JPH0690921A - Magnetic resonance imaging device - Google Patents

Abstract

PURPOSE:To obtain a T1 emphasizing image effective in clinically without strengthening an inclined magnetic field power source by dividing an encoder into a center part and a peripheral part, and switching an encoder inclined magnetic field quantity from applied level control to the applied time control in the center part. CONSTITUTION:By driving the inclined magnetic field power source 22 of the inclined magnetic field coil 21 wound in the triaxial direction in an inclined magnetic field generating system 14 in accordance with an instruction from a sequencer 12, the inclined magnetic field in the triaxial direction is applied to an examinee 6. Subsequently, by a signal processing system 16, the processing of an image reconstitution, etc., is executed with respect to data inputted from a reception system 15, and the image of an arbitrary cross section is displayed on display 28. In this case, applied quantity control of an encoder inclined magnetic field is executed by combining two methods for which the control is executed by maintaining an applied time in a prescribed time and changing an applied level, and the other control executed by maintaining the applied level in the prescribed value and changing the applied time. Also, at the time of changing applied time, the applied time of a lead-out inclined magnetic field is also changed simultaneously.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to improvement in image quality of a magnetic resonance imaging apparatus for imaging a desired portion of a subject by utilizing magnetic resonance.

[0002]

2. Description of the Related Art Magnetic resonance imaging apparatus (hereinafter referred to as MRI
The device will be used to measure the nuclear spin density distribution, relaxation time distribution, etc. at a desired inspection site in the subject using the nuclear magnetic resonance phenomenon, and display the cross section of the subject as an image from the measured data. It is a thing. The nuclear spins of the subject placed in the uniform and strong static magnetic field generator perform precession about the direction of the static magnetic field at a frequency (Larmor frequency) determined by the strength of the static magnetic field. Therefore, when a high-frequency pulse having a frequency equal to this Larmor frequency is externally applied, spins are excited and transition to a high energy state (nuclear magnetic resonance phenomenon). When this irradiation is stopped, the spin returns to the original low energy state with a time constant corresponding to each state, and at this time, an electromagnetic wave (NMR signal) is emitted to the outside. This is detected by the high frequency receiving coil tuned to the frequency. The MRI apparatus applies a triaxial gradient magnetic field corresponding to the slice direction, the encoding direction, and the readout direction to the static magnetic field space in order to add position information to the detected signal. As a result, it is possible to separate and identify a signal from each position in the subject as frequency information.

An image reconstruction method in the MRI apparatus will be described here. FIG. 6 is an explanatory diagram of a pulse sequence in a generally used spin echo (SE) method. There are two types of irradiation pulses, 90 degrees and 180 degrees, and they are applied to the subject together with the slice gradient magnetic field to excite nuclear spins in the cross section to be imaged. A readout gradient magnetic field is applied between these two kinds of irradiation pulses to promote phase diffusion of the excited spins, and then a 180 degree pulse is applied to reverse the spin diffusion direction, and the readout gradient magnetic field is applied again. When applied, the spin converges and 90
A sharp echo signal is generated at twice the time between -180 degree pulses. This time is called the echo time Te. The echo signal obtained here has one-dimensional projected image information in the read-out direction, but a two-dimensional image cannot be constructed by this alone. Therefore, during the application of the readout gradient magnetic field that gives the phase diffusion, the gradient magnetic field is applied in the encode direction which is the other axis to give the phase rotation at the position, and the information of the encode direction is superimposed on the echo signal as the phase information. Furthermore, the encode gradient magnetic field amount is applied while being changed, and the echo signal is repeatedly measured. This repetition time is referred to as Tr, and the encode amount is given a positive or negative number with the encode No. zero being the encode gradient magnetic field amount of zero, and the encode gradient magnetic fields having different polarities are applied.

When the echo signal train thus obtained is analyzed by the two-dimensional Fourier transform means, two-dimensional image information can be obtained. The image reconstruction method by the two-dimensional Fourier transform will be referred to as a 2DFT method below.

Image enhancement that is clinically important for MRI images will be described below. The proton spin has two types of relaxation phenomena called longitudinal relaxation and transverse relaxation, which change depending on the existing environment. Due to this relaxation phenomenon, the signal strength S is calculated by the following equation.

S = ρ · (1-exp (−Tr / T1)) · (exp (−Te / T2)) where ρ is the density of existing protons. Also, T
1 and T2 are constants unique to each tissue and are values determined by the environment in which protons exist. With this formula,
The second term represents the recovery process of the signal strength with respect to the repetition period Tr, which is the vertical relaxation phenomenon (depending on the T1 value). The third term is the attenuation process of the signal intensity with respect to the echo measurement time Te and is due to the lateral relaxation phenomenon (depending on the T2 value).

FIG. 7 illustrates the relationship between the measured echo signal intensities depending on Tr and Te. As an example, as shown in FIG. 7C, the outside is the tissue A and the inside is the tissue B. Considering that, the signal intensity due to the relaxation phenomenon has the characteristics shown in FIGS. 7A and 7B, respectively. When capturing a T1-weighted image that reflects the T1 value, Te as short as possible (short Te, S
Is displayed), and Tr uses an optimum value (displayed as S, about 300 ms in the human body) so that T1 difference due to tissue can be visualized. In the case of this example, the tissue B has a shorter T1 value, so that an image in which signal recovery by vertical relaxation is faster and a high signal is obtained can be obtained. When capturing the T2-weighted image, a sufficiently long Tr (long Tr,
The signal of each tissue is recovered by using "L" (usually 1500 ms or more), and Te is set to be long (display L, about 100 ms) to perform imaging. In this example, organization A has T
Since the binary value is long, an image whose contrast is inverted from that of the T1-weighted image can be obtained. When Tr is set to be long and Te is set to be short, an image that is not affected by the T1 value and T2 value of each tissue can be obtained. This image is called a proton density image because it becomes an image based on the proton density of the tissue. Since the T1 and T2 values differ depending on the same tissue or its state (for example, tumor), it is used for identifying a lesion site. Above, MR
I described the outline of I, but the details are "NMR Medicine" (basic and clinical) (Nuclear Magnetic Resonance Medical Society, published by Maruzen Co., Ltd.
Issued January 20, 9).

[0008]

The above-mentioned image enhancement is important for obtaining a clinically effective image. In particular, in order to obtain a good T1-weighted image, Te should be set as short as possible, and high gradient magnetic field strength is required. Hereinafter, this principle will be described.

In the pulse sequence of FIG. 6, it is necessary to shorten the time during which the encode gradient magnetic field and the read-out gradient magnetic field are applied in order to shorten Te.
However, since the applied amount of the gradient magnetic field is determined by the product of the application time and the application level, when the application time is shortened,
That is, the same gradient magnetic field amount is not obtained unless the applied level is increased. Therefore, the shortest Te that can be set by the pulse sequence is limited by the maximum gradient magnetic field strength of the apparatus, that is, the maximum capacity of the gradient magnetic field power supply. The gradient magnetic field coil that generates the gradient magnetic field is an inductive load for the gradient magnetic field power source, and it is very difficult to control a large current with high accuracy. Further, such a large changing magnetic field is not preferable in terms of influence on the human body, and it can be said that the smaller the gradient magnetic field strength is, the more preferable.

The present invention solves the problem caused by such gradient magnetic field strength, and has a short Te even with a small gradient magnetic field power source capacity.
MRI that can be set to obtain a good T1-weighted image
The purpose is to provide a device.

[0011]

As described above, in order to shorten Te, it is necessary to shorten the application time of the encode gradient magnetic field and the read-out gradient magnetic field, and this means will be described with reference to FIG. In normal imaging by the 2DFT method, the encoding No. is set to ± 128 to obtain the image resolution.
Although it is set to about ± 256, here, in order to simplify the description, it is assumed that the image is reconstructed by performing 13 measurements from +6 encode to −6 encode including zero.

First, regarding the encoding gradient magnetic field, in the conventional pulse sequence (shown by a broken line), the application time is kept constant and the application level is increased or decreased to control the gradient magnetic field amount (this method is referred to as application level control hereinafter). I'll call it). Therefore, Te is constant over the entire measurement, the repetition time Tr does not change, and the total measurement time is 13 times the Tr. The maximum encoding gradient field in this conventional method
The application time is determined according to the capacity of the gradient magnetic field power supply so that (No. + 6 and No.-6) can be applied, and as a result, the shortest Te is limited.

By the way, in the signal measurement by the 2DFT method, the collected information differs depending on the range of the encoding No. This situation will be described with reference to FIG.

It is assumed that a normal image can be obtained by performing a reconstruction process using all measurement data of 13 times. here,
When the reconstruction process is performed using only seven pieces of encoding information from +3 to -3 in the central portion (hereinafter, some low encoding portions including zero encoding are referred to as the central portion),
Since only the coarse information in the encoding direction (corresponding to the low frequency component) is used, the contrast of light and dark of the image can be obtained, but fine subdivision information such as contours cannot be obtained, and the image becomes unclear. On the other hand, the peripheral part (hereinafter,
Reconstructed by inserting zero data in the central part into 6 encodes (corresponding to high frequency data) of +4 to +6 and -4 to -6, which are data of the encoded part other than the central part are referred to as peripheral parts). When the processing is performed, an image in which contour information is extracted is obtained. From this, it is understood that it is possible to almost separate the image into the contrast component and the contour component according to the used encoded data range of the reconstruction. Therefore, even in the central part that determines the contrast of the image,
If Te is set short, a T1-weighted image can be obtained.

Therefore, according to the present invention, as shown by the solid line in FIG. 1, the encoding gradient magnetic field amount is controlled by the application level control in the peripheral portion which is relatively unrelated to the image enhancement as usual, and the application level is constant in the central portion. The application time is reduced by switching to the control of changing the application time while maintaining the value (hereinafter, referred to as application time control).
To shorten the measurement. Further, the application time of the read-out gradient magnetic field needs to be shortened accordingly. However, in this case, since Te is shortened, the sampling time of the received signal is shortened and data is lost.
Although this needs to be corrected, the details will be described with reference to FIG. In the conventional sequence (a), signal sampling centered on the peak of the echo signal is performed, and it is possible to fill the entire reception data space created by collecting sampling data in each encoding. On the other hand, in the sequence (b) according to the present invention, the peak position of the echo signal moves to the front of the sampling time due to the shortening of Te, and the time margin in the first half decreases, so that the number of sampling points decreases. That is, asymmetric sampling is performed. The loss of data due to the decrease in the number of sampling points is a problem because the resolution in the read-out direction decreases. Therefore, the missing data is estimated.

The received NMR signals are in a complex conjugate relationship with each other at diagonal positions in the data space. If all the data after the peak are sampled, the first half can be estimated and interpolated by calculation. is there. Using this method, the first half is estimated from the actual missing first half sampling data so that the reception data space is filled. As shown in the figure, the zero-encoding part has the largest estimated portion, and the longer Te, the longer the sampling time and the smaller the estimated portion.

As described above, according to the present invention, the application time of the encode gradient magnetic field and the read-out gradient magnetic field shown in FIG. 1 is controlled so as to be reduced in the central portion, whereby Te and Tr are shortened and good T1 is obtained. The enhanced image can be captured in a shorter time than the conventional method. The position at which the encoding gradient magnetic field is switched from the application level control to the application time control depends on the total number of encoding points, the required image enhancement degree, the gradient magnetic field power supply capacity of the apparatus, etc., and the optimum position is selected.

[0018]

According to the present invention, the control means of the encode gradient magnetic field is switched from the level control to the time control at the central portion, and at the same time, the application time of the read-out gradient magnetic field is shortened to shorten Te and to estimate the sampling data. It is possible to capture a good T1-weighted image together with the processing in a shorter time than in the past, and efficiently obtain a good M with a small gradient magnetic field current.
An RI image can be taken.

[0019]

Embodiments of the present invention will now be described in detail with reference to the accompanying drawings. FIG. 5 is a block diagram showing an example of the overall configuration of the MRI apparatus according to the present invention. This MRI apparatus obtains a tomographic image of the subject 6 by utilizing a nuclear magnetic resonance (NMR) phenomenon, and includes a static magnetic field generating magnet 10, a central processing unit (hereinafter referred to as CPU) 11, a sequencer 12, and a transmitter. The system 13 includes a gradient magnetic field generation system 14, a reception system 15, and a signal processing system 16.

The static magnetic field generating magnet 10 gives a strong and uniform static magnetic field to the subject 6, and a permanent magnet type, a normal conducting type or a superconducting type is provided in a certain space around the subject 6. A magnetic field generating means is arranged. The sequencer 12 operates under the control of the CPU 11 and sends various commands necessary for collecting tomographic image data of the subject 6 to the transmission system 13, the gradient magnetic field generation system 14, and the reception system 15.

The transmission system 13 includes a high frequency oscillator 17, a modulator 18, a power amplifier 19, and an irradiation coil 20 on the transmission side.
The high frequency pulse output from the high frequency oscillator 17 is modulated by the modulator 18 in accordance with the instruction from the sequencer 12, and the modulated irradiation pulse is supplied to the power amplifier 19
After being amplified by, the electromagnetic waves are applied to the subject 6 by supplying them to the irradiation coil 20 arranged close to the subject 6.

The gradient magnetic field generation system 14 is composed of a gradient magnetic field coil 21 wound in three axial directions of X, Y and Z, and a gradient magnetic field power source 22 for driving each coil, and a command from the sequencer 12 is given. By driving the gradient magnetic field power sources 22 of the respective coils in accordance with the above, gradient magnetic fields Gx, Gy, Gz in the three axial directions of X, Y, Z are applied to the subject 6. The slice plane for the subject 6 can be set by the method of applying the gradient magnetic field.

The receiving system 15 comprises a receiving coil 2, a receiving circuit 23, a quadrature detector 24 and an A / D converter 25, and the subject 6 is irradiated with the electromagnetic waves emitted from the transmitting side irradiation coil 20. The response electromagnetic wave (NMR signal) is detected by the receiving coil 2 arranged close to the subject 6, is input to the quadrature phase detector 24 via the receiving circuit 23, and the quadrature phase detector 24 causes the high frequency oscillator 17 The waveform is shaped in synchronism with the output signal, and the signals of the two systems of the sin component and the cos component are separated and output. Then, the signals of these two systems are converted into digital signals by the A / D converter 25 and the signals are output. It is designed to be sent to the processing system 16.

The signal processing system 16 includes a CPU 11, a recording device such as a magnetic disk 26 and an optical disk 27, and a CRT.
The CPU 11 performs processing such as Fourier transform, correction coefficient calculation, image reconstruction, etc. on the data input from the receiving system 15, and performs an appropriate calculation on a signal intensity distribution of an arbitrary cross section or a plurality of signals. The signal intensity distribution obtained by performing the above is imaged and displayed on the display 28. In this figure, the irradiation coil 20, the receiving coil 2 and the gradient magnetic field coil 21 are
Static magnetic field generating magnet 10 arranged in the space around the subject 6
It is located in the magnetic field space.

Here, the pulse sequence according to the present invention will be described by taking the SE method as an example. FIG. 6 is a timing chart of the conventional SE method sequence. An encode gradient magnetic field and a read-out gradient magnetic field are applied between the irradiation pulses of 90 degrees and 180 degrees. The measurement is performed while changing the applied amount of the encode gradient magnetic field 256 to 512 times with positive and negative polarities.
Normally, the application time t is kept constant and controlled by changing the application level. Of course, the maximum encode gradient field level must be within the capacity of the gradient field power supply. For example, if the applied time t is 10 ms and the required phase rotation in the encoding direction is obtained when the gradient magnetic field coil current is 100 A, then 200 A of coil current needs to be passed in order to reduce t to 5 ms. There is. That is, the minimum value of the application time is determined by the current capacity of the gradient magnetic field power supply. This also determines the minimum Te limit.

FIG. 4 shows a timing chart when the present invention is applied to the SE method sequence. This figure is encoded No.
Regarding the +4 and +3 parts, this is an example in which the central part within ± 3 is switched to the application time control. The application level control is performed up to No. +4, and the application time t is constant. In No. +3, the application time control is changed, and the application time is reduced from t to t3 without changing the application level. As a result, the irradiation pulse interval can be narrowed, and Te, which is twice as long, can be shortened to Te3. In this case, the application time of the read-out gradient magnetic field also needs to be shortened similarly, but this effect will be described with reference to FIG. The phase of each spin is diffused by the read-out gradient magnetic field applied between the irradiation pulses, the direction is reversed by the 180-degree pulse, and then the diffused phase is converged by the application of the read-out gradient magnetic field again. The peak of the echo signal occurs at the time when the same amount of convergent gradient magnetic field as the amount is applied. Then, a signal is received in the applied state of the convergent gradient magnetic field and data sampling is performed. Therefore, the application time of the convergent gradient magnetic field is twice the application time of the diffusion gradient magnetic field.

In this way, symmetrical sampling of data centered on the peak of the echo signal can be performed. However, if the application time of the diffusion gradient magnetic field is shortened, the direction is reversed during the diffusion and the converging gradient magnetic field is applied,
The peak of the echo signal occurs earlier than in the previous example.
That is, Te is shortened, but as can be seen from FIG. 2, data sampling is asymmetric sampling in which the first half is omitted. If the number of sampling points is small like this, the resolution of the image will decrease.
This is a problem as it is. It is conceivable to increase the application level of the readout gradient magnetic field to compensate for the decrease in the application time and use symmetrical sampling. However, this increases the bandwidth of the received signal and is disadvantageous to the SN ratio. is not. Therefore, the NMR signal is complex-conjugated at a diagonal position in the received data space, and the property that the data has symmetry is used to estimate the missing portion of sampling by calculation and collect data.

As described above, in the pulse sequence according to the present invention, Te can be shortened at the central portion of the encoding No. that determines the contrast of the image, so that an image with sufficient T1 emphasis can be obtained. it can. In this enhancement effect, it is important to set the point at which the encode gradient magnetic field is switched from the application level control to the application time control, and it is necessary to set it at the optimum point depending on the required image enhancement state and the gradient magnetic field power supply capacity of the device. There is. Further, at the time of image reconstruction, received data having different Tes will be combined, so that it is necessary to consider for data level correction and the like.

Furthermore, in the sequence according to the present invention, T
Since Tr can be shortened as e is shortened, the total measurement time can be shortened. Further, since there is an optimum value for Tr in T1-weighted imaging, Tr may be set in accordance with this optimum value, and the time margin due to shortening Te may be used to increase the number of images to be picked up.

The present invention can be applied not only to the SE method but also to other pulse sequences. For example, FIG. 10 is a timing diagram of the gradient echo (GE) method. By applying an encode gradient magnetic field and a read-out gradient magnetic field immediately after the irradiation pulse and reversing the polarity of the read-out gradient magnetic field, , The diffused phase is converged to obtain an echo signal. Normally, the control of the amount of the encoding gradient magnetic field is based on the application level control, so that the application time t'is constant.

FIG. 9 shows a timing chart when the present invention is applied to the GE method. In the GE method as well, in the same way as the SE method, the control of the amount of the encoding gradient magnetic field is switched from the application level control to the application time control at the central portion, so that the reduction of Te can be realized. In this case as well, asymmetrical sampling is performed (see FIG. 2), so data estimation processing is necessary. GE
According to the method, shortening Te is not only an effect on image enhancement but also an important technique in high-performance sequences such as angio imaging, so the present invention is effective.

[0032]

As described above, according to the present invention, the control of the encode gradient magnetic field amount is switched from the application level control to the application time control at the central portion of all the encodes, thereby shortening this time and shortening the minimum Te. Achieved shortening without strengthening the gradient magnetic field power source, enabling clinically effective T1-weighted images to be captured in a shorter time than before, and further expanding the Te setting range in high-performance sequences to obtain good MRI images. It has the effect of being obtained.

[Brief description of drawings]

FIG. 1 is an explanatory view of applying a gradient magnetic field and Te, Tr according to the present invention.

FIG. 2 is an explanatory diagram of an asymmetric sampling method.

FIG. 3 is an explanatory diagram of signal estimation by partial signal sampling.

FIG. 4 is an explanatory diagram of a spin echo method sequence according to the present invention.

FIG. 5 is an overall configuration diagram of an MRI apparatus.

FIG. 6 is an explanatory diagram of a spin echo method sequence.

FIG. 7 is an explanatory diagram of image contrast according to relaxation time.

FIG. 8 shows changes in the reconstructed image depending on the measurement data range.

FIG. 9 is an explanatory diagram of a gradient echo method sequence according to the present invention.

FIG. 10 is an explanatory diagram of a gradient echo method sequence.

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Claims (2)

[Claims] 1. A magnetic circuit for applying a static magnetic field to a subject, a gradient magnetic field coil for applying a slice gradient magnetic field, a read-out gradient magnetic field and an encode gradient magnetic field to the subject, and an atom constituting the tissue of the subject. An irradiation coil that repeatedly applies an irradiation pulse that causes magnetic resonance to the atomic nuclei in a predetermined pulse sequence, a receiving coil that detects a magnetic resonance signal, and an image that represents the physical properties of the target object using the detection signal In a magnetic resonance imaging apparatus including an image reconstructing means, the application amount control of an encoding gradient magnetic field of a pulse sequence in a series of image data measurement is performed by maintaining an application time at a predetermined value and changing the application level. While combining the two methods, one is to maintain the level at a predetermined value and the other is to change the application time. When the application time is changed, a means for changing the application time of the readout gradient magnetic field is provided accordingly, the echo signal acquisition time is changed, the signal is received, and the received signals are combined to reconstruct the image. Magnetic resonance imaging device. 2. The magnetic recording medium according to claim 1, wherein in the case of changing the read-out gradient magnetic field application time, the asymmetry of the received signal resulting from the shortening of the echo signal sampling time is corrected by the estimating means. Resonance imaging device.