JPH03193034A - Magnetic resonance imaging device - Google Patents

Abstract

PURPOSE:To shorten the photographing time, and also, to miniaturize the device by constituting the device so that two axial directions of an inclined magnetic field coil being orthogonal to three axial directions generate a higher magnetic field than that of the other axial direction. CONSTITUTION:Magnetic field strength in Y0 and Z0 axial directions is that which is larger by about three folds or more than magnetic field strength in the Z0 axial direction. For instance, in the case Y0 and Z0 faces being orthogonal to the X0 axis are set as a slice face, an inclined magnetic field coil for the Z0 axis is formed so as to allow a large current to flow as a read-out inclined magnetic field GR, and to the inclined magnetic field coil for the X0 axis and the inclined magnetic field coil for the Y0 axis, a slice inclined magnetic field GS and an encoding inclined magnetic field GE are applied, respectively. Also, not only three orthogonal cross sections but also a desired slice face having an arbitrary inclined face of single-of-leak, and also, double-of-leak, etc., can be set. Since the inclined magnetic field which can form a high magnetic field is limited two axes, an inclination coil and an inclined magnetic field power source can be miniaturized.

Description

DETAILED DESCRIPTION OF THE INVENTION [Object of the Invention] (Industrial Application Field) The present invention relates to a magnetic resonance imaging apparatus for obtaining magnetic resonance images of a subject (usually a patient).

(Prior art) A subject is placed in a static magnetic field, and a high-frequency magnetic field is applied in a direction perpendicular to the static magnetic field to generate a magnetic resonance phenomenon, and an X-ray is applied so as to be superimposed on the static magnetic field. shaft.

A magnetic resonance imaging apparatus is known in which gradient magnetic fields are applied in three directions, the Yo axis and the Zo axis, and the MR multiplied signal probe obtained from the subject is detected and imaged.

Data collection with the above device is performed with one data collection time TR of 20 to 100 m5, which is 128 times or 2 times.
All necessary data was collected by repeating 56 times. An example of this pulse sequence is shown in FIG.

The sequence shown in the figure is a one-time data collection sequence, and is actually repeated the number of times described above. here,
After selective excitation by irradiating a 90° RF pulse without applying the slicing gradient magnetic field G, the encoding gradient magnetic field G is subsequently applied to perform position encoding while reversing the readout gradient magnetic field GR. The MR multiplied signal is received after a time T after the RF pulse is irradiated.

FIG. 8 shows details of an example of the waveform of the readout gradient magnetic field GR. In the conventional case, for example, the rise time t1 is 1 ms, the current a1 is 150A, and the readout gradient magnetic field G
If the inductance L of the coil for R generation is 500μH, the output voltage V required for the gradient magnetic field power supply is 7 from the formula I V-L□ ... (+, ).
A gradient magnetic field power source is required that can output 5V and a peak output of 1.2kW or more in each of the three axial directions of the Xo, Yo, and Zo axes. In the sequence shown in FIG. 7, it ultimately takes several tens of seconds to several minutes for one photographing time.

Furthermore, in order to obtain a high-quality MR image, if images are taken several times and processing such as averaging is performed, it takes about 1 to 10 minutes. During this imaging, the patient must remain stationary, which is painful for the patient. Furthermore, the respiratory heartbeat movement, which cannot be stopped, obstructs the data and causes a virtual image, that is, an artifact.

In order to solve the above-mentioned drawbacks, an ultra-high-speed imaging method called the echo precursor method invented by Mansfield F. et al. has become well known.

This imaging method uses the transverse relaxation time T, ! of hydrogen atoms, as shown in the sequence of Figure 2. The readout gradient magnetic field G I is reversed by the number of repetitions described above in a short time to the extent that it is negligible, data is collected at the time of reversal, and an image is reconstructed.

The problem here is the power supply that drives the gradient coil 22 for the readout gradient magnetic field GR.

Since one MR multiplication signal must be collected in a time of 1-/3 to ]/4 compared to normal times, the spins of the hydrogen nuclei, which are the signal sources, must be aligned in phase within that short time. For this reason, a gradient magnetic field strength of 3 to 4 times the normal strength is required, and therefore a current of 3 to 4 times as much as the gradient magnetic field power source is required.
L1 is required. Further, the reversal time T of the read gradient magnetic field G is several hundred to 1-000 μs, and the necessary rise time tlo is about 100 μs. a 1
l-1-(3 to 4) a+ , tlo-1/], , Ot
When these are substituted into equation (1) as l, the required voltage v' is V'-3000 to 4000 V, and therefore the peak power is about 100 times that of the conventional device.

A gradient magnetic field power supply capable of outputting such a high voltage becomes extremely large and expensive.

(Problem to be solved by the invention) The conventional device has a problem in that it takes a long time to take an image.

There is also a method of increasing the gradient magnetic field strength in each of the three axial directions to increase the speed, but there is a problem that the device would be too large and expensive.

The present invention has been made in view of the above circumstances, and it is an object of the present invention to provide a magnetic resonance imaging apparatus that shortens the imaging time and is miniaturized.

[Structure of the Invention] (Means for Solving the Problems) The apparatus of the present invention places a subject in a static magnetic field generated by a static magnetic field generating means, and supplies current to each gradient magnetic field coil from a gradient magnetic field power supply. A magnetic resonance image of a desired slice plane of the subject is obtained by superimposing controlled magnetic fields that are orthogonal to each other in three axial directions on the static magnetic field as a slicing gradient magnetic field, an encoding gradient magnetic field, and a readout gradient magnetic field. In the magnetic resonance imaging apparatus, each of the gradient magnetic field coils is configured to generate a higher magnetic field in two axial directions than in the other axial directions. In addition, if double oblique is required, in order to switch the direction of the high magnetic field, a coil rotation mechanism that can rotate each of the gradient magnetic field coils by a predetermined angle around one of the three axes, and a coil rotation mechanism that can rotate the high magnetic field The apparatus further includes a control means for controlling the coil rotation mechanism and the gradient magnetic field power source so that the read gradient magnetic field can be formed by the generated biaxial gradient magnetic fields.

(for production) The operation of the apparatus having the above configuration will be explained below.

The control means of the present apparatus controls the gradient magnetic field power source and the coil rotation mechanism so that a read gradient magnetic field can be formed by a biaxial gradient magnetic field capable of generating a high magnetic field.

(Example) Examples of the present invention will be described in detail below.

FIG. 1 is a schematic block diagram of an apparatus 1 according to an embodiment of the present invention.

This device 1 includes a static magnetic field coil 21. Excitation coil 24, detection coil 23. A magnet arrangement 2 with a gradient coil 22 and an excitation coil 24 . Detection coil 23° Gradient coil 22 with axis Z
. A coil rotation mechanism 3 capable of rotating ±900 around A receiver 6 for receiving data, a sequencer 8 for controlling the gradient magnetic field power supply 4° transmitter 5 under the control of a computer system 7, a display such as a CRT display (not shown), a keyboard (not shown), etc. It has a console 9.

The excitation coil 24. Detection coil 23. Gradient coil 22
is fixed to the coil support tube 20.

Note that the Xo, Yo, and Z axes represent a stationary coordinate system, where the Xo axis is horizontal, the Yo axis is vertical, and the ZO axis is the magnet device 2.
shall indicate the axial direction of

The gradient coil 22 includes gradient coils (not shown) for the xo, yo, z, and axial directions.

The coil rotation mechanism 3 includes an optical detector 31 that detects the rotational position of the coil support tube 20, and a roller 32a that guides the rotation of the coil support tube 20.

32b, and an ultrasonic motor 33 that transmits rotational power to the shaft of the roller 32a. The ultrasonic motor 33 rotates the coil support tube 20 by a predetermined angle based on a rotation command signal sent from the computer system 7.
After rotation, the angle can be maintained until the next rotation command signal is sent.

The gradient magnetic field power source 4 supplies current to each of the gradient coils in the Xo, Yo, Z, and axial directions. Axial gradient magnetic field power supply 41. Y,] Axial gradient magnetic field power supply 42. Z, uniaxial gradient magnetic field power supply 43 is provided.

Of these, Y. Axial gradient magnetic field power supply 42. ZO axis gradient magnetic field power supply 4
3 is a large-capacity power supply, and the gradient magnetic field in the Y 11 + 211 axis direction is a high magnetic field.

FIG. 2 shows the pulse sequence of this device 1.

Under the control of the sequencer 8, the transmitter 5 sends a high frequency wave to the excitation coil 24, and the excitation coil 24 sends 9 waves to the subject.
Send a 0° RF pulse. At the same time, a slicing gradient magnetic field G is applied from the gradient coil 22. continue,
The encoding gradient magnetic field G is sequentially increased in steps of 128 or 2
A variable magnetic field of 56 steps is applied. This encoding gradient magnetic field G.

Immediately after applying the readout gradient magnetic field GR, a pulsed magnetic field GR is repeatedly applied for about 1 ms, and the MR multiplied signal from the subject is detected by the detection coil 23. Therefore, even if the readout gradient magnetic field G8 is repeated 256 times and then repeated three times for averaging processing, the imaging is completed within one second.

Figure 3 shows the details of the gradient magnetic field GR waveform. For example, the rise time tlO is 0.1m+, the current (all) is about 50OA, and the voltage () is the inductance L of 500μH.
Then, from the above equation (1), it becomes 2.5 kV. In the end, the maximum output of the gradient magnetic field power supply is 1 to 50 kW or more. Axial gradient magnetic field power supply 42. Zo-axis gradient magnetic field power supply 43
The power consumption is about 100 times that of conventional technology. Other X. The axial gradient magnetic field power source 41 has a power of 12 kW or more, which is about the same as the conventional one.

0 In other words, the area S1 of the gradient magnetic field GR waveform shown in FIG. 8 of the conventional device is the area S1 of the gradient magnetic field GR waveform shown in FIG.
By making it equal to IO, an MR multiplied signal can be detected.

For example, for the area S1 shown in FIG. 8, al is 150A,
If t2 is 4ms, it becomes 600m5A, and the area SIO shown in Figure 3 is a+o500A.

If t20 is 1.2 ms, it becomes 600 mtA, which can be made into the same area.

The operation of the present device 1 will be explained with reference to FIG. 4 as well.

First, X. , Yo, and Zo axes will be described.

When a desired slice plane is set by operating the keyboard on the console 9, the sequencer 8 causes the gradient magnetic field power supply 4 to supply a controlled current to the gradient coil 22 based on this slice plane setting information (under the control of the computer system 7). .

FIG. 4(a) shows the gradient magnetic field power source 4. It shows a vector of magnetic field strength that can be generated by the gradient coil 22.

The magnetic field strength in the Yo, Z, and axial directions is approximately three times or more larger than the magnetic field strength in the X, axial direction.

When the 11 planes of Y Q Z perpendicular to the X t) axis are set as slice planes, the Z
A large current is made to flow through the gradient magnetic field coil around the O axis as a readout gradient magnetic field GR, and a gradient magnetic field GS for slicing is applied to the gradient magnetic field coil for the XC1 axis.

An encoding gradient magnetic field GL is applied to the Yo-axis gradient magnetic field coil. In this case, the uses of the Yo axis and the 71° axis may be interchanged.

In addition, when the X, Y planes perpendicular to the Zo axis are set as slice planes, as shown in Figure 4(c), a large current is applied to the gradient magnetic field coil of the Y axis piece as the reading gradient magnetic field G8. A gradient magnetic field G for slicing is applied to the gradient magnetic field coil for the Z3 axis.

An encoding gradient magnetic field GL is applied to the gradient magnetic field coil of the shaft piece.

Next, X perpendicular to the Yo axis. When plane 71 is set as the slice plane, as shown in Fig. 4(d), a large current is passed through the gradient magnetic field coil for the Zo axis as the reading gradient 1 magnetic field GR, and the gradient magnetic field of Yl and the axis piece is The coil is provided with a slicing gradient magnetic field Gs.

An encoding gradient magnetic field GE is applied to the Xo-axis gradient magnetic field coil.

If a plane inclined to a certain axis (also called a single oblique) is set in addition to the above-mentioned orthogonal slice plane, as shown in FIGS. 4(e) and 4(f), the X shown in FIG. 4(C). If plane 71 is rotated by θ around the Y Ll axis as slice planes X and Y, then if the static coordinate system is rotated by θ around the Y t + axis, the rotated coordinate system after the θ rotation is X + =XoCosθ−Zn3! n θ −(2)
Yl-Yo...(3)Z
, -X. sinθ+ZoCO8θ (4) Here, the gradient magnetic field G for slicing in FIG. 2(B).

When applying the slicing gradient magnetic field Gs in the direction of the 71-axis direction to determine the slice plane as shown in FIG. 4(e), as shown in FIG. 4(f), Let the gradient magnetic field in the axial direction be the encoding gradient magnetic field G, and let the gradient magnetic field in the Y1 axis direction be the reading gradient magnetic field G□. As shown in equation (2), Y
□Axis direction, ie Y. Since the gradient magnetic field in the axial direction is compatible with high magnetic fields, high-speed reading can be performed.

Similarly, as shown in Fig. 4 (g) and (h,), when the X[, 71 plane shown in Fig. 4 (d) is set as Xl, the slice planes X and Z are rotated by θ around the axis. , the stationary coordinate system is X.

When θ is rotated around the axis, the rotational coordinate system after θ rotation is Xl-Xo...(5)Y
l=Zos+nθ+YocO8θ−(6)
Zl--Z. cos θ-Y, sin θ (7) Here, when applying the slicing gradient magnetic field G5 in FIG.
is applied to determine the slice plane as shown in FIG. 4(g), and then the gradient magnetic field in the X1-axis direction is set as the encoding gradient magnetic field G1, and the gradient magnetic field in the Y-axis direction is Let the gradient magnetic field be the readout gradient magnetic field GR. Zl is the Zo axis direction and Y, as shown in equation (7). Since the gradient magnetic field in the axial direction corresponds to a magnetic field as high as 3-4, high-speed reading can be performed.

In this way, if there is a single oblique, high-speed reading is possible by appropriately selecting the gradient magnetic field GL for encoding and the gradient magnetic field GI for reading. There are cases where it is desired to photograph a surface (this is called a double oblique), but in this case the situation is different from the case of the single oblique. For example, the fourth
X shown in Figure (C).

Rotate the 78 plane by θ about the axis Yl, and then rotate it to X. Rotate by φ about the axis, and the coordinates after rotation are (x2゜Y2.Z2
), then X2-Xocos θ-Z. sin θ・(8
)Y2=Xosin φsin θ−YIJS1n φ
10Z, sinφcosθ ・(9)Z2−X
ocos φsin θ−Yos+n φten Zocos
It is expressed as φcoSθ (10).

Here, the slicing direction is the axis Z2 direction, and the axis X2 and the axis Y
2 corresponds to either the reading gradient magnetic field GR or the encoding gradient magnetic field G. However, neither X is compatible with high magnetic fields. Since the shaft is included,
In neither case can the readout gradient magnetic field GR be selected.

Therefore, in order to solve this problem, the gradient coil 22 (coil support tube 20) is rotated by the coil rotation mechanism 3.

That is, the Xo axis and the Y axis. The axis is rotated about the zo axis. Here, the gradient coil 22 is at Z. η around the axis
The coordinate system after rotation is (xo', Yo,
zo/ ), then YO=-XOtin η+YOCO8η ・(11)
Xo=Xo cos η+Yosin 1 −(1
2) zo = zo −(13
), and by further transforming these equations (11) to (13), Xo=Xo'cosrl-Yo' sin rl
-(14) Yo=Xo' sin 77+Y, C
O81-(15) Zo=Zo'
−(16). Here, in order to solve the above problem, if the axis X2 direction shown in equation (8) is set as the reading gradient magnetic field GR direction, then x2:Xo'CO5θ・cos η Yo CoS θsin η− Zo sinθ...
(17) By setting η-±90°, X.

can be erased, Yo' and Z. ′, the axis X2 direction can correspond to the reading gradient magnetic field G8, which is a high magnetic field, and high-speed reading can be performed.

This also applies to other double obliques. For example, if the direction of the readout gradient magnetic field GR is as shown in equation (8), AXo + BYo + CZl) ・ (1
8), then the gradient coil 22 is Z.

Rotate by η around the axis, (A coB+B si+B)Xo'+(-A
sinη+B coSη)Yo'1C2o'...(
19) Here, if r)=-tan-”A/B, then A
cos η+B+in r)=O, the term Xo' can be set to zero, and this direction can be made compatible with a high magnetic field.

Since one embodiment of the present invention is configured as described above,
It has the following effects.

Zo axis direction and Y. Since the two axial axes are configured to generate a high magnetic field, it is possible to set a desired slice plane not only with three orthogonal cross sections but also with any inclined plane such as a single oblique or even a double oblique. can.

Since the gradient magnetic field capable of forming a high magnetic field is limited to two axes, the gradient coil and the gradient magnetic field power source can be downsized.

Although one embodiment has been described above, the present invention is not limited to this, and various modifications can be made without changing the gist thereof.

For example, the coil rotation mechanism may be configured as shown in FIGS. 5 and 6.

A gear 44 is installed on the outer periphery of the gradient coil 22, and a drive motor 43 and gears 42a, 42b are installed at the bottom of the magnet device 2.
is installed. These eight drive motors 43 and gears 42a, 42b can be finely adjusted up and down by a vertical fine adjustment device (not shown), thereby making it possible to accurately maintain their positional relationship with respect to the coil support tube 20. Furthermore, if the drive motor is a non-magnetic ultrasonic motor,
The drive source can be directly connected to the gear 42a without removing it from the magnet device 2. The rotation angle is detected by a rotation angle detection sensor 41 installed on the side or bottom of the magnet device 2.

This structure is, for example, as shown in FIG. 6, in which a Citro 1 on which a light-reflecting object is pasted in the form of a blind is attached to the circumferential surface of the gradient coil 22. The rotation angle detection sensor 41 is LE
It consists of a light source 45 such as D and a photodetector 46, and when the gradient coil 22 rotates, the sheet 61. moves, the light emitted from the light source 45 is reflected by the six sheets, and the photodetector 46 generates pulses in response to the rotation angle. By counting the number of pulses in the computer system 7, it becomes clear how many times the rotation has occurred.

Alternatively, the rotation angle may be detected by attaching an encoder directly to the drive motor 43.

[Effects of the Invention] According to the present invention described in detail above, the imaging time is shortened because two axial directions of the gradient magnetic field coils perpendicular to the three axial directions are configured to generate a higher magnetic field than the other axial directions. However, it is possible to provide a magnetic resonance imaging device that is more compact.

[Brief explanation of drawings]

FIG. 1 is a schematic block diagram of an apparatus according to an embodiment of the present invention, and FIG.
The figure shows the pulse sequence of this device, and Figure 3 shows the gradient magnetic field GR.
4(a) to 4(h) are explanatory diagrams showing the operation of this device. FIGS. 5 and 6 are explanatory diagrams showing other examples of the coil rotation mechanism. The figure shows the pulse sequence of the conventional device, and Figure 8 shows the gradient magnetic field G + of the conventional device.
It is an explanatory diagram showing details of a + waveform. 1... Magnetic resonance imaging device, 3... Coil rotation mechanism, 4... Gradient magnetic field power supply, 7
... Computer system (control means), 9 0 22 ... Gradient coil, GR ... Gradient magnetic field for reading, G, ... Gradient magnetic field for encoding, G, ... Gradient magnetic field for slicing.

Claims (4)

[Claims] (1) The subject is placed in a static magnetic field generated by a static magnetic field generating means, and the gradient magnetic fields perpendicular to each other in three axes are generated by controlling the current supplied to each gradient magnetic field coil from a gradient magnetic field power source. In a magnetic resonance imaging apparatus that obtains a magnetic resonance image of a desired slice plane of a subject by superimposing a slicing gradient magnetic field, an encoding gradient magnetic field, and a readout gradient magnetic field on the static magnetic field, each of the gradient magnetic field coils is placed in the coil. A magnetic resonance imaging apparatus characterized in that it is configured to generate a higher magnetic field in two axial directions than in other axial directions. (2) a coil rotation mechanism capable of rotating each of the gradient magnetic field coils by a predetermined angle around any one of the three axes;
The magnetic resonance imaging apparatus according to claim 1, further comprising a control means for controlling the coil rotation mechanism and the gradient magnetic field power source so that the readout gradient magnetic field can be formed by the biaxial gradient magnetic field that generates the high magnetic field. . (3) The magnetic resonance imaging apparatus according to claim 2, wherein the drive source of the coil rotation mechanism is an ultrasonic motor. (4) The magnetic resonance imaging apparatus according to claim 2 or 3, wherein the rotation angle of the gradient magnetic field coil by the coil rotation mechanism is ±45° or more.