JPH0751245A - Magnetic resonance imaging device - Google Patents

Abstract

PURPOSE:To improve the S/N of the image by making the multification value when the first echo signal is taken in at K times, and making the first echo signal 1/k times prior to operating the echo signal at each slice surface, in the system to photograph by using the doubling method of the slice sheet number. CONSTITUTION:In an MRI device, the output signal of a receiver coil 14 to detect an MRI signal from a body to examine 7 is input to a CPU 1 by converting into a collecting data in two series by a DC phase detector 16, in the timing by the instruction from a sequencer 2 through an amplifier 15, the DC phase detector 16, and an A/D converter 17. In this case, the amplifier 15 can convert the amplification level into two stages, for example, and the amplification value when the first echo signal is taken in is made k times of the amplification value when the other echo signal is taken in. And prior to the operation of the echo signal at each slice surface, the first echo signal is made 1/k times, so as to reduce the influence of the noise superposed to the second echo signal which has a small strength.

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 for performing imaging using a so-called double slice number method.

[0002]

2. Description of the Related Art In a so-called doubling of the number of slices in a magnetic resonance imaging apparatus, when an echo signal is taken out by sequentially changing phase encoding, one slice is obtained by phase encoding of one of odd number and even number. And the echo signals from the other slice are amplified in the same phase and extracted as the first echo signal, and further, the echo signals from the one slice and the other slice are extracted by the phase encoding of the other number of times. The phase is reversed, the second echo signal is amplified and taken out, and the image on each slice plane is obtained by the arithmetic processing of each echo signal.

A magnetic resonance imaging apparatus using such a method of doubling the number of slices is disclosed in, for example, [Journal of Japanese Radiation Technology; 1992, Vol. 48, No.
3, P.I. 544, abstract 342].

[0004]

However, the magnetic resonance imaging apparatus configured as described above amplifies the signals from two slices with the same amplification degree when the signals are acquired with the same phase and when the signals are acquired with the opposite phases. Then, these signals were arithmetically processed.

Therefore, it has been pointed out that a tomographic image with an insufficient S / N is obtained.

That is, the first echo signal obtained when the signals of the two slices have the same phase is a relatively strong signal because it is the addition signal of the signals from the two slices. The second echo signal obtained when the phases are made different becomes a relatively weak signal because it becomes a differential signal of the signals from the two slices.

Therefore, even if it can be said that the influence of noise superimposed on one strong signal is relatively small, the influence cannot be ignored when the same noise is superimposed on the other weak signal. Therefore, the S / N ratio of the image becomes insufficient due to this noise.

Therefore, the present invention has been made under such circumstances, and an object thereof is to provide a magnetic resonance imaging apparatus capable of improving the S / N of an image. In particular.

[0009]

In order to achieve such an object, the present invention, when the echo signal is taken out by sequentially changing the phase encoding, the phase encoding which is one of an odd number and an even number is performed. Means for reconstructing an image using these two types of signals by performing different excitation methods on the other number of times of phase encoding, and as a result, performing a measurement method in which the intensities of the two types of signals obtained by both excitations are different In the magnetic resonance imaging apparatus including the above, the amplification value when capturing the first echo signal is k times the amplification value when capturing the other echo signal (k is a positive number)
And a means for multiplying the first echo signal by 1 / k before calculating the echo signal on each slice plane.

Further, according to the present invention, when the echo signal is taken out by sequentially changing the phase encode, the signal from one slice plane and the other slice plane is obtained by the phase encode of one of odd number and even number. The first echo signal having the same phase is amplified and extracted, and the second echo signal having the opposite phase of the signals from the one slice plane and the other slice plane is further amplified by phase encoding of the other number of times. And a magnetic resonance imaging apparatus having means for calculating the image of the one slice plane and the image of the other slice plane based on the first echo signal and the second echo signal and separately obtaining them. In the above, the means for increasing the amplification value when the first echo signal is captured to a value k times (k is a positive number) the amplification value when capturing the other echo signal, and Wherein prior to calculating the echo signal first
And a means for multiplying the echo signal of 1 / k times.

[0011]

In the magnetic resonance imaging apparatus thus constructed, first, the second echo signal whose intensity is smaller than that of the first echo signal is amplified more than the amplification of the first echo signal.

As a result, the influence of noise superimposed on the second echo signal having a small intensity as has been conventionally seen can be significantly reduced. This is because the peak value of the amplified second echo signal becomes extremely large with respect to the superimposed noise.

As a result, it becomes possible to almost ignore the noise that has entered the circuit until the calculation for image construction after the amplification of each echo signal by the amplifier, as in the case of the first echo.

In the calculation for the image construction, the amplification of the second echo signal is returned to the original by an amount larger than the amplification of the first echo signal in order to eliminate the trouble. .

As described above, since the influence of noise can be reduced, the S / N ratio of the image can be improved.

[0016]

FIG. 2 is an overall block diagram showing an embodiment of a magnetic resonance imaging apparatus according to the present invention.

The magnetic resonance imaging apparatus shown in FIG.
When roughly classified, a central processing unit (CPU) 1, a sequencer 2, a transmission system 3, a static magnetic field generating magnet 4, a reception system 5,
It is composed of a signal processing system 6.

The central processing unit (CPU) 1 controls each of the sequencer 2, the transmission system 3, the reception system 5, and the signal processing system 6 according to a predetermined program. The sequencer 2 operates based on a control command from the central processing unit 1 and transmits various commands necessary for collecting data of a tomographic image of the subject 7, a transmission system 3, a gradient magnetic field generation system 21 of a static magnetic field generation magnet 4, It is designed to be sent to the receiving system 5.

The transmission system 3 has a high frequency oscillator 8, a modulator 9 and an irradiation coil 11 as a high frequency coil. The high frequency pulse from the high frequency oscillator 8 is amplitude-modulated by the modulator 9 according to a command from the sequencer 2. The amplitude-modulated high-frequency pulse is amplified through the high-frequency amplifier 10 and supplied to the irradiation coil 11, so that the subject 7 is irradiated with a predetermined pulsed electromagnetic wave.

The static magnetic field generating magnet 4 is adapted to generate a uniform static magnetic field around the subject 7 in an arbitrary direction. Inside the static magnetic field generating magnet, in addition to the irradiation coil 11, a gradient magnetic field coil 13 for generating a gradient magnetic field and a receiving coil 14 of the receiving system 5 are arranged. The static magnetic field generation system 21 has a gradient magnetic field coil 13 having a configuration capable of independently applying a gradient magnetic field in mutually orthogonal Cartesian coordinate axis directions, a gradient magnetic field power supply 12 that supplies a current to the gradient magnetic field coil 13, and the gradient magnetic field. The sequencer 2 controls the power supply 12.

The receiving system 5 includes a receiving coil 14 as a high frequency coil and an amplifier 1 connected to the receiving coil 14.
5, a quadrature detector 16, and an A / D converter 17. When the receiving coil 14 detects the NMR signal from the subject 7, the signal is converted into a digital amount via the amplifier 15, the quadrature phase detector 16, and the A / D converter 17, and at the timing instructed by the sequencer 2. The quadrature detector 16 converts the sampled data into two series of collected data and sends the collected data to the central processing unit 1.

In this embodiment, in particular, the amplifier 15 in the receiving system 5 can switch its amplification degree in two stages, for example.
The switching of the amplification degree is made by a command from the CPU 1. The switching in this case will be described in detail later.

The signal processing system 6 has an external storage device such as a magnetic disk 20 and an optical disk 19 and a display 18 such as a CRT. When data from the receiving system 5 is input to the central processing unit 1, The central processing unit 1 executes processing such as signal processing and image reconstruction, and the subject 7
The desired cross-sectional image is displayed on the display 18 and recorded on the magnetic disk 20 or the like of the external storage device.

The sequencer 2 stores a sequence as shown in FIG. 3, for example. This figure shows a sequence called a so-called spin echo method.

RF indicates the application timing of a high frequency magnetic field pulse for exciting spins, Gs, Gp, and Gf are gradient magnetic field pulses for slice position selection and gradient magnetic field pulses for position determination in the phase encoding direction, respectively. The application timing of the gradient magnetic field pulse for position determination in the frequency encoding direction is shown, and Signal shows an echo signal.

In such a pulse sequence, first, 9
After the 0 ° pulse 28 is applied, when the echo time is Te, the 180 ° pulse 29 is added after a time of Te / 2.
After the 90 ° pulse 28 is applied, each spin starts rotating in the XY plane at an inherent velocity proportional to the magnetic field strength that it receives, so that a phase difference occurs between the spins over time. Here, when the 180 ° pulse 29 is applied, each spin is inverted symmetrically with respect to the x-axis, and since it continues to rotate at the same speed and at the same speed, the spin converges again at time Te and forms an echo signal.

In such signal measurement, in order to form a tomographic image, the spatial distribution of the signal must be obtained, and for this purpose, a gradient magnetic field having a spatially linear intensity distribution is used. Since the spin rotation frequency is proportional to the magnetic field strength that the spin receives, the spin rotation frequency is spatially different in the state in which a gradient magnetic field is applied.
Therefore, the position of each spin can be known by examining this frequency. A phase encode gradient magnetic field 36 and frequency encode gradient magnetic fields 37 and 38 are used for this purpose.

With such a sequence as a basic unit, the intensity of the phase-encoding gradient magnetic field is changed every time, and a predetermined number of times, for example, 256 times are repeated at a constant repetition time (TR). Further, the spatial distribution of macroscopic magnetization can be obtained by subjecting the measurement signal thus obtained to two-dimensional Fourier transform.

Although several methods have been proposed to realize the doubling method of the number of slices, the present invention can be applied to any of these methods, and in the present embodiment, it is disclosed in Japanese Patent Laid-Open No. 4-3.
Take the method described in 57935 as an example. In this method, two types of different excitation wave envelopes shown in FIGS. 4A and 4B are used. FIG. 4A is a result of inverse Fourier transform of a rectangular area when frequency is represented on the horizontal axis as shown in FIG.
For example, it substantially matches three cycles.

Y = sinX / X (1) On the other hand, in FIG. 4B, the opposite signs adjacent to each other when the frequency is represented on the horizontal axis as shown in FIG. 5B. (2), which is the result of the inverse Fourier transform of the two regions, substantially coincides with several cycles, for example, three cycles.

Y = (1-cosX) / X (2) The state of the NMR signal obtained by the high frequency magnetic field pulse thus irradiated will be described with reference to FIG.

In the same figure, the Real axis shows the intensity of the NMR signal generated from the excitation region, and the Z axis shows the slice position direction or the frequency direction of the excitation high-frequency magnetic field pulse.

First, considering the case where the first high-frequency magnetic field pulse (hereinafter referred to as RF1) having the envelope shown in FIG. 4A is irradiated, the obtained signal is as shown in FIG.
As shown in (a).

[0034] When the reference frequency of the RF1 and f _0, the width corresponding to the waveform data and the high frequency pulses and gradient field strength that is the same applied around the reference frequency f ₀ (e.g., ±
The region with Δf) is excited. At this time, the signal A-1 generated from the excited region (f ₀ −Δf) and the region (f ₀
The phase of the signal A-2 generated from + Δf) is the same,
For example, both are on the positive side of the real axis as shown in the figure.

On the other hand, the second high frequency magnetic field pulse (hereinafter referred to as RF2) having the envelope shown in FIG.
6), the obtained signal is as shown in FIG. 6 (b).

Assuming that the reference frequency of RF2 is f ₀ , a region having a width of ± Δf around the reference frequency f ₀ is excited as in the case of RF1, but the region (f ₀ -Δ
The phase of the signal B-1 generated from f) and the phase of the signal B-2 generated from the area (f ₀ + Δf) are opposite to each other. For example, if the positive side of the real axis of the signal B-1 is used as shown in FIG. -2 is the negative side of the real axis.

Here, the signal A obtained in the first measurement can be expressed by the following equation.

A = ∫dz [PA (z) · ρ (z)] ≈k · Aa + l · Ba (3) The signal B obtained by the second measurement can be expressed by the following equation.

B = ∫dz [PB (z) · ρ (z)] ≈m · Aa + n · Ba (4) However, where z is the coordinate in the slice direction ρ (z), the density of magnetization of the object to be imaged PA is the thrust profile of the first measurement. PB is the thrust profile of the second measurement. Aa is a double slice. Absolute value of the signal generated from one slice of the object to be imaged. Ba is the absolute value of the signal generated from the other slice. It is assumed that k, l, m, and n are complex numbers whose absolute value is 1 or less, and the cross product of the vectors (k, l) and (m, n) is not zero.

Using equations (3) and (4), the slices Aa and Bb can be separated as follows.

Aa = (m · A−1B) / (n · k · l · m) (5) Ba = − (m · A−k · B) / (n · k−1 · m) ) (6) Here, considering an ideal case, since k = 1, l = 1, m = 1, and n = -1, equations (5) and (6) are as follows. Become.

Aa = (A + B) / 2 (7) Ba = (A−B) / 2 (8) As shown in these two equations, the sum and difference of two types of signals A and B are calculated. By obtaining it, it becomes possible to separate the images of the two slices.

Here, the signal A expressed by the above equation (3) is used.
Has a waveform as shown in FIG. 1 (a), and the signal B expressed by the equation (4) has a waveform as shown in FIG. It is found that the signal B has a significantly smaller peak value than the signal A. This is because the signals are obtained by applying high-frequency magnetic field pulses whose phases are different by 180 °.

In this case, the signals A and B (see FIG.
(Amplified by the amplifier 15 shown in 1) is used as it is to form an image of a slice, there arises a problem that the S / N of the image becomes insufficient due to superposition of noise.

That is, the signal A shown in FIG.
When the noise shown in (d) is superimposed, there is no particular problem because the peak value of the signal A is sufficiently large, but the signal B in FIG.
When the noise shown in (e) is superimposed, a problem occurs because the peak value of the signal B is small.

Therefore, in this embodiment, in particular, this signal B
1, the amplification degree is switched in the amplifier 15 shown in FIG.
Amplification is performed as shown in (c), and the peak value thereof is made substantially equal to the peak value of the signal A, for example.

In this case, even when the noise shown in FIG. 1 (f) is superimposed on the amplified signal B ',
Since the peak value of the signal B ′ is sufficiently large, there is no particular problem.

In FIG. 1, (d),
The noises shown in (e) and (f) are usually the same in many cases.

In the calculation for the image configuration in the CPU 1, it is difficult to use the amplified signal B'as it is, so that the peak value of the signal B'is returned to the original signal B. Has become.

FIG. 7 is an explanatory view showing a procedure for applying the above-mentioned high frequency pulse in the sequence in the above-mentioned spin echo method. In the figure, the state of application of the high-frequency excitation pulse when the number of integration times is 1
~ It shows up to the fourth Echoondo. RF1 is applied in the first encoding, and RF2 is applied in the second encoding.
Similarly, RF1 is applied to the odd-numbered encodes and RF2 is applied to the even-numbered encodes. In addition,
The figure also shows that an echo signal is generated for each encoding.

Further, FIG. 8 is an explanatory view showing the application states of the high frequency excitation pulse when the number of times of integration is n times, from the first to the fourth encoding. In the first encoding, RF1 is set to n
It is applied repeatedly. RF2 in the second encoding
Is repeatedly applied n times. Below, RF1 is used for odd-numbered encoding and R is used for even-numbered encoding.
By applying F2 n times and adding and averaging the signals of n times, even if the number of times of integration is n, the conventional number of slices is 2
A signal similar to that of the excitation method of the doubling method can be obtained.

According to the magnetic resonance imaging apparatus of the above-described embodiment, first, the first echo signal (the high frequency magnetic field pulse having one frequency and the high frequency magnetic field pulse having another frequency are applied with the same phase). A second echo signal (high frequency magnetic field pulse having one frequency and high frequency magnetic field pulse having another frequency) whose intensity is smaller than that of the echo signal obtained by The echo signal obtained by applying so as to have a phase of 90 ° is amplified more than the amplification of the first echo signal.

As a result, the influence of noise superimposed on the second echo signal having a small intensity as seen in the past can be greatly reduced. This is because the peak value of the amplified second echo signal becomes extremely large with respect to the superimposed noise.

As a result, it is possible to almost ignore the noise that has entered the circuit until the calculation for image construction after the amplification of each echo signal by the amplifier, as in the case of the first echo.

In the calculation for the image construction, the amplification of the second echo signal is made larger than the amplification of the first echo signal in order to eliminate the trouble. .

As described above, since the influence of noise can be reduced, the S / N ratio of the image can be improved.

In the above-mentioned embodiment, when the second echo signal is amplified, the peak value thereof is made substantially the same as that of the first echo signal, but the present invention is not limited to this. The point is that it may be set to be larger than the amplification factor of the first echo signal.

[0058]

As is apparent from the above description,
According to the magnetic resonance imaging apparatus of the present invention, the S / N of an image can be improved.

[Brief description of drawings]

1A to 1F are schematic explanatory views showing an embodiment of a magnetic resonance imaging apparatus according to the present invention.

FIG. 2 is an overall configuration diagram showing an embodiment of a magnetic resonance imaging apparatus according to the present invention.

FIG. 3 is an explanatory diagram showing an example of a sequence of the magnetic resonance imaging apparatus according to the present invention.

4 (a) and 4 (b) are diagrams showing excitation waveforms used in the slice number doubling method in the magnetic resonance imaging apparatus according to the present invention.

5 is a diagram showing a frequency distribution of the high frequency magnetic field pulse envelope shown in FIG. 4 before inverse Fourier transform.

6A and 6B are diagrams showing macroscopic magnetization in a subject obtained by using the slice number doubling method in the magnetic resonance imaging apparatus according to the present invention.

FIG. 7 is an explanatory diagram showing an example of a procedure of applying a high frequency pulse in a sequence of the magnetic resonance imaging apparatus according to the present invention.

FIG. 8 is an explanatory view showing another embodiment of the procedure for applying a high frequency pulse in the sequence of the magnetic resonance imaging apparatus according to the present invention.

[Explanation of symbols]

 1 CPU 15 Amplifier

Claims (2)

[Claims] 1. When extracting an echo signal by sequentially changing phase encoding, different excitation methods are performed for one of the odd-numbered and even-numbered phase encoding and the other number of phase encoding, and as a result, both excitations are performed. In a magnetic resonance imaging apparatus equipped with a means for reconstructing an image using these two types of signals by performing a measuring method in which the intensities of the two types of signals obtained by A means for setting the amplified value to a value k times (k is a positive number) the amplified value when the other echo signal is taken in, and the first echo signal is set to 1 before calculating the echo signal on each slice plane. / K times, and a magnetic resonance imaging apparatus. 2. When the echo signals are extracted by sequentially changing the phase encode, the signals from one slice plane and the other slice plane have the same phase in one of the odd and even phase encodes. Amplifying and extracting the first echo signal, further amplifying and extracting the second echo signal in which the signals from the one slice plane and the other slice plane have opposite phases in the phase encoding of the other number of times, A magnetic resonance imaging apparatus comprising means for calculating the image of the one slice plane and the image of the other slice plane based on the first echo signal and the second echo signal and separately obtaining them. Means for setting the amplification value when capturing the first echo signal to a value k times (k is a positive number) the amplification value when capturing the other echo signal, and the echo signal at each slice plane Magnetic resonance imaging apparatus characterized by comprising a means for the first echo signal to 1 / k times before operation.