gradient magnetic field pulse for slicing in the image data acquisition sequence is applied by the second gradient magnetic field application means, and the gradient magnetic field pulse for phase encoding and the gradient magnetic field for reading in the acquisition sequence. Since the pulse is applied by the first gradient magnetic field applying means, the linear gradient region of the gradient magnetic field in the slice direction becomes wider than the gradient magnetic field by the first gradient magnetic field applying means. Therefore, the linearity of the gradient magnetic field does not deteriorate at the edge of the imaging region.

Further, in the magnetic resonance imaging apparatus according to any one of claims 5 to 7, a predetermined gradient magnetic field pulse is applied by the second gradient magnetic field applying means by the saturation excitation means which is one of the constituent elements of the control means, In the state in which the gradient magnetic field pulse is applied, the high-frequency magnetic field applying means executes the first sequence in which the high-frequency magnetic field pulse having the frequency corresponding to the area adjacent to the imaging area is executed. After the end, the signal collecting means automatically executes the second sequence for imaging. That is, the second sequence for imaging is executed after the region adjacent to the imaging region is saturated and excited by the first sequence.

Further, in the magnetic resonance imaging apparatus according to the present invention, a predetermined gradient magnetic field pulse is applied with a predetermined polarity by the first gradient magnetic field applying means during execution of the second sequence for imaging. In synchronization with the application timing of the gradient magnetic field pulse, the second gradient magnetic field applying means applies a predetermined gradient magnetic field pulse with a polarity opposite to the polarity. As a result of applying the gradient magnetic fields of opposite polarities, the gradient magnetic field of the imaging region becomes substantially constant, while the gradient magnetic field of a certain gradient is generated in the region adjacent to the imaging region, so that part is saturated with excitation. To be done.

Furthermore, in the magnetic resonance imaging apparatus according to the ninth to tenth aspects, x, y, which are provided so as to cover a local imaging region of the subject placed in a static magnetic field.
The z-coil of the z-coil is composed of paired coils arranged opposite to each other along the z-direction parallel to the direction of the magnetic flux of the static magnetic field, and the y-coils are arranged opposite to each other along the y-direction orthogonal to the z-direction. And a plurality of pairs of coils, wherein the x-coil is composed of a plurality of pairs of coils arranged to face each other along an x-direction orthogonal to the z-direction and the y-direction. A current for applying a gradient magnetic field for phase encoding and reading is supplied.

Before the imaging sequence in the acquisition sequence is executed, one of the pair coils forming the z coil, one of the plurality of pair coils forming the y coil, and the x coil are operated by the saturation excitation current supply means. Supply so that the direction of the saturation excitation current that flows in at least one of the paired coils of the pair of coils that constitutes the above is opposite to the direction of the gradient magnetic field application current that flows in one of the paired coils. Therefore, the magnetic field applied from the entire coil pre-saturates and excites the region adjacent to the imaging region.

In the magnetic resonance imaging apparatus according to the eleventh to thirteenth aspects, the saturation excitation x, y, and z coils are arranged around the imaging region in order to preliminarily saturate the adjacent region of the imaging region. . X for this saturation excitation,
Of the y and z coils, the z coil is a pair of coils arranged opposite to each other along the z direction parallel to the direction of the magnetic flux of the static magnetic field, and the y coils are opposite to each other along the y direction orthogonal to the z direction. Consisting of multiple pairs of coils arranged,
The x-coil is composed of a plurality of pairs of coils that are arranged so as to face each other along an x-direction that is orthogonal to the z-direction and the y-direction.

Before the imaging sequence in the acquisition sequence is executed, one of the paired coils forming the saturated excitation z coil and the plurality of paired coils forming the saturated excitation y coil are provided by the saturation excitation current supply means. And the direction of the current flowing through at least one of the plurality of pairs of coils forming the saturation excitation x-coil constitutes the imaging x, y, z coils corresponding to the execution of the imaging sequence. Since the electric current is supplied in the opposite direction to the direction of the current flowing through one of the pair coils, the magnetic field applied from the entire coil pre-saturates and excites the region adjacent to the imaging region.

[0030]

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

(First Embodiment) FIG. 1 shows the schematic arrangement of a magnetic resonance imaging apparatus according to the first embodiment. This magnetic resonance imaging apparatus includes a magnet unit for generating a static magnetic field, a gradient magnetic field unit for adding position information to the static magnetic field, a transmitter / receiver unit for selective excitation and MR signal reception, system control and image reconstruction. It is equipped with a control / arithmetic unit.

The magnet section is provided with, for example, a superconducting magnet 1, and is adapted to generate a static magnetic field H ₀ in the Z-axis direction of the cylindrical diagnostic space into which the subject P placed on the bed 2 is inserted. ing.

The gradient magnetic field section generates a first gradient magnetic field coil 3 and a second gradient magnetic field coil 4 which generate two gradient magnetic fields having different linear regions in the respective X, Y and Z axis directions shown in the figure.
It has. Among them, the first gradient magnetic field coil 3 is a small-sized gradient magnetic field coil system in which the diameter of the coil is reduced and the axial length thereof is shortened, and the second gradient magnetic field coil 4 is used for normal whole body imaging. For use.

That is, as shown in FIG. 2, the gradient magnetic field coils 3 are three sets (three channels) of the gradient magnetic field coils 3x to 3z (hereinafter referred to as small G) incorporated in the magnet 1. Coils 3x to 3z). The small G coil 3x is provided with two pairs of coils 3x and 3x. The paired coils 3x, 3x are composed of a pair of coils 3x, 3x arranged with the diagnostic space in between. In addition, G
Similarly, the coil 3y is composed of two pairs of coils 3y and 3y, and the G coil 3z is composed of a pair of circular coils 3z and 3z. A part of the small G coil 3 is embedded in the bed 2 as shown in FIG.

The gradient coil 4 is, as shown in FIG.
Similarly, three sets (3 channels) of gradient magnetic field coils 4x to 4z in the X-, Y-, and Z-axis directions, which are also incorporated in the magnet 1, are provided. The gradient magnetic field coil 4x includes two paired coils 4x,
It has 4x. The pair coil is composed of a pair of coils 4x, 4x which are arranged so as to sandwich the diagnostic space. In addition,
Similarly, the G coil 4y is also made up of two pairs of pair coils 4y and 4y, and the G coil 4z is made up of a pair of circular pair coils 4z and 4z.

Of the gradient magnetic fields generated by the small G coils 3x to 3z and the gradient magnetic field coils 4x to 4z, the magnetic field distribution in the Z-axis direction formed by the small G coils 3z and 3z and the gradient magnetic field coils 4z and 4z is shown. 4 shows. In addition, small G
The solid line represents the magnetic field distribution in the Z-axis direction generated by the coils 3z and 3z, and the broken line represents the magnetic field distribution in the Z-axis direction generated by the gradient magnetic field coils 4z and 4z. This Figure 4
According to the above, the gradient magnetic field distribution generated by the gradient magnetic field coils 4z, 4z is different from the gradient magnetic field distribution generated by the small G coils 3z, 3z, and its linear region is the linear region of the small G coils 3z, 3z. It is wider than the linear region of the magnetic field distribution. This relationship is related to the small G coil 3x,
Magnetic field distribution generated by 3x and gradient magnetic field coil 4x,
Magnetic field distribution generated by 4x and small G coil 3
Magnetic field distribution generated by y, 3y and gradient magnetic field coil 4
The same applies to the magnetic field distribution generated by y and 4y. The imaging region (imaging region) is determined according to the linear gradient region of the magnetic field distribution by the small G coils 3x to 3z.

Further, the gradient magnetic field section is composed of a gradient magnetic field power source 5
(5a to 5f) and the gradient magnetic field sequencer 6 are provided. Of the gradient magnetic field power supplies 5a to 5f, the gradient magnetic field power supply 5
a to 5c are respectively connected to the small G coils 3x to 3z, and the gradient magnetic field power supplies 5d to 5f are respectively connected to the linear gradient magnetic field coils 4x to 4z.
Electric current is supplied to 3z and the gradient magnetic field coils 4x to 4z. The number of the gradient magnetic field power supplies 5a to 5f can be increased or decreased according to the number of channels of the small G coil 3.

The gradient magnetic field sequencer 6 has a computer, and receives a signal instructing a desired data acquisition sequence from a controller 7 (having a computer) for controlling the entire apparatus. Then, the gradient magnetic field sequencer 5 receives the signal sent from the controller 7 to individually control the currents sent to the respective gradient magnetic field power supplies 5a to 5f in accordance with the commanded sequence, and the small G coils 3x to 3z. The linear gradient magnetic field coils 4x to 4z can be individually controlled. That is, the gradient magnetic field based on the small G coils 3x to 3z and the gradient magnetic field based on the linear gradient magnetic field coils 4x to 4z can be simultaneously generated, or can be generated at independent timings. In this way, the gradient magnetic field sequencer 5 controls the application and strength of each gradient magnetic field in the X, Y and Z axis directions. The generated gradient magnetic field is superimposed on the static magnetic field H ₀ .

Further, the gradient magnetic field sequencer 5 can stop the operation of the gradient magnetic field power supplies 5a to 5f which are not used, that is, make the output current non-flow, in order to avoid the abnormal operation due to the coupling, or the channel is not used. It is also possible to perform control so as to avoid current waveform interference due to coupling between the two.

It should be noted that, in this embodiment, 3 which are orthogonal to each other are used.
The gradient magnetic field in the Z-axis direction within the axis is the gradient magnetic field G _s for slicing.
Let it be the read (read) gradient magnetic field G _r in the X-axis direction, and let it be the phase encoding gradient magnetic field G _{e in} the Y-axis direction. Of the slicing gradient magnetic fields G _s , the gradient magnetic field in the Z-axis direction based on the small G coils 3z and 3z is defined as G _s1, and the gradient magnetic field in the Z-axis direction based on the linear gradient magnetic field coils 4z and 4z is defined as G _s2. And so on,
A gradient magnetic field G _r1 for reading the gradient magnetic field in the X-axis direction based on the small G coils 3x, 3x, and linear gradient magnetic field coils 4x, 4
The gradient magnetic field G for reading the X-axis gradient magnetic field based on x
_r2 , the gradient magnetic field in the Y-axis direction based on the small G coils 3y and 3y is a gradient magnetic field for phase encoding G _e1 , and the gradient magnetic field in the Y-axis direction based on the linear gradient magnetic field coils 4y and 4y is a gradient magnetic field for phase encoding G _r2 . .

On the other hand, the transmitter / receiver unit is a radio frequency coil (RF coil) 8 arranged near the subject P in the diagnostic space in the magnet 1, a transmitter 9T and a receiver connected to the RF coil 8. 9R, and an RF sequencer 10 (having a computer) that controls the operation timings of the transmitter 9T and the receiver 9R and the transmission frequency transmitted from the transmitter 9T to the RF coil 8. This transmitter 9T and receiver 9
Under the control of the RF sequencer 10, the R supplies the RF coil 8 with an RF current pulse of a required frequency for exciting nuclear magnetic resonance (NMR) in spins of a required portion in the subject P, while RF The MR signal (nuclear magnetic resonance signal, high frequency signal) received by the coil 8 is subjected to various kinds of signal processing to form a digital signal.

Further, in addition to the controller 7 described above, the control / arithmetic unit receives the digital signal formed by the receiver 9R and performs arithmetic processing such as Fourier transform to generate reconstructed image data. A storage unit 12 for storing the generated image data, a display 13 for displaying an image (reconstructed image), and an input device 14 for data input. Further, the controller 7 controls the operation contents and operation timing of the gradient magnetic field sequencer 6 and the RF sequencer 10 while synchronizing them.

Next, the overall operation of this embodiment will be described.

For example, when diagnosing the head of the subject P, the operator positions the subject P so that the head is located inside the small G coil 3.
Then, when the positioning is completed, imaging is performed.

That is, the gradient magnetic field sequencer 6 and RF
When the controller 7 commands the acquisition sequence based on the execution of imaging, the sequencer 10 controls application of a gradient magnetic field to the subject P and transmission / reception of high-frequency signals according to the sequence. In addition, here, SE (Spin E
It is assumed that the pulse sequence according to the cho) method is commanded as shown in FIG.

That is, the gradient magnetic field sequencer 6 uses the linear gradient magnetic field coil 4 from the gradient magnetic field power source 5f to selectively excite the positioning slice position in the Z direction in the imaging region.
The slicing gradient magnetic field G _s2 is applied for a certain period of time via z and 4z. In addition, the small G coil 3 corresponding to the inside of the imaging region
The point of this embodiment is that the linear gradient magnetic field coils 4z and 4z are used instead of z and 3z.

Then, the RF sequencer 10 applies the slice gradient magnetic field G _s2 to the frequency band (center frequency, frequency width) that determines the slice position and slice thickness from the transmitter 9T via the RF coil 8. ) With R
Send an F pulse. As a result, the slice plane of the position and thickness corresponding to the frequency band of the RF pulse is excited. Since this RF pulse has a flip angle of 90 degrees, it is called a π / 2 pulse.

Next, the gradient magnetic field sequencer 6 outputs the read gradient magnetic field G _r1 from the gradient magnetic field power source 5a through the small G coils 3x and 3x for a certain period at the timing when the application of the slice gradient magnetic field G _s2 is completed. Apply. After the application is completed, the slice gradient magnetic field G _s2 is applied again from the gradient magnetic field power source 5f via the linear gradient magnetic field coils 4z and 4z.

Then, the RF sequencer 10 has the same frequency band as the π / 2 pulse from the transmitter 9T via the RF coil 8 when TE / 2 has elapsed since the π / 2 pulse was transmitted. Π with a flip angle of 180 degrees (π)
A pulse is transmitted to excite the same slice plane.

In this state, the gradient magnetic field sequencer 6 sends the gradient magnetic field power source 5b through the small G coils 3y and 3y.
The phase-encoding gradient magnetic field G _e1 is applied, and then the readout gradient magnetic field G _r1 is applied from the gradient magnetic field power source 5a through the small G coils 3x and 3x.

As a result, after transmitting the π / 2 pulse, T
When E has passed, an MR signal is generated from the slice plane selectively excited. This MR signal is received via the RF coil 8 and collected as frequency-encoded and phase-encoded pixel data.

By executing these series of sequences while changing the phase encoding gradient magnetic field G _r2 , two-dimensional image data is collected.

That is, in this embodiment, the linear gradient magnetic field coil 4z is used as the slice gradient magnetic field G _s , not the gradient magnetic field G _s1 based on the small G coils 3z and 3z adapted to the imaging region (inside the small G coil 3). , 4z based gradient magnetic field G _s2 is applied. And the gradient magnetic field sequencer 6
The gradient G _r, as G _e, both small G coils 3x, applies a 3x and 3y, gradient G _r1 and G _e1 based on 3y.

That is, the gradient magnetic field G _s for slice selective excitation, which does not require a relatively high speed for rising, is a gradient magnetic field G _s2 based on the linear gradient magnetic field coils 4z, 4z, that is,
Since the gradient magnetic field is wide in the linear region, the linearity of the gradient magnetic field in the imaging region (inside the small G coil 3) does not deteriorate. Therefore, the spatial nonuniformity of the slice thickness and the non-linearity of the slice shift amount are improved as compared with the conventional case.

The present embodiment can be applied not only to the SE method but also to other imaging methods for performing selective excitation of slices.

For example, FIG. 6 shows a pulse sequence by the FE (Field Echo) method. According to FIG. 6, similarly to the SE method, in order to selectively excite the positioning slice position in the Z direction in the imaging region, the gradient magnetic field sequencer 6 receives the gradient magnetic field power source 5f via the linear gradient magnetic field coils 4z and 4z. ,
The slicing gradient magnetic field G _s2 is applied for a certain period.

Then, the RF sequencer 10 applies the slice gradient magnetic field G _s2 to the frequency band (center frequency, frequency width) that determines the slice position and slice thickness from the transmitter 9T via the RF coil 8. ) With R
Transmit F pulse (flip angle α ° <90 °).

In this state, the gradient magnetic field sequencer 6 outputs the gradient magnetic field power source 5b through the small G coils 3y and 3y.
While applying the phase encoding gradient magnetic field G _r2 , a negative readout gradient magnetic field G _r1 is applied for a certain period of time from the gradient magnetic field power source 5a via the small G coils 3x and 3x.

Subsequently, the gradient magnetic field sequencer 6 applies a positive reading gradient magnetic field G _r1 , that is, a gradient magnetic field G _r1 with reversed polarity from the gradient magnetic field power source 5a through the small G coils 3x, 3x for a certain period. To do. As a result, an MR signal is generated from the slice plane selectively excited when TE has elapsed after transmitting the RF pulse. That is, the MR signal is obtained by inverting the gradient magnetic field G _r1 without using the π pulse.

The resulting MR signal is received via the RF coil 8 and collected as frequency-encoded and phase-encoded pixel data.

By executing these series of sequences while changing the phase encoding gradient magnetic field G _r2 , two-dimensional image data is collected.

Also in the FE method described above, since the slice plane is excited by the gradient magnetic field having a wide linear region, the same effect as the SE method described above can be obtained.

(Second Embodiment) The hardware structure of the second embodiment is the same as that of the first embodiment, and the description thereof is omitted.

Here, again, the magnetic field distribution in the Z-axis direction of the small G coil 3z in the configuration of FIG. 1 is shown in FIG. 7A, and the magnetic field distribution in the z-axis direction of the linear gradient magnetic field coil 4z is shown in FIG. FIG. 7C shows the frequency distribution of the saturation excitation RF pulses P1 and P2. This FIG.
In (a), the point “a” is the non-imaging area A on the low frequency side.
Although included in out1, since the point “b” in the imaging area Ain and the z component of the magnetic field are equal, the signal (unnecessary signal) generated from the spin at the point “a” is represented by the point “on the reconstructed image”. The signal of "b" corresponds to the same point and overlaps (cause of the artifact described in the related art). It should be noted that this phenomenon is also a phenomenon that occurs between the non-imaging area Aout2 and the imaging area Ain on the high frequency side.

The present embodiment is made to reduce the overlap of the unnecessary signals (images). Here, the pulse sequence (FE method) of the present embodiment is shown in FIG.

In this pulse sequence, before the imaging sequence (described in the first embodiment) in the pulse sequence of the normal FE method is executed, the non-imaging area A described above is used.
A saturation excitation (presaturation) sequence is executed in advance for out1 and the non-imaging area Aout2.

That is, the gradient magnetic field sequencer 6 applies the slice gradient magnetic field G _s2 from the gradient magnetic field power supply 5f via the linear gradient magnetic field coils 4z, 4z for a certain period. Note that this slicing gradient magnetic field G _s2 has a linear gradient magnetic field region including an imaging region Ain, a non-imaging region Aout1, and a non-imaging region Aout2, as shown in FIG. 7B, as in the first embodiment. ing.

In this state, the RF sequencer 10 first saturates the outside imaging area Aout1 on the low frequency side,
As shown in FIG. 7C, the transmitter 9T is controlled to control the frequency band (center frequency f _l , which corresponds to the non-imaging area Aout1).
An RF pulse P1 having a frequency width ± Δf) is generated. The RF pulse P1 is transmitted to the subject P via the RF coil 8 under the control of the RF sequencer 10.
As a result, the non-imaging area Aout1 is saturated and excited by the RF pulse P1. The state at this time is shown in FIG. 9 (Aout1 (hatched portion) in the figure is saturated and excited).

Subsequently, the RF sequencer 10 also transmits the transmitter 9T in order to saturate and excite the non-imaging region Aout2 on the high frequency side in the state where the slice gradient magnetic field G _s2 is also applied.
Is controlled to generate an RF pulse P2 having a frequency band (center frequency f _h , frequency width ± Δf) corresponding to the non-imaging area Aout2. The RF pulse P2 is transmitted to the subject P via the RF coil 8 under the control of the RF sequencer 10. As a result, the non-imaging area Aout2 is RF
It is saturatedly excited by the pulse P2. The state at this time is shown in FIG. 9 (Aout2 (hatched portion) in the figure is saturated and excited).

After saturating and exciting the non-imaging regions Aout1 and Aout2, the gradient magnetic field sequencer 6 applies the phase encoding gradient magnetic field G _r1 from the gradient magnetic field power source 5b through the small G coils 3y and 3y for a certain period, At the same time, the read gradient magnetic field G _r1 is applied from the gradient magnetic field power source 5a through the small G coils 3x, 3x for a certain period.

When the application is finished, this time,
A slicing gradient magnetic field G _s1 is applied for a certain period from the gradient magnetic field power source 5c through the linear gradient magnetic field coils 3z and 3z, and R
The F sequencer 10 transmits an RF pulse via the RF coil 8. Hereinafter, the above-mentioned FE method pulse sequence is repeated.

At this time, since the non-imaging areas Aout1 and Aout2 are saturated and excited, the generation of signals from this portion is drastically reduced. Therefore, in the MR signal obtained by the RF coil 8, the above-described phenomenon of overlapping of unnecessary signals is hardly seen.

Therefore, the image I1 based on the MR signal obtained as a result is compared with the image I2 based on the MR signal which is not saturated and excited, as shown in FIG.
The artifacts that had occurred in the are not seen, and the visibility is greatly improved.

In the present embodiment, the case where the non-imaging area is selectively saturated and excited is shown. However, the method for saturation and exciting a certain specific area is, for example, for the purpose of blood flow artifact countermeasure. It can also be used to selectively saturate specific regions inside and outside the imaging region as well.

(Third Embodiment) The hardware configuration of the third embodiment is similar to that of the first embodiment, and the description thereof is omitted.

This embodiment is also an embodiment for reducing the overlap of unnecessary signals (images) described in the second embodiment.
Here, the pulse sequence (SE method) of this embodiment is shown in FIG.
It is shown in FIG.

This pulse sequence is substantially the same as the pulse sequence of the first embodiment until the π pulse is transmitted under the control of the RF sequencer 10 (however, the slicing gradient magnetic field G _s is small. It is assumed that a slice gradient magnetic field G _s1 based on the G coil 3z is applied).

Application of π pulse and gradient magnetic field G for slicing
_In a state where the application of _{s1 for} a certain period is completed and the slice plane is selectively excited, the gradient magnetic field sequencer 6 causes the gradient magnetic field power supply 5b to pass the linear gradient magnetic field coils 3y and 3y through the phase encoding gradient magnetic field G _e1. And a gradient magnetic field G _s1 (positive polarity) for slicing (this gradient magnetic field pulse is shown as G _s1 (A) in FIG. 11) from the gradient magnetic field power source 5 c via the linear gradient magnetic field coils 3 z and 3 z. It is applied for a certain period of time, and at the same time as the slicing gradient magnetic field G _s1 , from the gradient magnetic field power source 5 f via the linear gradient magnetic field coils 4 z and 4 z, the slicing gradient magnetic field G _s2 (negative polarity, that is, the polarity is inverted). This gradient magnetic field pulse is represented by G _s2 (A) in FIG.
Is applied for a certain period of time.

FIG. 12A shows the distribution of the Z-direction component of the magnetic field in this state (a state in which the slice gradient magnetic field G _s1 and the slice gradient magnetic field G _s2 are applied). FIG.
The solid line in (a) is the slice gradient magnetic field G _s1 , and the broken line is
It is a gradient magnetic field G _s2 for slicing. The magnetic field distribution of the phase encode gradient magnetic field G _e1 is omitted.

[0080]

[Outside 1]

As can be seen from FIG. 12C, the two gradient magnetic field pulses G _s1 (A) and G shown in FIG.
The gradient magnetic field of the synthetic magnetic field based on _s2 (A) is
It can be seen that there is almost no effect inside (the inclination is almost "0"), but in the non-imaging regions Aout1 and Aout2, a gradient magnetic field is applied, that is, the spoil effect is obtained. Therefore, when the pulse sequence as shown in FIG. 11 is used, the signal from the non-imaging region in the slice direction is drastically reduced, and the above-mentioned unnecessary signal (image) overlap is reduced. As a result, an image without artifacts can be obtained as in the second embodiment, and the visibility is improved. The sequence for generating the saturated excitation magnetic field shown in FIG. 8 and the sequence for generating the saturated excitation magnetic field shown in FIG. 11 can be used together.

The method of reducing the development in which the signal of the non-imaging region in the slice direction overlaps the image has been described above.
By changing _s2 to G _e2 or G _r2 , it is possible to reduce the similar overlapping phenomenon that occurs in the phase encode direction or the read direction. In addition, in FIG.
Two gradient magnetic field pulses G _s1 (A) and G _s2 (A) are applied in the Y-axis direction (gradient magnetic field pulses G _e1 (A) and G
_e2 (A)) or can be applied in the X-axis direction (as gradient magnetic field pulses G _r1 (A) and G _r2 (A)) to reduce the similar overlapping phenomenon occurring in the phase encode direction or the read direction. Can be made. Then, the gradient magnetic field pulses G _s1 (A), G _s2 (A), the gradient magnetic field pulses G _e1 (A), G _e2 (A) and the gradient magnetic field pulse G _{r1 described above.}
By simultaneously applying (A) and G _r2 (A), the overlapping phenomenon in all directions can be reduced. It should be noted that this relationship also holds in the second embodiment.

Further, changing G _s2 in FIG. 8 to G _e2 or G _r2, and the gradient magnetic field pulse G in FIG.
_e1 (A), G _e2 (A) or gradient magnetic field pulse G
_r1 (A) and G _r2 (A) are applied, that is, the phase encoding gradient magnetic field or the readout gradient magnetic field is reduced to a small size G
By generating both the coil 3 and the gradient magnetic field coil 4 in combination, data regarding the linear region of the gradient magnetic field based on the gradient magnetic field coil 4 and the linear region of the gradient magnetic field based on the small G coil 3 can be transmitted via the gradient magnetic field sequencer 6. Is obtained by the controller 7. Therefore, based on the above data, from the comparison between the non-linear region of the gradient magnetic field based on the small G coil 3 and the linear region of the gradient magnetic field based on the gradient magnetic field coil 4, the nonlinearity of the linear gradient magnetic field in the phase encode direction or the read direction is shown. It is possible to obtain data that compensates for In other words, it is possible to visualize the non-linearity and obtain data that compensates for the non-linearity. For example, the controller 7 sends this data to the arithmetic unit 11 as compensation data at the time of image reconstruction and is processed to correct the distortion in the phase encoding direction or the lead direction of the image obtained by the small G coil 3. It is also possible to do so.

On the other hand, the RF pulses P1 and P in FIG.
2 may be used as the selective saturation excitation pulses “tag1” and “tag2” in the imaging region for the purpose of tagging.

The state of this tagging is shown in FIGS. In the case of axial tomography, under the control of the RF sequencer 10, the selective saturation excitation pulse “t” is transmitted via the transmitter 9T.
By controlling the frequency bands of "ag1" and "tag2" to selectively saturate the slice planes corresponding to S1 and S2 in FIG. 14 to preliminarily create a low (no) signal band (tag), the FE When the imaging by the method is performed, low signal regions (tag) T1 and T2 appear in the obtained image as shown in FIG. By observing the partial movement amount of this tag, the velocity of blood flow and the like can be measured.

(Fourth Embodiment) In the hardware configuration of the fourth embodiment, in the configuration of the first embodiment, the gradient magnetic field coil 4 and the gradient magnetic field power source 5 (5d) connected to the coil 4 are used.
5f) is removed.

In addition, in the structure of this embodiment, the two pairs of small G coils 3x are the independent pair of loop coils 3x1, 3x3, and loop coils 3x2,
It is composed of 3x4. Similarly, the two pairs of small G coils 3y are the independent pair of loop coils 3y1 and 3y3 and loop coil 3y2.
, 3y4, and the pair coil 3z of the small G coil 3z is composed of a pair of independent loop coils 3z1 and 3z2.

As shown in FIG. 16, the pair of loop coils 3x1 and 3x3 are individually connected to the gradient magnetic field power sources 5x1 and 5x3, and the loop coils 3x2 and 3x4 are also individually gradient magnetic field power sources 5x2. , 5x
Connected to 4.

Similarly, a pair of loop coils 3y1 and 3y
3 are individually connected to the gradient magnetic field power supplies 5y1 and 5y3, and the loop coils 3y2 and 3y4 are also
The gradient magnetic field power supplies 5y2 and 5y4 are individually connected.
Further, the pair of loop coils 3z1 and 3z2 are individually connected to the gradient magnetic field power supplies 5z1 and 5z2.

Gradient magnetic field power source 5x1 to 5x4, 5y1 to 5
y4 and 5z1 to 5z2 are connected to the gradient magnetic field sequencer 6, respectively. And the gradient magnetic field sequencer 6
Is a pair coil 3 through the gradient magnetic field power supplies 5x1 and 5x3.
It is possible to independently control the direction and intensity of the current flowing in each of the x1 and 3x3 coils and the other coil. In addition, other pair coils 3x2, 3
x4, 3y1, 3y3, 3y2, 3y4 and 3z1, 3
z2 can be similarly controlled via the corresponding gradient magnetic field power supply.

Since other configurations are substantially the same as the hardware configuration of the MRI apparatus of the first embodiment shown in FIG. 1, the description thereof will be omitted.

This embodiment is also an embodiment for reducing the overlap of unnecessary signals (images) described in the second embodiment.

Here, the pulse sequence (F
Method E) is shown in FIG.

In this pulse sequence, before the imaging sequence in the normal FE method (described in the first embodiment) is executed, the above-mentioned non-imaging area Aout1 and non-imaging area Aout2 are saturated in advance (presaturation). You are running the sequence.

That is, the gradient magnetic field sequencer 6 includes gradient magnetic field power supplies 5x1 to 5x4, 5y1 to 5y4 and 5z1.
Controlling 5z2, small G coils 3x1 to 3x4, 3y1
To 3y4 and 3z1 to 3z2, respectively, to supply saturation excitation gradient magnetic fields G _r1 ′, G _e1 ′ and G _s1 ′ (these gradient magnetic fields (pulses are collectively referred to as G ′)) simultaneously for a certain period. Apply.

Each small G coil 3x1 to 3x4 at this time
The directions of the currents sent to 3y1 to 3y4 and 3z1 to 3z2 are shown in FIGS.

FIG. 18A shows the direction of the current flowing through the small G coils 3x and 3x for generating the normal read gradient magnetic field G _r1 , and FIG. 18B shows the saturation excitation gradient of this embodiment. Small G coil 3 for generating magnetic field G _r1 ′
The directions of the currents flowing in x and 3x are shown. The direction of the arrow in the figure represents the direction of current flow.

Similarly, FIG. 19A shows the direction of the current flowing through the small G coils 3y, 3y for generating the normal phase encoding gradient magnetic field G _e1 , and FIG. 18B shows the direction of this embodiment. FIG. 2 shows the directions of currents flowing through the small G coils 3y and 3y for generating the saturation excitation gradient magnetic field G _e1 ′.
0 (a) is the direction of the current flowing through the small G coils 3z and 3z for generating the normal slicing gradient magnetic field G _s1 .
FIG. 20B shows the gradient magnetic field for saturation excitation G _s1 ′ according to this embodiment.
The direction of the current flowing through the small G coils 3z and 3z for generating is shown.

As shown in FIGS. 18 to 20, the gradient magnetic field sequencer 6 simultaneously applies the saturation excitation gradient magnetic fields G _r1 ′, G _e1 ′ and G _s1 ′ to each gradient magnetic field power supply 5x.
1 to 5x4, 5y1 to 5y4 and 5z1 to 5z2 are controlled respectively, and the paired coils 3x1 and 3x3 (3
x2, 3x4, 3y1, 3y3, 3y2, 3y4, 3z
Current flowing in one of the coils (1, 3z2) of the normal gradient magnetic field (the gradient magnetic field for reading G _r1 , the gradient magnetic field for phase encoding G _e1 and the gradient magnetic field for slice G _s1 ) Allow it to flow in the opposite direction. The direction of the current flowing through the other coil is the same as the direction of the current flowing when generating a normal gradient magnetic field (reading gradient magnetic field G _r1 , phase encoding gradient magnetic field G _e1 and slice gradient magnetic field G _s1 ). Is.

As a result, the saturation excitation gradient magnetic fields G _r1 ′, G _e1 ′ and G _s1 ′ shown in FIGS. 21 and 22 are generated in the Z direction (slice direction) in the diagnostic space.

Saturation excitation gradient magnetic field G _r1 ′ shown in FIG.
The magnetic field strength distribution of {G _e1 ′ is also the same} (the magnetic field strength B _x (B _y ) in the imaging area Ain is maintained at a substantially constant strength in the small G coil, and the strength is maintained in the non-imaging areas Aout1 and Aout2. The intensity gradually decreases as the distance from the center of the imaging area Ain (the center of the coil) increases.
Similarly, in the magnetic field intensity distribution of the gradient magnetic field for saturation excitation G _s1 ′ shown in FIG. 22, the magnetic field intensity B _z in the imaging region Ain is maintained at a substantially constant intensity, and the intensity is imaged in the non-imaging regions Aout1 and Aout2. The strength gradually decreases with increasing distance from the center of the area Ain (center of the coil).

That is, the non-imaging area Aout1 in the slice direction
Therefore, a gradient magnetic field having a certain gradient is applied to Aout2.

Therefore, with the saturation excitation gradient magnetic fields G _r1 ′, G _e1 ′ and G _s1 ′ applied simultaneously,
The RF sequencer 10 uses the non-imaging area Aout1 to the imaging area A
When the RF pulse P3 in the frequency band including in to Aout2 is generated, the non-imaging areas Aout1 and Aout2 (spins thereof) are saturated and excited.

After saturating the non-imaging areas Aout1 and Aout2, the gradient magnetic field sequencer 6 operates the gradient magnetic field power supplies 5x.
The small G coils 3x1, 3x3, 3x2, 1x5x4, 5y1-5y4 and 5z1-5z2 are controlled respectively.
3x4, 3y1, 3y3, 3y2, 3y4, 3z1, 3
Through z2, the readout gradient magnetic field G _r1 , the phase encoding gradient magnetic field G _r1 and the slice gradient magnetic field G _s1 (all normal gradient magnetic fields) are simultaneously applied for a certain period (the direction of current to the coil is shown in FIG. a) to 20 (a)).

When the application is finished, this time,
From the gradient magnetic field power source 5c to the linear gradient magnetic field coils 3z1, 3z
The gradient magnetic field G _s1 for slicing is applied for a certain period of time via the RF coil 2, and the RF sequencer 10 passes the R
Send an F pulse. Hereinafter, the above-mentioned FE method pulse sequence is repeated.

At this time, since the non-imaging areas Aout1 and Aout2 are saturated and excited, the signal generation from this portion is drastically reduced. Therefore, in the MR signal obtained by the RF coil 8, the above-described phenomenon of overlapping of unnecessary signals is hardly seen.

Therefore, as shown in FIG. 10, the image I1 based on the MR signal obtained by the saturated excitation does not show the artifacts generated in the image I2 based on the MR signal not subjected to the saturated excitation. , The visibility is greatly improved.

Further, as described above, since the artifacts can be reduced, it is possible to avoid the occurrence of the artifacts, which was a matter of concern in making the small G coil smaller, so that the small G coil can be avoided. Can contribute to improving the performance of.

In this embodiment, a current in the direction opposite to the direction of the current for generating a normal linear gradient magnetic field is passed through one of the pair of small G coils to saturate the non-imaging region. However, the present invention is not limited to this, and for example, small G coils 3x1 to 3x4, 3y.
Small G coil 3 which is the same as 1 to 3y4 and 3z1 to 3z2
Another set of x1a to 3x4a, 3y1a to 3y4a and 3z1a to 3z2a (generally referred to as saturation excitation small G coils 3a) is prepared, and the small G coils 3a are placed close to the small G coils 3 and, for example, in the xy section. It is provided so as to be on the concentric circles (see FIG. 23). A current for generating a normal linear gradient magnetic field is passed through the small G coil 3 (see FIG. 18a).
(See FIG. 20a), the small G coil 3a for saturation excitation includes:
A current as shown in FIGS. 18b to 20 (b) described above may be supplied. That is, instead of controlling so as to change the direction of the current, another set of coils is provided to generate a magnetic field distribution as shown in FIG. 21 and FIG. May be.

In this embodiment, the case where the non-imaging region is selectively presaturated (pre-saturation excitation) is shown. However, the method of presaturating a certain region is aimed at, for example, blood flow artifact countermeasure. In such a case, it can also be used to selectively and selectively excite specific regions inside and outside the imaging region.

Further, in this embodiment, the gradient excitation coil 4 and the gradient magnetic field power source 5 connected to the coil 4 can be used to perform the above-described saturation excitation sequence as in the first embodiment. .

Although the first to fourth embodiments have been described above, these embodiments can be used in combination as much as possible. In particular, applying the slicing gradient magnetic field of the first embodiment by the gradient magnetic field coil 4 can be performed in all the other embodiments, and the spatial nonuniformity of the slice thickness and the slice nonuniformity described in the first embodiment. It is possible to improve the nonlinearity of the slice shift amount and reduce artifacts.

[0113]

As described above, according to the magnetic resonance imaging apparatus of the first to fourth aspects, the first gradient magnetic field having the gradient magnetic field coil provided so as to cover the local imaging region of the subject. A first gradient magnetic field applied by the applying means, a second gradient magnetic field applied by the second gradient magnetic field applying means having another gradient magnetic field coil arranged so as to include the gradient magnetic field coil, and Since the application of the high-frequency magnetic field can be controlled, the gradient magnetic field in the slice direction can be, for example, the second gradient magnetic field having a wider gradient (linearity of the gradient magnetic field) than the first gradient magnetic field. Therefore, the deterioration of the linearity of the gradient magnetic field, which is a problem in the case of using the first gradient magnetic field applying unit having the gradient magnetic field coil having a size that covers the local imaging region of the subject, can be eliminated. Therefore, image distortion due to deterioration of linearity at the peripheral portion of the imaging region, read, distortion in the encoding direction, and image nonuniformity due to slice thickness are suppressed. As a result, the image quality of the image is improved, and accurate image reading can be performed without extra burden on the image reader.

According to the magnetic resonance imaging apparatus of the fifth to thirteenth aspects, an area (non-imaging area) adjacent to an imaging region (imaging area) is
Alternatively, since saturation excitation (presaturation) can be performed in advance during execution, a signal that folds back from the non-imaging region and mixes into the imaging region is significantly suppressed. Therefore, the artifact generated in the reconstructed image due to the folding signal can be significantly reduced, and the visibility of the reconstructed image can be improved.

[Brief description of drawings]

FIG. 1 is a block diagram showing a schematic configuration of a magnetic resonance imaging apparatus according to first and second embodiments of the present invention.

FIG. 2 is a diagram showing a schematic configuration of a gradient magnetic field coil that constitutes the first gradient magnetic field coil (small G coil) in FIG.

FIG. 3 is a diagram showing a schematic configuration of a gradient magnetic field coil which constitutes a second gradient magnetic field coil in FIG.

FIG. 4 is a diagram illustrating gradient magnetic field distributions generated by a first gradient magnetic field coil and a second gradient magnetic field coil.

FIG. 5 is a sequence diagram of the spin echo method in which the linearity of the gradient magnetic field distribution in the slice direction is improved.

FIG. 6 is a sequence diagram of the field echo method in which the linearity of the gradient magnetic field distribution in the slice direction is improved.

7A is a diagram showing a magnetic field distribution in the Z-axis direction of a small G coil, FIG. 7B is a diagram showing a magnetic field distribution in the Z-axis direction of a linear gradient magnetic field coil, and FIG. Of the frequency distribution of the pulse for use.

FIG. 8 is a sequence diagram of the field echo method when the non-imaging region is saturated and excited in advance.

FIG. 9 is a diagram illustrating a state when the non-imaging region is saturated and excited.

FIG. 10 is a view showing a comparison between a reconstructed image in a case where the non-imaging region is not saturated with excitation and a reconstructed image in the case where saturated excitation is performed.

FIG. 11 is a sequence diagram of a spin echo method when the non-imaging region is saturated and excited during sequence execution.

12A is a diagram showing gradient magnetic field distributions of a gradient magnetic field pulse G _s1 (A) and a gradient magnetic field pulse G _s2 (A) (solid line → gradient magnetic field pulse G _s1 (A), broken line → gradient magnetic field pulse). G
_s2 (A)) and (b) are diagrams showing magnetic field distributions when the respective gradient magnetic fields of the gradient magnetic field pulse G _s1 (A) and the gradient magnetic field pulse G _s2 (A) are synthesized, and (c) is a diagram The figure which shows the magnetic field gradient distribution immediately after applying the gradient magnetic field pulse _Gs1 (A) and the gradient magnetic field pulse _Gs2 (A) at the time of performing the sequence of FIG.

FIG. 13 is a sequence diagram of a spin echo method for the purpose of tagging.

FIG. 14 is a diagram illustrating a tagging state.

FIG. 15 is a diagram showing a reconstructed image when tagging is performed.

FIG. 16 is a diagram showing a schematic configuration of a small G coil according to a fourth embodiment.

FIG. 17 is a sequence diagram of the field echo method when the non-imaging region is saturatedly excited in advance.

FIG. 18 (a) is a diagram for explaining the directions of currents flowing in the loop coils 3x1, 3x3 and the loop coils 3x2, 3x4 when a normal read gradient magnetic field is generated, and FIG. When generating the gradient magnetic field for excitation, the loop coils 3x1, 3x3 and the loop coils 3x2, 3
It is a figure explaining the direction of the electric current which flows into x4.

19 (a) is a diagram for explaining the directions of currents flowing through the loop coils 3y1, 3y3 and the loop coils 3y2, 3y4 when a normal phase-encoding gradient magnetic field is generated, and FIG. Loop coils 3y1, 3y3, and loop coil 3y when generating a gradient magnetic field for saturation excitation
It is a figure explaining the direction of the electric current which flows into 2, 3y4.

FIG. 20 (a) is a diagram for explaining the directions of currents flowing through the loop coils 3z1 and 33z2 when generating a normal slice gradient magnetic field, and FIG. 20 (b) shows a saturation excitation gradient magnetic field. It is a figure explaining the direction of the electric current which flows into loop coils 3z1 and 3z2 at the time.

FIG. 21 shows loop coils 3x1 and 3x3, and loop coils 3x2 and 3x4 (or loop coils 3y1 and 3).
FIG. 3 is a diagram showing a gradient magnetic field distribution for saturation excitation generated by y3 and loop coils 3y2, 3y4).

FIG. 22 is a diagram showing a gradient magnetic field distribution for saturation excitation generated by the loop coils 3z1 and 3z2.

FIG. 23 is a diagram showing a schematic configuration of a magnetic resonance imaging apparatus when a small G coil 3a that generates a gradient magnetic field for saturation excitation is provided.

FIG. 24 is a diagram showing artifacts that occur in a reconstructed image when a conventional small G coil is used.

[Explanation of reference numerals] 1 magnet 2 bed 3 first gradient magnetic field coil 3x to 3z small G coil 3x1 to 3x4 loop coil 3y1 to 3y4 loop coil 3z1 to 3z2 loop coil 3a small G coil for saturation excitation 4 second gradient magnetic field Coil 4x-4z Gradient magnetic field coil 5a-5f Gradient magnetic field power supply 5x1-5x4 Gradient magnetic field power supply 5y1-5y4 Gradient magnetic field power supply 5z1-5z2 Gradient magnetic field power supply 6 Gradient magnetic field sequencer 7 Controller 9R receiver 9T transmitter 10 RF sequencer 11 Arithmetic unit 12 Memory unit 13 Display unit 14 Input unit

Claims (13)

[Claims] 1. A first gradient magnetic field which is provided so as to cover a local imaging region of a subject placed in a static magnetic field and includes a slice magnetic field, a phase encoding magnetic field, and a readout magnetic field for the imaging region. First gradient magnetic field applying means having a gradient magnetic field coil to be applied, and high frequency magnetic field applying means for applying a high frequency magnetic field of a required frequency to a region including the imaged region are provided, and the high frequency magnetic field is applied along an image data acquisition sequence A magnetic resonance imaging apparatus configured to collect a magnetic resonance signal from the imaging region based on a gradient magnetic field and a high-frequency magnetic field, and reconstruct an image based on the magnetic resonance signal. Second gradient magnetic field applying means having another gradient magnetic field coil which is arranged so as to include a coil and which applies a second gradient magnetic field to a region including the imaging region, Magnetic resonance imaging apparatus characterized by comprising a control means for controlling the application and application of the high frequency magnetic field of the serial first gradient and the second gradient magnetic field. 2. The control means applies a gradient magnetic field pulse for slicing in the acquisition sequence by a second gradient magnetic field applying means, and a gradient magnetic field pulse for phase encoding and a readout gradient magnetic field in the acquisition sequence. The magnetic resonance imaging apparatus according to claim 1, wherein a pulse is applied by the first gradient magnetic field applying means. 3. The magnetic resonance imaging apparatus according to claim 1, wherein the acquisition sequence is a spin echo method. 4. The magnetic resonance imaging apparatus according to claim 1, wherein the acquisition sequence is a gradient field echo method. 5. The control means comprises saturation excitation means for performing a first sequence for saturation excitation of a region adjacent to the imaging region, and signal acquisition means for performing a second sequence for imaging. The magnetic resonance imaging apparatus described. 6. The magnetic resonance imaging apparatus according to claim 5, wherein the control unit automatically and continuously executes the second sequence after the end of the first sequence. 7. The saturation excitation means applies a predetermined gradient magnetic field pulse to the second gradient magnetic field applying means, and the high-frequency magnetic field applying means applies the gradient magnetic field pulse to the imaging region. 7. A magnetic resonance imaging apparatus according to claim 6, further comprising means for applying a high-frequency magnetic field pulse having a frequency corresponding to the adjacent region of. 8. The saturation excitation means is a second image pickup device.
Means for applying a predetermined gradient magnetic field pulse with a predetermined polarity by the first gradient magnetic field applying means, and the second gradient magnetic field application in synchronization with the application timing of the gradient magnetic field pulse during the execution of the sequence 6. The magnetic resonance imaging apparatus according to claim 5, further comprising means for applying a predetermined gradient magnetic field pulse with a polarity opposite to the polarity by the means. 9. A gradient magnetic field, which is provided so as to cover a local imaging region of a subject placed in a static magnetic field, and which is applied to the imaging region, including various magnetic fields for slicing, phase encoding, and reading. A gradient magnetic field applying means having x, y, z coils and a high frequency magnetic field applying means for applying a high frequency magnetic field of a required frequency to a region including the imaging region are provided along with a sequence of collecting image data. In a magnetic resonance imaging apparatus configured to collect a magnetic resonance signal from the imaging region based on the gradient magnetic field and the high frequency magnetic field, and reconstruct an image based on the magnetic resonance signal, the imaging sequence in the acquisition sequence is Before being executed, a saturation excitation current is supplied to the x, y, and z coils for saturation excitation current that pre-saturates a region adjacent to the imaging region. Magnetic resonance imaging apparatus characterized by comprising a supply unit. 10. The z-coil comprises a pair of coils arranged opposite to each other along a z-direction parallel to the direction of the magnetic flux of the static magnetic field, and the y-coil is a y-coil orthogonal to the z-direction.
The pair of coils are arranged in a plurality of pairs facing each other along a direction, and the x-coil is composed of pairs of coils arranged in a pair along the x-direction orthogonal to the z-direction and the y-direction. Each pair coil is provided with a supply means for supplying the current for slicing, phase encoding, and reading gradient magnetic field, while the saturation excitation current supply means is one of the pair coils constituting the z coil, One of a plurality of pairs of coils forming the y coil and one of a plurality of pairs of coils forming the x coil have a direction of a saturation excitation current supplied to at least one coil, and the direction of saturation excitation current flowing to one of the pair of coils. The magnetic resonance imaging apparatus according to claim 9, wherein the direction is the direction opposite to the direction of the current for applying the gradient magnetic field. 11. A gradient magnetic field, which is provided so as to cover a local imaging region of a subject placed in a static magnetic field, and which has various kinds of gradient magnetic fields including slicing, phase encoding, and reading to the imaging region. A gradient magnetic field applying means having x, y, z coils for imaging and a high frequency magnetic field applying means for applying a high frequency magnetic field of a required frequency to a region including the imaging region are applied along an image data acquisition sequence. A magnetic resonance imaging apparatus configured to collect a magnetic resonance signal from the imaging region based on the gradient magnetic field and the high frequency magnetic field, and reconstruct an image based on the magnetic resonance signal, in a region adjacent to the imaging region. Saturated excitation x, y, z coils for pre-saturated excitation are arranged around the imaging region, and saturated excitation electrodes are placed in the excitation x, y, z coils. The magnetic resonance imaging apparatus characterized by having a saturation excitation current supplying means for supplying before imaging sequence is executed in the acquisition sequence. 12. The z coil is a pair of coils arranged opposite to each other along the z direction parallel to the direction of the magnetic flux of the static magnetic field, and the y coil is a y coil orthogonal to the z direction.
The pair of coils are arranged in a plurality of pairs facing each other along a direction, and the x-coil is composed of a pair of pairs of coils arranged in a pair along the x-direction orthogonal to the z-direction and the y-direction. The saturation excitation current supply means includes one of a pair of coils forming the z-coil for saturation excitation, one of a plurality of pair coils forming the y-coil for saturation excitation, and a plurality of pairs of coils forming the x-coil coil. The direction of the current supplied to at least one of the coils is opposite to the direction of the current flowing through one of the paired coils forming the imaging x, y, and z coils corresponding to the execution of the imaging sequence. The magnetic resonance imaging apparatus according to claim 11, wherein the magnetic resonance imaging apparatus is configured as described above. 13. A first coil group consisting of the imaging x, y, z coils and a second coil group consisting of the saturation excitation x, y, z coils are concentric in the xy cross section. 13. The magnetic resonance imaging apparatus according to claim 12, wherein the magnetic resonance imaging apparatus is arranged at.