JP3163125B2 - MRI equipment - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to nuclear magnetic resonance (hereinafter referred to as "N").
MRI (Magnetic Resonance Imaging) that obtains a tomographic image of a desired part of a subject (human body) using the phenomenon
More particularly, the present invention relates to an MRI apparatus capable of measuring a large number of images having a small slim thickness and photographing the images in a short time.

[0002]

2. Description of the Related Art A magnetic resonance imaging apparatus measures the density distribution, relaxation time distribution, and the like of nuclear spins (hereinafter, simply referred to as spins) at a desired examination site in a subject by utilizing an NMR phenomenon. An image of an arbitrary cross section of the subject is displayed based on the data. The conventional magnetic resonance imaging apparatus includes a static magnetic field generating magnet that applies a static magnetic field to a subject, a magnetic field gradient generating system that applies a gradient magnetic field to the subject, and an atomic nucleus of an atom constituting a living tissue of the subject. A sequencer that repeatedly applies a high-frequency pulse that causes nuclear magnetic resonance in a predetermined pulse sequence, and a high-frequency magnetic field that causes the atomic nuclei of the atoms constituting the living tissue of the subject to generate nuclear magnetic resonance by the high-frequency pulse from the sequencer A transmission system for irradiating the nuclear magnetic resonance, a reception system for detecting an echo signal emitted by the above-described nuclear magnetic resonance, and a signal processing system for performing an image reconstruction operation using the echo signal detected by the reception system. Measurement of echo signals emitted by magnetic resonance is repeatedly performed to obtain a tomographic image.

In this apparatus, a subject is placed in a static magnetic field generating magnet that generates a static magnetic field of about 0.02 to 2 Tesla. At this time, the spins in the subject perform precession at a frequency determined by the strength H ₀ of the static magnetic field with the direction of the static magnetic field as an axis. This frequency is called the Larmor frequency, and the Larmor frequency ν ₀ is

(Equation 1) Here, there is a unique value for each type of nucleus represented by H ₀ : static magnetic field strength γ: gyromagnetic ratio. When the angular velocity of the Larmor precession is ω ₀ , since ω ₀ = 2πν ₀ ,

(Equation 2) Given by

When a high-frequency magnetic field (electromagnetic field) having a frequency f ₀ equal to the Larmor frequency ν ₀ of the nucleus to be measured by a high-frequency irradiation coil in the transmission system is applied, spins are excited and transit to a high energy state. When the high-frequency magnetic field is terminated, the spin returns to the original low energy state with a time constant corresponding to each state. The electromagnetic wave emitted at this time is received by a high-frequency receiving coil in the receiving system, the image is reconstructed by the CPU, and a tomographic image of the subject is displayed on a display (hereinafter, referred to as a CRT). The magnetic resonance imaging apparatus includes a gradient magnetic field coil for generating a gradient magnetic field for obtaining positional information in a space, in addition to the static magnetic field and the high-frequency magnetic field. These gradient magnetic field coils generate a gradient magnetic field.

Here, the imaging principle of the magnetic resonance imaging apparatus will be described with reference to FIG. First, FIG.
As shown in (a), the nucleus placed in the static magnetic field H ₀ in the Z direction behaves like a single bar magnet in classical physics, and at the Larmor frequency ν ₀ described above, the Z axis Precession around. This frequency is given by [Equation 2] and is proportional to the strength H _{0 of the} static magnetic field. Γ in [Equation 1] and [Equation 2] is called a gyromagnetic ratio and has a value specific to an atomic nucleus. In general, the number of nuclei to be measured is enormous, and each of them rotates at an arbitrary phase. Therefore, when viewed as a whole, components in the XY plane cancel each other out, and macroscopic magnetization only in the Z-direction component is lost. Remains. As shown in FIG. 4 (b) in this state, applying a high frequency magnetic field H ₁ of frequency equal to the Larmor frequency [nu ₀ in the X direction, the macroscopic magnetization begins falling in the Y direction. This fall angle is proportional to the product of the amplitude and application time of the high frequency magnetic field H _1, the high-frequency magnetic field H ₁ by 90 ° pulse when fall 90 ° with respect to pulse application time point, the RF magnetic field H ₁ at which fall 180 ° 180 ° Called pulse. Note that X, Y, and Z in FIGS.
The three axes are Cartesian coordinate axes orthogonal to each other.

A method generally used for imaging in such a magnetic resonance imaging apparatus is a two-dimensional Fourier imaging method. FIG. 5 is a timing diagram schematically showing a pulse sequence of a typical spin echo method in the two-dimensional Fourier imaging method. FIG.
3A shows an irradiation timing of a signal of a high-frequency magnetic field and an envelope for selectively exciting a slice position of a subject. (B) diagram shows the timing of the application of the gradient magnetic field G _z in the slice direction, indicating that the measuring by changing the application timing and amplitude (c) drawing the phase encode direction gradient magnetic field G _y, (d ) the figure shows the timing of the application of the frequency encoding direction gradient magnetic field G _x. (E) shows the measured echo signal (NMR signal).

In the pulse sequence shown in FIG. 5, first,
After the application of 90 ° pulse, it added 180 ° pulse at the time of substantially T _e / 2 when the echo time was set to T _e. After the 90 ° pulse is applied, each spin starts to rotate in the XY plane at its own speed, and a phase difference occurs between the spins with the passage of time. Here, when a 180 ° pulse is applied, each spin reverses symmetrically to the X axis as shown in FIG. 6, and thereafter continues to rotate at the same speed.
Spin converge again at time T _e shown in, to form an echo signal as shown in FIG. (E). Although the signal is measured as described above, the spatial distribution of the signal must be obtained in order to form a tomographic image. 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, since the spin rotation frequency is proportional to the magnetic field strength, the spin frequency of each spin is spatially different when a gradient magnetic field is applied. To perform two-dimensional imaging, phase information and frequency information are given to spins, and the position of each spin can be known by examining them. For this purpose, the gradient magnetic field G shown in FIG.
_y and a frequency encoding direction gradient magnetic field _Gx are used.

[0008] as a basic unit a pulse sequence described above, every predetermined repetition time T _r while changing the strength of the phase encoding direction gradient magnetic field G _y each, repeated a predetermined number of times, for example 256 times. By subjecting the measurement signal thus obtained to two-dimensional inverse Fourier transform, a spatial distribution of macroscopic magnetization shown in FIG. 4A is obtained. In the above description,
If the three types of gradient magnetic fields do not overlap each other, X, Y, Z
Or a combination thereof. The basic principle of the magnetic resonance imaging described above is described in detail in "NMR Medicine (Basic and Clinical)" (edited by the Society for Nuclear Magnetic Resonance Medicine, Maruzen Co., Ltd., issued on January 20, 1984). .

[0009]

However, in the two-dimensional Fourier imaging method, since the selected area reconstructs one image as it is, when the slice thickness is reduced, the intensity of the echo signal becomes weaker, the SN ratio decreases, and the number of additions decreases. Must be increased. Furthermore, in order to reduce the slice thickness, the gradient magnetic field strength at the time of selecting a region increases, and the burden on hardware increases.

On the other hand, a three-dimensional Fourier imaging method shown in FIG. 7 is an effective imaging method for measuring an image having a small slice thickness. In this imaging method, the sequence of FIG. 5 is repeated several times for the selected region (slab) while changing the intensity of the slice encoding gradient magnetic field _Gz by the number as shown in A. Thereby, the slice direction is encoded, and an image of a slice having a small slice encode number is obtained by the FFT in the slice direction. The S / N of three-dimensional Fourier imaging is proportional to the square root of the slab thickness, and a high S / N can be obtained even with a small slice thickness by increasing the number of slice encodes.

A comparison between slice selection in the three-dimensional Fourier imaging method and multi-slice in the two-dimensional Fourier imaging method will be described with reference to FIGS. 8 (a) and 8 (b). FIG. 8A is a diagram showing slimming by the two-dimensional Fourier imaging method. Each portion indicated by oblique lines indicates a slice, and FIG. 8A shows an example of four multi-slices. FIG. 8B shows slimming by the three-dimensional Fourier imaging method, in which the selected area is one slab.
As can be seen from FIG. 8A, continuous slices are realized. That is, the three-dimensional Fourier imaging method is a method of exciting a spin in a region (slab) corresponding to a plurality of slices in the two-dimensional Fourier imaging method to obtain a plurality of gapless tomographic images. Slice thickness d
Can be considered in the same way as phase encoding in two-dimensional Fourier imaging,

[0012]

D = π / γ · G _sE · T γ: Proton gyromagnetic ratio G _sE : Maximum slice encode gradient magnetic field intensity T: Slice encode application time Therefore, as for the slice thickness, a thin slice of 1 to 2 mm can be obtained by increasing the number of slice encodes without changing the applied amount (intensity × time) of the slice selection gradient magnetic field, and the slice gap becomes completely zero. , Without overlooking small lesions.

However, in the above-described three-dimensional Fourier imaging method, since the sequence of the two-dimensional Fourier imaging method is repeated several times, the imaging time is several times the slice encoding time of the two-dimensional Fourier imaging method. This causes pain in the subject and causes a problem in diagnosis.

According to the present invention, when the three-dimensional Fourier imaging method is adopted, the MRI can shorten the imaging time.
It is intended to provide a device.

[0015]

According to the present invention, a gradient magnetic field in the slice direction and a frequency encoding gradient magnetic field are applied alternately in positive and negative directions while changing their magnitudes, and a plurality of echo signals generated thereby are measured. And

[0015]

According to the present invention, for each arbitrary gradient magnetic field in the phase encoding direction, a positive / negative magnetic field is alternately applied in the slice encoding direction by changing the gradient magnetic field in the series of slice directions. A gradient magnetic field is applied, and a plurality of echo signals in the slice direction can be detected. If this operation is performed for the gradient magnetic fields in all the phase encoding directions, an image can be obtained by the three-dimensional Fourier imaging method. Thus, the imaging time can be reduced.

[0017]

FIG. 2 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 using a nuclear magnetic resonance (NMR) phenomenon, and as shown in FIG. 2, a static magnetic field generating magnet 2, a magnetic field gradient generating system 3, a transmission 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 the body axis direction or in a direction perpendicular to the body axis. A permanent magnet type, normal conduction type or superconducting type magnetic field generating means is arranged. 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 each gradient magnetic field coil, and according to a command from a sequencer 7 described later. By driving the gradient magnetic field power supplies 10 of the respective coils, gradient magnetic fields G _x , G _y , and G _z in the three axes of X, Y, and 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 atoms constituting the living tissue of the subject 1 in a predetermined pulse sequence.
It operates under the control of the CPU 8 and sends various commands necessary for data collection of tomographic images of the subject 1 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 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 12
, A high-frequency amplifier 13 and a high-frequency coil 14a on the transmission side. After being amplified by 13 and supplied to the high-frequency coil 14a arranged close to the subject 1, the subject 1 is irradiated with electromagnetic waves. Receiving system 5
Is for detecting an echo signal (NMR signal) emitted by nuclear magnetic resonance of an atomic nucleus of a living tissue of the subject 1.
The reception-side high-frequency coil 14b, the amplifier 15, the quadrature-phase detector 16, and the A / D converter 17 are provided. ) Is detected by the high-frequency coil 14b arranged close to the subject 1, and the amplifier 1
5 and input to an A / D converter 17 via a quadrature detector 16 to be converted into a digital quantity.
The data is converted into two series of collected data sampled by the quadrature phase detector 16 at a timing according to a command from the controller, and the signal is sent to the signal processing system 6. The signal processing system 6 includes a CPU 8, a recording device such as a magnetic disk 18 and a magnetic tape 19, and a display 20 such as a CRT. The CPU 8 performs processing such as Fourier transform and correction coefficient calculation image reconstruction. A signal intensity distribution of an arbitrary section or a distribution obtained by performing an appropriate operation on a plurality of signals is imaged and displayed on the display 20 as a tomographic image. In FIG. 1, the high-frequency coils 14a and 14b on the transmission side and the reception side and the gradient coil 9
Is installed in the magnetic field space of the static magnetic field generating magnet 2 arranged in the space around the subject 1. Here, in the present invention, the sequencer 7 can excite the imaging region of the subject with the high-frequency magnetic field before applying the high-frequency magnetic field or the pulse of each axis gradient magnetic field and measuring the echo signal.

Next, the measurement sequence of the imaging method in the magnetic resonance imaging apparatus according to the present invention will be described with reference to FIGS.
This will be described in (f). (A) shows an irradiation timing of a signal of a high-frequency magnetic field and an envelope for selectively exciting a slice position of a subject. (B) diagram shows the timing of the application of the gradient magnetic field G _z slice direction and rice indicates that measured by changing the application timing and amplitude (c) drawing the phase encode direction gradient magnetic field G _y, ( d) The figure shows the gradient magnetic field G in the frequency encoding direction.
The timing of _x application is shown. (E) shows the measured echo signal (NMR signal).
(F) shows the time sequence divided into sections IX. A sequence for performing the high-speed three-dimensional Fourier imaging method in the spin echo method will be described with reference to FIG. In section I, a 90 ° selective excitation pulse is irradiated as a high-frequency magnetic field as shown in FIG. 7A, and a slice-direction gradient magnetic field G _z1 is applied as shown in FIG. Next, in section II, applying a phase encoding magnetic field gradient G _y as shown in (c) FIG. Section I
In II, a 180 ° selective excitation pulse is radiated as shown in (a), and a slice-direction gradient magnetic field G _z2 is applied as shown in (b).

The operations in the above sections I to III are the same as in the conventional two-dimensional Fourier imaging method. In the present embodiment, in the following sections VI to | G _z3 | <| G _z4 | <
It is characterized in that a series of gradient magnetic fields in the slice direction set as | G _z5 | <... <| G _zn | The positive / negative is alternately used in order to repeat the recombination of spins, thereby shortening the echo generation timing. Also, _let G _z be | G _z3
<| G _z4 | <…
This is because, when a plurality of echo signals are obtained by one excitation, the size of an echo generated later becomes small, which is to be prevented. Here, the magnitude of the magnetic field is (intensity x application time)
It is a value determined by Therefore, in section IV, (b)
As shown in the figure, a gradient magnetic field G _z3 in the slice direction is applied. In the section V, the frequency encoding gradient magnetic field G _x1 shown in (d) is applied and the echo signal S ₁ shown in (e) generated by G _z3 is measured. Next, in the section VI, as shown in (b), the sign is inverted with respect to G _z3 , and the application time of the gradient magnetic field G _z4 in the slice direction is longer than that in the section IV. In the section VII, the frequency encoding gradient magnetic field G _x2 shown in FIG. (D) is applied, and the echo signal S ₂ shown in FIG. (E) generated by G _z4 is measured. Next, after section VIII, V _{z is changed} to V _z5 , V _z
z6 ,..., the same operation as in section IV to section V is performed for the number of times (n
/ 2-1) is repeated, and the echo signals S ₃ , S ₄ ,..., _Sn are measured. The encoding number n is 2, 4, 8, 16,
Various settings such as 32, 64 and 128 are possible. Information obtained by encoding the phase in the slice direction is obtained from the n echo signals.

The encoding of the phase in the Y direction is performed by G shown in FIG.
_By changing the intensity of _y as G _y1 , G _y2 ,..., the section I to the section IX are repeated several times for phase encoding. Thus, a three-dimensional Fourier imaging method in a shorter time than before can be realized. The signal in this embodiment is expressed by K _x -K
_When expressed by _z (corresponding to ω _x −ω _z on Fourier space), it is shown in FIG. FIG. 3B shows a conventional three-dimensional Fourier imaging measurement. FIG. 3 (a),
(B) Both arrows correspond to echo signals. In FIG. 3A, the arrow _points from the inside to the outside because | G _z3 <|
G _z4 | corresponds to the fact you have set <... as G _z, the arrow is in the right direction → left → right direction → left → ... is to apply a G _z in the positive and negative alternately Corresponding to being.
Therefore, FIG. 3A shows an example of detection of ten echo signals. On the other hand, in FIG. 3 (b), an arrow is a one-way, which corresponds to that applied only positive direction G _z without sign applied, and, FIG. 3 seven in (b) 4 shows an example of detection of an echo signal.

Another embodiment will be described with reference to FIG. A fast three-dimensional Fourier imaging method using a gradient echo will be described. In the gradient echo method, the pulse of the selective excitation pulse by the high frequency magnetic field shown in FIG. The sections I to III shown in FIG. 1 correspond to the sections I to II in FIG. 9 and are measured by one excitation. Obtaining an image by the spin echo method as well as a slice magnetic gradient G _z and the frequency encode gradient magnetic field G _x measured signal by applying alternately positive and negative three-dimensional imaging method shown above.

[0023]

As described above, the present invention realizes high-speed three-dimensional Fourier imaging by reducing the number of repetitions of a sequence by alternately and positively and negatively applying a gradient magnetic field in the slice direction and a frequency encoding gradient magnetic field to measure a signal. did.

[Brief description of the drawings]

FIG. 1 is a timing diagram schematically illustrating a time sequence of signal measurement in an imaging method in a magnetic resonance imaging apparatus of the present invention.

FIG. 2 is a block diagram showing the overall configuration of the present invention and a conventional magnetic resonance imaging apparatus.

FIG. 3 is a diagram showing a K-space representation of magnetic resonance imaging.

FIG. 4 shows the behavior of macroscopic magnetization in magnetic resonance imaging.

FIG. 5 is a timing diagram schematically showing a time sequence of signal measurement in an imaging method in a conventional magnetic resonance imaging apparatus.

FIG. 6 is a diagram showing the behavior of spins when generating an echo signal in magnetic resonance imaging.

FIG. 7 is a comparison diagram between a three-dimensional Fourier imaging method and a two-dimensional Fourier imaging method.

FIG. 8 is a timing diagram schematically illustrating a time sequence of signal measurement in a three-dimensional imaging method in a conventional magnetic resonance imaging apparatus.

FIG. 9 is a timing diagram schematically illustrating a time sequence of signal measurement in an imaging method according to another embodiment of the present invention.

[Explanation of symbols]

G _z slice direction gradient magnetic field G _y phase encoding direction gradient magnetic field G _x frequency encode direction gradient magnetic field

──────────────────────────────────────────────────続 き Continuation of the front page (58) Field surveyed (Int.Cl. ⁷ , DB name) A61B 5/055 JICST file (JOIS)

Claims (1)

(57) [Claims] 1. An MRI apparatus for performing three-dimensional measurement by changing a slice direction gradient magnetic field and changing a phase encode direction gradient magnetic field, and applying a slice direction gradient magnetic field and a high-frequency magnetic field for selective excitation; Means for applying an arbitrary gradient magnetic field in the phase encoding direction after the selective excitation, and slicing a positive / negative alternating gradient magnetic field in a slice direction and a positive / negative alternating gradient magnetic field in a frequency direction for reading after applying the phase encoding direction gradient magnetic field. An MRI apparatus comprising: means for continuously applying a number of times corresponding to the number of times; and means for detecting an echo signal when a gradient magnetic field is applied alternately in the frequency direction.