JPH0531099A - Magnetic resonance imaging device - Google Patents

Abstract

PURPOSE:To provide a magnetic resonance imaging device capable of measuring a perfusion at high sensitivity with an IVCM model. CONSTITUTION:The first pulse sequence symmetrically applying the gradient magnetic field with constant intensity at the front and rear of the 180 deg.-pulse on at least one axis and the second pulse sequence applying the gradient magnetic field with the same intensity and the opposite polarity to the first pulse sequence at the front and rear of the 180 deg.-pulse on at least one axis are executed in the pulse sequence of the spin echo method, parameters reflected with the phase difference are detected from the magnetic resonance signals collected in the first and second sequences, thus the blood flow coherently moved in a voxel is imaged.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a magnetic resonance imaging (MRI) apparatus for performing blood flow imaging in a living tissue by utilizing the magnetic resonance phenomenon.

[0002]

2. Description of the Related Art In recent years, diffusion (diffusion), which is a motion of water in a living tissue due to molecular motion (Brownian motion), and perfusion (perfusion), which is a flow of blood in a capillary, are performed by utilizing a magnetic resonance phenomenon. A method for quantitatively or qualitatively imaging fusion) has been actively studied.

A Stejscal-Tanner pulse sequence is known as a pulse sequence for measuring diffusion. In order to emphasize the phase shift of the spin due to the fine movement of water molecules due to diffusion, a 180 ° pulse as shown in FIG. 12 (b) is applied to either axis, for example, the read direction (frequency encode direction) by the spin echo method. Gradient magnetic field (Motion Probing Gradient) with the same applied time d and amplitude G at symmetrical positions (± I / 2) across
Hereinafter, a sequence in which an MPG is applied (hatched in the figure) is applied and a sequence in which the MPG is not applied as shown in FIG. 7A are executed, and the velocity is calculated from the difference between the amplitude and phase of the echo signals obtained by the two. Calculate the parameters for
A gradient factor: B (here, as a parameter indicating the degree of phase shift (dephase) due to motion
B is γ ² G ² d ² (Approximately equal to (I + 2d / 3)) is defined. Here, γ is the gyromagnetic ratio.

On the other hand, the measurement of perfusion is known as tissue blood flow imaging. As a method of tissue blood flow imaging, a tracer method that uses a tracer and a method that captures minute movements of water molecules (water spins) (Moti
on Detection method, abbreviated as MD method here). The tracer method is classified into various types according to the modality of imaging, and the one using MRI is Gd-DT.
Sustainability contrast agents such as PA and dynamic MR
There is a method using I. Tracer method includes SPECT, P
Some use modalities such as ET and xenon CT, but these all have a drawback in 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 it is detected by amplitude change or phase change is determined by IVIM (Intra Voxel Incoherent Motio) as a model of proton movement.
n) model or IVCM (Intra Voxel Cohere
nt Motion) model is used.

These models will be described with reference to FIG. FIG. 13 is a macroscopic concept viewed from the voxel order (1 side: 1 to 2 mm) used in normal imaging of MRI. The IVCM model is a movement of protons in one direction on average in a voxel, as shown in (a) of the figure, and corresponds to blood flow in a large blood vessel. I
In the VCM model, since 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 is also unchanged for a perfect IVCM (ie when all spins move in one direction at constant velocity), but usually Has a velocity distribution,
Amplitude also usually 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 random directions within a voxel, as shown in FIG. 2B, which has no directional movement on average, and a gradient magnetic field is applied. If so, the average phase shift will not occur and the amplitude will decay depending on the magnitude of the motion. The IVIM model is a model of water molecule diffusion due to Brownian motion, and the capillary model is also a model close to IVIM.

In the current report, at least at the level of clinically important cerebral blood flow, a method of quantifying the blood flow is considered as an IVIM model of 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 a few percent in the brain, which was close to the noise level, and due to the influence of motion artifacts. Further, the IVIM model has a contribution of diffusion, and it was not easy to detect the diffusion and the perfusion separately.

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

The influence of motion artifacts is affected to the same degree 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 a small movement, but in the IVCM model the diffusion is affected. Does not contribute to the phase, and only perfusion can be detected with high sensitivity.

It should be noted that perfusion is not limited to either one of the IVCM model and the IVIM model, and although the ratio varies depending on the type of tissue, 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 and diffusion. The IVCM model is relatively sensitive to blood flow, but has not yet been sufficiently sensitive.

Furthermore, as a conventional method for measuring perfusion, the phase contrast (Phase Cont
rast) method, but this method also has the drawback of being insensitive to slow blood flow.

[0013]

SUMMARY OF THE INVENTION The present invention has been made to address the above-mentioned circumstances, and an object thereof is to provide a magnetic resonance imaging apparatus capable of highly sensitively measuring perfusion using an IVCM model. . Another object of the present invention is to provide a magnetic resonance imaging apparatus in which contribution of eddy current is small and accuracy is improved.

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

[0015]

In a magnetic resonance imaging apparatus according to the present invention, a gradient magnetic field having a constant intensity is applied to at least one of X-axis, Y-axis and Z-axis, and a 180 ° pulse is used in the case of a pulse sequence of spin echo method. Before and after the timing of
Means for performing a first pulse sequence that is applied symmetrically and a second pulse sequence that applies the gradient magnetic field to the at least one axis with opposite polarities; and the first and second pulse sequences. And a means for detecting a parameter reflecting a phase difference from a magnetic resonance signal, and imaging a blood flow which moves coherently in a voxel.

In another magnetic resonance imaging apparatus according to the present invention, in the above-mentioned configuration, the pulse sequence executing means, when applying the gradient magnetic field of the constant intensity to the read axis in the frequency encoding direction, the imaging read gradient magnetic field. Is reversed in the first and second pulse sequences so as to be in phase with the gradient magnetic field having the constant intensity.

Further, in another magnetic resonance imaging apparatus according to the present invention, in the above-mentioned configuration, the pulse sequence execution means further executes a normal third pulse sequence in which the gradient magnetic field of the constant intensity is not applied, The parameter detecting means further detects a parameter reflecting an amplitude difference from the magnetic resonance signals collected in the first to third pulse sequences, and also images incoherent movement in a voxel.

[0018]

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

Further, according to the present invention, the effect of the eddy current is compensated by reversing the polarity of the imaging lead gradient magnetic field by the first and second pulse sequences so as to be in phase with the gradient magnetic field of constant intensity. Therefore, the accuracy can be improved.

Furthermore, according to the present invention, a normal pulse sequence is added to the first and second pulse sequences, and
By executing this 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 a magnetic resonance imaging apparatus according to the present invention will be described below with reference to the drawings. First, FIG.
FIG. 1 is a block diagram showing a schematic configuration of one embodiment. Inside 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 the static magnetic field generator is configured by using, for example, a superconducting coil or a normal conducting coil. X axis, Y
The axis and Z axis gradient magnetic field coils 2 are coils 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 transmitter / receiver coil 3 generates a high frequency pulse,
And used to detect magnetic resonance (MR) signals generated by magnetic resonance. The subject P on the bed 13 is inserted into an imageable region (a spherical region where a magnetic field for imaging is formed and a diagnosis can be made only in this region) in the gantry 20 by the movement of the top plate of the bed 13. To be done.

The static magnetic field magnet 1 is driven by the static magnetic field controller 4. The transmitter / receiver coil 3 is driven by the transmitter 5 when magnetic resonance is excited, and is coupled to the receiver 6 when MR signals are detected. The X-axis, Y-axis and Z-axis gradient magnetic field coils 2 are driven by an X-axis gradient magnetic field power source 7, a Y-axis gradient magnetic field power source 8 and a Z-axis gradient magnetic field power source 9.

The X-axis gradient magnetic field power source 7, the Y-axis gradient magnetic field power source 8, the Z-axis gradient magnetic field power source 9 and the transmitter 5 are driven by a sequencer 10 in a predetermined sequence, and an X-axis gradient magnetic field Gx, a Y-axis gradient magnetic field Gy, A Z-axis gradient magnetic field Gz and a radio 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, for example, as the read gradient magnetic field Gr, the phase encoding gradient magnetic field Ge, and the slice gradient magnetic field Gs. The computer system 11 drives and controls the sequencer 10, captures 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, and the display unit 12 displays the tomographic image. indicate. Next, the principle of the present invention will be described. The signal value in one voxel i of the MRI image is expressed as a complex number 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,
The spin echo method shifts mainly by movement in the direction of the gradient magnetic field, and the field echo method further contributes by nonuniformity of the magnetic field, susceptibility, etc. in addition to the contribution by movement.

An object of the present invention is to detect only a phase shift due to blood flow (perfusion) in the living body with high sensitivity separately from diffusion and with high precision selectively. That is, the present invention measures perfusion by detecting a phase change in a portion that is considered to include many IVCM models. In the conventional example, in both the IVIM model and the IVCM model,
Two pulse sequences with and without MPG are used in combination, and perfusion is obtained by the difference in amplitude of echo signals of both sequences in the IVIM model and the difference in phase in the IVCM model.

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

FIG. 2 shows changes in amplitude and phase shift between two pulse sequences due to coherent movement when the pulse sequences of FIGS. 1 (a) and 1 (b) are executed.
They are shown in (a) and (b), respectively. As described above, in the present invention, since two sequences in which the polarities of MPGs are opposite are used, the phase difference Δφ for the same motion is doubled compared to the conventional sequence. Thereby, the phase difference can be detected with high sensitivity by the IVCM model.

The actual procedure for creating an IVCM image will be described with reference to FIG. The phase difference Δφ may be directly calculated from the two sequences, but a parameter reflecting the phase difference may be simply obtained like the MR angiogram of the PC method.

First, each sequence is executed, and vectors Pa = Aa.exp (-j.phi.a) and Pb = Ab.multidot.
Calculate exp (-jφb). Next, the vectors Pa and P
Parameter Δφ reflecting phase difference from b, | Pa −Pb
Ask for |. The relationship between these parameters is shown in FIG. 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 more intuitive parameter speed v 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 However, 0 ≦ Δφ ≦ 2π. Here, α and β are constants determined by the shape of the inclined MPG.

Since the velocity v is a periodic function having a period of 2π as in the above equation, it is necessary to keep the dynamic range of the target velocity v within 2π. 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 calculated in advance for the gradient magnetic field MPG of the pulse sequence to be used by calculation or phantom experiment, and tabulated. In other words, the velocity v can be easily calculated from the phase difference Δφ.

As described above, according to the first embodiment, the combination of the dephase / rephase sequences makes the polarities of the MPGs opposite to each other, so that the sensitivity of the phase difference to the same flow velocity is 2. Double. In addition, IVC
Since it is measured as an M model, there is no contribution of diffusion and only perfusion can be detected.

Although the eddy current was not taken into consideration in the first embodiment, the phase shift may be caused by the eddy current which is one of the problems on the device side. It can be compensated by using it. Further, when the phase difference is obtained by the calculation between the two images in the pulse sequence, it can be offset by controlling the application method of the gradient magnetic field so that both have the same contribution.
FIG. 6 shows a pulse sequence as a modified example of the first embodiment capable of canceling this eddy current. 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
Although the two sequences have the same polarity, in FIG.
Similar to the polarity of PG, the polarity of the imaging read gradient magnetic field is opposite in the 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 in the same polarity and in the same order, so that the contribution of the eddy current is canceled in the phase difference between the two sequences. . Therefore, the eddy current can be compensated to some extent without using the shield type gradient magnetic field coil. In this case, since the images are reversed in the read direction, one of them is reversed during the inter-image calculation. Further, as seen in the PC method, the gradient magnetic field for imaging may have a magnetic field waveform for at least the first-order flow rephase as shown in FIG. Another embodiment will be described below.

In the second embodiment, as shown in FIG.
In addition to the two sequences shown in (1) and (2) in which the opposite polarity MPG is applied, a normal sequence in which no MPG is applied at all ((c) in the figure) is used as a set. , IVCM parameter Δφ, | Pa −Pb |
In addition, the IVIM parameter ADC defined as follows can be obtained. ADC = {1 / (Ba, b-Bc)} ln [| Pc | / {(| Pa | + | Pb |) / 2}] where Ba and b represent Ba or Bb, and ln is a natural logarithm. Indicates. 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 of (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 of using one of | Pa | or | Pb |. Regarding the gradient factor, Ba is almost 0,
Bb is set almost equal to Bc.

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

FIG. 11 shows a pulse sequence for 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 by being inverted in the middle of one sequence. The positive and negative polar 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. Further, 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, it is also possible to measure the parameter reflecting the phase difference from the two sequences to which the MPG is applied, as in the case of the ordinary field echo method pulse sequence. .

The present invention is not limited to the above-mentioned embodiments, but can be implemented with various modifications. For example, MPG is applied only to the axis in the lead direction, but MPG may be applied to the axis along the flow direction and is not necessarily limited to the lead axis. Furthermore, the voltage may be applied not only in the uniaxial direction but also biaxially and triaxially. Further, the imaging method is not limited to the above method, and other methods may be used. Further, the specific example of imaging the parameter reflecting the phase difference is not limited to the above example, and may be appropriately changed.

[0040]

As described above, according to the present invention, I
A magnetic resonance imaging apparatus capable of highly sensitively measuring perfusion using a VCM model is provided. Further, according to the present invention, there is provided a magnetic resonance imaging apparatus in which contribution of eddy current is small and accuracy is improved. Also, according to the present invention, in addition to the parameters of the IVCM model, I
A magnetic resonance imaging apparatus is provided that can also determine the parameters of a VIM model.

[Brief description of drawings]

FIG. 1 is a first magnetic resonance imaging apparatus according to the present invention.
The figure which shows the pulse sequence of an Example.

FIG. 2 is a diagram schematically showing the movement of spins 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 diagram for explaining phase shift in the two sequences of the first embodiment.

FIG. 6 is a diagram showing a pulse sequence of a modified example 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 of a second embodiment of the present invention.

FIG. 9 is a diagram for explaining phase shifts in the three sequences of the second embodiment.

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

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

FIG. 12 is a diagram illustrating an IVCM model and an IVIM model for blood flow imaging.

FIG. 13 is a diagram 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 ... Transmitting / receiving 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.

─────────────────────────────────────────────────── ─── Continuation of the front page (51) Int.Cl. ⁵ Identification code Internal reference number FI technical display location 9118-2J G01N 24/08 Y

Claims (5)

[Claims] 1. A magnetic resonance imaging apparatus for executing a pulse sequence of a spin echo method, wherein a gradient magnetic field having a constant intensity is symmetrically provided before and after a 180 ° pulse timing in at least one of X-axis, Y-axis and Z-axis. A first pulse sequence applied to
Means for executing a second pulse sequence for applying the gradient magnetic field to the axis in opposite polarity, and means for detecting a parameter reflecting a phase difference from the magnetic resonance signals collected by the first and second pulse sequences. And a magnetic resonance imaging apparatus for imaging blood flow that moves coherently in a voxel based on the parameter. 2. The pulse sequence executing means, when applying the gradient magnetic field to the read axis in the frequency encoding direction, sets the polarity of the read gradient magnetic field for imaging to be the same as the polarity of the constant intensity gradient magnetic field. The magnetic resonance imaging apparatus according to claim 1, wherein the first and second pulse sequences are inverted. 3. The pulse sequence execution means further executes a normal third pulse sequence in which the gradient magnetic field is not applied, and the parameter detection means is magnetic resonance collected in the first to third pulse sequences. The magnetic resonance imaging apparatus according to claim 1, wherein a parameter reflecting an amplitude difference is further detected from the signal, and incoherent motion in the voxel is also imaged. 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 X axis, Y axis and Z axis in a positive direction and a negative direction. And a means for executing a second pulse sequence for applying to the at least one axis a gradient magnetic field having the same strength as that of the first pulse sequence but having the same intensity as that of the first pulse sequence, and the first and second pulse sequences. And a means for detecting a parameter reflecting a phase difference from the magnetic resonance signals collected in 1., and a magnetic resonance imaging apparatus for imaging a blood flow having coherent movement in a voxel based on the parameter. 5. The pulse sequence executing means according to claim 1, wherein the gradient magnetic field for imaging includes a gradient magnetic field for at least a first-order flow rephase.
The magnetic resonance imaging apparatus according to.