JPH11113880A - Static magnetic field unifying method in magnetic resonance imaging device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a method quickly measuring and correcting a static magnetic field distribution within a three-dimensional area of an object in a medical magnetic resonance imaging device. SOLUTION: In this method, after magnetization outside an interest area of an object is previously searched, the magnetization is energized by a high frequency pulse 11 with a small flip angle and phase encoding 14, 16 by 3 steps each in the orthogonal two directions are carried out, and then, a signal 19 is measured with a lead-out inclination magnetic field 18 reversed continuously. Four-dimensional Fourier transformation analysis of the signal is carried out, and a static magnetic field distribution per a voxel of the object is provided from a frequency of a water proton of the voxel. This magnetic field distribution is approximatively developed while using primary and secondary terms of a spherical harmonic function, and a shim electric current value is obtained. For a high speed MRSI method is executed by means of a small number of encodes, a static magnetic field distribution within the three- dimensional area of the object can be obtained in about two seconds, and as a result, the whole of a static magnetic field distribution unifying process can be finished in several seconds. Consequently, high quality image for an EPI method or a spectroscopic imaging method can be provided.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a method for equalizing a static magnetic field in a magnetic resonance imaging (hereinafter, referred to as "MRI") apparatus. The present invention relates to a method for uniformizing.

[0002]

2. Description of the Related Art A typical image construction method of MRI is 2
In the D or 3D Fourier transform method, a subject is placed in a static magnetic field oriented in a spatially uniform and constant direction (this is defined as a z direction), and the nuclear spin of the subject is applied to a horizontal plane ( (x, y planes) and then apply a predetermined combination of gradient magnetic fields to generate a distribution corresponding to the spatial position in the nuclear spin phase (reference direction, for example, azimuth from y-axis direction). Measure the signal.

[0003] However, when the static magnetic field is not uniform, phase rotation occurs due to this, which causes false images, positional distortions, etc., and deteriorates image quality. The inhomogeneity of the static magnetic field is caused firstly by the limitations of the device of the static magnetic field generating magnet, and furthermore, since the living body itself has a slightly different magnetic permeability for each tissue, a distribution is generated in the static magnetic field inside the subject.

In MRI, shimming (uniform static magnetic field) is important in reducing such artifacts.
In particular, since correction of the non-uniformity of the static magnetic field due to the living body (hereinafter referred to as living body shimming) needs to be performed for each subject,
Processing must be completed in a short time.

In general, in living body shimming, a static magnetic field distribution induced by a magnetic permeability distribution in a subject is measured in advance by an MRI method, and a correction shim current is calculated based on the measured magnetic field distribution.
This current is passed through the shim coil to generate a magnetic field in the direction opposite to the non-uniformity, thereby flattening the static magnetic field distribution. The main imaging is performed under this shimming.

As a method of measuring a static magnetic field distribution in a living body, a method of obtaining a static magnetic field distribution in a subject from two complex images having different development times of magnetization (Japanese Patent Laid-Open No. 4-288136 and US Pat. No. 5,168) , 232) or a method in which the static magnetic field distribution is obtained separately from the chemical shift by measuring the image while suppressing the specific chemical shift (Japanese Patent Publication No. 6-44).
904) has been proposed by Glover et al.
In each of these methods, the generation time of the RF spin echo and the generation time of the gradient magnetic field echo are made different from each other,
A phase distribution due to the static magnetic field distribution is generated in the signal, and the static magnetic field distribution is obtained from the phase distribution.

[0007] Since the static magnetic field distribution measurement is performed for each patient as a preliminary imaging prior to the main imaging, as a method for increasing the speed, a method of obtaining a phase distribution image with a single excitation (for example,
"Dynamic Field Changes and EPI Geometric in Lung and Heart"
Distortions in the Chest and Heart) S.Kanayama, SK
uhara and K. Satoh, Proceedings of the Society of
Magnetic Resonance, 746 (1995)) has also been proposed.

As a technique for obtaining a static magnetic field distribution from these phase distributions, if the EPI method or the FSE method is used, the measurement of the static magnetic field distribution in the slice can be performed in a short time, for example, about several seconds. Also, the image matrix can be made relatively large, for example, about 34 to 128, comparable to that of a normal image. Therefore, it can be said that this is an effective method for measuring the static magnetic field distribution in the slice.

However, in order to obtain a static magnetic field distribution in a three-dimensional volume, it is necessary to repeat phase encoding in the thickness direction. The benefits are lost. Conversely, if the matrix in the thickness direction is made smaller, it becomes less meaningful to make only the matrix in the slice larger. Another problem associated with the use of a phase is the problem of phase unwrapping. This is a problem of how to find a main value when a phase difference of 2π radians or more occurs between pixels.
Complicated processing is required.

As a static magnetic field distribution measuring method without using a phase, there is a method using a chemical shift. This is 3D /
4D-chemical shift imaging (CSI: Chemical
This method measures the spectrum of a multi-pixel or voxel by the ShiftImaging method, and directly obtains the local magnetic field strength from the frequency of the spectral line of a specific molecule such as water proton. From the frequency of the spectrum, the local static magnetic field strength is obtained by equation (1).

(Equation 1) This static magnetic field distribution measurement method is applied to correct a spectral position other than water based on the static magnetic field strength and static magnetic field uniformity of a pixel.

[0011]

However, 3D / 4D-
When the CSI method is applied as preliminary imaging of living body shimming, there is a disadvantage that measurement requires a long time. For example, 3
For the measurement of three-dimensional volume, iterative loop of phase encoding is 3
Because of the overlap, the excitation must be performed three times the number of matrices. As an example, if the repetition time (TR) is 1 second and the matrix is 16 × 16 × 16, 1 second × 16 × 16 × 16
= About 68 minutes are required.

[0012]

According to the present invention, there is provided a living body shimming method including a static magnetic field distribution measurement and a shim current determination.
A three-dimensional high-speed MR spectroscopic imaging (MRSI) method is adopted. (Spatial Mapping of t.
he Chemical Shift in NMR), Magnetic Resonance in Med. 1,370-3
86 (1984) and a paper by Matsui et al.
R-Spectroscopy ”(High-SpeedSpatially Resolv
ed High-Resolution NMR Spectroscoy), J. Am. Chem. So
c., 107, 2817-2818 (1985)). In determining the shim current, the expansion of the static magnetic field distribution is limited to the second order. This provides a practical and quick static magnetic field homogenization method.

More specifically, the static magnetic field homogenizing method of the present invention is as follows: 1) Excitation of a desired volume portion of an object by irradiation of a high-frequency magnetic field, then application of a phase encoding gradient magnetic field in two directions, and continuous application in the readout direction A multi-voxel spectrum is obtained for a desired volume portion of a subject by a three-dimensional MRSI method in which a basic sequence of applying a reversed gradient magnetic field is repeated, and 2) a resonance frequency of a specific peak of each voxel spectrum is converted into a static magnetic field intensity. To obtain a static magnetic field distribution in a volume portion, 3) develop this static magnetic field distribution with a magnetic field distribution generated by a plurality of shimming means, and 4) apply the shimming means to each shimming means so as to minimize the static magnetic field distribution. By optimizing the flowing current, the static magnetic field distribution in a desired volume portion of the subject is made uniform.

The three-dimensional MRSI method is a high-speed spectroscopic imaging method in which the repetition loop is reduced by one dimension by continuously inverting the gradient magnetic field in the readout direction and superimposing and acquiring spatial information and spectral information. Thus, a multi-voxel spectrum can be obtained in a short time.

It is preferable to use the resonance frequency of water protons as the specific spectrum peak. Thereby, the static magnetic field intensity can be measured over almost all voxels in the living body. In addition, even when the repetition time (TR) is shortened to shorten the time, a high signal can be obtained by using a pulse having a low flip angle as the high-frequency magnetic field. Further 3D M
When the sequence based on the spin echo (SE) method is adopted as the RSI method and the TE is set long, it is possible to obtain a sharp spectrum in which a signal from bound water or the like is eliminated using the difference in T2, The reading accuracy of the resonance frequency is improved.

In a preferred embodiment of the present invention, when the static magnetic field distribution is developed in step 2), one of the spherical harmonics
The static magnetic field in the desired volume is corrected using at least eight shimming means, which are developed using the next and second order terms and generate first and second order terms of the approximate spherical harmonic function. In the present invention, as the shimming means, a gradient magnetic field coil having a first-order magnetic field distribution and a shim coil having a higher-order magnetic field distribution of approximately second-order or more are used.

By approximating the static magnetic field distribution using only the quadratic and lower terms, the expansion term can be determined from the static magnetic field measurement data of a small number of voxels, and the static magnetic field measurement can be performed within a practically short time. it can. Further, by setting the second order or less, it is possible to avoid the difficulty of manufacturing a shim coil which generates only a specific higher order as a shimming means. In general, since the third order or higher order component is small, correction with sufficient accuracy for practical use is performed. It can be performed.

Since the number of terms less than quadratic is eight, the number of voxels may be eight or more in principle. For example, 3 × 3 × 3
(= 27). In this case, static magnetic field measurement can be performed with nine repetitive loops in the three-dimensional MRSI method.

Further, as a preferred embodiment of the present invention, a three-dimensional M
In the RSI method, the number of matrices in the lead-out direction is set larger than the remaining two directions. This makes it possible to obtain highly accurate information in the spectral direction without substantially extending the static magnetic field measurement time.

[0020]

DESCRIPTION OF THE PREFERRED EMBODIMENTS Hereinafter, the present invention will be described in detail based on embodiments. FIG. 4 is a schematic configuration diagram of an MRI apparatus to which the present invention is applied. This apparatus includes a static magnetic field generating magnetic circuit 402 for generating a uniform static magnetic field B0 inside a subject 401, and orthogonal x, y. Gradient magnetic field generating system 40 for giving a magnetic field gradient whose intensity changes linearly in three directions of z and z
3. Transmission system 40 for generating high-frequency magnetic field in subject 401
4, a detection system 405 for detecting a nuclear magnetic resonance signal generated from the subject, a signal processing system 406 for processing and storing the detected signal, a gradient magnetic field generation system 403, a transmission system 404, and a detection system 405. A sequencer 407 for controlling the apparatus and a computer 408 for controlling the entire apparatus and performing various operations in a signal processing system.
And an operation unit 421 for giving an instruction to the user.

The static magnetic field generating magnetic circuit 402 includes an electromagnet or a permanent magnet (not shown) for generating a uniform static magnetic field and a shim coil 43 for correcting non-uniformity of the static magnetic field.
0 and a shim power supply 431. Shim coil 43
0 is used in combination with a secondary coil such as z2, xz, xy, yz, x2-y2 or a tertiary coil such as z3, z2x, z2y. The primary term is corrected by using a gradient magnetic field coil 409 that generates a linear gradient magnetic field in the x, y, and z directions, and superimposing and supplying a shim correction current to an original gradient magnetic field generation current. In the present invention, the shimming means includes the shim coil 430 and the gradient coil 409. FIG. 6 shows an example of a magnetic field distribution generated by the shim coil. The distribution in the yz plane is indicated by a solid line on the positive side and a contour line by a broken line on the negative side.

Next, an outline of the operation of the apparatus will be described. In the transmission system 404, the high frequency generated by the synthesizer 411 is modulated by the modulator 412, amplified by the power amplifier 413, and supplied to the transmission coil 414 a, so that the subject 4
A high-frequency magnetic field is generated inside 01 to excite nuclear spins. Usually, the target is ¹ H, but the target may be other nuclei having nuclear spin, such as ³¹ P and ¹² C.

At this time, gradient magnetic fields Gx, Gy, Gz are generated by the gradient magnetic field coil 409 in order to add positional information to a nuclear magnetic resonance signal generated from the subject. Each of the three-axis gradient magnetic field coils 409 is supplied with a current from a power supply 410 and is driven according to a predetermined pulse sequence controlled by a sequencer together with irradiation of a high-frequency magnetic field.

The nuclear magnetic resonance signal emitted from the subject 401 is received by the receiving coil 414 b in the detection system 405, passes through the amplifier 415, is subjected to quadrature phase detection by the detector 416, and at a timing according to a command from the sequencer 407. Computer 408 via A / D converter 417
Is input to The high-frequency coils 414a and 414b may be one coil for both transmission and reception, or may be separate. After the signal processing, the computer 408 displays an image corresponding to the nuclear spin density distribution, relaxation time distribution, spectrum distribution, and the like on a display 428 such as a CRT. Data in the middle of calculation or final data is stored in the memory 425. Gradient magnetic field generation system 403, transmission system 404, detection system 40
5 are all controlled by a sequencer 407, which is controlled by a computer 408.

The present invention relates to the measurement of the non-uniformity (static magnetic field distribution) of the static magnetic field of the MRI apparatus having such a configuration, particularly to the measurement of the non-uniformity of the static magnetic field caused by the living body and the control of the shim coil based on the measurement. A step of measuring a static magnetic field distribution in a state where the subject is placed; 2) a step of developing the measured static magnetic field distribution with a magnetic field distribution generated by a shim coil;
Obtaining a correction current to be supplied to the shim coil.

1) Static Magnetic Field Distribution Measurement Step In the static magnetic field distribution measurement step, a pulse sequence based on the high-speed MRSI method is employed, and a spectrum for each voxel is obtained by this. Here, as an example, a case where a spectrum is obtained for a 3 × 3 × 3 (the number of voxels = 27) matrix 22 of the subject's head 21 as shown in FIG. 2 will be described. First, a combination of a general region-selective RF pulse and a diffusion gradient magnetic field is applied a plurality of times to pre-saturate the outside of the region of interest and suppress the generation of a signal from outside the region of interest. Subsequently, the high-speed MRSI method shown in FIG. 1 is performed.

This embodiment uses a gradient echo (Gr).
A high-speed MRSI method based on the E) method is adopted, and an RF pulse 11 for selectively exciting a region of interest is applied together with a gradient magnetic field (here, a gradient magnetic field in the z direction) 13 and then in two directions (here, In this case, the echo signal 19 is measured while applying the phase encoding gradient magnetic fields 14 and 16 in the z and y directions) and repeating the inversion of the gradient magnetic field 18 in the readout direction (x direction).

Here, the flip angle α of the RF pulse 11 is 9
Although it may be 0 °, in order to perform living body shimming in a short time,
Preferably less than 90 °. An example is (90 ° × TR) / T1. Here, T1 is an approximate T1 value of the target tissue, and T1 is a repetition time TR.
If a value longer than 1 is used, replace it with T1 / TR = 1. By using a pulse having a small flip angle, a reduction in signal strength can be suppressed and the repetition time TR can be shortened. The method of setting the flip angle may be constant as described above, but may be sequentially increased for each excitation, and in that case, the generated transverse magnetization component can be constant.

Such a sequence is repeated for a time TR.
By repeatedly executing the phase encoding loop in two directions, four-dimensional data of x-axis, y-axis, z-axis and time (δ) axis is obtained. A spoiler gradient magnetic field 21 is added in the readout direction to remove coherence of the residual transverse magnetization in the repetition of the sequence, and a rewind gradient magnetic field 2 is added in the phase direction to prevent accumulation of phase encoding.
0 is added.

The number of steps in the phase encoding direction (y, z directions) corresponds to the number of voxels in that direction. Here, encoding is repeated every three steps. Therefore, the measurement is completed by repeating 3 × 3 times. The number of sampling points in the lead-out direction corresponds to the number of voxels in that direction. Here, the number of sampling points is three. The number of reversals of the gradient magnetic field in the readout direction is one of the matrix number in the time axis direction.
/ 2, which can be about several tens.

FIG. 3 shows the k-space arrangement of data measured by this sequence. FIG. 3 depicts a trajectory on the kx-kδ plane for specific phase steps ky and kz (the kδ axis is a time axis).
This indicates that sampling is performed at two sampling points 32, 33, and 34. Kx, k
A four-dimensional FFT is performed on y, kz, and kδ to obtain a proton spectrum for each voxel as shown in FIG. From the spectrum of each voxel, the peak frequency of water proton 2
3 and the static magnetic field distribution is obtained from this frequency. By using the resonance frequency of water protons, the static magnetic field intensity can be measured over almost all voxels in a living body.

As an example, suppose that a gradient magnetic field in the read-out direction is repeatedly applied 26 times with a rectangular wave having a period of 6 ms, and 52 echoes are acquired. Generally, spectrum measurement band L
δ is, when the data interval in the kδ direction is Δt (FIG. 3),
Since there is a relation of Lδ = 1 / (△ t), here 1 /
(6 ms) = 1.67 Hz. This is 1.5T M
Since it corresponds to 2.6 ppm in RI and is narrower than the spectrum band (5 ppm) of various metabolite protons present in the living body, the peak outside the measurement band folds back into the band, but the intensity of the usual metabolite is Because it is three to four orders of magnitude smaller than water, it does not hinder the identification and location of water peaks. Subcutaneous fat or the like may be comparable to water, but this does not hinder the peak of fat (for example, 24 in FIG. 2) because it is separated from water. In general, aliasing is an obstacle to identification and quantification when trying to quantify metabolites, but here the advantage is that it is only necessary to measure the frequency of water present in large quantities.

When the static magnetic field of the voxel deviates from the reference value by 2.6 ppm or more, it is difficult to determine the true frequency because the water peak is folded twice or more. Large deviations are unlikely to occur. Further, if the reversal period of the read-out gradient magnetic field is set shorter than 6 ms and the data interval Δt is set to be short, the return can be avoided. With a high magnetic field MRI apparatus compatible with EPI, it is possible to shorten the inversion cycle to about 2 ms.

In FIG. 3, for the sake of simplicity, positive and negative echoes of the gradient magnetic field are individually processed to obtain individual spectra. In that case, Δt is shorter than that in FIG. 3 (Δt /
2) That is, the measurement band Lδ is extended.

The frequency resolution Δf of the spectrum has a relationship of Δf = 1 / T with the measurement time T. In the above example, T = 1
56 ms (= 6 ms × 26) and Δf = 6.4 Hz. This is 0.1 M in 1.5T MRI. 1 ppm. Therefore, the measurement accuracy of the static magnetic field is also at most 0. lppm. The uniformity of the static magnetic field at which no remarkable artifacts occur in EPI is about 0.3 ppm or less, which is sufficient for measurement accuracy.

In the above embodiment, the number of sampling points in the lead-out direction has been described as three. However, the number of sampling points can be increased without extending the measurement time. In this case, the average value of oversampled data is recorded. It is of course possible to do so.

The peak frequency of the water proton is automatically detected by a computer from the proton spectrum of each voxel obtained by 4D-FFT as described above. As a peak detection method, there is a method in which a Lorentz curve is fitted around a standard water position, and the position, width, and height are finely adjusted. Also, phase correction is used together as necessary.

Next, the static magnetic field strength is calculated from the peak position of water for each voxel according to equation (2).

(Equation 2) In the equation, δw (x, y, z) is the shift (ppm) of the water peak from the reference position, and Be (x, y, z) is non-uniform in the static magnetic field.

The reference position of the water peak of a certain voxel (4.
Since the deviation δw from 7 ppm) indicates inhomogeneity of the static magnetic field averaged in the voxel, the static magnetic field distribution can be mapped for the entire region of interest by obtaining the static magnetic field strength for each voxel. However, if the distribution of water in the voxels is uneven, the average weighted by the distribution of water in the voxels is obtained.

2) Step of expanding the measured static magnetic field distribution with the magnetic field distribution generated by the shim coil In this step, the static magnetic field distribution over the previously obtained 3 × 3 × 3 voxels is expanded with the spherical harmonic function. I do. In MRI, in many cases, a shim coil is designed to generate a magnetic field having a substantially spherical harmonic function. Therefore, by expanding with a spherical harmonic function, a shim current can be immediately obtained from its coefficient.

The spherical harmonic function has a functional form as shown in Table 1. In this embodiment, the static magnetic field distribution is approximated using only the second-order and lower-order terms. This is because, in general, higher-order components of the third order or higher are small, and in practice, only the second-order or higher-order items can sufficiently achieve a uniform static magnetic field. In the function form in Table 1, the first-order terms are the three terms x, y, and z, and the second-order terms are z ² , zx, x ² −y ² , x
There are five terms, y and zy.

[Table 1] For determining the coefficients of each harmonic function, for example, the gradient method is used, and the residual I of the equation (3) is minimized.

(Equation 3) Where Pi (x, y, z) is the spherical harmonic and Cj is its coefficient. The summation regarding v is performed over each coordinate (x, y, z) in the volume to be equalized (here, 27 voxels).

Here, when the center of the measured volume coincides with the center of the shim coil, the determination of the coefficient is easy. If they do not match, the measured static magnetic field distribution Be (x, y,
z) and the magnetic field distribution Pi (x, y,
A coefficient is determined after performing a conversion process so that the coordinate system of z) matches.

3) Step of Determining Shim Current In this step, a correction current flowing through the shim coil is determined from the spherical harmonic function coefficient obtained in the previous step. The correction current ideally generates a correction magnetic field distribution having the same amplitude as the measured static magnetic field distribution and an opposite direction. As described above, when the shim coil is designed to generate a magnetic field having a substantially spherical harmonic function, the shim current can be immediately obtained from the coefficient of the spherical harmonic function. In this case, the unique magnetic field generation efficiency for each coil of the shim is considered.

As another method, a correction current may be obtained from a magnetic field distribution (shim characteristics) generated by each shim coil by a matrix operation shown in equation (4).

(Equation 4) In the equation, A represents a shim characteristic matrix (Ajk = δBj / δIk), and an element is represented by a change in the static magnetic field of the j-th pixel with respect to a small change in the k-th channel shim current. ΔI is a shim current vector, B0 is a vector obtained by one-dimensionally rearranging the static magnetic field deviation for each pixel, and At is a transposed matrix of A. Shim characteristics
Using a homogeneous sample such as water in advance, by measuring the change in the static magnetic field distribution per unit current flowing through the shim coil,
Can be obtained and stored in the memory in the form of a matrix or an expansion function expanded by a harmonic function.

This method can be applied to a case where the shim coil generates a magnetic field different from the spherical harmonic function or a composite magnetic field of a plurality of spherical harmonic functions.

The above three steps are performed as a prescan prior to the main imaging, and the primary shim value of the shim current obtained in these steps is set as an offset of the gradient magnetic field.
The secondary shim value is set for the secondary shim coil. Then, the main imaging is performed.

According to this method, living body shimming can be performed in a very short time, and a high-quality image can be obtained in imaging requiring high static magnetic field uniformity such as the EPI method or the spectroscopic imaging method. . As an example, the number of voxels in the first step is 3 × 3 × 3.
(= 27), RF pulse flip angle 10 °, sequence repetition time TR = 160 ms, 160 m
The static magnetic field correction can be performed with s × 9 = 1.44 second measurement time + calculation time. Therefore, the entire shimming process including the calculation of the shim current value can be completed in about several seconds.

In the above embodiment, the number of voxels in the first step is (3.times.3.times.3) on the assumption that up to the second-order terms are used as expansion terms when the static magnetic field distribution is expanded.
However, as described above, the number of voxels is determined by the number of phase encoding steps and the number of sampling points in the frequency encoding direction, and by changing these, the number of voxels and the shape of the voxel can be adjusted to an arbitrary region of interest. .

It is also possible to include higher-order terms than third-order, fourth-order, etc. as expansion terms, thereby enabling high-precision shimming. In that case, it is necessary to increase the number of pixel matrices in the volume corresponding to the number of terms.However, the number of image matrices in the readout direction can be increased without extending the measurement time. It is effective to use a third-order or higher expansion term.
Therefore, it is effective to select a direction in which higher-order unevenness is expected, such as the body axis direction of the living body, as the readout direction.

Further, if the number of phase encodes is set to 4 or more and the matrix is expanded in the phase direction, a third-order or higher-order expansion term can be used, and spatial uniformization with higher accuracy can be achieved. Needless to say. In this case, the measurement time increases with the number of phase encode steps, but the temporal superiority of the high-speed MRSI method employed in the present invention becomes more significant as the number of encode steps increases.

Further, in the above embodiment, the gradient echo (GrE) type high-speed MRSI method shown in FIG. 1 is employed as a sequence for measuring the static magnetic field distribution, but the SE type high-speed MRSI method shown in FIG. 5 is employed. Is also good. The sequence shown in FIG. 5 is different from that in which the high-frequency pulse 51 is irradiated and the high-frequency pulse 510 for inverting the spin after TE / 2 is irradiated is that the phase encodes 54 and 55 are applied in two directions and the read-out direction is applied. Reversing gradient magnetic field 58
At the end of the repetition time.
1 and the spoiler 512 respectively are shown in FIG.
Is the same as

When the GrE type sequence is compared with the SE type sequence, the former causes a swell of the spectrum baseline because the signal in the low frequency region is sacrificed. However, detection of the water peak does not become a major obstacle. The measurement time of the SE type is longer than that of the GrE type,
Since the data of the center of the echo can be acquired, the swell of the baseline can be prevented.

Also, as a feature of the SE type, a well-known T2
Can be used for spectrum editing. That is, by setting the echo time TE to be as long as several tens of ms, it is possible to obtain a spectrum in which the spectrum of the short T2 substance such as fat is eliminated. The half width ν1 / 2 of the spectral line is T2
And ν1 / 2 = 1 / (πT2), the short T2
The peak of the substance is broad, and the spectrum excluding the short T2 substance becomes sharp. Therefore, the position of the peak can be detected with high accuracy.

[0054]

As described above, according to the present invention, when performing living body shimming, a sequence by the high-speed MRSI method is executed in a small encode step to measure the static magnetic field inhomogeneity, and the obtained static By developing the magnetic field distribution with a lower-order term harmonic function, it is possible to achieve a uniform static magnetic field in the volume in a short time according to the distribution subject, and to improve the image quality of the EPI method or an imaging method based on the EPI method. be able to.

[Brief description of the drawings]

FIG. 1 is a diagram showing an embodiment of a static magnetic field measurement sequence according to the present invention.

FIG. 2 is a diagram showing a concept of measuring a static magnetic field distribution by using a spectrum.

FIG. 3 is a diagram showing a k-space trajectory of the high-speed MRSI method.

FIG. 4 is a diagram showing the overall configuration of an MRI apparatus to which the present invention is applied.

FIG. 5 is a diagram showing another embodiment of the static magnetic field measurement sequence of the present invention.

FIG. 6 is a diagram showing a magnetic field distribution of a shim coil.

[Explanation of symbols]

 401 subject 414a transmission RF coil 414b detection RF coil 430 shim coil 431 shim power supply

Claims (3)

[Claims] 1. A magnetic resonance imaging apparatus for imaging a region of a subject placed in a static magnetic field based on a nuclear magnetic resonance signal generated from a desired portion of the subject, wherein a static magnetic field distribution of a desired volume portion of the subject is provided. 1) Exciting a desired volume portion of the object by irradiating a high-frequency magnetic field, then applying a phase encoding gradient magnetic field in two directions, and further generating a gradient magnetic field that is continuously inverted in the readout direction. Obtaining a multi-voxel spectrum for a desired volume portion of the subject by a three-dimensional spectroscopic imaging method in which a basic sequence to be applied is repeated, and 2) converting a resonance frequency of a specific peak of each voxel spectrum into a static magnetic field intensity. To obtain a static magnetic field distribution in the volume part, and 3) a magnetic field generated by a plurality of shimming means. 4) A static magnetic field homogenization method characterized by optimizing a current flowing through each shimming means so as to minimize the static magnetic field distribution. 2. The method according to claim 1, wherein the static magnetic field distribution calculated from the multi-voxel spectrum of the volume portion is expanded using first and second order terms of a spherical harmonic function to generate at least 8 terms of first and second order terms of the approximate spherical harmonic function. 2. The static magnetic field equalizing method according to claim 1, wherein the static magnetic field in the volume is corrected using two shimming means. 3. The homogenization of an in-volume static magnetic field according to claim 1, wherein in the three-dimensional spectroscopic imaging method, the number of matrices in the readout direction is set to be larger than the remaining two directions. Method.