JPH04343833A - Magnetic resonance imaging device capable of correcting primary item of ununiformity of static magnetic field - Google Patents

Abstract

PURPOSE:To perform accurate measurement of an amount of correction in order to prevent images from being affected adversely by correcting the primary item of the ununiformity of a static magnetic field if such a phenomenon occurs. CONSTITUTION:While the tilting magnetic fields Gx, Gy, Gz of X-, Y- and Z-axes are applied in sequence, RF pulses 30,31,32 are applied whereby a small area including the center of a static magnetic field is excited. Thereafter, echo signals are obtained by using preciples similar to those of gradient echo. The difference in the location among peaks in the direction of phase encoding is measured from data obtained and thereby an amount of correction can be found. Therefore, the primary item of the ununiformity of the static magnetic field can be eliminated and so images without problems such as spacial distortion and signal deterioration can be obtained.

Description

[Detailed description of the invention]

0001

[Industrial Application Field] The present invention relates to nuclear magnetic resonance (hereinafter referred to as "N"
This invention relates to a magnetic resonance imaging apparatus that obtains a tomographic image of a desired part of a subject by utilizing the phenomenon (abbreviated as "MR"), and particularly relates to a magnetic resonance imaging apparatus that has means for measuring and correcting non-uniformity of a static magnetic field at any time. It is.

0002

2. Description of the Related Art A magnetic resonance imaging apparatus uses NMR phenomena to measure the density distribution, relaxation time distribution, etc. of nuclear spins (hereinafter simply referred to as spins) at a desired inspection site in a subject. It displays an image of an arbitrary cross section of the object from the measurement data. In the conventional magnetic resonance imaging apparatus, as shown in FIG.
a static magnetic field generating magnet 2 that applies a static magnetic field to the subject 1; a magnetic field gradient generating system 3 that applies a gradient magnetic field to the subject 1; and a high-frequency pulse that causes nuclear magnetic resonance in the nuclei of atoms constituting the living tissue of the subject 1. A sequencer 7 that repeatedly applies the pulses in a predetermined pulse sequence, and a high-frequency magnetic field to cause nuclear magnetic resonance in the nuclei of atoms constituting the living tissue of the subject 1 using the high-frequency pulses from the sequencer 7. a transmitting system 4 that irradiates the radio waves, a receiving system 5 that detects the echo signals emitted by the above-mentioned nuclear magnetic resonance, and a signal processing system that performs image reconstruction calculations using the echo signals detected by the receiving system 5. 6, and was designed to obtain tomographic images by repeatedly measuring echo signals emitted by nuclear magnetic resonance.

In this device, as shown in FIG.
Static magnetic field generating magnet 2 that generates a static magnetic field of ~2 Tesla
The subject 1 is placed inside. At this time, the spins in the subject 1 precess around the direction of the static magnetic field at a frequency determined by the strength H0 of the static magnetic field. This frequency is called the Larmor frequency, and the Larmor frequency ν0 is here H0
: Static magnetic field strength γ : Each type of atomic nucleus has a unique value expressed by the gyromagnetic ratio. Also, if the angular velocity of Larmor precession is ω0, then ω
Since there is a relationship of 0=2πν0, ω0=γ・H0
It is given by (2).

[0004] Then, a high frequency irradiation coil 14 in the transmission system 4
Larmor frequency ν of the atomic nucleus to be measured by a
When a high frequency magnetic field (electromagnetic wave) with a frequency f0 equal to 0 is applied, the spins are excited and transition to a high energy state. When this high-frequency magnetic field is discontinued, the spins return to their original low energy state with a time constant depending on each state. The electromagnetic waves emitted at this time are received by the high-frequency receiving coil 14b in the receiving system 5, amplified and waveform-shaped by the amplifier 15, digitized by the A/D converter 17, and central processing unit 8 (
(hereinafter referred to as CPU). The CPU 8 performs image reconstruction calculations based on this data, and displays a tomographic image of the subject 1 on a display 20 (hereinafter referred to as CRT). The above-mentioned high-frequency magnetic field is generated by amplifying the signal sent out by the sequencer 7 controlled by the CPU 8 by a high-frequency transmitting coil power supply (not shown) and transmitting it to the high-frequency transmitting coil 14.
It can be obtained by sending it to a.

[0005] In addition to the static magnetic field and high-frequency magnetic field described above, the magnetic resonance imaging apparatus described above is equipped with a gradient magnetic field coil 9 to generate a gradient magnetic field for obtaining positional information in space. There is. These gradient magnetic field coils 9 are
A gradient magnetic field is generated by supplying current from a gradient magnetic field power supply 10 operated by a signal from the sequencer 7.

The imaging principle of the magnetic resonance imaging apparatus will now be explained with reference to FIG. 4. First, Figure 4 (a
), an atomic nucleus placed in a static magnetic field H0 in the Z direction behaves like a bar magnet from a classical physics perspective, and ages around the Z axis at the Larmor frequency ν0 mentioned earlier. Performing differential movement. This frequency is given by the above equation (2) and is proportional to the strength H0 of the static magnetic field. γ in Equations (1) and (2) is called the gyromagnetic ratio, and has a value specific to the atomic nucleus. In general, there are a huge number of atomic nuclei to be measured, each rotating with an arbitrary phase, so when viewed as a whole, the components in the X-Y plane cancel each other out, and the macroscopic magnetization of only the Z-direction component is remain. In this state, as shown in FIG. 4(b), when a high frequency magnetic field H1 having a frequency equal to the Larmor frequency ν0 is applied in the X direction, the macroscopic magnetization begins to fall in the Y direction. This angle of inclination is proportional to the product of the amplitude of the high-frequency magnetic field H1 and the application time, and the high-frequency magnetic field H1 when the high-frequency magnetic field H1 tilts at 90 degrees with respect to the pulse application time is a 90° pulse, and the high-frequency magnetic field H1 when the high-frequency magnetic field H1 tilts by 180 degrees is a 180° pulse. call. In addition, X, Y in Figures 4 (a) and (b)
, Z are orthogonal Cartesian coordinate axes.

[0007] A two-dimensional Fourier imaging method is a commonly used method for imaging in such a magnetic resonance imaging apparatus. FIG. 5 is a timing diagram schematically showing a pulse sequence of the spin echo method, which is a typical two-dimensional Fourier imaging method. Figure 5
In this figure, (a) shows the irradiation timing of the high-frequency magnetic field signal and the envelope for selectively exciting the slice position of the subject. The figure (b) shows the timing of applying the gradient magnetic field Gz in the slice direction, and the figure (c) shows the timing of applying the gradient magnetic field Gy in the phase encoding direction and measurement by changing the amplitude. d) The figure shows the timing of application of the gradient magnetic field Gx in the frequency encoding direction. Moreover, the figure (e) shows the echo signal (NMR signal) to be measured.

In the pulse sequence shown in FIG. 5, first,
After applying a 90° pulse, a 180° pulse is applied at Te/2, where Te is the echo time. After applying the 90° pulse, each spin starts rotating in the X-Y plane at its own unique speed, so a phase difference occurs between the spins over time. When a 180° pulse is applied here, each spin is reversed symmetrically around the X axis as shown in Figure 6, and in order to continue rotating at the same speed, the spins converge again at time Te shown in Figure 5. An echo signal is formed as shown in (e). Although the signals are measured as described above, the spatial distribution of the signals must be determined in order to construct a tomographic image. For this purpose, a linear gradient magnetic field is used. A spatial magnetic field gradient is created by superimposing a gradient magnetic field on a uniform static magnetic field. As mentioned earlier, the rotational frequency of the spins is proportional to the magnetic field strength, so when a gradient magnetic field is applied, the rotational frequency of each spin differs spatially. Therefore, by examining this frequency, the position of each spin can be determined. For this purpose, a phase encoding gradient magnetic field Gy and a frequency encoding gradient magnetic field Gx shown in FIG. 5 are used.

Using the above-described pulse sequence as a basic unit, it is repeated a predetermined number of times, for example, 256 times, at a constant repetition time Tr while changing the intensity of the gradient magnetic field Gy in the phase encoding direction each time. By subjecting the measurement signal thus obtained to two-dimensional inverse Fourier transformation, the spatial distribution of macroscopic magnetization shown in FIG. 4(a) can be obtained. In the above explanation,
If the three types of gradient magnetic fields do not overlap each other, X, Y, Z
It may be any one of these, or a combination thereof. The above basic principles of magnetic resonance imaging are explained in detail in "NMR Medicine (Basic and Clinical)" (edited by the Nuclear Magnetic Resonance Medicine Study Group, published by Maruzen Co., Ltd., January 20, 1980). .

[0010] Here, the non-uniformity of the static magnetic field will be briefly explained. In an MRI apparatus, a static magnetic field is required to have extremely high uniformity. This is because, as described above, the position of the MR image is discriminated based on the difference in resonance frequency, and if the uniformity does not meet the required specifications, spatial distortion, positional shift, and phase distortion of the image occur. The first-order term of static magnetic field inhomogeneity is
The influence on the image differs depending on which of the three axes of slice, phase encoding, and frequency encoding the linear term occurs along. The effect also differs depending on whether the pulse sequence used is the spin echo method or the gradient echo method. This will be classified and explained below.

[1] Slice axis: (1) Slice thickness variation: The resonance frequency for the spatial position in the slice axis direction is determined by the strength of the gradient magnetic field applied when selecting a slice, as shown in FIG. 6-A. If a non-uniform first-order term is superimposed on the slice axis, the spatial width corresponding to the excitation frequency band given by the RF pulse will vary, and the slice thickness will change. Both the spin echo method and the gradient echo method are affected in the same way. Specifically, R.F.
When the excitation frequency band given by the pulse is Δf, the slice thickness Δd is expressed as follows: where Gs: Gradient magnetic field strength at the time of slice selection,
The slice thickness Δd' when the non-uniform linear term Gu is superimposed is as follows. Therefore, from equations (3) and (4), a slice having the thickness is selected as follows.

(2) Slice position shift: For the same reason as (1), the spatial position corresponding to the excitation frequency given by the RF pulse changes and the slice position shifts. Both the spin echo method and gradient echo method are affected. Specifically, R
When the excitation frequency given by the F pulse is f, the slice position d is expressed as follows, and the slice position d' when the nonuniform linear term Gu is superimposed is as follows. Therefore, Equations (6) and (7) Therefore, a shifted position will be selected.

(3) Oblique of the frequency encode axis in the direction of the slice axis: When sampling an echo signal, it is necessary to change the frequency with respect to the position only on the frequency encode axis; in other words, the frequency It is necessary to apply a gradient magnetic field only to the encode axis. If a slice axis gradient magnetic field is applied (Figure 6-B) during signal sampling, the vector sum of the slice axis gradient magnetic field and the frequency encode axis gradient magnetic field will actually be superimposed on the static magnetic field. As a result, the frequency encode axis obliques in the slice axis direction. Spin echo method, gradient echo method
Both laws are affected. This oblique angle θ is θ=t
an-1 (Gu/Gf) (9) G here
f: Represented by frequency encoded gradient magnetic field strength at the time of data sampling.

(4) Discrepancy in the amount of rephasing (rephasing or phase return) of spins arranged in the slice axis direction: Ideally, the RF pulse for slice selection should select all the spins in the slice plane. There is no problem if the magnetic field can be excited instantly and the slicing gradient magnetic field can be applied only at that moment. However, the actual application time of the RF pulse is a non-negligible time of about several ms, and the effective excitation time (symmetrical sinc
If the function is an envelope, the gradient magnetic field in the slice direction applied after approximately the center of the total application time
The phase of spins arranged in the slice direction is diffused in proportion to the total applied area (intensity x time). In order to restore this phase diffusion, a gradient magnetic field with reversed polarity is applied to the same slice axis in an area equal to the above-mentioned total area to perform rephasing (rephasing or phase return). This is because the phase of all spins in the slice direction is aligned by setting the area sum of gradient magnetic field application to 0 after excitation until echo signal measurement. The first-order term of the static magnetic field inhomogeneity is equivalent to superimposing an extra gradient magnetic field, so it affects the area summation, but as a result it does not affect the image obtained by the spin echo method, and the gradient It affects only the echo method (Figure 6-C). In other words, in the spin echo method, the time from the application of the 90° pulse to the application of the 180° pulse, and the time from the application of the 180° pulse
°The time from pulse application to echo peak is te
/2, and since the gradient magnetic field is effectively reversed by the spin phase reversal caused by applying a 180° pulse, Gu・te
/2-Gu·te/2=0 (10), and the gradient magnetic field application areas given by the nonuniform first-order terms before and after the 180° pulse cancel each other out and are not affected. On the other hand, in gradient echo, the phase spreads in proportion to the time te from excitation to echo signal measurement. That is, the phase rotation amount Φ of the slice position d due to the non-uniform linear term Gu is Φ=γ・d・Gu・te
(11), the phases of the spins in the slice direction are not aligned, causing problems such as a decrease in signal strength.

(5) Shift in the amount of phase rotation of spins moving due to blood flow, etc.: The amount of phase rotation Φ that spins moving due to blood flow, etc. in a magnetic field undergoes due to the non-uniform linear term Gu is If the position is x0 and the moving speed is v, then in the case of a spin echo, it becomes as follows. On the other hand, in the case of gradient echo, it becomes. The first term of equation (13) is the same as equation (11),
The phase rotation caused by the movement is the second term. Therefore, for the same TE, the gradient echo is more strongly affected.

[2] Phase encode axis: (6) Oblique of frequency encode axis in the phase encode axis direction: (
For the same reason as 3), if a gradient magnetic field of the phase encode axis is applied (Fig. 7-A) during signal sampling, the frequency encode axis will oblique in the direction of the phase encode axis as a result. I will do it. The oblique angle θ follows equation (9). Both the spin echo method and gradient echo method are affected.

(7) Echo peak shift in the phase encoding direction in k-space: As mentioned above, the phase encoding changes the intensity stepwise for each TR, and changes the echo peak in the phase axis direction. It gives a phase rotation according to the position of
What determines the amount of phase rotation is the total area to which the gradient magnetic field is applied in each step. Therefore, for the same reason as (4), the first-order term of static magnetic field inhomogeneity affects the sum of areas, but as a result, it does not affect the spin echo method, but only the gradient echo method. (Figure 7-B). In gradient echo, the time t from excitation to echo signal measurement
During e, the area to which the gradient magnetic field is applied changes by Gu·te due to the nonuniform first-order term Gu. However, since this change is always constant for each TR, as shown in FIG. 8, it is the same as if each step of phase encoding was directly added with an offset. In other words, 1 of the static magnetic field inhomogeneity
If the following term exists, the echo signal has a phase encode of 0
There is no peak when , but a peak appears at a position shifted in the phase encoding direction in k-space. The number of steps s for the shift is determined by the phase encode application time te.
If nc is the gradient magnetic field strength of one step, Genc is,

(8) Shift in the amount of phase rotation of moving spins such as blood flow: Same effect as in (5) above.

[3] Frequency encode axis: (9) Image size variation in the frequency encode direction: Due to the strength of the gradient magnetic field applied during signal sampling, the resonance frequency with respect to the spatial position in the frequency encode axis direction changes. As it is decided,
As shown in FIG. 9-A, when a non-uniform first-order term is superimposed on the frequency axis, the frequency with respect to the spatial position varies. On the other hand, since the sampling pitch of the data is set so that there is no non-uniform linear term, the position of the signal source obtained when Fourier transform is performed in the frequency direction is
It will be different from the actual one. As a result, the image appears as a reduction in expansion and contraction with respect to the center of the static magnetic field in the frequency axis direction. Both the spin echo method and the gradient echo method are affected in the same way. Specifically, if the resonant frequency of the spin at a position d from the center is f, then the resonant frequency f' when the nonuniform linear term Gu is superimposed is, and the resonant frequency f' at a position d from the center is The spin is at the resonant frequency f
It is recognized as being at the location d' corresponding to '. From equation (16), d' is expressed as follows.

(10) Shift in echo peak position: As shown in (4) and (7) above, the static magnetic field Uneven 1
The next term has an effect, but the effect is on the gradient echo.
On the frequency encode axis, it appears as an echo peak shift, as shown in FIG. 9B. The appearance time te' of the shifted echo peak is as follows.

(11) Shift in the amount of phase rotation of moving spins such as blood flow: Same effect as in (5) above.

[0022] As mentioned above, the static magnetic field uniformity, which is extremely important in an MRI apparatus, is adjusted so that it falls within specifications when the apparatus is installed. However, in an MRI apparatus that uses a permanent magnet as a static magnetic field generation mechanism, the static magnetic field uniformity is also affected by temperature fluctuations of the magnet. That is, as is well known, the magnetomotive force of a permanent magnet changes with temperature. For example, when a Nd-Fe-B magnet is used, the temperature coefficient of the magnet alone can reach -900 ppm/°C. If the magnetic circuit is a self-shielded type using a yoke, the thermal inertia of the yoke will be added to reduce this number, but in any case, temperature fluctuations will occur locally or unevenly with respect to the magnet. This causes local fluctuations in the static magnetic field strength and changes the static magnetic field uniformity. For example, if the N and S poles are configured to face each other vertically,
If there is a temperature difference between the upper and lower magnets, the temperature will change in the vertical direction (static magnetic field direction:
This results in the formation of a magnetic field with an inclination to the Z-axis). This non-uniform component corresponds to the first-order term of Z (hereinafter referred to as Z1 term). In actual permanent magnet type MRI equipment, JP-A-63-43
649 and JP-A-63-278310, by covering the magnetic circuit with a heat insulating material and keeping the inside at a constant temperature, the static magnetic field non-uniformity described above is controlled so as not to occur. The generation of static magnetic field inhomogeneity in permanent magnet MR devices has been explained above, but
Even in the R device, static magnetic field non-uniformity will occur if some kind of malfunction occurs in the static magnetic field correction by the shim coil. Regarding the first-order non-uniformity of the static magnetic field mentioned above, please refer to Japanese Patent Application No. 2-19042.
As mentioned in , measure the amount of non-uniformity in advance,
Correction can be made by adding an offset gradient magnetic field to cancel it during the actual measurement.

[0023]

[Problem to be Solved by the Invention] In the above-mentioned conventional technology, even if first-order static magnetic field inhomogeneity occurs due to temperature fluctuations in the magnet in a permanent magnet system or shim defects in normal conducting/superconducting equipment, it is not possible to There was no way to automatically correct it. Furthermore, in the case of the correction according to Japanese Patent Application No. 2-19042, no consideration was given to the pulse sequence for improving the correction accuracy. In particular, the gradient echo method, which can be used as a sequence for measuring the correction amount, is more sensitive to the homogeneity of the static magnetic field than the spin echo method, and care must be taken to ensure the accuracy of the measurement of the correction amount. . In the gradient echo method, echoes are formed only by reversing the gradient magnetic field, and phase diffusion due to non-uniformity of the static magnetic field cannot be prevented. Since phase diffusion progresses in proportion to the passage of time after spin excitation, the longer the echo time TE, the wider the width of the echo signal in the time axis direction and the lower the peak value. The detection accuracy deteriorates, and the accuracy of correction amount measurement also decreases. on the other hand,
When the TE is shortened, the influence of the first order term of static magnetic field inhomogeneity on the spin is reduced, and the inhomogeneous component cannot be captured with sufficient accuracy. In other words, as already mentioned, if a static magnetic field inhomogeneity linear term exists in the phase encoding direction, the echo signal does not peak when the phase encoding is 0, and the echo signal does not reach a peak when the phase encoding is 0. - A peak appears at a position shifted in the direction. Nonuniformity can be calculated by measuring this amount of deviation, but since the amount of deviation is proportional to TE, if TE is shortened, the amount of deviation becomes smaller, and as a result, the correction amount cannot be measured accurately. SUMMARY OF THE INVENTION An object of the present invention is to measure first-order static magnetic field inhomogeneity with high precision and to perform measurements with sufficient correction of static magnetic field uniformity during measurement.

[0024]

[Means for Solving the Problems] In order to achieve the above object, in the magnetic resonance imaging apparatus according to the present invention, an offset gradient magnetic field is always applied during sequence operation in order to cancel first-order non-uniformity. The intensity of the offset gradient magnetic field to be applied is calculated from measurement data obtained in a prior correction amount measurement sequence. Here, the correction data measurement sequence excites only a small area at the center of the static magnetic field that is sufficiently homogeneous in order to measure the correction amount with high precision, and takes a relatively long echo time to eliminate the first-order distortion. We decided to extract only the uniform terms efficiently.

[0025]

[Operation] The magnetic resonance imaging apparatus configured as described above first applies a high-frequency magnetic field and gradient magnetic field pulses to each axis by controlling the pulse sequence by a sequencer as a preliminary measurement, and then applies a high-frequency magnetic field and gradient magnetic field pulses to each axis. After selective excitation of only the region, the echo signal is measured by the gradient echo method. The CPU calculates the first-order term of the static magnetic field inhomogeneity from the measured echo signal, and sets the correction gradient magnetic field strength to be applied during the main measurement for each of the X, Y, and Z axes. In this measurement, a high-frequency magnetic field and gradient magnetic field pulses for each axis are applied according to a desired pulse sequence under control by a sequencer to obtain a tomographic image. At this time, each axis gradient magnetic field is echo-
This is the sum of the above-mentioned correction gradient magnetic field strength and the gradient magnetic field strength required for signal measurement.

[0026]

DESCRIPTION OF THE PREFERRED EMBODIMENTS Hereinafter, embodiments of the present invention will be described in detail with reference to the accompanying drawings. FIG. 3 is a block diagram showing the overall configuration of a magnetic resonance imaging apparatus according to the present invention. This magnetic resonance imaging apparatus obtains a tomographic image of a subject by using the nuclear magnetic resonance (NMR) phenomenon, and as shown in FIG. system 4, receiving system 5, signal processing system 6, sequencer 7,
It is equipped with a CPU8. The static magnetic field generating magnet 2 generates a uniform static magnetic field around the subject 1 in the body axis direction or in a direction perpendicular to the body axis.
A magnetic field generating means of a permanent magnet type, a normal conduction type, or a superconducting type is arranged in a spacious space around the . The magnetic field gradient generation system 3 consists of gradient magnetic field coils 9 wound in the three axis directions of X, Y, and Z, and a gradient magnetic field power supply 10 that drives each gradient magnetic field coil. By driving the gradient magnetic field power supply 10 of each coil according to the command, gradient magnetic fields Gx, Gy, and Gz in the three axis directions of X, Y, and Z are applied to the subject 1. A slice plane for the subject 1 can be set by applying this gradient magnetic field.

The sequencer 7 repeatedly applies a high-frequency magnetic field pulse in a predetermined pulse sequence to cause nuclear magnetic resonance in the nuclei of atoms constituting the living tissue of the subject 1, and operates under the control of the CPU 8. Various commands necessary for data collection of tomographic images of the subject 1 are sent to the transmission system 4, the magnetic field gradient generation system 3, and the reception system 5. The transmission system 4 irradiates a high-frequency magnetic field using high-frequency pulses sent out from the sequencer 7 to cause nuclear magnetic resonance in the nuclei of atoms constituting the living tissue of the subject 1, and includes a high-frequency oscillator 11 and a modulator 12. , a high frequency amplifier 13, and a transmitting side high frequency coil 14a. The high frequency pulse output from the high frequency oscillator 11 is amplitude-modulated by a modulator 12 according to instructions from the sequencer 7. The amplitude-modulated high-frequency pulse is amplified by a high-frequency amplifier 13 and then supplied to a high-frequency coil 14a placed close to the subject 1, so that the subject 1 is irradiated with electromagnetic waves.

The receiving system 5 detects an echo signal (NMR signal) emitted by nuclear magnetic resonance of the atomic nucleus of the biological tissue of the subject 1, and includes a high frequency coil 14b, an amplifier 15, and a quadrature phase detector on the receiving side. 16 and an A/D converter 17. The electromagnetic wave (NMR signal) of the response of the subject 1 due to the electromagnetic wave irradiated from the high frequency coil 14a on the transmitting side is transmitted to the high frequency coil 14b placed close to the subject 1.
The signal is detected by the A/D converter 17 via the amplifier 15 and the quadrature phase detector 16, and is converted into a digital quantity. Furthermore, two series of collected data are sampled by the quadrature phase detector 16 at the timing according to the command from the sequencer 7, and the signals are sent to the signal processing system 6. This signal processing system 6 includes a CPU 8,
A recording device such as a magnetic disk 18 and a magnetic tape 19;
It consists of a display 20 such as a CRT, and the CPU 8
Processes such as Fourier transformation, correction coefficient calculation, and image reconstruction are performed using the , and the signal intensity distribution of an arbitrary cross section or the distribution obtained by performing appropriate calculations on multiple signals is converted into an image and displayed on the display 2.
0 as a tomographic image. In addition, in FIG. 3, the high frequency coils 14a on the transmitting side and the receiving side,
14b and the gradient magnetic field coil 9 are installed in the magnetic field space of the static magnetic field generating magnet 2 placed in the space around the subject 1.

In the present invention, the sequencer 7 controls normal measurement and uneven linear term correction amount measurement using respective pulse sequences. During normal measurement, an offset gradient magnetic field is added to the original gradient magnetic field pulse to correct the first-order term for each physical axis (X, Y, Z axis), and this application starts from the start of measurement. Lasts until the end of measurement. This situation is shown in FIG. That is, if the offset gradient magnetic field intensities corresponding to the nonuniform first-order terms of the X, Y, and Z axes are X1, Y1, and Z1, then this gradient magnetic field is always applied in a DC manner during the measurement period. Even during oblique measurement, this is applied with a fixed intensity on the physical axis.

Next, a method of measuring the static magnetic field inhomogeneity first-order term in the magnetic resonance imaging apparatus according to the present invention will be explained with reference to FIGS. 1 and 10. FIG. 1 is a timing diagram schematically showing a static magnetic field inhomogeneous first-order term measurement sequence according to the present invention. In FIG. 1, (a) shows the irradiation timing of a high-frequency magnetic field signal and an envelope for selectively exciting a slice position of a subject. (b
) The figure shows the timing of application of the gradient magnetic field Gx in the X direction, the figure (c) shows the timing of the application of the gradient magnetic field Gy in the Y direction, and the figure (d) shows the timing of the application of the gradient magnetic field Gz in the Z direction. Figure (e) shows that measurement is performed by changing the timing and amplitude of the application of the gradient magnetic field Gp in the phase encoding direction, and Figure (f) shows the timing of application of the gradient magnetic field Gf in the frequency encoding direction. It shows. Also,(
Figure g) shows the echo signal (NMR signal) to be measured, and Figure h) shows the pulse sequence divided into sections 1 to 10. Here, Gx, Gy, Gz and Gs,
The relationship between Gp and Gf is one to one.

In FIG. 1, in section 1, (a) and (b)
As shown in the figure, the first excitation region in FIG. 10 is excited by applying the 90° selective excitation pulse and the X-axis gradient magnetic field pulse Gx1. In section 2, Gx1 is continued to be applied, and the Y-axis gradient magnetic field pulse Gy1 is also applied. Gx1,
By applying Gy1, the phase of spins distributed in the X and Y axis directions is diffused. In section 3, the application of Gx1 is finished, and only Gy1 is continued to be applied. As a result, the phase of only the spins distributed in the Y-axis direction is diffused. In section 4, Gy1 is continued to be applied and 180°
Apply selective excitation pulses. As a result, the second excitation region shown in FIG. 10 is excited, and although the longitudinal magnetization is reversed in regions other than the region overlapping with the first excitation region, transverse magnetization does not occur, and when measuring the echo signal in section 9, no echo signal is generated. Does not emit. Only the spins present in the first excitation region in section 1 and in the second excitation region in section 4 have transverse magnetization reversed by this excitation. The spins existing in this overlapping region were excited at 90° and 180°, similar to the normal SE method. In section 5, Gy1 is continued to be applied, and X- and Z-axis gradient magnetic field pulses Gx2 and Gz1 are applied. By applying Gx2 and Gy1, the phases of the spins (inverted transverse magnetization) aligned in the X and Y axis directions in the overlapping region are aligned at the end of section 5. At this time, the phase of the spins that are transversely magnetized outside the overlapping region, that is, the spins that have been excited only by the 90° pulse, is not reversed and is therefore further diffused. On the other hand, regarding the Z-axis direction due to Gz1,
The phase is diffused.

In section 6, Gz1 is continued to be applied, and a 180° selective excitation pulse is also applied. This excites the third excitation region in FIG.
Inverted transverse magnetization does not occur in areas other than the overlapping area of , and no echo signal is generated when measuring the echo signal in section 9. Having transverse magnetization reversed by this excitation is
Only spins exist in the first excitation region in section 1, in the second excitation region in section 4, and in the third excitation region in section 4. The spins existing in the overlapping region of these three regions are 90°-180°-1, which is the same as in the normal multi-echo method.
This means that it was excited at 80°. In section 7, Gz1 is continued to be applied. By applying Gz1, the phases of the spins (inverted transverse magnetization) arranged in the Z-axis direction in the overlapping region are aligned at the end of section 7. At this time, the phase of the spins that are transversely magnetized outside the overlapping region, that is, the spins excited only by the 90° pulse or only the 90°-180° pulse, is not reversed and is therefore further diffused. In section 8, a phase encoding gradient magnetic field pulse Gp is applied, and a negative gradient magnetic field pulse Gf1 is applied in the frequency encoding direction. In section 9, a positive gradient magnetic field pulse Gx4 is applied in the frequency encoding direction, and an echo signal 90 is measured. In section 10, no pulse is applied, and there is a waiting time until the next repetition. With such a pulse sequence, the first to third pulses in FIG.
Only from the overlapping cubes (or cuboids) of the excitation region,
Resonant signals can be extracted.

By the way, in the gradient echo method, as mentioned above, phase diffusion occurs due to static magnetic field inhomogeneity. Even in the non-uniform linear term measurement sequence shown in Figure 1,
The pulse sequence is constructed to be similarly influenced. The state of phase diffusion due to static magnetic field inhomogeneity in both of these sequences is shown in FIGS. 11 and 12. First, in the gradient echo, as shown in FIG. 11, the phase is diffused due to non-uniformity between the excitation time of the α° pulse and the echo time TE. On the other hand, in the pulse sequence according to the present invention shown in FIG. 12, the phase spread by the 180° pulse 31 once converges, and completely converges when the 180° pulse 32 is applied again. Here, the phase is reversed by 180 degrees, but the phase is not diffused, so if TE' is the time from this point to the time when the echo peak appears, then if TE = TE', the amount of phase rotation for both is equal. Become. A correction amount can be calculated from this amount of phase rotation.

As mentioned above, sections 1 to 10 shown in FIG.
The pulse sequence is first repeated a predetermined number of times while changing the intensity of the gradient magnetic field Gp in the phase encoding direction shown in FIG. The CPU 8 searches the raw data for a data string that gives the highest signal strength, and calculates how far that data string deviates from the center, that is, the 0 phase encode, in either the positive or negative direction.

Next, a procedure for determining the first-order term of static magnetic field inhomogeneity from the calculated maximum signal strength shift in the phase encoding direction will be explained. FIG. 2 shows the main parts related to phase encoding and the non-uniform first-order term in the non-uniform first-order term measurement sequence shown in FIG. 1. Goff shown in the figure
If a non-uniform linear term of set exists, the actual magnetic field will be as shown by the solid line with respect to the 0 level shown by the broken line. The total amount S of this extra magnetic field applied is: S=Goffset×TE'
(19), and the value Gpeak, which is obtained by inverting the sign of the value obtained by dividing this by the phase encoding application time tenc, becomes the phase encoding level that gives the highest signal strength of the echo. Gpeak=-S/tenc
(20) From actual measurements, Gpeak has already been determined as an integral multiple of the gradient magnetic field strength of one step of phase encoding, so from equations (19) and (20), the correction amount Gcor is as follows. By always applying this correction amount during normal measurement, the influence of the first-order term of the static magnetic field can be removed.

[0036]

[Effects of the Invention] Since the present invention is configured as described above,
By controlling the pulse sequence by the sequencer 7,
During the actual measurement, the first-order term of static magnetic field inhomogeneity is always corrected with high precision. This makes it possible to obtain an image free from problems such as spatial distortion, positional deviation, variation in slice thickness, and S/N reduction.

[Brief explanation of drawings]

FIG. 1 is a timing diagram schematically representing a pulse sequence for measuring the first-order term of static magnetic field inhomogeneity in the magnetic resonance imaging apparatus of the present invention.

[Fig. 2] A diagram for explaining the principle of calculating the correction amount by the static magnetic field inhomogeneity measurement sequence according to the present invention.

FIG. 3 is a block diagram showing the overall configuration of the present invention and a conventional nuclear magnetic resonance imaging apparatus.

[Fig. 4] An explanatory diagram showing the behavior of nuclear spin to explain the imaging principle of a magnetic resonance imaging device.

[Figure 5] Timing diagram schematically representing the pulse sequence of the spin echo method, which is a typical two-dimensional Fourier imaging method in a general magnetic resonance imaging device.

[Figure 6] An explanatory diagram showing the influence that appears on images when there is primary non-uniformity in the slice axis.

[Figure 7] An explanatory diagram showing the influence that appears on images when there is first-order non-uniformity in the phase encode axis.

[Figure 8] An explanatory diagram showing how the influence appears when a non-uniform linear term is superimposed on the phase encode axis

[Figure 9] An explanatory diagram showing the influence that appears on an image when there is first-order non-uniformity on the frequency encode axis.

FIG. 10 is an explanatory diagram showing the region excited by the static magnetic field nonuniform linear term measurement sequence of the present invention.

[Figure 11] Explanatory diagram showing phase diffusion caused by static magnetic field inhomogeneity in the gradient echo method

FIG. 12 is an explanatory diagram showing phase diffusion caused by static magnetic field non-uniformity in the static magnetic field non-uniform linear term measurement sequence according to the present invention.

[Explanation of symbols]

1 Subject 2 Magnetic field generator 3 Magnetic field gradient generation system 4 Transmission system 5　　　　　　　Reception system 6 Signal processing system 7 Sequencer 8 CPU 9 Gradient magnetic field coil 10 Gradient magnetic field power supply 14a Transmitting side high frequency coil 14b Receiving side high frequency coil

Claims (6)

[Claims] 1. Static magnetic field generating means for applying a static magnetic field to a subject; gradient magnetic field generating means for applying a gradient magnetic field to the subject; a sequencer that repeatedly applies high-frequency pulses in a predetermined pulse sequence; a transmission system that irradiates a high-frequency magnetic field to cause nuclear magnetic resonance in the nuclei of the biological tissue of the subject using the high-frequency pulses from the sequencer; A receiving system for detecting echo signals emitted by nuclear magnetic resonance, a signal processing system for performing image reconstruction calculations using the echo signals detected by the receiving system, and means for displaying the obtained image. In a nuclear magnetic resonance imaging apparatus that obtains tomographic images by repeatedly measuring echo signals emitted by nuclear magnetic resonance, the sequencer performs a pulse sequencer that measures the first-order term of the static magnetic field inhomogeneity prior to the main measurement. - Operate the sensor and obtain the data obtained from the measurement.
1. A magnetic resonance imaging apparatus, comprising means for calculating a non-uniformity correction amount from data and applying a gradient magnetic field to which the non-uniformity correction amount is added during main measurement. 2. When measuring the correction amount, only a predetermined region located at the center of the static magnetic field of the subject is selectively excited, and the correction amount is measured using only the echo signal from the region. 2. The magnetic resonance imaging apparatus according to claim 1, wherein the magnetic resonance imaging apparatus improves correction accuracy. 3. The selection of the correction amount measurement area is realized by sequentially selectively exciting three planes orthogonal to each other and generating echo signals only from the rectangular parallelepiped in which excitation overlaps. Item 2. The magnetic resonance imaging apparatus according to item 2. 4. The magnetic resonance imaging apparatus according to claim 2, wherein the selection of the correction amount measurement area is performed by applying a plurality of presaturation pulses. 5. The selective excitation region has a static magnetic field uniformity of 10p.
3. The magnetic resonance imaging apparatus according to claim 2, wherein the magnetic resonance imaging apparatus is limited to a region below pm. 6. The selective excitation region has a size of 1 cm x 1 cm x 1
3. The magnetic resonance imaging apparatus according to claim 2, wherein the magnetic resonance imaging apparatus has a cubic shape of cm.