JP3144840B2 - Magnetic resonance imaging equipment - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001] The present invention relates to a magnetic resonance imaging (MRI) apparatus for performing blood flow imaging and the like in a living tissue using the magnetic resonance phenomenon.

[0002]

2. Description of the Related Art In recent years, diffusion (diffusion), which is the movement of water due to molecular motion (Brownian motion) in living tissue, and perfusion (perfusion), which is the flow of blood in capillaries, utilizing the magnetic resonance phenomenon. Techniques for quantitatively or qualitatively imaging fusion have been actively studied.

As a pulse sequence for measuring diffusion, a Stejscal-Tanner pulse sequence is known. In order to emphasize the phase shift of the spin due to the minute movement of the water molecule due to the diffusion, a 180 ° pulse is applied to any axis, for example, the read direction (frequency encoding direction) as shown in FIG. A gradient field (Motion Probing Gradient) having the same application time d and amplitude G at a position (± I / 2) symmetrical with respect to
A sequence in which an MPG is applied (hereinafter, referred to as an MPG) and a sequence in which an MPG is not applied as shown in FIG. Calculate the parameters related to
As a parameter representing the degree of phase shift (dephase) due to motion, a gradient factor: B (where,
B is γ ² G ² d ² (Approximately equal to (I + 2d / 3)). Here, γ is the gyromagnetic ratio.

On the other hand, the measurement of perfusion is known as tissue blood flow imaging. As a means of tissue blood flow imaging, a tracer method using a tracer and a method that captures minute movements of water molecules (water spin) (Moti
on Detection method, here abbreviated as MD method). The tracer method is classified variously according to the modality of imaging, and the method using MRI is Gd-DT.
Susceptibility contrast agent such as PA and dynamic MR
There is a method using I. SPECT, P
Some use modalities such as ET and xenon CT, but all have the drawback that it is necessary to irradiate the subject with radiation. The MD method is classified into two types depending on whether it is detected by an amplitude change or a phase change.
Whether to detect by amplitude change or phase change is determined by IVIM (Intra Voxel Incoherent Motio
n) Use a model or IVCM (Intra Voxel Cohere
nt Motion) model.

[0005] These models will be described with reference to FIG. FIG. 13 is a macro concept viewed in the order of voxels (one side: 1 to 2 mm) used in MRI ordinary imaging. The IVCM model is, as shown in FIG. 1A, the movement of protons in one direction on average in a voxel, and corresponds to blood flow in a large blood vessel. I
In the VCM model, it moves in a certain direction in the gradient magnetic field,
A corresponding average phase shift (φ ≠ 0) occurs within the echo time TE, and the average amplitude does not change in full IVCM (ie, when all spins move in one direction at a constant speed), but usually Has a velocity distribution,
Usually the amplitude also decreases. However, at t = 0, the phase φ = 0 and the amplitude A = 1.

On the other hand, the IVIM model is a model of movement in a random direction in a voxel as shown in FIG. However, no average phase shift occurs, and the amplitude is attenuated according to the magnitude of the motion. The IVIM model is a model of diffusion of water molecules by Brownian motion, and the capillary model is also a model close to IVIM.

[0007] At present, at least at the level of cerebral blood flow, which is clinically important, a method of quantifying blood flow is considered as an IVIM model by Le. Bihan et al. ("Separation of Diffusion and Perfusion in intr
avoxel incoherent motion MR Imaging ", Radiology, v
ol 499, No. 168, pp. 401-407) could not be measured correctly because the blood volume was as small as several percent in the brain, close to the noise level, and due to the effects of motion artifacts. In addition, the IVIM model has a contribution of diffusion, and it is not easy to separate and detect the diffusion and the perfusion.

[0008] On the other hand, a method of assuming an IVCM model by IR Young et al. ("Assessment of Brain Perfusion Wi
th MR Imaging ", Journal of Computer Asisted Tomogr
aphy, 12 (5), pp. 721-727, September / October), if perfusion is coherent, the flow velocity on the order of tissue blood flow is on the order of 0.5 to 2 mm / sec. Can be captured.

The effects of motion artifacts are affected to the same extent because both use a pulse sequence to which a large gradient magnetic field MPG is applied in order to emphasize the phase shift due to minute movements. Does not contribute to the phase, and only perfusion can be detected with high sensitivity.

[0010] The ratio of perfusion is not limited to either the IVCM model or the IVIM model, and the ratio varies depending on the type of tissue, but it is considered that both components are mixed.

As described above, the conventional IVIM model has low sensitivity to blood flow, and it is difficult to separate perfusion from diffusion. Although the IVCM model has relatively high sensitivity to blood flow, sufficient sensitivity has not yet been obtained.

Further, as a conventional method for measuring perfusion, a phase contrast of MRA is used.
There is also a drawback in that this method is insensitive to slow blood flow.

[0013]

SUMMARY OF THE INVENTION An object of the present invention is to provide a magnetic resonance imaging apparatus capable of measuring perfusion with high sensitivity using an IVCM model. . Another object of the present invention is to provide a magnetic resonance imaging apparatus in which the contribution of eddy current is small and the accuracy is improved.

It is another object of the present invention to provide a magnetic resonance imaging apparatus capable of obtaining parameters of an IVIM model in addition to parameters of an IVCM model.

[0015]

According to the present invention, there is provided a magnetic resonance imaging apparatus for executing a pulse sequence of a spin echo method.
Axis, Y-axis, and the first pulse sequence symmetrically applied before and after the timing of the at least one axis to the gradient of 180 ° pulses of constant intensity in the Z-axis, wherein the at least one
Means for executing a second pulse sequence for applying the gradient magnetic field in the opposite polarity to the axis of the first axis, and a first parameter reflecting a phase difference from the magnetic resonance signals collected in the first and second pulse sequences. means for detecting a said first
Coherent motion in voxels based on parameters
Means for imaging the pulse sequence,
Means for executing the gradient magnetic field in the frequency encoding direction.
When adding to the lead axis, use the lead gradient magnetic
So that the polarity of the field is the same as that of the gradient magnetic field of the constant intensity
It is characterized in that the first and second pulse sequences are inverted .

In another magnetic resonance imaging apparatus according to the present invention, a pulse sequence of the spin echo method is implemented.
X-axis, Y-axis
A gradient magnetic field of constant intensity is applied to at least one of the
Apply symmetrically before and after the timing of the 180 ° pulse
A first pulse sequence and the at least one axis
A second pulse sequence for applying the gradient magnetic field in reverse polarity
And a third pulse sequence not applying the gradient magnetic field.
Executing the first and second pulse sequences.
The first method that reflects the phase difference from the magnetic resonance signals collected in the
Means for detecting the parameters of the first or second
Acquisition with pulse sequence and the third pulse sequence
Second parameter reflecting the amplitude difference from the extracted magnetic resonance signal
Means for detecting data, based on the first parameter.
Imaging of coherent motion in voxels
Means, and within the voxel based on the second parameter.
Means for imaging incoherent movement.
It is characterized by

Further, in another magnetic resonance imaging apparatus according to the present invention , the pulse sequence of the spin echo method is used.
In a magnetic resonance imaging apparatus for performing
At least one of the axes, Y-axis, and Z-axis, with a constant inclination
Magnetic field is marked symmetrically before and after the 180 ° pulse timing.
A first pulse sequence to be applied and said at least one
A second pulse system for applying the gradient magnetic field to the axis
And means for executing the first and second pulses.
Reflect phase difference from magnetic resonance signals acquired in sequence
Means for detecting the first parameter obtained, and the first parameter
Coherent movement in voxels based on parameters
Means for imaging, said pulse sequence
The execution means includes a gradient magnetic field for imaging of at least one.
Note that it includes a gradient magnetic field for the next flow rephase.
Sign. Furthermore, another magnetic resonance image according to the invention
In a scanning device, the pulse echo of the field echo method is used.
In a magnetic resonance imaging apparatus that performs cans,
X-axis, Y-axis, Z-axis at least one of the inclination of a certain strength
A first pulse that applies a gradient magnetic field symmetrically in the positive and negative directions
Pulse sequence and the same strength as the first pulse sequence.
The gradient magnetic field, the polarity of which is inverted in degrees,
Means for executing a second pulse sequence applied to
The magnetic flux collected in the first and second pulse sequences
Means for detecting a parameter reflecting a phase difference from a sound signal
And vice versa, and based on the parameters,
It is characterized by imaging healing movement.
You.

[0018]

According to the magnetic resonance imaging apparatus of the present invention, the phase change can be doubled, so that the perfusion based on the IVCM model can be measured with high sensitivity.

Furthermore, according to the present invention, the influence of the eddy current is compensated by inverting the polarity of the imaging read gradient magnetic field by the first and second pulse sequences to make the polarity of the gradient magnetic field in phase with the constant intensity gradient magnetic field. And accuracy can be improved.

Further, according to the present invention, the normal pulse sequence is added to the first and second pulse sequences,
By executing one sequence, the parameters of the IVIM model can be obtained in addition to the parameters of the IVCM model.

[0021]

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS A first embodiment of the magnetic resonance imaging apparatus according to the present invention will be described below with reference to the drawings. First, FIG.
1 is a block diagram showing a schematic configuration of one embodiment. In the gantry 20, a static magnetic field magnet 1, an X-axis, a Y-axis, a Z-axis gradient magnetic field coil 2, and a transmission / reception coil 3 are provided. The static magnetic field magnet 1 as a static magnetic field generator is configured using, for example, a superconducting coil or a normal conducting coil. X axis, Y
The axis / Z-axis gradient magnetic field coil 2 is a coil for generating an X-axis gradient magnetic field Gx, a Y-axis gradient magnetic field Gy, and a Z-axis gradient magnetic field Gz. The transmitting / receiving coil 3 generates a high-frequency pulse,
And it is used for detecting a magnetic resonance (MR) signal generated by the magnetic resonance. The subject P on the couch 13 is inserted into an imageable area (a spherical area where an imaging magnetic field is formed, and a diagnosis can be made only within this area) in the gantry 20 due to the movement of the couch top of the couch 13. Is done.

The static magnetic field magnet 1 is driven by a static magnetic field control device 4. The transmitting and receiving coil 3 is driven by the transmitter 5 when exciting magnetic resonance, and is coupled to the receiver 6 when detecting an MR signal. The X-axis, Y-axis, and Z-axis gradient magnetic field coils 2 are driven by an X-axis gradient magnetic field power supply 7, a Y-axis gradient magnetic field power supply 8, and a Z-axis gradient magnetic field power supply 9.

The X-axis gradient magnetic field power supply 7, the Y-axis gradient magnetic field power supply 8, the Z-axis gradient magnetic field power supply 9, and the transmitter 5 are driven by a sequencer 10 according to a predetermined sequence, and the X-axis gradient magnetic field Gx, the Y-axis gradient magnetic field Gy, A Z-axis gradient magnetic field Gz and a high frequency (RF) pulse are generated in a predetermined pulse sequence. In this case, the X-axis gradient magnetic field Gx, the Y-axis gradient magnetic field Gy,
The Z-axis gradient magnetic field Gz is mainly used as, for example, a readout gradient magnetic field Gr, a phase encoding gradient magnetic field Ge, and a slice gradient magnetic field Gs. The computer system 11 drives and controls the sequencer 10, and also takes in an echo signal as an MR signal received by the receiver 6 and performs predetermined signal processing to generate a tomographic image of the subject. indicate. Next, the principle of the present invention will be described. The signal value in one voxel i of the MRI image is represented by a complex number expression as follows. Si = Ai.exp (-j.phi.i) where Ai is the amplitude and .phi.i is the phase.

Although the phase changes due to various factors, for example,
In the spin echo method, the shift is mainly caused by the movement in the gradient magnetic field direction. In the field echo method, the contribution due to the non-uniformity of the magnetic field, the susceptibility, and the like are added in addition to the contribution due to the movement.

It is an object of the present invention to selectively detect only a phase shift due to a blood flow (perfusion) in a living body with high sensitivity and selectively with high accuracy. That is, the present invention measures perfusion by detecting a phase change in a portion considered to contain many IVCM models. In the conventional example, in both the IVIM model and the IVCM model,
Two pulse sequences with / without MPG are used in combination, and the perfusion is determined based on the difference between the amplitudes of the echo signals of both sequences in the IVIM model, and the phase difference in the IVCM model.

On the other hand, in the present invention, as shown in FIGS. 1A and 1B, in order to obtain a fine phase shift due to perfusion movement with higher sensitivity, the same intensity and the same applied voltage are used. A time gradient magnetic field MPG is applied together, but scanning is performed using two sequences having opposite polarities as a pair, and a phase difference is directly or indirectly obtained from two echo signals.

FIG. 2 shows an amplitude change and a phase shift between two pulse sequences due to coherent motion when the pulse sequences shown in FIGS. 1A and 1B are executed.
These are shown in (a) and (b), respectively. As described above, in the present invention, since two sequences having the opposite polarities of the MPG are used, the phase difference Δφ for the same motion is doubled as compared with the conventional sequence. Thus, the phase difference can be detected with high sensitivity using the IVCM model.

The procedure for creating an actual IVCM image will be described with reference to FIG. Although the phase difference Δφ may be calculated directly from the two sequences, a parameter that reflects the phase difference may be easily obtained, such as an MR angiogram of the PC method.

First, the respective sequences are executed, and the vector Pa = Aa.exp (-j.phi.a), Pb = Ab.
exp (−jφb) is obtained. Next, vectors Pa and P
b, the parameter Δφ reflecting the phase difference, | Pa−Pb
| FIG. 5 shows the relationship between these parameters. Δφ = arg (Pa / Pb) = φb−φa | Pa−Pb | = ｛(Ra−Rb) ² + (Ia-Ib) ² ｝ ^1/2 where Pa = Ra + jIa and Pb = Rb + jIb. A velocity image is created from these parameters.

As described above, only the absolute value of the phase difference or the vector difference may be used as the perfusion image by the IVCM, but the speed v, which is a more intuitive parameter, may be obtained. An example of a method for obtaining the speed from the phase difference will be described. The phase difference Δφ and the velocity v have a substantially linear relationship as follows. v = f (Δφ) = α · Δφ + β where 0 ≦ Δφ ≦ 2π. Here, α and β are constants determined by the shape of the gradient MPG.

Since the velocity v is a periodic function having a period of 2π as in the above equation, the dynamic range of the target velocity v needs to be suppressed to 2π or less. Since the dynamic range of the velocity v is determined by the shape of the gradient magnetic field MPG and other gradient magnetic fields, α and β are determined in advance for the gradient magnetic field MPG of the pulse sequence to be used by calculation or phantom experiment, and are tabulated. In other words, the velocity v can be easily obtained from the phase difference Δφ by conversion.

As described above, according to the first embodiment, the combination of the dephase / rephase sequence
By the polarity of the opposite polarity to each other in both the sensitivity of the phase difference is doubled for the same flow rate. In addition, IV
Since the measurement is performed as a CM model, there is no contribution of diffusion, and only perfusion can be detected.

Although the eddy current is not considered in the first embodiment, the phase shift may be caused by the eddy current as one of the problems on the device side. It can be compensated by using it. Further, when a phase difference is obtained by an operation between two images in a pulse sequence, the difference can be canceled by controlling the method of applying a gradient magnetic field so that both have the same contribution.
FIG. 6 shows a pulse sequence as a modified example of the first embodiment that can cancel the eddy current. FIG. 6 differs from FIG. 1 in the polarity of the imaging lead gradient magnetic field. In FIG. 1, the polarity of the imaging read gradient magnetic field is 2
6 have the same polarity, but in FIG.
Like the polarity of the PG, the polarity of the imaging read gradient magnetic field is reversed in two sequences. As a result, when viewed from the imaging lead gradient magnetic field, the residual gradient magnetic field due to the eddy current overlaps with the same polarity and in the same order, so that the contribution of the eddy current is offset in the phase difference between the two sequences. . For this reason, the eddy current can be compensated to some extent without using a shield type gradient coil. In this case, since the images are reversed in the read direction, one of the operations is performed in reverse during the calculation between the images. Also, as shown in the PC method, the imaging gradient magnetic field may be a magnetic field waveform for at least the first-order flow rephase as shown in FIG. Hereinafter, another embodiment will be described.

In the second embodiment, as shown in FIG.
In addition to the two sequences (FIGS. (A) and (b)) in which MPGs of the opposite polarities are applied as shown in (a), a normal sequence (FIG. (C)) in which no MPG is applied is used as a set. , IVCM parameter Δφ, | Pa−Pb |
In addition, an IVIM parameter ADC defined as follows can be obtained. ADC = {1 / (Ba, b-Bc)} ln [| Pc | / {(| Pa | + | Pb |) / 2}] where Ba, b represents Ba or Bb, and ln is a natural logarithm. Is shown. FIG. 9 shows the relationship between the IVCM parameter and the IVIM parameter.

In the IVIM model, the amplitude decay is shown in FIG.
The sequences (a) and (b) are equivalent and | Pa |
= | Pb |. Also, (| Pa | + | Pb |) / 2
Is used, the S / N ratio is doubled as compared with the case where either | Pa | or | Pb | is used. For the gradient factor, Ba is almost 0,
Bb is set substantially equal to Bc.

The above embodiment relates to a normal (normal) imaging method. An embodiment relating to an ultra-high-speed imaging (echo planar imaging) method will be described below.
FIG. 10 shows a pulse sequence of echo planar imaging by the spin echo method as the third embodiment. MP
The polarity of G is changed as shown by a solid line and a broken line in the first and second sequences.

FIG. 11 shows a pulse sequence of echo planar imaging by the field echo method as the fourth embodiment. In the field echo method, since the 180 ° pulse is not used, the gradient magnetic field MPG is applied in a reversed manner in one sequence. The positive and negative polarity components of the gradient magnetic field MPG have the same application time and the same intensity as in the case of the spin echo method. The application timing of the gradient magnetic field MPG is not limited, and may be any time within the period from the application of the RF pulse to the start of data collection.

Although not shown, a parameter reflecting a phase difference can be measured from two sequences to which MPG is applied in the same manner as in FIG. 11 in the pulse sequence of the ordinary field echo method. .

The present invention is not limited to the embodiments described above, but can be implemented with various modifications. For example, although MPG is applied only to the axis in the lead direction, MPG need only be applied to the axis along the flow direction, and is not necessarily limited to the lead axis. Further, the voltage may be applied not only in one axis direction but also in two axes and three axes. Further, the imaging method is not limited to the method described above, and another method may be used. Furthermore, the specific example of imaging the parameter reflecting the phase difference is not limited to the above-described example, and may be appropriately changed.

[0040]

As described above, according to the present invention, I
A magnetic resonance imaging apparatus capable of measuring perfusion with high sensitivity using a VCM model is provided. Further, according to the present invention, there is provided a magnetic resonance imaging apparatus in which the contribution of eddy current is small and the accuracy is improved. Further, according to the present invention, in addition to the parameters of the IVCM model,
Provided is a magnetic resonance imaging apparatus capable of determining parameters of a VIM model.

[Brief description of the drawings]

FIG. 1 shows a first embodiment of a magnetic resonance imaging apparatus according to the present invention.
FIG. 4 is a diagram showing a pulse sequence according to the embodiment.

FIG. 2 is a diagram schematically showing the movement of a spin in the first embodiment.

FIG. 3 is a block diagram showing the configuration of the first embodiment.

FIG. 4 is a flowchart showing the operation of the first embodiment.

FIG. 5 is a view for explaining a phase shift in two sequences according to the first embodiment.

FIG. 6 is a diagram showing a pulse sequence according to a modification of the first embodiment.

FIG. 7 is a diagram showing a pulse sequence of another modification of the first embodiment.

FIG. 8 is a diagram showing a pulse sequence according to a second embodiment of the present invention.

FIG. 9 is a diagram illustrating phase shifts in three sequences according to the second embodiment.

FIG. 10 is a diagram showing a pulse sequence according to a third embodiment of the present invention.

FIG. 11 is a diagram showing a pulse sequence according to a fourth embodiment of the present invention.

FIG. 12 is a view for explaining an IVCM model and an IVIM model for blood flow imaging.

FIG. 13 is a view showing a pulse sequence for conventional perfusion imaging.

[Explanation of symbols]

DESCRIPTION OF SYMBOLS 1 ... Static magnetic field magnet, 2 ... X-axis, Y-axis, Z-axis gradient magnetic field coil, 3 ... Transmission / reception coil, 4 ... Static magnetic field control device, 5 ... Receiver, 6 ... Transmitter, 7 ... X-axis gradient magnetic field power supply, 8 ... Y axis gradient magnetic field power supply, 9 ... Z axis gradient magnetic field power supply, 10 ... Sequencer,
11 computer system, 12 display unit.

──────────────────────────────────────────────────続 き Continuation of the front page (58) Field surveyed (Int.Cl. ⁷ , DB name) A61B 5/055 JICST file (JOIS)

Claims (4)

(57) [Claims] 1. A magnetic resonance imaging apparatus for executing a pulse sequence of a spin echo method, wherein a gradient magnetic field of a constant intensity is symmetrically applied to at least one of an X-axis, a Y-axis, and a Z-axis before and after a 180 ° pulse timing. A first pulse sequence applied to the at least one
Means for executing a second pulse sequence for applying the gradient magnetic field to the axis in the opposite polarity; and a first parameter reflecting a phase difference from the magnetic resonance signals collected in the first and second pulse sequences. Means for detecting , and coherence in the voxel based on the first parameter.
Means for imaging an unexpected movement, wherein the pulse sequence executing means converts the gradient magnetic field to a frequency.
When adding to the lead axis in several encoding directions,
The polarity of the lead gradient magnetic field for
The first and second pulse sequences so that
A magnetic resonance imaging apparatus characterized in that the magnetic resonance imaging apparatus performs reversal in the magnetic field. 2. A pulse sequence of the spin echo method is implemented.
In a magnetic resonance imaging apparatus to be performed, a tilt having a constant intensity is applied to at least one of the X axis, the Y axis, and the Z axis.
The gradient magnetic field is symmetrical before and after the 180 ° pulse timing.
A first pulse sequence to be applied;
Second pulse for applying the gradient magnetic field to two axes with opposite polarities
A sequence and a third pulse not applying the gradient magnetic field
Means for executing a sequence, and magnetic recordings collected in the first and second pulse sequences.
Detecting a first parameter reflecting a phase difference from a sound signal;
Means, and the first or second pulse sequence and the third pulse sequence.
The amplitude difference from the magnetic resonance signals collected in the pulse sequence.
Means for detecting a reflected second parameter, and a coherent in the voxel based on the first parameter.
Means for imaging unintended movements, and an incognito in voxel based on the second parameter.
Means for imaging the movement
A magnetic resonance imaging apparatus comprising: 3. A pulse sequence of the spin echo method is implemented.
In a magnetic resonance imaging apparatus to be performed, a tilt having a constant intensity is applied to at least one of the X axis, the Y axis, and the Z axis.
The gradient magnetic field is symmetrical before and after the 180 ° pulse timing.
A first pulse sequence to be applied;
Second pulse for applying the gradient magnetic field to two axes with opposite polarities
Means for executing the first and second pulse sequences.
Detecting a first parameter reflecting a phase difference from a sound signal;
Means for cohering in the voxel based on said first parameter.
Means for imaging an unintended movement, wherein the pulse sequence executing means comprises an imaging tilting means.
The gradient magnetic field is tilted for at least the first-order flow rephase.
Magnetic resonance imaging apparatus characterized by including a gradient magnetic field
Place. 4. A magnetic resonance imaging apparatus for executing a pulse sequence of a field echo method, wherein a gradient magnetic field having a constant intensity is symmetrically applied to at least one of an X-axis, a Y-axis, and a Z-axis in a positive direction and a negative direction. Means for executing a first pulse sequence to be applied, and a second pulse sequence to apply a gradient magnetic field having the same intensity as that of the first pulse sequence and having an inverted polarity to the at least one axis; Means for detecting a parameter reflecting a phase difference from a magnetic resonance signal collected by the pulse sequence of the above, and a coherent motion in the voxel is imaged based on the parameter.