JPH05329126A - Magnetic resonance imaging apparatus - Google Patents

Abstract

PURPOSE:To improve the total SN ratio by increasing the frequency of integrating high echoes without so much increase in image pickup time. CONSTITUTION:In a magnetic resonance imaging apparatus which is adapted to reconstruct an image from an NMR signal to be obtained based on the repetition of a specified pulse sequence, the pulse sequence is so arranged to generate at least two echo signals or more while regarding the magnitude of the intensity of a phase encode inclined magnetic field and that of an inclined magnetic field of a phase offset which is applied immediately before the application of a reading inclined magnetic field, and the number of encodes on a space K becomes less with higher echoes on the space K. At the same time, the frequency of integration is increased with all the higher echoes.

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 obtaining a tomographic image of a subject by utilizing a magnetic resonance (NMR) phenomenon.

[0002]

2. Description of the Related Art A magnetic resonance imaging apparatus utilizes a magnetic resonance phenomenon to measure the density distribution, relaxation time distribution, etc. of nuclear spins (hereinafter simply referred to as "spins") at a desired examination site in a subject. The cross section of the subject is displayed as an image from the measurement data.

In this device, as shown in FIG.
The subject 7 is placed in a static magnetic field generator 4 that generates a static magnetic field of about 2 to 2 tesla. At this time, the spins in the subject perform a precession motion with the direction of the static magnetic field as an axis at a frequency determined by the strength H ₀ of the static magnetic field. This frequency is called the Larmor frequency. Where the Larmor frequency ν ₀ is

[0004]

## EQU1 ## ν ₀ = γ / 2πH ₀ (1) Here, γ is a gyromagnetic ratio and has a unique value for each type of nucleus. If the angular velocity of the Larmor precession is ω ₀ ,

[0005]

Since there is a relationship of ω ₀ = 2πν ₀ (2),

[0006]

## EQU3 ## ω ₀ = γH ₀ (3)

Here, when a high frequency pulse (electromagnetic wave) having a frequency equal to the Larmor frequency ν ₀ of the atomic nucleus to be measured by the high frequency irradiation coil 11 is applied, spins are excited and a transition is made to a high energy state. When this high-frequency pulse is cut off, the spin returns to its original low energy state. The electromagnetic wave emitted at this time is received by the high-frequency receiving coil 14, amplified by the amplifier 15, and waveform-shaped.
It is digitized by the / D converter 17 (hereinafter referred to as ADC) and sent to the central processing unit 1 (hereinafter referred to as CPU).
The CPU 1 performs reconstruction calculation based on this data, and the calculated data is displayed on the display 18 as a tomographic image of the subject 7. The above high frequency pulse is generated by the CPU1.
The signal sent from the sequencer 2 controlled by the above is amplified by the high-frequency transmission coil amplifier 10 and sent to the high-frequency transmission coil 11.

In addition to the static magnetic field 4 and the high frequency pulse described above, the MRI apparatus is provided with a gradient magnetic field coil group 21 for producing a gradient magnetic field for obtaining positional information in space. These gradient magnetic field coils 13 are supplied with current from the gradient magnetic field coil power supply 12 which operates by a signal from the sequencer 2 to generate a gradient magnetic field.

Here, the imaging principle of the MRI apparatus will be described. As shown in FIG. 10A, an atomic nucleus placed in a static magnetic field H ₀ in the Z direction behaves like a single bar magnet in classical physics, and has the Larmor frequency ν ₀ described above. Precessing around the Z axis. This frequency is given by the equation (2) and is proportional to the strength of the static magnetic field.
Γ in the equations (1) and (3) is called a gyromagnetic ratio and has a value unique to the nucleus. In general, the number of nuclei to be measured is enormous, and each rotates in an arbitrary phase. Therefore, when viewed as a whole, the components in the XY plane cancel each other out, and the macroscopic magnetization of only the Z direction component occurs. Remain. In this state, when a high frequency magnetic field H ₁ having a frequency equal to the Larmor frequency ν ₀ is applied in the X direction (FIG. 10 (b)), the macroscopic magnetization begins to fall in the Y direction. This tilting angle is approximately proportional to the product of the amplitude of H ₁ and the application time, and H ₁ when tilting 90 ° with respect to the pulse application time is called a 90 ° pulse, and H ₁ when tilting 180 ° is called a 180 ° pulse.

Now, a two-dimensional Fourier imaging method is a method which is generally used in imaging by an MRI apparatus at present. FIG. 9 shows a schematic pulse sequence of a typical spin echo method among these methods. In this pulse sequence, first, a 90 ° pulse 28 is applied, and then a 180 ° pulse 29 is applied after a time of Te / 2 when the echo time is Te. After the 90 ° pulse 28 is applied, each spin starts rotating in the XY plane at its own velocity, 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 thereafter, the spin is focused at time Te to continue rotating at the same speed, and an echo signal is formed.

Further, FIG. 11 shows a schematic diagram of a pulse sequence called so-called multi-echo measurement. In this multi-echo measurement, a 180 ° pulse is further applied to converge spins at times 2Te, 3Te, ... And the echo signal is measured.

Although the signals are sequentially measured as described above, the spatial distribution of the signals must be obtained in order to form a tomographic image. For this purpose, a linear gradient magnetic field is used. A spatial magnetic field gradient can be created by superimposing a gradient magnetic field on a uniform static magnetic field. As described above, since the spin rotation frequency is proportional to the magnetic field strength, the spin rotation frequencies are spatially different in the state where a gradient magnetic field is applied. Therefore, the position of each spin can be known by examining this frequency. For this purpose, a phase encode gradient magnetic field 31 and frequency encode gradient magnetic fields 33 and 34 are used.

Using the pulse sequence described above as a basic unit, the intensity of the phase-encoding gradient magnetic field is changed each time, and a predetermined number of times, for example 256 times, are repeated at a constant repetition time (TR). The spatial distribution of macroscopic magnetization can be obtained by subjecting the measurement signal thus obtained to two-dimensional inverse Fourier transform. The basic principle of such MRI is detailed in, for example, "NMR Medicine" (Basic and Clinical) (edited by Nuclear Magnetic Resonance Medicine, Maruzen Co., Ltd., issued January 20, 1984).

Next, the high speed spin echo method will be described. The high-speed spin echo method distributes each echo of multi-echo measurement to the K space of raw data, thereby further shortening the imaging time. Here, a sequence diagram of the high speed spin echo is shown in FIG. In FIG. 8, the high-speed spin echo method is performed by applying a 180-degree pulse a plurality of times as in multi-echo measurement. Then, the signal of each echo is distributed in the phase direction in the raw data space by applying a predetermined gradient magnetic field in the phase direction for a predetermined time for each echo signal measurement. As shown in FIG. 7, the distribution method is such that the data of each echo occupies almost the same region in the K space ((a) in the same figure). This is because when the four-echo measurement is performed as described in the spin echo method, four data of one echo, two echoes, three echoes, and four echoes are sequentially measured as a set. The order of arrangement is determined by the echo number used and the echo number determining the image quality. To decide, the direct current part of the K space (near the center) is the data of the desired echo time, and the adjacent part is
Use adjacent echo numbers. In addition, it is assumed that the end number on the K space and the other end have the same or adjacent echo numbers.

FIG. 7 shows how data is arranged in the K space when the echo time of 4 echoes is emphasized by using up to 4 echoes. The data of each echo number is almost the same. First, the desired four-echo data 27 is allocated to the low range (near the center). The adjacent area has three echo data 26, and further has two echo data 25 and one echo data 24 as it goes to a higher region (end portion). As a result, the highest region (end) has one echo and one echo, and the same echo number. When the desired echo number is an echo with another echo time, barrel shift of FIG. 7 may be performed in the kx direction so that the desired echo number is centered. Here, the number of times of integration is the same for each echo ((c) in the same figure).

[0016]

However, in the magnetic resonance imaging apparatus constructed as described above, the SN ratio in the fast spin echo method is determined by the SN ratio of the echo data used. In other words, if a little data with a bad SN ratio is added, the overall SN ratio will decrease.

Therefore, as described above, particularly when the SN ratio of the high echo is low (FIG. 7B), this results in a decrease in the overall SN ratio. Furthermore, only high echo S
Increasing only the number of times the high echo data is accumulated to increase the N ratio is meaningless because the low echo data is discarded. This is because when measuring high echo data, low echo data is also measured as a pair, and this increases the total number of integrations, and the imaging time increases in proportion to the number of integrations. Will be done.

Therefore, the present invention has been made under such circumstances, and an object of the present invention is to increase the number of integration of high echoes without significantly increasing the imaging time, It is an object of the present invention to provide a fast spin echo method that can improve the overall SN ratio.

[0019]

In order to achieve such an object, the present invention is basically a magnetic resonance imaging apparatus for reconstructing an image from an NMR signal obtained based on repetition of a predetermined pulse sequence. In the above, the pulse sequence generates at least two echo signals, and the magnitude of the phase encode gradient magnetic field strength and the magnitude of the phase offset gradient magnetic field applied immediately before the read gradient magnetic field are applied. , The higher the number of echoes in the K space, the smaller the number of encodings in the K space, and the higher the number of echoes, the greater the number of integrations.

[0020]

In the magnetic resonance imaging apparatus configured as described above, the number of times of integration is increased as the echo becomes higher. Therefore, the high echo obtained in the state where the SN ratio is low can improve the SN ratio by increasing the number of times of integration.

Therefore, since the K space can be constructed by the low echo having a high SN ratio and the high echo having an improved SN ratio, the overall SN can be improved.

Moreover, as described above, in order to form the K space by making the number of times of integration different between the high echo and the low echo, the number of repetitions of the sequence may be slightly increased to obtain many echoes. Therefore, the overall SN ratio can be improved without significantly increasing the imaging time.

[0023]

DESCRIPTION OF THE PREFERRED EMBODIMENTS A schematic configuration of a magnetic resonance imaging apparatus to which the present invention is applied will be described with reference to FIG.

This magnetic resonance imaging apparatus is roughly classified into a central processing unit (CPU) 1, a sequencer 2 and
The transmission system 3, the static magnetic field generating magnet 4, the reception system 5, and the signal processing system 6 are provided.

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 data acquisition of a tomographic image of the subject 7 to the transmission system 3 and the gradient magnetic field generation system 21 of the static magnetic field generation magnet 4. It is 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, and a modulator 9 amplitude-modulates a high frequency pulse from the high frequency oscillator 8 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 for generating 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 installed. The gradient magnetic field generation system 21 controls the gradient magnetic field coil 13 and the gradient magnetic field power supply 12 which supplies a current to the gradient magnetic field coil 13 and the gradient magnetic field power supply 12 which are configured to independently apply the gradient magnetic fields in the Cartesian coordinate axis directions orthogonal to each other. It is configured by the sequencer 2.

The receiving system 5 has a receiving coil 14 as a high frequency coil, an amplifier 15 connected to the receiving coil 14, a quadrature detector 16 and an A / D converter 17, and the NMR from the subject 7 is detected. When the receiving coil 14 detects the signal,
The signal is converted into a digital amount through the amplifier 15, the quadrature detector 16, and the A / D converter 17, and the quadrature detector 1 is operated at the timing instructed by the sequencer 2.
The data is converted into the two series of collected data sampled by 6 and sent to the central processing unit 1.

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 the 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.

Next, a first embodiment of the present invention will be described.
In this embodiment, the sequencer 2 captures a tomographic image by the so-called high-speed spin echo method. The sequence is, for example, as shown in FIG.

The NMR signal obtained by repeating such a sequence is stored in the K space as shown in FIG.

That is, in FIG. 1, for example, the number of times of integration of one echo is set to 1, and the number of times of integration of 2 echoes, 3 echoes, and 4 echoes is set to 2, 3, and 4, respectively, and data is stored in the K space. .. Here, if the number of times of integration is two, the number of data finally obtained will be halved by averaging the data obtained twice. Similarly, when the number of times of integration is 3, the number of data finally obtained becomes 1/3 by adding and averaging the data obtained 3 times. For this reason, the ratio of the data areas of 1 echo, 2 echoes, 3 echoes, and 4 echoes in the K space is 4: respectively.
It becomes 3: 2: 1.

From this, it is possible to increase only the number of times of high echo integration. Further, in the pulse sequence, the kx direction of the K space takes discrete values. In FIG. 1, at the center in the kx direction, the integrated value of the gradient magnetic field in the phase direction sensed by the spin before the measurement of the echo at the center becomes zero. At a position deviated from the center by one in the kx direction, a gradient magnetic field in the phase direction is felt so that the spin makes one rotation at both ends of the imaging region (FOV). In the formula,

[0034]

## EQU00004 ## 2.pi. =. Gamma..SIGMA. (Gpn.tn. (FOV)) (4). Here, γ is the gyromagnetic ratio, Gpn is the intensity of the nth phase-encoding direction gradient magnetic field (however, since it is the intensity felt by the spin, the sign is inverted once for each 180 ° pulse), and tn is The application time of the nth phase direction gradient magnetic field, Σ, n, is the number of gradient magnetic fields applied in the phase direction until the echo signal of the target echo number is measured, and (FOV) is the imaging The length of one side of the area. Similarly, the mth data from the center in k space is

[0035]

(5) 2πm = γΣ (Gpn · tn · (FOV)) (5) The gradient magnetic field strength in the phase encode direction and the application time are determined so as to hold.

The raw data thus picked up is reconstructed by the two-dimensional Fourier transform method in the conventional manner.

The arrangement order changes depending on the number of echoes used and the echo number of the desired echo time. It goes without saying that the application time, intensity, etc. of the gradient magnetic field in the phase direction are changed according to the arrangement.

In the magnetic resonance imaging apparatus according to the above-described embodiment, the number of times of integration is increased as the echo becomes higher. Therefore, the high echo obtained in the state where the SN ratio is low increases the SN due to the increase in the number of times of integration.
The ratio can be improved.

Therefore, since the K space can be composed of the low echo having a high SN ratio and the high echo having an improved SN ratio, the overall SN can be improved.

The image measuring time is, for example, 4 in 1 measurement in the conventional idea of obtaining an image of 4 times integration with 256 × 256 pixels.
Even if a pulse sequence for obtaining echoes was used, 256 times of measurement were required, but in the present embodiment, 4 echoes are made 4 times, 3 times.
Since the echo is obtained 3 times, the 2 echoes 2 times, and the 1 echo is obtained once, it is possible to obtain 10 times of the conventional 256 times of measurement.
The data of 1 pixel is obtained by the measurement of / 16 for 160 times.

FIG. 3 shows the relationship between the arrangement order of data on the K space and the number of times of integration according to another embodiment of the present invention. In this case, the number of times of integration of one echo and the number of two echoes are the same, that of three echoes is twice the number of times of integration, and four echoes are the number of times of integration of three times. In this case, the ratio of the areas on the K space is determined according to the reciprocal of the number of times of integration.

FIG. 4 shows the relationship between the arrangement order of data on the K space and the number of times of integration according to another embodiment of the present invention. In this case, the number of times of integration of 1 echo and 2 echoes is the same, and the number of times of integration of 3 echoes and 4 echoes is doubled. In this case, the ratio of the areas on the K space is determined according to the reciprocal of the number of times of integration.

Further, FIG. 6 shows K of another embodiment of the present invention.
The relationship between the order of data in space and the number of times of integration is shown. As it is, only the four-echo measurement is shown, but this is an example performed on the data of the two-echo measurement.

Further, as the arrangement order, only the number of echoes to be used is shown as 2, 4 echoes, but it goes without saying that the same can be done for 5 or more echoes and 3 echoes. In addition, the fast spin echo method can be applied even when the central echo changes.

In the present invention, the fast spin echo method shown in FIG. 8, that is, spin excitation is performed at 90 ° -180 ° -1.
80 ° -... Applicable to other than those performed in series. Figure 1
2 is a schematic diagram of a pulse sequence called the CPMG method. In this CPMG method, a 90 ° pulse is applied together with the application of a slice direction gradient magnetic field to selectively excite spins, and a high frequency pulse for generating a slice direction gradient magnetic field and an echo signal after a predetermined time after spin diffusion, for example, Te / 2. Is applied, the spin is rotated by 180 ° in the 90 ° excitation plane, then the phase is encoded, the signal is read out in the frequency direction, and then, as shown in FIG. 12, an echo signal is generated by applying a high frequency pulse and a gradient magnetic field. Is read in sequence. The present invention can also be applied to such a CPMG method.

[0046]

As is clear from the above description,
According to the magnetic resonance imaging apparatus of the present invention, it is possible to increase the number of times of high echo integration and improve the overall SN ratio without significantly increasing the imaging time.

[Brief description of drawings]

FIG. 1 is an explanatory diagram of an embodiment of the present invention and is an explanatory diagram of a data arrangement order in a K space using data of up to 4 echoes.

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

FIG. 3 is an explanatory diagram of another embodiment of the present invention and is an explanatory diagram of a data arrangement order in the K space using data of up to 4 echoes.

FIG. 4 is an explanatory diagram of another embodiment of the present invention, and is an explanatory diagram of a data arrangement order in the K space using data of up to 4 echoes.

FIG. 5 is a schematic explanatory diagram of a pulse sequence using data of up to 4 echoes according to an embodiment of the present invention.

FIG. 6 is an explanatory diagram of another embodiment of the present invention, and is an explanatory diagram of a data arrangement order in the K space using data of up to two echoes.

FIG. 7 is an explanatory diagram of a data arrangement order on a K space in a conventional high-speed spin echo method.

FIG. 8 is a schematic explanatory diagram of a pulse sequence of a conventional high-speed spin echo method.

FIG. 9 is a schematic explanatory diagram of a pulse sequence of the spin echo method.

FIG. 10 is an explanatory diagram of macroscopic magnetization.

FIG. 11 is a schematic explanatory diagram of a pulse sequence of the multi-echo method.

FIG. 12 is a schematic explanatory diagram of a pulse sequence of the CPMG method.

[Explanation of symbols]

2 ... Sequencer, 7 ... Subject, 8 ... High-frequency oscillator, 1
2 ... gradient magnetic field power supply, 13 ... gradient magnetic field coil, 21 ... gradient magnetic field generation system, 24 ... first echo signal, 25 ... second echo signal, 26 ... third echo Signal, 27 ... fourth echo signal.

Claims (1)

[Claims] 1. A magnetic resonance imaging apparatus for reconstructing an image from an NMR signal obtained based on repetition of a predetermined pulse sequence, wherein the pulse sequence generates at least two echo signals, and The magnitude of the phase encode gradient magnetic field strength or the magnitude of the phase offset gradient magnetic field applied immediately before the application of the read gradient magnetic field, or the higher the echo in the K space, the smaller the number of encodes in the K space, In addition, the magnetic resonance imaging apparatus is characterized in that the number of times of integration is increased as the echo becomes higher.