JPH01115349A - Nuclear magnetic resonance imaging apparatus - Google Patents

Abstract

PURPOSE:To set NMR receiving sensitivity so as to make the input to an A/D converter optimum, by performing preparatory measurement and carrying out control so that a nuclear magnetic resonance (NMR) signal in final measurement becomes max. at a time when phase encoding quantity is 0 by reversely correcting the measured variation quantity. CONSTITUTION:When the change of amplitude at the time origin in the measure ment of an NMR signal in such a case that the magnitude of the NMR signal is taken on a vertical axis and phase encoding quantity is taken on a transverse axis, the phase encoding quantity is not 0 and signal quantity becomes max. at a point Gy0. Then, when the quantity of Gy0 is predetermined by measurement and the offset current corresponding to -Gy0 is allowed to reversely flow during final measurement as the inclined magnetic field in a phase encoding direction from the obtained primary variation component Gy0, said primary variation component is negated and the NMR signal received by a receiving system 15 takes the max. value when the phase encoding quantity is 0. Therefore, when the input attenuation quantity of an A/D converter 25 is adjusted in this state, a measured signal does not overflow.

Description

DETAILED DESCRIPTION OF THE INVENTION [Industrial Application Field] The present invention relates to nuclear magnetic resonance (hereinafter abbreviated as NMR) I!
The present invention relates to an NNR imaging apparatus that obtains a tomographic image of a subject using a single image.

[Conventional technology]

In an NMR imaging device, the uniformity of the static magnetic field is important because it affects the geometric distortion and phase distortion of the reconstructed image. However, conventionally, the uniformity of the static magnetic field is rarely adjusted except during installation or periodic inspections every few years.

When a superconducting system (electromagnet) is adopted as the static magnetic field generator, once the static magnetic field uniformity is adjusted, the static magnetic field uniformity will not deviate significantly from the practical allowable limit for imaging. However, when a permanent magnet is used in a static magnetic field generator, not only the central magnetic field strength but also the static magnetic field uniformity can easily change due to changes in the temperature of the surrounding environment. especially,
When a Nd-Fe-B magnet is used as the main magnet, the temperature coefficient is larger than that of other magnets, so the change is noticeable.

By the way, if the static magnetic field uniformity is sufficiently ensured,
The optimal sensitivity for tomographic imaging is usually found at a phase encode of 0 in the pulse sequence implemented for tomographic imaging.
It is determined based on the signal strength at However, if the homogeneity of the static magnetic field is insufficient, a first-order variational component of the magnetic field will appear, so it cannot be said that the signal strength is maximum when the phase encode is 0.

Now, the direction of the static magnetic field is taken as two axes, and each spatial axis is X.

The first-order variations in the y and z directions are respectively δB2/δ8°δ
Assuming B°2/δy and δB2/δ2, the static magnetic field B assuming a certain space is B ” B z as a first-order approximation.
o+δB, /δx”X+δB, /δ76 y+δB2/
It is expressed as δ2.2 (1). Here, Bzo indicates the zero-order Z-direction static magnetic field strength.

In equation (1), good static magnetic field uniformity means: When the static magnetic field uniformity deteriorates, the relationship in equation (2) no longer holds true.

This first-order variational component, δBz/δ8. δB2/δa.

Since δB2/δ2 has exactly the same dimension as the gradient magnetic field strength, it can be canceled by the gradient magnetic field.

For example, when it is assumed that the first-order variation δB2/δ2 of the static magnetic field in the Z-axis direction is not 0, if the phase encoding direction is taken in the Z-axis direction, it cannot be said that the signal strength is maximum when the phase encoding is 0. Rather, the NMR signal strength is at its maximum when the phase encode amount exactly matches -δBz/δ2.

[Problem that the invention seeks to solve]

From this, when the static magnetic field homogeneity collapses and the first-order variation of the static magnetic field can no longer be ignored, it is correct to optimally set the NMR signal reception strength based on the signal strength at phase encode 0. do not have. With this setting, when the phase encode amount matches -δB2/δ2, the signal becomes maximum, and this detection signal is converted into an analog/digital converter (hereinafter referred to as A
/D converter) will overflow.

As described above, the conventional technology does not determine the optimal reception sensitivity by taking into account the first-order variation of the static magnetic field in the phase encoding direction based on static magnetic field inhomogeneity.

In the present invention, when the static magnetic field inhomogeneity collapses and the first-order variation of the static magnetic field cannot be ignored, the magnitude of the first-order variation is prepared in advance of cross-sectional surface photography (hereinafter referred to as main measurement). By performing measurement and inversely correcting the measured variation amount, the input to the A/D converter is optimized by implementing control such that the NMR signal in one measurement is maximized when the phase encode is 0. The purpose of the present invention is to provide a technique that can set the NMR reception sensitivity so that

[Means for solving problems]

The object is to provide a means for applying a static magnetic field to a subject, a means for applying a slice direction gradient magnetic field, a frequency encoding gradient magnetic field, and a phase encoding gradient magnetic field to the subject; A nuclear magnetic resonance imaging apparatus comprising means for applying a high frequency pulse that causes magnetic resonance, means for detecting a nuclear magnetic resonance signal, and means for Fourier transforming the obtained nuclear magnetic resonance signal to reconstruct an image. In addition to the main measurement pulse sequence, a slice direction gradient magnetic field and a frequency encoding gradient magnetic field are applied immediately before the start of measurement, and 90' and 180° selective excitation high frequency pulses are repeatedly applied at predetermined time intervals, and the repetition time is A means is provided for generating a preliminary measurement pulse sequence for determining the phase encode amount that gives the maximum value of the nuclear magnetic resonance signal by changing the phase encode gradient magnetic field at each time, and executes the preliminary measurement pulse sequence at the time of the main measurement. This is achieved by adding the phase encode amount determined by the method to the phase encode gradient magnetic field.

[Effect]

In the preliminary measurement scan, when one modulus of the static magnetic field is generated due to static magnetic field inhomogeneity, if phase encoding is taken in this direction, the NMR The signal never reaches its maximum value.

For example, in this state, the NM for the main measurement (tomography)
If the R signal receiving sensitivity is set optimally, the NMR signal will take the maximum value when the phase encode is other than 0, so the A/D
The input to the converter will overflow and correct image reconstruction will not occur. However, the virtual phase encoding O
Even if the NMR signal reaches its maximum value elsewhere,
Since the phase encode amount that gives the maximum value can be known by the preliminary measurement scan, in one measurement, if the phase encode amount that gives the maximum value of the NMR signal in the preliminary measurement is added to the phase encode gradient magnetic field with the sign reversed, , it is possible to prevent the input to the A/D converter from overflowing during the main measurement.

〔Example〕

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

FIG. 1 is a block diagram showing an example of the overall configuration of a nuclear magnetic resonance imaging apparatus according to the present invention. This nuclear magnetic resonance imaging apparatus uses a nuclear magnetic resonance (NMR) phenomenon to obtain a tomographic image of a subject 1. For this purpose, a static magnetic field generating magnet 1o with a sufficiently large bore diameter and a center Processing bag r11 (hereinafter referred to as CPU) 11, sequencer 12, transmission system 13, and magnetic field gradient generation system 1
4 and.

It consists of a receiving system 15 and a signal processing system 16. The static magnetic field generating magnet 10 generates a strong and uniform static magnetic field around the subject 1 in the body axis direction or in a direction perpendicular to the body axis, and is used to generate a strong and uniform static magnetic field around the subject 1 in a certain expanse of space around the subject 1. A magnetic field generating means of a permanent magnet type, a normal conduction type, or a superconducting type is arranged in the magnetic field. The sequencer 12 is a CP
It operates under the control of the UII and sends various commands necessary for data collection of tomographic images of the subject 1 to the transmission system 13, magnetic field gradient generation system 14, and reception system 15. Above transmission system 1
3 is a high frequency oscillator 17, a modulator 18, and a high frequency amplifier 1
9 and a transmitting-side high-frequency coil 20a, the high-frequency pulse outputted from the high-frequency oscillator 17 is amplitude-modulated by a modulator 18 according to the command from the sequencer 12, and this amplitude-modulated high-frequency pulse is amplified by a high-frequency amplifier 19. After that, the high frequency coil 20 placed close to the subject 1
By supplying the electromagnetic waves to a, the subject 1 is irradiated with electromagnetic waves. The magnetic field gradient generation system 14 is
It consists of gradient magnetic field coils 21 wound in the three axial directions of X, Y, and Z, and gradient magnetic field power supplies 22 that drive each coil, and drives the gradient magnetic field power supplies 22 of each coil according to commands from the sequencer 12. By doing so, x
, y, and z, gradient magnetic fields Gx, Ga, and Gz are applied to the subject 1. Depending on how this gradient magnetic field is applied, a slice plane for the subject 1 can be set. The receiving system 15 includes a receiving side high frequency coil 20
b, amplifier 23, quadrature phase detector 24, and A/D converter 2
5, and electromagnetic waves (NMR
signal) is a high-frequency coil 2 placed close to the subject 1.
0b, is input to the A/D converter 25 via the amplifier 23 and the quadrature phase detector 24, is converted into a digital quantity, and is further sampled by the quadrature phase detector 24 at the timing according to the command from the sequencer 12. Two series of collected data are collected, and the signals are sent to the signal processing system 16. This signal processing system 16 is a CPU
The CPU II is composed of a recording device such as a magnetic disk 26 and a magnetic tape 27, and a display 28 such as a CRT.The CPU II performs processing such as Fourier transform, correction coefficient calculation, image reconstruction, etc. The distribution obtained by performing appropriate calculations on a plurality of signals is converted into an image and displayed on the display 28. In addition, the first
In the figure, high frequency coils 20a on the transmitting side and receiving side,
20b and the gradient magnetic field coil 21 are arranged in the magnetic field space of the static magnetic field generating magnet 10 arranged in the space around the subject 1.

FIG. 2 schematically represents a time sequence in a typical spin echo measurement. In FIG. 2, RF indicates the timing of irradiation of radio frequency signals and the envelope for selective excitation. G2 indicates the timing of applying a gradient magnetic field in the slice direction. Gy
indicates that measurement is performed by changing the timing and amplitude of the gradient magnetic field application in the phase encoding direction. G8 indicates the timing of application of the frequency encoded gradient magnetic field, Signal indicates the NMR signal to be measured, and the bottom row is the time sequence divided into sections 1 to 6.

Note that the three axes x, y, and z are Cartesian coordinate axes that are orthogonal to each other. In section 1 in FIG. 2, a 90 degree selective excitation pulse is irradiated and a gradient magnetic field in the slice direction is applied. In section 2, a gradient magnetic field in the phase encoding direction is applied to add position-dependent rotation of the nuclear spins in the Y direction.

Furthermore, in section 2, a frequency encoding gradient magnetic field is applied. This is to dephase (invert the phase) the nuclear spins in advance so that the time origin is at the center of section 6 when measuring the NMR signal in section 6. In section 3, no signal is issued. In section 4, a 180 degree selective excitation pulse is irradiated and a gradient magnetic field in the slice direction is applied.

In section 5, no signal is issued. In section 6, a frequency encode gradient magnetic field is applied and an NMR signal is measured.

To perform NMR imaging, as described above, an RF pulse is irradiated with a gradient magnetic field applied to the static magnetic field, and the gradient magnetic field is applied to encode the NMR signal emitted from the examination area of the subject 1 as spatial information. is applied, NMR
After a total of d1 signals.

Reconstruct the image.

To encode space, a gradient magnetic field is used, which takes advantage of the fact that the nuclear magnetic resonance frequency ω has a linear relationship with the magnetic field strength. In other words, if the linearity of the gradient magnetic field is maintained spatially, the relationship between the spatial position and frequency in the target region will be linear. The image is reconstructed by taking advantage of the location information obtained.

Specifically, the image is reconstructed using a two-dimensional Fourier transform method, but below we will explain how to encode the space after the nuclear spins in a region with a certain thickness in the slice direction are excited by selective excitation. do.

In order to add rotation to the nuclear spin of a two-dimensional plane region with a certain thickness by an amount corresponding to the spatial coordinates,

It is encoded separately in two directions of Y. According to FIG. 2, the X direction is divided into a frequency encoding direction, and the Y direction is divided into a phase encoding direction.

In the frequency encoding direction, when reading a spin echo signal, the phase must be shifted by Nπ at both ends of the field of view, and if the frequency encoding time is Tx.

The following relationship must be satisfied: γGX-D-Tx=N7C-(3). Here, γ: gyromagnetic ratio of proton, which is the target nucleus (2,6751X 10'rad
/see/Gauss) Gx: Strength of gradient magnetic field in frequency encoding direction D: Field of view diameter N: Number of measurement samples.

In addition, in the phase encoding direction, if phase encoding is performed M times, the phase at both ends of the field of view is at most Mπ
Therefore, when the phase encode pulse application time is Ty, γGy-D-Ty=Mπ
...(4) must be satisfied. Here, Gy: Maximum value of the gradient magnetic field in the phase encoding direction M: ゛Phase encoding number. In addition, the field of view was set to a square area.

The gradient magnetic field in the frequency encoding direction is applied with the same intensity for each phase encoder, and the spatial coordinate in the X direction is encoded on the frequency axis. On the other hand, in the phase encoding direction,
The phase encode amount γGy'D-Ty is set so that the gradient magnetic field strength for each encode is as follows: (5)
The spin echo signal is measured by changing Gy so that it changes by π.

In this way, we have N samples in the x direction and M samples in the y direction.
Two-dimensional measurement data with samples is collected. usually,
QPD (Quad-ratu) is used for NMR signal measurement.
Since the real part and the imaginary part are collected simultaneously using the ve Phase Detection method, complex data of NXM samples is obtained, and an image can be obtained by performing two-dimensional Fourier transform on this.

By the way, a permanent magnet may be used as the static magnetic field generating magnet 10. Various materials have been considered for the material, but rare earth magnets (NdF
e-B) has the highest maximum energy product and can generate a strong magnetic field, but on the other hand has a large temperature coefficient.

Generally, as the ambient temperature rises, the generated static magnetic field weakens, which is what is called a negative temperature coefficient. - As an example, there are those whose temperature coefficient reaches 11000 ρpm/'C.

In this case, when the ambient temperature increases by 1°C, the static magnetic field strength decreases by 11000 pp. For example, 1000 Gauss
The static magnetic field strength of s corresponds to I Gauss.

Because the static magnetic field uniformity tends to collapse due to slight temperature changes, it is usually done by covering the magnetic circuit with a heat insulating material or installing a constant temperature control device.
The temperature coefficient of the entire magnetic circuit can be reduced.

Even if the above-mentioned measures are taken, in a magnetic circuit using permanent magnets configured in a vertical magnetic field type, a temperature difference may occur between the upper and lower magnets. This is because the lower part of the magnetic circuit is in contact with the floor, which has a large heat capacity, whereas the upper part is not.

When a temperature difference occurs between the upper and lower magnets, the uniformity of the static magnetic field within the target magnetic field space collapses, and a variational term of the static magnetic field appears in the vertical direction. In reality, higher-order variational components may occur, but the dominant component is the lower-order component.6 Also, the dominant component is the first-order component. From now on, the discussion will be limited to the first-order component.

When the direction of the static magnetic field of the static magnetic field generating magnet 10 is set as two axes,
In a vertical magnetic field type permanent magnet magnetic circuit, the vertical direction is also the vertical direction, and when a first-order variational term appears, the static magnetic field B
If the uniformity-adjusted magnetic field is BO, then B=Bo+δBZ/δ2・Z...(
6). The second term in equation (6) is a first-order variational term.

By the way, since the second term in equation (6) is the same expression as the gradient magnetic field generated by the gradient magnetic field coil 21, it can be said that correction is possible if only the amount of the coefficient δB2/δ2 of the second term is known. Therefore, a method for determining this amount will be explained below.

Assume now that the uniformity of the static magnetic field has been adjusted and the second term in equation (6) does not appear. In this state, when the gradient magnetic field Gy in the phase encoding direction is not applied, that is, when the phase encoding amount is 0, the measured NM
The R signal takes the maximum value.

Figure 3 shows the measurement sequence at this time.
The difference from FIG. 2 is that the phase encode gradient magnetic field Gy is not applied in section 2.

By the way, if the second term in equation (6) cannot be ignored, it cannot be said that the NMR signal takes the maximum value at the time of phase encoding to. Therefore, the second term of equation (6) cannot be ignored, and the A/D that digitizes the NMR signal
If the attenuation amount is adjusted to optimize the input amount to the converter 25, the input amount will overflow when the true maximum value is captured. The problem is that a signal amount other than the true maximum value is regarded as the maximum value in the actual measurement. To solve this problem, it is necessary to capture the true maximum value of the NMR signal measured in advance.

FIG. 4 shows the change in amplitude at the time origin in NMR signal measurement, where the vertical axis represents the magnitude of the NMR signal and the horizontal axis represents the phase encode amount. In FIG. 4, the phase encode amount is not 0, but the signal amount is maximum at the point Gyo.

Correction is possible if the amount of Gyo is known.

Therefore, a preliminary measurement will be performed before the main measurement. At this time, if the phase encode amount is measured in accordance with equation (5) more finely than the amount of change in Gy, for example with a step width of a fraction, the accuracy of capturing the amount of Gyo will be improved. The measurement sequence at this time is carried out according to FIG. Also, since the static magnetic field uniformity does not usually deviate significantly,
For example, the phase encode amount can be switched in the preliminary measurement.

Approximately 1/16 of the number M of phase encodes is sufficient.

From the first-order variational component Gyo obtained in this way, if an offset current corresponding to Gyo is passed during the actual measurement as a gradient magnetic field in the phase encoding direction, this negative modulus component will be canceled, so the phase When the encoding amount is 0, the NMR signal received by the receiving system 15 takes the maximum value. Therefore, if the amount of attenuation of the input to the A/D converter 25 is adjusted in this state, the measured signal will not overflow.

FIG. 5 shows -Gy when the offset current is applied.

This shows the sequence of measuring while applying the component. Thereby, it is possible to correct the first-order variational component caused by static magnetic field inhomogeneity, and to perform measurement without overflow of the maximum value. The pulse sequence for preliminary measurement is C
Needless to say, it is generated by PUI 1 and sequencer 12.

In the above detailed description, the X+ yvZ axes are interchangeable, and the content of the present invention is applicable even when imaging an oblique section where the tomographic plane has an arbitrary inclination with respect to these axes. I would like to add that.

〔Effect of the invention〕

According to the present invention, when the static magnetic field homogeneity collapses and a first-order variational component of the magnetic field appears, a total of i11'l preparatory scans are performed, and once the amount of variation is known, a component that corrects it can be used. By applying as a gradient magnetic field, it is possible to optimize the amount of input to the analog/digital converter that digitizes the NMR signal without causing overflow. This has the effect of maximizing the effective bit length of a given analog/digital converter and increasing the precision of the reconstructed image.

[Brief explanation of the drawing]

Fig. 1 is a block diagram showing an example of the overall configuration of an NMR imaging apparatus according to the present invention, Fig. 2 is a diagram showing a sequence of NMR signal measurement in two-dimensional Fourier transform imaging, and Fig. 3 shows no phase encoding gradient magnetic field applied. FIG. 4 is a diagram showing a measurement data profile viewed in the phase encoding direction, and FIG. 5 is a diagram showing a measurement sequence in which a correction component in the phase encoding direction is applied. 1... Subject, 10... Static magnetic field generating magnet, 11...
・Central processing unit, 12... Sequencer, 13 Transmission system, 14... Magnetic field gradient generation system, 15... Receiving system, 16
...Signal processing system, 17...High frequency oscillator, 18...
- Modulator, 19... High frequency amplifier, 20a... Transmitting side high frequency coil, 20b... Receiving side high frequency coil, 2
1... Gradient magnetic field coil, 22... Gradient magnetic field power supply, 2
3...Amplifier. 24... Quadrature phase detector, 25... A/D converter,
26...Magnetic disk, 27...Magnetic tape, 28
...$ 2 Mouth IδQ. $5 Mouth 180゜Maya 5 Eyes 780゜

Claims (1)

[Claims] 1. means for applying a static magnetic field to a subject; means for applying a slice direction gradient magnetic field, a frequency encoding gradient magnetic field, and a phase encoding gradient magnetic field to the subject; In a nuclear magnetic resonance imaging apparatus, the nuclear magnetic resonance imaging apparatus includes means for applying a high-frequency pulse that causes a nuclear magnetic resonance signal to occur, a means for detecting a nuclear magnetic resonance signal, and a means for reconstructing an image by Fourier transforming the obtained nuclear magnetic resonance signal, In addition to the main measurement pulse sequence, immediately before the start of the main measurement, a slice direction gradient magnetic field and a frequency concord gradient magnetic field are applied, and 90° and 180° selective excitation high-frequency pulses are repeatedly applied at predetermined time intervals, and at each repetition time. A means is provided for generating a preliminary measurement pulse sequence that changes the phase encode gradient magnetic field to obtain the phase encode amount that gives the maximum value of the nuclear magnetic resonance signal, and the phase encode amount obtained by executing the preliminary measurement pulse sequence during the main measurement is provided. A nuclear magnetic resonance imaging device characterized by adding the following to a phase encoding gradient magnetic field.