JPH088916B2 - Magnetic resonance imager - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION Object of the Invention (Industrial field of application) The present invention utilizes a magnetic resonance (hereinafter referred to as [NMR]) phenomenon to determine a specific atomic nucleus density distribution of each tissue in a living body. The present invention relates to a diagnostic NMR device that noninvasively measures from the outside to obtain information for medical diagnosis.

(Prior Art) Magnetic Resonance Imaging (MRI) is, as is well known, when a group of intrinsic spins and associated nuclear magnetic efficiencies are placed in a uniform static magnetic field of intensity Ho. , Resonantly absorbs the energy of a high-frequency magnetic field that rotates at an angular velocity determined by ωo = γHo (γ is called a gyromagnetic ratio, which is a constant unique to the type of nucleus) in a plane perpendicular to the direction of the static magnetic field. It is a method that makes it possible to obtain chemical and physical microscopic information of molecules by utilizing this fact. As a method of imaging the spatial distribution of specific nuclei (for example, hydrogen nuclei in water and fat) in a subject using this magnetic resonance imaging method, a projection reconstruction method by a rotor bar (Lauterbur), a Kumar ( Kumar), Welti, Ernst
And the Fourier method by Hutchison, etc., which is a modification of the Fourier method, and the like have been devised.

Further, since these image reconstruction methods take a long time to image, an echo planer method by Mans field and a fast Fourier method have been devised as a method for high-speed imaging.

Figures 14 and 15 show the pulse sequences of the echo planar method and the fast Fourier method, respectively. As described above, in order to perform high-speed imaging, it is necessary to switch the gradient magnetic field at high speed. However, when the conventional gradient coil for image reconstruction is used for high-speed imaging as it is, the inductance L is too large and the conventional method is used. High-speed drive is not possible with the power supply that was used. Therefore, there is a disadvantage that an expensive power source with high output is required.

(Problems to be Solved by the Invention) As described above, the conventional apparatus has a problem that an expensive power source with high output is required to perform high-speed imaging.

The present invention has been made in view of the above circumstances, and an object of the present invention is to provide a magnetic resonance imaging apparatus that does not require an expensive power source to perform high-speed imaging.

[Structure of Invention]

(Means for Solving Problems) In order to solve such a drawback, in the present invention, in addition to the main gradient magnetic field coil, a local gradient coil smaller than the main gradient magnetic field coil is provided. Is used by switching to the main gradient magnetic field coil as needed to obtain an image.

(Operation) With such a configuration, since the coil is small,
Since the inductance L is small and the generated magnetic field intensity per unit current is also large, high-speed switching is possible with the conventional gradient magnetic field power supply for image reconstruction. In addition, with a local gradient coil, only a partial region of the subject can be imaged, but in order to solve this point, a separate type gradient coil that is not integrated with the device main body including the uniform static magnetic field generating coil or the bed or the like. To use. Furthermore, high-resolution local high-speed imaging is possible by combining with a surface coil.

Embodiments The present invention will be described in detail below with reference to the drawings.

FIG. 1 shows an embodiment of a magnetic resonance imaging apparatus according to the present invention. In FIG. 1, the static magnetic field generating coil 22
Is excited by energization from the excitation power supply 23, and thereby generates a uniform static magnetic field in the imaging region of the subject (living body) 21. On the other hand, from the RF pulse generator 25 controlled by the pulse sequencer 3, rectangular, Gaussian or si
An RF pulse modulated in a nc shape or the like is output, amplified to a predetermined level by the RF amplifier 26, and then applied to the coil 28 via the duplexer 27 to induce magnetic resonance in the subject 21. A rotating magnetic field for forming is generated. Transverse magnetization generated by the application of this rotating magnetic field is induced at both ends of the coil 28 as a magnetic resonance signal. In this example, the single coil 28 is shared by the transmitting coil for generating the rotating magnetic field and the receiving coil for receiving the magnetic resonance signal.

The magnetic resonance signal induced in the coil 28 is a duplexer.
The signal is input to the RF amplifier 29 via 27, amplified to a predetermined level, and then detected by a phase sensitive detection circuit 30 such as quadrature phase detection to become a signal in the video band. The output signal of the phase sensitive detection circuit 30 is voltage-amplified by the video amplifier 31, and the high-pass noise component is removed by the low-pass filter 32. The output signal of the low pass filter 32 is A / D
After being converted into a digital signal by the converter 6, it is taken into the electronic computer 1 through the interface 13 and stored as image reconstruction data. In addition, electronic computer 1
Also controls the pulse sequencer 3 via the interface 2.

Determining the slice plane within the subject 21, phase encoding (conversion of position information within the subject 21 to the phase of the magnetic resonance signal)
Is performed by switching the gradient magnetic field and applying it in a pulsed manner. The timing of switching the gradient magnetic field is controlled by the pulse sequencer 3.

On the other hand, the gradient magnetic field intensity and pulse shape are controlled by the gradient magnetic field controller 37. That is, the gradient magnetic field controller 37 controls the power amplifiers 38a to 38c corresponding to the magnetic field gradients in the x, y, and z directions, and the gradient magnetic field generating coil 39 is driven by these power amplifiers 38a to 38c, whereby a predetermined strength is obtained. , A gradient magnetic field that changes with time is generated in the vicinity of the imaging region of the subject 21.

The present invention is characterized in that, in such a configuration, a local gradient magnetic field generating coil (hereinafter referred to as a local gradient coil) 19 is provided in addition to the gradient magnetic field generating coil (main gradient magnetic field coil) 39. 2 and 3 show an example of a local gradient coil used in the magnetic resonance imaging apparatus, which is a rectangular current loop (Rectanguler Cu).
rrent Loop) is an example using a gradient coil.
Since this gradient coil can be composed of two parallel coil mounting plates 36 as shown in FIG. 2, it does not require a circumferential frame unlike the gradient coil formed by the combination of Maxwell pair and Golay coil, and it can It is convenient for pushing in a part and taking an image. That is, the distance between the two coil mounting plates 36 in the configuration of the local gradient coil 19 is fixed to one side of the frame 18 and the other end is left open, and the body 18 is attached to the body. As shown in Fig. 3, the mounting method can be either from the side of the measurement site according to the type of the static magnetic field generating coil or from the direction from the front, and the gradient according to each shape. This can be achieved by preparing a coil. At this time, the shape is determined so that the center of the coil and the center of the imaging region can be easily aligned. For example, two coil mounting plates 36
If the thicknesses of the two are not the same and one of them is thicker instead of the spacer, it can be aligned with the center only by mounting the imaging region. Further, a spacer having a variable thickness can be used for imaging different parts.

Various other modifications are possible. For example, one end may not be opened and both ends may be fixed. Again 2
Instead of supporting the coil mounting plates 36 on each other, the coil mounting plates 36 may be supported by a bed or another structure. Also, the type of gradient coil is not Rectanguler Current Loop,
A combination of Maxwell pair and Golay coil may be used. In this case, the mounting is done along the axis of the coil. Alternatively, a part of the circumferential support frame of the coil may be attached so that it can be opened and closed. Also, the imaging is
When combined with a surface coil, high S / N and high resolution high speed imaging can be achieved. Further, as an imaging method, various combinations are possible, such as selectively exciting a target site with the main gradient magnetic field coil 39 and a selective excitation pulse, and performing only high-speed imaging by performing the reversal of the gradient magnetic field with the small gradient coil 19. is there.

In any case, according to the above configuration, an expensive power source is not required to perform high speed imaging, and high speed imaging can be performed with the power source of the conventional apparatus.

By the way, in order to obtain an image as described above, by using 90 ° selective excitation pulse 24 and 180 ° selective excitation pulse 33 and gradient magnetic fields 15, 16, 17 as shown in FIG.
An image is obtained by acquiring the spin echo, subjecting it to A / D conversion, and then processing it with a computer. In this case, since the spin echo changes abruptly around the peak, it is necessary to find the peak position and data value of the spin echo by interpolation. That is, in order for the computer to obtain a correct reconstructed image from the A / D converted values, the points before and after the maximum value are used to calculate the peak position x _p of the spin echo by least square approximation calculation, etc. as shown in FIG. Need to calculate. Also, as shown in FIG. 6, the peak position x _p and the maximum value position max _p
It is created by interpolation of the position where the data point is shifted by Dx _p from the AD point using Dx _p of the difference of. Specifically, a quadratic function that passes through two points before and after (a total of three points) for one data point is used to calculate the data value of the point that is deviated by Dx _p on the time axis, and this data value is used by the electronic computer. Was performing a reconstruction operation.

However, since the vicinity of the peak of the echo changes rapidly as described above, it is necessary to perform AD conversion or high-order interpolation calculation at short time intervals in order to perform correct peak position and data interpolation.

However, if A / D conversion is performed at a short time interval, a large amount of data will be needed, and a large amount of memory will be required in the computer, so it will be possible to perform correct peak position detection and interpolation from the limit of memory capacity. It was difficult to secure the amount of data to be used. In addition, when the high-order interpolation calculation is performed, there is a problem that the processing time of the electronic computer becomes long and the time until reconstructing and obtaining an image becomes very long.

In order to solve such a problem, in the embodiment of the present invention, before and after the peak value of the magnetic resonance signal is detected, and the sampling frequency of the A / D converter is increased only before and after the peak value.
Write to the memory of the computer. That is, the time interval of data written in the memory of the A / D-converted computer becomes fine only near the peak. Using this data, the computer interpolates the position of the peak of the magnetic resonance signal and data interpolation. Note that the spin echo peak sharply decreases as it shifts from the center of the image. That is, the data interpolation is performed only on the resonance signal obtained by the encoding step near the center.

With such a configuration, since the data interval near the peak of the magnetic resonance signal is short, the correct peak position and the correct data interpolation can be performed by the low-order interpolation operation even if the data near the peak changes rapidly. .

Since the sampling clock of the A / D converter is high only in the vicinity of the peak of the resonance image signal, the memory of the computer does not change from the conventional method, and high-speed data processing is possible because low-order interpolation operation is performed.

That is, the correct peak position and interpolation can be performed at high speed regardless of the memory capacity of the computer.

FIG. 7 shows a block diagram for explaining the above configuration in detail.

As shown in FIG. 8, the electronic computer 1 in the figure predicts the expected peak position of the magnetic resonance signal, and the value at the point P1 at which the signal abruptly changes is passed through the interface 2 to the register 4 in the pulse sequencer 3 Write in. Further, the value at the point P2 at which the change of the magnetic resonance signal shown in FIG. 8 is reduced is written into the register 5 in the pulse sequencer 3 through the interface 2. The electronic computer 1 issues an oscillation command of the oscillator 7 that creates a sampling clock of the AD converter 6 to capture the magnetic resonance signal. Output of oscillator 7 is 1 / n
It is input to the frequency divider 8, and the output of this frequency divider 8 is input to the AND gate 9. By the output of the counter 10 which monitors the position of the data to be AD converted, if the data before P1 is A /
If D-converted, the AND gate 9 opens and the clock divided by 1 / n is input to the OR gate 11 and becomes the sampling clock of the A / D converter 6. At the same time, this clock is input to the counter 10 and the sampling position is constantly monitored. The contents of this counter 10 are
It is compared with the contents of registers 4 and 5. If the content of the counter 10 exceeds P1, the AND gate 12 opens and the oscillator 7
Is input to the A / D converter 6 through the OR gate 11. That is, the A / D converter 6 is supplied with a clock having a frequency n times that of the above clock. When the content of the counter exceeds the content (P2) of the register 5, a clock with a cycle of 1 / n is input to the A / D converter 6 by the same operation as described above. Interface 13 for AD converted data
It is input to the electronic computer 1 through and is stored in the memory in the electronic computer 1. In this way, the memory in the electronic computer 1 stores the data having short time intervals only near the peak of the magnetic resonance signal, the peak position is accurately detected, and the data interpolation can be performed correctly.

Next, an embodiment in which the data interpolation is performed only on the center of the image will be described. A pulse sequence for obtaining a tomographic MR image using the spin warp method proposed by Hutchison et al. Is shown in FIG. The slice plane is determined by the gradient magnetic field G _S , and the phase encoding of the position information in the first direction is performed by the gradient magnetic field G _E.
The frequency encoding of the position information regarding the second direction orthogonal to the direction is performed by the gradient magnetic field G _R. In such a pulse sequencer 3, when the encoding step specifies the center position of the image, that is, when the gradient magnetic field G _E which changes according to the encoding step becomes a value indicating the center of the image, the peak value of the magnetic resonance signal (SIG). Becomes maximum, and the peak value of the magnetic resonance signal sharply decreases as the encoding step specifies a position away from the center position of the image. That is, the pulse sequencer 3 stores two encoding steps corresponding to before and after the center of the image, counts the encoding pulses input from the electronic computer 1, and compares them with the two encoding steps. If there are two stored encoding steps, the sequence controller finely AD-converts only around the peak value of the magnetic resonance signal according to the first example shown. In this way, data interpolation can be performed more efficiently.

Although data interpolation is performed only in the reading direction, data interpolation in the encoding direction may also be performed.

As described above, according to the above configuration, it is possible to reconfigure at high speed because there is no need to perform the high-order data interpolation without being limited by the memory capacity of the electronic computer in the device,
Correct image reconstruction is possible.

On the other hand, in the transmission / reception system of the magnetic resonance imaging apparatus as described above, the diodes 40, 40 'as shown in FIGS. 9 and 10 are used.
And ferrite cores 41, 41 '(magnetic parts),
A mixer (D.B.M.) is used. This mixer has a wide frequency band, is resistant to cross modulation, etc., and is extremely excellent in characteristics.

Under magnetic resonance imaging, this mixer is placed in or around an external magnetic field, such as a static or alternating magnetic field. Therefore, the magnetic field of the external magnetic field influences the magnetic field generated by the ferrite cores 41, 41 'built into the mixer, which causes defective characteristics and malfunction. Particularly, in a magnetic resonance imaging apparatus, it is a very big part together with the problems such as high-sensitivity detection of minute signals and improvement of SN ratio.

For such a problem, it is desirable to apply a magnetic shield to a part containing a magnetic material such as a ferrite core. In this way, the magnetic field of the external magnetic field does not pass inside the component.

FIG. 11 is a configuration diagram showing a specific example of the magnetic shield. The DBM case 51 contains a diode and a ferrite core (magnetic material). There is a magnetic shield 52 so as to surround it, and a substrate (both sides) 53 passing through it. 54 is a solder surface of the substrate, 55 is a component surface, and 56 is an external magnetic field having a certain strength (the strength is proportional to the thickness of the arrow). Now, the inside of the case 51 of the DBM is completely shielded from the external magnetic field 56 by the magnetic shield 52, so that the operation and characteristics of the DBM are not adversely affected. The magnetic field is applied, for example, in the direction of the arrow in the figure, but the effect is the same in any direction. In addition, although the magnetic shield 52 is used after being cut off for wiring, it does not affect the effect.

FIG. 12 is a configuration diagram showing a second embodiment of the present invention. 51 is a DBM, 52 is a first magnetic shield, 52 'is a second magnetic shield 53 is a substrate (both sides). 54 is the solder side of the board and 55 is the component side. 57 is an external magnetic field stronger than that in the first embodiment.

A description will be given with reference to FIG. Although the strength of the external magnetic field 57 is increased as compared with the first embodiment, the double (or more if necessary) shield and the first magnetic shield 5 are correspondingly increased.
2, the second magnetic shield 52 'causes the external magnetic field 5
7 does not reach.

FIG. 13 is a configuration diagram showing a third embodiment of the present invention. 51 is a DBM, 52 is a magnetic shield (thickness and material are different from those of the first and second embodiments), 53 is a substrate, 54 is a solder surface of the substrate, 55
Is the component side. 58 is an external magnetic field.

A description will be given with reference to FIG. The thickness of the magnetic shield 52,
It is generally known that the shielding effect due to a change in material is proportional to the thickness and that the material is large such as permalloy or supermalloy. Therefore, in addition to the first and second embodiments, the influence of the external magnetic field 58 on the DBM 51 can be prevented by changing the thickness of the magnetic shield 52 or selecting the material.

In this way, according to the above configuration, from the external magnetic field
There is no influence on the inside of the DBM, the operation of the transmission / reception system including the DBM is stable, and a magnetic resonance imaging apparatus with good reproducibility and high reliability can be obtained.

〔The invention's effect〕

As described above, according to the present invention, high-speed imaging is not required to perform high-speed imaging, and high-speed imaging can be performed with the power source of the conventional apparatus. In addition, since the coil is movable, high-speed imaging can be performed for the heart and abdomen that require high-speed imaging.
Others such as the conventional method can be used. Furthermore, high-speed imaging with high S / N and high resolution is possible by combining with a surface coil.

[Brief description of drawings]

FIG. 1 is a block diagram showing an embodiment of a magnetic resonance imaging apparatus according to the present invention, FIGS. 2 and 3 are embodiments of a gradient coil unit thereof, and FIG. 4 is a time chart showing a pulse sequence. 5 and 5 are diagrams for obtaining peaks of magnetic resonance signals, FIG. 6 is a diagram for performing data interpolation, FIG. 7 is a block diagram showing another embodiment of the present invention, and FIG. 8 is magnetic resonance. Fig. 9 shows the signal sampling method, Fig. 9 shows the inside of the double balance mixer, and Fig. 10 shows the double balance mixer.
Wiring diagram of the mixer, FIG. 11, FIG. 12, FIG. 13 is a diagram showing another embodiment of the present invention, FIG. 14 is a diagram showing a pulse sequence of the echo planar method, FIG. 15 is a high speed It is a figure which shows the pulse sequence of a Fourier method. 1 …… Computer, 25 …… Pulse generator 2,13 …… Interface, 27 …… Duplexer 3 …… Pulse sequencer, 28 …… Coil 4,5 …… Register, 30 …… Phase sensitive detection circuit 6 …… A / D converter, 37 ... Gradient magnetic field controller 7 ... Oscillator, 39 ... Gradient magnetic field generating coil 8 ... Divider, 40,40 ′ …… Diode 10 …… Counter, 41,41 ′ …… Ferrite core 15,16,17 …… Gradient magnetic field 19 …… Local gradient coil 21 …… Subject 22 …… Static magnetic field generating coil 23 …… Excitation power supply 24,33 …… Selective excitation pulse

Claims (4)

[Claims] 1. A uniform static magnetic field is applied to a subject, and a gradient magnetic field is switched and applied in a pulsed manner.
A magnetic resonance imaging apparatus for detecting and imaging a magnetic resonance signal from the inside of a subject, comprising a local gradient magnetic field coil that is smaller than the main gradient magnetic field coil and is installed in a part of the subject to provide a gradient magnetic field. The magnetic resonance imaging device is characterized by performing high-speed imaging of a specific part. 2. The magnetic resonance imaging apparatus according to claim 1, wherein the local gradient magnetic field coil is not fixed to the main body of the apparatus including the uniform static magnetic field generating coil, the bed, etc., but to an arbitrary portion of the subject. A magnetic resonance imaging device characterized in that the main body of the device is separated so that it can be moved and used. 3. The magnetic resonance imaging apparatus according to claim 1, wherein the local gradient magnetic field coil is a Rectangule.
A magnetic resonance imaging device characterized by using a current loop type gradient magnetic field coil. 4. A magnetic resonance imaging apparatus according to claim 1, wherein a local gradient magnetic field coil and a surface coil are used together for imaging.