JP3274879B2 - Magnetic resonance imaging equipment - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a magnetic resonance imaging apparatus, and more particularly, to a magnetic resonance imaging apparatus using a so-called variable field imaging method.

[0002]

2. Description of the Related Art A magnetic resonance imaging apparatus uses a magnetic resonance phenomenon to measure a nuclear spin density distribution, a relaxation time distribution, and the like at a desired cross-sectional site in a subject. Is displayed as an image.

In such a magnetic resonance imaging apparatus, for example, a variable field-of-view imaging method has come to be used in recent years because a measurement time required to display an image of a cross section becomes long.

According to this variable visual field imaging method, an image display area (field area) in which a tomographic image is displayed on a display surface in a region extending vertically from a central portion thereof, and corresponds to both sides of the area. In this case, the interval of the gradient of the gradient magnetic field in the phase encoding direction is increased, and the gradient magnetic field is repeatedly applied while being sequentially changed. The number of times may be reduced.

For this reason, it is possible to reduce the time in accordance with the decrease in the number of times of the repeated application of the gradient magnetic field, and to shorten the measurement time.

[0006]

However, in the above-described magnetic resonance imaging apparatus, for example, when the subject is positioned off the center of the measurement space, it corresponds to the center of the measurement space on the display surface. Since the visual field is set so that the subject is not hidden by the non-visual areas symmetrically present on both sides of the display surface with the center at the position, the tomographic image of the subject in the visual field is expressed in reverse. Are biased to one of the left and right sides, and an unnecessary field of view is to be formed on the other side.

Therefore, the number of times of application of the gradient magnetic field in the phase direction increases accordingly, and the measurement time cannot be sufficiently reduced.

[0008] Therefore, the present invention has been made in view of such circumstances, and an object of the present invention is to provide a magnetic resonance imaging apparatus capable of greatly reducing the measurement time. It is in.

[0009]

SUMMARY OF THE INVENTION In order to achieve the above object, the present invention basically provides a method of measuring a position shift of a subject in a measurement space and a maximum width of the subject by a gradient magnetic field in a phase encoding direction. The NMR signal is extracted by setting
First means for obtaining the relationship between the frequency and the signal amount obtained from the MR signal, and second means for narrowing the measurement field of view in the phase encoding direction based on the information on the maximum width obtained by the first means; And third means for performing image processing based on the NMR signal obtained by the second means and the positional deviation information obtained by the first means so that a sectional image of the subject is positioned at the center of the display surface. It is characterized by the following.

[0010]

In the magnetic resonance imaging apparatus thus constructed, first, the first means can determine the displacement of the subject in the measurement space and the maximum width of the subject.

The first means extracts an NMR signal while keeping the gradient magnetic field in the phase encoding direction constant, and can easily detect the NMR signal from the relationship between the frequency and the signal amount obtained from the extracted NMR signal. Can be done very quickly and easily.

The second means narrows the measurement field of view in the phase encoding direction based on the maximum width. That is, the interval between the gradients of the gradient magnetic field in the phase encoding direction can be increased, and the number of times of application of the gradient magnetic field, which is sequentially changed and repeated, can be reduced (therefore, the number of NMR signals required for image processing is also reduced).

In this case, the image processing is performed by the third means so that the cross-sectional image of the subject is positioned at the center of the display surface, so that the measurement field of view in the phase encoding direction can be minimized. Therefore, the measurement time can be significantly reduced from this.

[0014]

FIG. 4 is an overall schematic block diagram showing an embodiment of a magnetic resonance imaging apparatus according to the present invention.

The magnetic resonance imaging apparatus shown in FIG.
Static magnetic field generating magnet 2, magnetic field gradient generating system 3, transmitting system 4
, A receiving system 5, a signal processing system 6, a sequencer 7, and a central processing unit (CPU) 8.

The static magnetic field generating magnet 2 generates a uniform static magnetic field around the subject 1 in a body axis direction or a direction perpendicular to the body axis, and has a certain spread around the subject 1. A permanent magnet type, a normal conduction type, or a superconducting type magnetic field generating means is arranged in the space.

The magnetic field gradient generating system 3 includes a gradient magnetic field coil 9 wound in three directions of X, Y and Z, and a gradient magnetic field power supply 10 for driving the respective gradient magnetic field coils. By driving the gradient magnetic field power supplies 10 of the respective coils in accordance with the above-mentioned commands, the gradient magnetic fields Gx, Gy, Gz in the three axial directions of X, Y, Z are applied to the subject 1. A slice plane for the subject 1 can be set by how to apply the gradient magnetic field.

The sequencer 7 repeatedly applies a high-frequency magnetic field pulse for causing nuclear magnetic resonance to the nuclei of the atoms constituting the living tissue of the subject 1 in a predetermined pulse sequence. Various commands necessary for data collection of tomographic images of the specimen 1 are sent to the transmission system 4, the magnetic field gradient generation system 3, and the reception system 5.

The transmitting system 4 irradiates a high-frequency magnetic field in order to cause a nuclear magnetic resonance in the nuclei of the atoms constituting the living tissue of the subject 1 by the high-frequency pulse sent from the sequencer 7. , Modulator 1
2, a high-frequency amplifier 13 and a high-frequency coil 14a. The high-frequency pulse output from the high-frequency transmitter 11 is amplitude-modulated by the modulator 12 according to the command of the sequencer 7, and the amplitude-modulated high-frequency pulse is amplified by the high-frequency amplifier 13, and then placed close to the subject 1. By supplying the high frequency coil 14a, the subject 1 is irradiated with an electromagnetic wave.

The receiving system 5 emits an echo signal (NMR signal) emitted by nuclear magnetic resonance of the nucleus of the living tissue of the subject 1, and includes a high-frequency coil 14b and an amplifier 1
5, a quadrature phase detector 16 and an A / D converter 17. The electromagnetic wave (NMR signal) of the response of the subject 1 due to the electromagnetic wave emitted from the high-frequency coil 14a on the transmission side is detected by the high-frequency coil 14b arranged close to the subject 1, and is transmitted through the amplifier 15 and the quadrature phase detector 16. Via A
/ D converter 17 and is converted into a digital quantity,
Further, it is made into two series of collected data sampled by the quadrature phase detector 16 at a timing according to an instruction from the sequencer 7, and the signal is sent to the signal processing system 6.

The signal processing system 6 comprises a CPU 8, a storage device such as a magnetic disk 18 and an optical disk 19, and a CRT 20 as a display. The CPU 8 performs processing such as Fourier transform and correction coefficient calculation image reconstruction. The signal intensity distribution of the cross section or the distribution obtained by performing an appropriate operation on a plurality of signals is imaged and displayed on the CRT 20 as a tomographic image.

In FIG. 4, the high-frequency coils 14a and 14b on the transmitting and receiving sides and the gradient coil 9 are
It is arranged in the magnetic field space of the static magnetic field generating magnet 2 arranged in the space around the subject 1.

In the magnetic resonance imaging apparatus having the above-described configuration, the present embodiment particularly includes the following first means, second means, and third means. .

Hereinafter, each means will be described sequentially. (1). First means, that is, means for extracting an NMR signal while keeping the gradient magnetic field in the phase encoding direction constant, and calculating a positional shift of the subject in the measurement space and the maximum width of the subject from the relationship between the frequency and the signal amount obtained from the NMR signal. about.

First, as shown in FIG. 2, the phase encoding direction and the frequency encoding direction of the cross-sectional image are switched, the sequence is executed once without applying a phase encoding gradient magnetic field, and the NMR signal 20 obtained by this is measured. By performing the one-dimensional Fourier transform on the data, a relationship between the signal strength and the frequency is obtained as shown in FIG.

In the figure, a portion where the signal value indicating the signal strength is larger than the threshold value SL corresponds to the existence range of the subject. Then, the frequency f corresponding to the threshold SL
By using L and fH, the frequency fC at the center of the subject is obtained as follows:
{(FL + fH) / 2} will be obtained.

From the previously known resonance frequency f0 at the center of the static magnetic field and the above-mentioned frequency fC at the center of the subject, the shift amount of the center of the subject (center shift amount in the phase encoding direction) is expressed as ) = (F0−fC) × 2π / (γ × Gr) where π: pi γ: gyromagnetic ratio Gr: readout gradient magnetic field strength of the sequence. Can be obtained as

On the other hand, the size S of the subject is as follows: S = (fH−fL) × 2π / (γ × Gr), where π: circumference ratio γ: gyromagnetic ratio Gr: readout gradient magnetic field intensity of the sequence. . Can be obtained as

Therefore, it is possible to determine the displacement of the subject in the measurement space and the maximum width of the subject.

(2). Second means, that is, means for narrowing the measurement field of view in the phase encoding direction based on the information of the maximum width obtained by the first means.

A so-called variable visual field imaging method is used.
In the gradient magnetic field in the phase encoding direction, when the magnetic field is applied for each switching time (TR), the amount of imaging can be shortened by narrowing the imageable area by setting the increment amount.

Before describing the variable field-of-view imaging method, a description will first be given of the case of imaging without applying the variable field-of-view imaging method.

As shown in FIG. 5A, the increment of the gradient magnetic field in the phase encoding direction is Gpstep,
As a result, as shown in FIG. 2B, the gradient magnetic field is associated with the imaging region (or imaging field of view FOV) in FIG. 2C.

FIG. 6 is an explanatory view to which the variable visual field imaging method is applied. FIGS. 6A to 6C correspond to FIGS. 5A to 5C, respectively. As can be seen from these figures, the imaging field of view in the phase encoding direction is F
Assuming that OV is satisfied, the following equation holds: γtp · Gpstep · FOV = 2π (1) Therefore, in order to reduce the field of view in the phase encoding direction to 1 / m, the increment of the phase encoding may be multiplied by m. In other words, G'pstep which satisfies γtp · G′pstep · FOV / m = 2π (2) may be used.

From equations (1) and (2), G'pstep = m.Gpstep... (3) holds. Further, the imaging time is determined by the number of updates k of the phase encoding. In order to obtain the same spatial resolution, it is sufficient that k with respect to the imaging field is constant. Assuming that k in the normal imaging is k and k ′ is the variable visual field, FOV / k = (FOV / m) / k ′ (4) holds. From equation (4), k ′ / k = 1 / m (5) Since the imaging time is proportional to the number of updates of the phase encoding, a variable field of view can reduce the time by 1 / m.

Then, the FOV / m is determined by the first means as the maximum width S = (fH−fL) × 2π of the subject.
/ Γ · Gr
Set.

(3). Third means, that is, means for performing image processing based on the NMR signal obtained by the second means and the displacement information obtained by the first means so that the sectional image of the subject is positioned at the center of the display surface. about.

First, a sequence for performing this means will be described with reference to FIG. This pulse sequence has a so-called 2
It is a typical pulse sequence of the spin echo method in the dimensional Fourier imaging method.

As shown in the drawing, the pulse sequence is as follows. First, a 90 ° pulse 30 is applied, and when the echo time is Te, the 180 ° pulse 3 is applied after Te / 2.
One is added. After the 90 ° pulse 30 is applied, each spin starts to rotate in the XY plane at a unique speed, and a phase difference occurs between the spins with the passage of time. Here, when a 180 ° pulse is applied, the spins are inverted symmetrically with respect to the X ′ axis. Thereafter, the spins converge again at time Te in order to continue rotating at the same speed, thereby forming an echo signal 36. .

The principle of generating the echo signal 36 at this time will be described with reference to FIG. 8 based on the nuclear spin. In the same figure, after the 90 ° pulse irradiation (FIG. 8 (a)), the macroscopic magnetization that has fallen on the y-axis (FIG. 8 (b)) shows that as time passes (FIG. 8 (c)), the static magnetic field Due to the inhomogeneity of, the rotational phases of individual nuclear spins start to be disturbed,
The fan spreads in the xy plane as if a fan were spread (FIG. 8).
(D)). At this time, an FID signal is observed, but no signal appears in the receiving coil thereafter.

Next, 1 after the elapse of the time τ after the application of the 90 ° pulse,
An 80 ° pulse is applied (FIG. 8E). As a result, each spin that has caused a phase shift is 18
Rotated 0 ° counterclockwise. However, even in this state, the spin that had been clockwise on the xy plane before the application of the 180 ° pulse is now clockwise as it is, and the counterclockwise spin continuously changes its phase counterclockwise. Go (Fig. 8
(F)). Therefore, each spin converges to the −y axis 2τ after application of the first 90 ° pulse (FIG. 8).
(G)). Since the behavior of each of these nuclear spins is received as a signal by the coil placed on the xy plane, when focusing on the received signal, the FID signal starts to appear immediately before reaching 2τ after applying the 90 ° pulse, and the maximum after 2τ. Reach the value.

In this case, in order to form a tomographic image, the spatial distribution of signals must be obtained. For this purpose, a linear gradient magnetic field is used. By superimposing a gradient magnetic field on a uniform static magnetic field, a spatial magnetic field gradient can be created. As described above, the spin rotation frequency is proportional to the magnetic field strength, so that the spin rotation frequency is spatially different when a gradient magnetic field is applied. Therefore, by examining this frequency, the position of each spin can be known. For this purpose, a slice direction gradient magnetic field 32, a phase encoding gradient magnetic field 33, and frequency encoding gradient magnetic fields 34 and 35 are used.

In this case, as described above, the increment G'pstep in the phase encoding gradient magnetic field 33 is set to a value determined by the second means.

Using the above-described pulse sequence as a basic unit, the intensity of the phase-encoding gradient magnetic field (the increment is G'pstep) is changed every time, and is repeated a predetermined number of times, for example, 256 times, at a constant repetition time (TR). .
The spatial distribution of macroscopic magnetization is obtained by performing a two-dimensional inverse Fourier transform on the measurement signal thus obtained. In addition, regarding the above-mentioned basic principle of MRI, "NMR medicine" (basic and clinical) (edited by Nuclear Magnetic Resonance Medical Research Society, Maruzen Co., Ltd.,
(Issued January 20, 1984).

As described above, the NMR signal obtained by using the variable field-of-view imaging method in which the magnitude of the increment Gpsstep in the phase encoding direction is made to coincide with the maximum width of the subject is directly subjected to two-dimensional Fourier transform. The resulting image is shown in FIG.
As shown in (a), when there is a shift between the center 10A of the measurement space and the center 10B of the tomographic image, FIG.
As shown in the figure, the width S is centered on the center 10A of the measurement space.
Is formed, and as a result, the tomographic image displayed in the image display area is formed as an image that is folded in the phase encoding direction.

For this reason, in this embodiment, each data S (m, n) in the memory storing the NMR signal is phase-corrected to S '(m, n) using the following equation. .

S ′ (m, n) = S (m, n) exp {−2πi (mp / FOV ′)} where FOV ′ = FOV / k.

In each of the data in the memory, m is a sample number obtained sequentially in time when the gradient magnetic field in the phase encoding direction is constant, and n is a sequentially changing gradient magnetic field in the phase encoding direction. This is the projection number when performing this.

The image of the tomographic image thus obtained is
For example, an image is displayed as shown in FIG.

In the magnetic resonance imaging apparatus constructed as in the above-described embodiment, first, the position shift of the subject in the measurement space and the maximum width of the subject can be obtained by the first means.

The first means extracts the NMR signal with the gradient magnetic field in the phase encoding direction set to 0, and can easily detect the NMR signal from the relationship between the frequency and the signal amount obtained from the extracted NMR signal. Can be done very quickly and easily.

Further, the measurement field in the phase encoding direction based on the maximum width is narrowed by the second means. That is, the interval between the gradients of the gradient magnetic field in the phase encoding direction can be increased, and the number of times of application of the gradient magnetic field, which is sequentially changed and repeated, can be reduced (therefore, the number of NMR signals required for image processing is also reduced).

In this case, since the image processing is performed by the third means so that the sectional image of the subject is positioned at the center of the display surface, the measurement field of view in the phase encoding direction can be narrowed to the minimum. Therefore, the measurement time can be significantly reduced from this.

In the above-described embodiment, the tomographic image shift correction is performed only in the phase encoding direction. However, the present invention is not limited to this, and may be performed together in the frequency encoding direction. As shown in FIG. 9A, for example, when the center of the subject in the measurement space is separated by F in the frequency encoding direction, the detection frequency of the NMR signal is changed from the selective excitation frequency F0 of the cross section to F1 satisfying the following equation. Then, F1 = F0− (F / FOV) (FH−FL) As shown in FIG. 3B, an image shifted by F in the frequency encoding direction can be obtained.

In this case, although it has nothing to do with the object of the present invention, the tomographic image to be imaged is positioned at the center in both the vertical and horizontal directions of the display surface, and is extremely easy to observe. The effect is obtained.

In the above-described embodiment, the first means executes the sequence once without applying the phase-encoding gradient magnetic field to obtain the NMR signal. However, the present invention is not limited to this. It goes without saying that the same effect can be obtained even if the sequence is executed once and the NMR signal is obtained with the phase encoding gradient magnetic field kept constant.

[0057]

As is apparent from the above description,
According to the magnetic resonance imaging apparatus of the present invention, the measurement time can be significantly reduced.

[Brief description of the drawings]

FIG. 1 is an explanatory diagram for explaining an embodiment of a third means of a magnetic resonance imaging apparatus according to the present invention.

FIG. 2 is a sequence diagram for explaining an embodiment of the first means of the magnetic resonance imaging apparatus according to the present invention.

FIG. 3 is an explanatory diagram for explaining an embodiment of the first means of the magnetic resonance imaging apparatus according to the present invention.

FIG. 4 is an overall schematic configuration diagram showing an embodiment of a magnetic resonance imaging apparatus according to the present invention.

FIG. 5 is an explanatory diagram necessary for clarifying the description of FIG. 6;

FIG. 6 is an explanatory view showing an embodiment of the second means of the magnetic resonance imaging apparatus according to the present invention.

FIG. 7 is an explanatory diagram of a sequence used for the second means of the magnetic resonance imaging apparatus according to the present invention.

FIG. 8 is an explanatory diagram for explaining movement of nuclear spins in the magnetic resonance imaging apparatus according to the present invention.

FIG. 9 is an explanatory view showing another embodiment of the magnetic resonance imaging apparatus according to the present invention.

[Explanation of symbols]

 10A: Center of measurement space, 10B: Center of tomographic image.

Claims (1)

(57) [Claims] An NMR signal is extracted from a positional deviation of a subject in a measurement space and a maximum width of the subject with a gradient magnetic field in a phase encoding direction being fixed, and is obtained from a relationship between a frequency and a signal amount obtained from the NMR signal. A second means for narrowing the measurement field of view in the phase encoding direction based on the information of the maximum width obtained by the first means.
A third means for performing image processing such that a sectional image of the subject is positioned at the center of the display surface based on the NMR signal obtained by the second means and the positional deviation information obtained by the first means. A magnetic resonance imaging apparatus comprising: