JP5413875B2 - Magnetic resonance imaging and magnetic resonance imaging apparatus - Google Patents

Description

  The present invention relates to a magnetic resonance imaging method and a magnetic resonance imaging apparatus.

  Magnetic resonance imaging (hereinafter abbreviated as “MRI” if necessary) is widely used in the medical field as a precision examination technique for human bodies and the like, and many documents on MRI have already been published ( For example, see Patent Document 1).

In recent years, magnetic resonance functional imaging (functional MRI; hereinafter, abbreviated as “fMRI” as necessary) for measuring brain activity using MRI has attracted attention (see, for example, Patent Document 2). In this fMRI, activities in the human brain can be evaluated non-invasively with millimeter-level spatial resolution, and it is said that higher-order functions of the human brain that are difficult to evaluate by animal experiments can be known appropriately.
JP 2002-165774 A JP 2007-190120 A

  The present inventor is working on research and development that can visualize a human brain mechanism using an MRI technique such as fMRI. In the course of this research and development, the spin (magnetic moment) of protons in a specific direction within a minimal cube (hereinafter abbreviated as “voxel”) with a side of about 3 mm, which is the unit of the three-dimensional image analysis region of MRI. We focused on the importance and usefulness of phase gradient.

  For example, by causing the subject to perform a hand opening and closing task called two-hand tapping and one-hand tapping and measuring the above-mentioned phase gradient, the corpus callosum (left and right brains) that could not be observed with the conventional fMRI technique was measured. We succeeded in visualizing the activity of a nerve fiber bundle (a group of nerve cells axons).

  Therefore, the present inventor expects that a method using the above-described phase gradient as the contrast of an image will be widely used as a standard imaging method of MRI in the future. Development of a magnetic resonance imaging apparatus capable of displaying such a phase gradient as image contrast can be expected to develop functional research of the human brain, and thus breakthrough progress in diagnostic technology in brain medicine.

  In the conventional MRI technology, the main focus is on creating the contrast of various images using proton density information, and reports focusing on the importance and usefulness of phase gradients (research papers and patents). Is not found as far as the inventor knows.

  The present invention has been made in view of such circumstances, and magnetic resonance imaging that can derive the phase gradient of the magnetic moment of protons in a predetermined direction in a voxel, and such a phase gradient. An object of the present invention is to provide a magnetic resonance imaging apparatus capable of displaying as image contrast.

In order to solve the above problems, the present invention is a magnetic resonance imaging method using nuclear magnetic resonance of protons,
The following formula (1) formulated with the phase difference between the center of the voxel and the end of the voxel as an independent variable (X) and the signal intensity correlated with the density of the proton as a dependent variable (S) Is used to provide a magnetic resonance imaging method capable of deriving a phase gradient of the magnetic moment of the proton in a predetermined direction in the voxel.

S = S ₀ Sinc (X) (1)
However, S ₀ : Signal intensity when there is no phase shift A technique using such a phase gradient as the contrast of an image is expected to be widely used as a standard MRI imaging method in the future. The development of functional research, and, in turn, breakthrough advances in diagnostic technology in brain medicine.

  That is, in the magnetic resonance imaging method of the present invention, the phase gradient at the end of the voxel with respect to the center of the voxel is the variable term of the independent variable (X).

  Thereby, a phase gradient can be appropriately derived using the above equation (1).

  In the magnetic resonance imaging method of the present invention, a gradient magnetic field pulse is applied in the predetermined direction, the phase at the end of the voxel in the predetermined direction is shifted with respect to the center of the voxel, and the independent variable (X) is It is preferable to emphasize the difference of the dependent variable (S) corresponding to the change of the independent variable (X) by including the shifted phase.

  As a result, a significant signal intensity difference occurs in the detection of the phase gradient, and the phase gradient can be detected appropriately.

Further, in the magnetic resonance imaging method of the present invention, the phase at the end in the predetermined direction of the voxel is linearly shifted by π / 2 with respect to the center of the voxel, and the independent variable (X) is shifted by the above. The dependent variable (S) when including a phase is the first signal strength,
The dependent variable (S) when the phase at the end in the predetermined direction of the voxel is linearly shifted by −π / 2 with respect to the center of the voxel, and the independent variable (X) includes the shifted phase. ) As the second signal strength,
The difference between the first signal strength and the second signal strength and the sum between the first signal strength and the second signal strength are taken, and the difference is divided between the sum and the sum. ,
It is preferable to derive the phase gradient using a relationship between the signal intensity ratio obtained by the division and the phase gradient.

Thereby, noise (for example, machine noise of the magnetic resonance imaging apparatus or noise derived from the subject) included in the signal intensity “S ₀ ” can be removed, and the phase gradient can be calculated appropriately. Moreover, the phase gradient is conveniently arranged in a simple mathematical expression using the signal intensity ratio by the expansion of the Sinc function.

In the magnetic resonance imaging method of the present invention, the gradient magnetic field pulse is applied in the slice selection direction, and the phase at the end of the voxel in the slice selection direction is linearly shifted with respect to the center of the voxel as π / 2. ,
A gradient magnetic field pulse may also be applied in a direction opposite to the slice selection direction, and the phase at the end of the voxel in the slice selection direction may be shifted linearly by −π / 2 with reference to the center of the voxel.

  Thereby, the phase gradient of the magnetic moment of protons in the slice selection direction in the voxel can be derived.

Further, in the magnetic resonance imaging method of the present invention, in the readout direction to which the gradient magnetic field pulse is applied, k-space is acquired by an imaging scan,
An image corresponding to the first signal intensity and an image corresponding to the second signal intensity are obtained by performing a Fourier transform on each half of the k space divided with respect to the center of the k space,
An image corresponding to the gradient of the phase may be acquired using each of the acquired images.

  Thereby, the phase gradient of the magnetic moment of the proton in the lead-out direction in the voxel can be derived. In addition, since an image corresponding to the first signal intensity and an image corresponding to the second signal intensity can be obtained simultaneously in the same imaging scan, noise related to the imaging scan can be appropriately reduced.

Further, the present invention provides a static magnetic field generating means capable of generating a uniform static magnetic field in a space where a subject is placed,
A gradient magnetic field generating means capable of generating a gradient magnetic field in three orthogonal directions in the space;
High-frequency pulse generating means capable of irradiating a high-frequency pulse causing nuclear magnetic resonance to protons in a static magnetic field constituting the subject's living tissue;
Detection means for detecting a signal depending on the magnetic moment state of the proton in which the nuclear magnetic resonance has occurred;
There is provided a magnetic resonance imaging apparatus comprising: an image processing apparatus that derives the phase gradient described above based on the signal and displays a contrast image of the phase gradient for each voxel.

  With the development of a magnetic resonance imaging apparatus that can display such a phase gradient as the contrast of an image, the development of functional research of the human brain, and thus the breakthrough of diagnostic techniques in brain medicine can be expected.

  According to the present invention, there is provided a magnetic resonance imaging method capable of deriving a phase gradient of a magnetic moment of protons in a predetermined direction in a voxel, and a magnetic resonance imaging apparatus capable of displaying such a phase gradient as image contrast. can get.

Hereinafter, first and second embodiments of the present invention will be described with reference to the drawings.
(First embodiment)
FIG. 1 is a block diagram showing a configuration example of a magnetic resonance imaging apparatus (MRI apparatus) according to the first embodiment of the present invention.

  As shown in FIG. 1, an MRI apparatus 1 includes a static magnetic field generating magnet 2 (static magnetic field generating means) that can generate a uniform static magnetic field in a space where a subject X is placed, and an inclination in three orthogonal directions in the space. And gradient magnetic field coils 5 and 5 (gradient magnetic field generating means) capable of generating a magnetic field. The gradient magnetic field coils 5 and 5 are supplied with driving power by the gradient magnetic field generator 11.

  The high frequency signal output from the high frequency generator 8 (high frequency pulse generating means) of the MRI apparatus 1 is modulated into an appropriate signal by the modulator 9, amplified by the amplifier 10, and supplied to the high frequency transmission coil 3. Thereby, the high frequency transmission coil 3 can irradiate the high frequency pulse which can cause nuclear magnetic resonance to the nucleus (proton) in the static magnetic field which comprises the biological tissue of the subject X.

  A magnetic resonance signal such as an echo signal radiated using proton nuclear magnetic resonance is received by the high-frequency receiving coil 4 of the MRI apparatus 1 and amplified by the amplifier 12, and then the phase detector 13 ( Signal detection means) detects an analog signal. The analog signal is converted into measurement data of a digital signal by the AD converter 14 and sent to the computer 6.

  The computer 6 of the MRI apparatus 1 includes a storage device 15 (RAM, ROM, etc.), a display device 16 (liquid crystal monitor, etc.), an operation input unit (not shown), a CPU, etc. (not shown). It functions as an image processing apparatus capable of executing computation and creation and display of a nuclear magnetic resonance (MR) image (hereinafter abbreviated as “MR image” if necessary). That is, the CPU reads out the image processing program stored in the storage device 15, processes the measurement data of the digital signal received from the AD converter 14 according to the processing procedure by the image processing program, and outputs the processing result to the display device 16. Display as an MR image.

  The sequencer 7 of the MRI apparatus 1 is constituted by a microprocessor, for example, and is connected to the computer 6. The sequencer 7 sends commands necessary for performing magnetic resonance imaging in accordance with a pulse sequence (imaging sequence) to be described later, a magnetic field generation system (such as the gradient magnetic field coils 5 and 5 and the gradient magnetic field generator 11) of the MRI apparatus 1 and a high frequency. Send to transmission / reception system (high frequency generator 8, modulator 9, amplifiers 10, 12, high frequency transmission coil 3, high frequency reception coil 4, phase detector 13, AD converter 14, etc.) and control these operations Yes.

  Next, a pulse sequence for capturing a tomographic image of the subject X using the above MRI apparatus 1 will be described with reference to FIG.

  FIG. 2 is a diagram showing an example of a pulse sequence used in the magnetic resonance imaging method of the first embodiment of the present invention. Several methods have been proposed for a pulse sequence for obtaining a nuclear magnetic resonance signal, and FIG. 2 shows a pulse sequence by a gradient echo method (hereinafter abbreviated as “GRE” if necessary). Has been.

  In the GRE method of FIG. 2, pulse configurations other than the gradient magnetic field pulse Gs ′ indicated by the thick solid line in FIG. 2 are known. In the present embodiment, the application of such a gradient magnetic field pulse Gs ′ emphasizes the phase gradient of the proton magnetic moment in the slice selection direction in the voxel.

  First, the known technique in the GRE method in FIG. 2 will be briefly described below, and the contents of the gradient magnetic field pulse Gs ′ will be described later.

  As shown in FIG. 2, simultaneously with the application of the slice gradient magnetic field pulse Gs, by applying an α degree high frequency pulse α ° (RF excitation pulse), protons on the slice plane of the subject X are selectively excited. This releases a free induction decay signal FID during the recovery process from the proton excited state. In this way, the slice plane in the slice selection direction, which is a two-dimensional image of the subject X, can be specified.

  Next, a gradient magnetic field pulse Gp is applied in a specific direction in the slice plane called a phase encoding direction. Further, the gradient magnetic field pulse Gr is applied in a direction orthogonal to the phase encoding direction in the slice plane (also referred to as a readout direction) (also referred to as a frequency encoding direction).

  Then, by applying a gradient magnetic field pulse Gp in the phase encoding direction, protons that have received a strong magnetic field rotate fast, and protons that have received a weak magnetic field rotate slowly, so after the disappearance of the gradient magnetic field pulse Gp, along the phase encoding direction Proton spins (magnetic moments) of different phases occur. Further, by applying the gradient magnetic field pulse Gr in the readout direction, echo signals Echo having different frequencies along the readout direction are obtained.

  With the above-described known operations, in the magnetic resonance imaging method, the slice plane of the subject X is specified and the spatial encoding in the slice plane is performed.

  Next, a method for deriving the phase gradient of the proton magnetic moment in the slice selection direction in the voxel using the gradient magnetic field pulse Gs ′ in FIG. 2 will be described with reference to the drawings.

  The MRI apparatus 1 observes various changes in the magnetic moment state of protons, and the signal intensity of the echo signal Echo described above is usually correlated (proportional) with the proton density.

Then, the phase difference of the magnetic moment of the proton between the center of the voxel and the end in the specific direction of the voxel is defined as an independent variable “X”, and the signal intensity of the echo signal Echo correlated with the density of the proton is defined as a dependent variable “ The following formula (1) formulated as “S” is already known (for example, Cary H. Glover: “3D z-Shim Method for Reduction of Susceptibility Effects in BOLD fMRI, Magnetic Resonance in Medicine 42: 290 -299 (1999)).

S = S ₀ Sinc (X) (1)

However, S ₀ : Signal strength of the echo signal Echo when there is no phase shift

The inventor of the present invention believes that using such equation (1), the phase gradient at the end of the voxel with respect to the center of the voxel can be derived as the difference in the signal intensity “S” of the echo signal Echo. ing.

In FIG. 3, the horizontal axis represents the phase difference of the proton magnetic moment, and the vertical axis represents the signal intensity of the echo signal Echo (however, the vertical axis is normalized by “S ₀ ”). = S ₀ Sinc (X)

Here, it is considered that the phase difference of the magnetic moment corresponds to the phase gradient (θ ₀ + θ). That is, the phase gradient (θ) of the phase gradient (θ ₀ + θ) is considered to be a variable term of the independent variable “X”.

  Such a phase gradient (θ) represents the amount of change contained in the magnetic moment state of protons constituting the biological tissue of the subject X (here, the brain tissue of the subject X). This is the target amount to be acquired in the imaging method.

  For example, it is considered that there is a possibility that the phase gradient (θ) obtained from the MRI apparatus 1 may change between a specific work by the subject X (brain activity) and a resting state. It can be expected that the detection of an accurate phase gradient (θ) will provide new knowledge in the functional study of the brain of the subject X.

Also, “θ ₀ ” of the phase gradient (θ ₀ + θ) is obtained by applying the gradient magnetic field pulse Gs ′ in FIG. 2 to the slice selection direction, and the voxel center and the end of the voxel in the slice selection direction. It is assumed that the phase is a component whose phase shifts linearly. That is, it is assumed that “θ ₀ ” in the phase gradient (θ ₀ + θ) is a phase shift intentionally given in the independent variable “X”.

  In such a situation, the present inventor is extremely advantageous in detecting the above-described phase gradient (θ) due to the nature of the Sinc function of Equation (1), in which the phase is linearly shifted by the gradient magnetic field pulse Gs ′. I realized that there was.

First, it is assumed that “θ ₀ ” due to the gradient magnetic field pulse Gs ′ does not occur (that is, θ ₀ = 0). Then, the following problems occur with respect to the detection of the phase gradient (θ).

  The phase gradient (θ) is the signal intensity “Sa” of the echo signal Echo corresponding to the “a” position in FIG. 3 and the signal intensity “Sb” of the echo signal Echo corresponding to the “b” position in FIG. Is detected as “Sa−Sb”. However, when the phase gradient (θ) takes a small value (the phase gradient (θ) usually takes a small value in many cases), “Sa-Sb” is obtained due to the nature of the Sinc function of the equation (1). Is small (that is, when X≈0, SincX≈0), there is a case where such a weak signal intensity difference of the echo signal Echo between “Sa” and “Sb” cannot be detected.

  Therefore, in the present embodiment, before acquiring the echo signal Echo (for example, at an appropriate time between the slice gradient magnetic field pulse Gs and the gradient magnetic field pulse Gp), the gradient magnetic field pulse Gs ′ is applied, and the voxel is based on the center of the voxel. The above-mentioned problem is appropriately solved by linearly shifting the phase at the end in the slice selection direction.

For example, as shown in FIG. 2 and FIG. 3, when the gradient magnetic field pulse Gs ′ is applied to the voxel and the phase is linearly shifted (ie, “θ ₀ ” ≠ 0), the phase gradient (θ) is The difference “Sc−Sd” between the signal intensity “Sc” of the echo signal Echo corresponding to the “c” position in FIG. 3 and the signal intensity “Sd” of the echo signal Echo corresponding to the “d” position in FIG. ”Is detected. This “Sc-Sd” takes a much larger value than the above-mentioned “Sa-Sb” due to the nature of the Sinc function of Equation (1).

  As a result, a significant signal intensity difference occurs in the echo signal Echo in detecting the phase gradient (θ), and the phase gradient (θ) can be detected appropriately.

  In this way, in the present embodiment, the independent variable “X” in Expression (1) includes the phase that is linearly shifted, thereby representing the contrast of the phase gradient (θ), that is, the phase gradient (θ). The difference in the signal intensity of the echo signal Echo corresponding to the change in the independent variable “X” (difference in the dependent variable “S”) is emphasized.

  However, the above-described contrast enhancement method of the phase gradient (θ) is only an example. For example, the gradient magnetic field pulse Gs ′ may be applied in the slice selection direction (see FIG. 2), and the gradient magnetic field pulse (−Gs ′) may be applied in the direction opposite to the slice selection direction.

Then, as shown in “e” position and “d” position in FIG. 3, the signal intensity “S ₁ ” (S ₁ = S ₀ Sinc (θ ₀ + θ)) of the echo signal Echo by the gradient magnetic field pulse Gs ′, The signal intensity “S ₂ ” (S ₂ = S ₀ Sinc (−θ ₀ + θ)) of the echo signal Echo by the gradient magnetic field pulse (−Gs ′) can be obtained. As a result, the phase gradient (θ) can be enhanced by using the difference in signal strength between the two (S ₁ −S ₂ = S ₀ (Sinc (θ ₀ + θ) −Sinc (−θ ₀ + θ)) as an image. By taking the difference in signal intensity (S ₁ -S ₂ ) in this way, appropriate noise cancellation is performed, and the contrast of the phase gradient (θ) itself is between “Sc” and “Sd”. This is convenient because it can be emphasized almost twice as compared with the case where the difference “Sc−Sd” in the signal intensity of the echo signal Echo is detected.

  The pulse sequence of the gradient magnetic field pulse (-Gs') can be easily understood with reference to FIG.

  Next, an example of a specific method for calculating the phase gradient (θ) described above will be described.

  Here, a calculation example of the phase gradient (θ) when the gradient magnetic field pulse Gs ′ is applied in the slice selection direction and the gradient magnetic field pulse (−Gs ′) is applied in the direction opposite to the slice selection direction will be described.

  First, the gradient magnetic field pulse Gs ′ is applied in the voxel slice selection direction so that the phase at the end of the voxel in the slice selection direction is linearly shifted in the plus direction with reference to the center of the voxel.

In this case, the gradient magnetic field pulse Gs ′ is set so that the phase shift between the center of the voxel and the end of the voxel is “θ ₀ ” = π / 2. Then, since the Sinc function of Formula (1) can be developed successfully, it is convenient (see Formula (3) described later).

In this way, the signal intensity “S ₁ ” (S ₁ = S ₀ Sinc (π / 2 + θ)) of the echo signal Echo by the gradient magnetic field pulse Gs ′ is obtained.

  Further, a gradient magnetic field pulse (-Gs') is also applied in the direction opposite to the voxel slice selection direction so that the phase at the end in the slice selection direction of the voxel is linearly shifted in the minus direction with reference to the center of the voxel.

In this case, the gradient magnetic field pulse (−Gs ′) is set so that the phase shift between the center of the voxel and the end of the voxel is “θ ₀ ” = −π / 2. Then, since the Sinc function of Formula (1) can be developed successfully, it is convenient (see Formula (3) described later).

In this way, the signal intensity “S ₂ ” (S ₂ = S ₀ Sinc (−π / 2 + θ)) of the echo signal Echo by the gradient magnetic field pulse (−Gs ′) is obtained.

Next, as shown in the following equation (2), the difference (S ₁ −S ₂ ) between the signal strength “S ₁ ” of the echo signal Echo and the signal strength “S ₂ ” of the echo signal Echo is taken. Further, the sum (S ₁ + S ₂ ) between the signal intensity “S ₁ ” of the echo signal Echo and the signal intensity “S ₂ ” of the echo signal Echo is taken. Then, division is performed between these differences (S ₁ −S ₂ ) and the sum (S ₁ + S ₂ ) (more precisely, the difference (S ₁ −S ₂ ) is divided by the sum (S ₁ + S ₂ )). ) Thereby, the signal intensity ratio “R” corresponding to the phase gradient (θ) is obtained by the division.

R = (S ₁ −S ₂ ) / (S ₁ + S ₂ ) (2)

In the above-described signal intensity ratio “R”, “S ₀ ” is appropriately canceled so as not to include the signal intensity “S ₀ ” of the echo signal Echo when there is no phase shift, as shown in the following equation (3). Has been. Further, the phase gradient (θ) is conveniently arranged in a simple mathematical expression using the signal intensity ratio “R” as shown in the following expression (3) by developing the Sinc function.

... (3)

Finally, the phase gradient (θ) can be derived from the following equation (4).

As a result, noise included in the signal intensity “S ₀ ” (for example, machine noise of the MRI apparatus 1 or noise derived from the subject X) can be removed, and the phase gradient (θ) can be calculated appropriately. it can.

Phase gradient (θ) = π / 2 × R (4)

  Hereinafter, an example of an experimental result in which a phantom simulating a human brain and measuring the phase gradient (θ) in the slice selection direction by the phantom will be described.

  In this embodiment, the computer 6 executes the following image processing according to the processing procedure by the image processing program.

  As shown in FIG. 2, before the echo signal Echo is read out, a gradient magnetic field pulse Gs ′ is applied in the slice selection direction, and the gradient magnetic field pulse Gs ′ has the phase at the center of both surfaces of the slice. It was set to be shifted by π / 2 and −π / 2 with respect to the reference.

First, the computer 6 acquires an image (hereinafter referred to as “S ₁ image”) corresponding to the signal intensity “S ₁ ” of the echo signal Echo in the above equation (3) by a normal imaging scan. _One image was displayed on the display device 16 as shown in FIG.

Next, before reading out the echo signal Echo, a gradient magnetic field pulse (−Gs ′) is applied in the direction opposite to the slice selection direction, and the computer 6 performs the echo signal Echo in the above equation (3) by a normal imaging scan. An image corresponding to the signal intensity “S ₂ ” (hereinafter referred to as “S ₂ image”) was obtained, and this S ₂ image was displayed on the display device 16 as shown in FIG.

Then, the computer 6 uses the S ₁ image, the S ₂ image, and the above equations (3) and (4) to obtain the phase gradient (θ) in the slice selection direction (hereinafter referred to as “PG in the slice selection direction ( The phase gradient (θ) contrast image for each voxel using the PG signal was displayed on the display device 16 as shown in FIG. 4C.

  Note that the experimental phantom is a skull-containing one, and in FIG. 4, a large change in the slice selection direction of the magnetic field strength due to the influence of the sinuses can be appropriately reproduced. Thereby, it is considered that the display result of FIG. 4 is appropriate.

As described above, the magnetic resonance imaging method of the present embodiment can derive the phase gradient (θ) of the proton magnetic moment in the slice selection direction in the voxel that is a unit of the three-dimensional image analysis region, and as a result. The contrast image for each voxel of the phase gradient (θ) in the slice selection direction can be appropriately displayed on the computer 6. Therefore, the visualization of the contrast image of such a phase gradient (θ) can be expected to lead to the development of functional research of the human brain, and thus breakthrough advances in diagnostic technology in brain medicine.
(Second Embodiment)
The hardware configuration of the MRI apparatus of this embodiment is the same as the hardware configuration of the MRI apparatus 1 described in the first embodiment (FIG. 1). Therefore, in the following description, a detailed description of the configuration of the MRI apparatus of this embodiment is omitted.

  That is, the MRI apparatus of this embodiment is different from the image processing program described in the first embodiment in the storage device 15 of the computer 6 (see FIG. 1). Although distinguished from the MRI apparatus 1, the MRI apparatus 1 can be used as it is on the hardware.

  In addition, since the formula (1), the formula (2), the formula (3), and the formula (4) described in the first embodiment can be used as they are in the present embodiment, the explanation of these formulas is given here. Is also omitted.

  FIG. 5 is a diagram showing an example of a pulse sequence used in the magnetic resonance imaging method of the second embodiment of the present invention. FIG. 6 is a schematic diagram for explaining an image construction method by magnetic resonance imaging according to the second embodiment of the present invention.

  In the present embodiment, the application of the gradient magnetic field pulse Gs ′ (see FIG. 2) described in the first embodiment is stopped, and the readout is more advanced than the conventional gradient magnetic field pulse Gr in the readout direction (shown by a dotted line in FIG. 5). The gradient magnetic field pulse Gr ′ (shown by a thick solid line in FIG. 5) in the readout direction is applied longer so that the resolution in the direction is doubled (that is, the readout time is doubled).

  More specifically, as shown in FIG. 6, two-dimensional frequency information (k-space) is obtained by imaging scanning in the lead-out direction, and as shown in FIGS. 6 and 7, In order to double the resolution, the length of the k-space in the readout direction is doubled compared to the length of the conventional k-space in the readout direction. Since the k space itself is well known in the MRI technique, a detailed description of the k space is omitted.

Next, each half of the k space (“1st half” and “2nd half” in FIGS. 6 and 7) divided based on the center of the k space is Fourier-transformed, so that An image corresponding to the signal intensity “S ₁ ” of the echo signal Echo (hereinafter abbreviated as “S ₁ image”) and an image corresponding to the signal intensity “S ₂ ” of the echo signal Echo in the above equation (3) "hereinafter referred to as" S ₂ image "" to build.

Finally, using the S ₁ image, the S ₂ image, and the above equations (3) and (4), a contrast image for each voxel of the phase gradient (θ) in the readout direction is obtained.

By adopting the above pulse sequence, the above-described S ₁ image and S ₂ image can be obtained simultaneously in the same imaging scan. Thereby, in this embodiment, there exists an effect that the noise reduction relevant to an imaging scan can be performed appropriately compared with the magnetic resonance imaging method of 1st Embodiment which requires two imaging scans.

  Hereinafter, an example of an experimental result obtained by measuring the phase gradient (θ) in the readout direction by the above phantom will be described.

  In this embodiment, as shown in FIG. 5, the computer 6 doubles the lead-out time so that the resolution of the lead-out is doubled compared with the resolution in the conventional lead-out direction when the echo signal Echo is read. And k-space was acquired in the lead-out direction (see FIG. 7). As a result, the length of the k space in the lead-out direction is double that of the conventional k space in the lead-out direction.

  Then, the computer 6 performed the following image processing according to the processing procedure by the image processing program.

First, the computer 6 obtains the above-described S ₁ image by performing Fourier transform on the right half of the k space, and displays this S ₁ image on the display device 16 as shown in FIG. Further, the computer 6 performs the Fourier transform on the left half of the k space to acquire the above-described S ₂ image, and displays this S ₂ image on the display device 16 as shown in FIG. 8B.

Then, the computer 6 uses the S ₁ image, the S ₂ image, and the above-described equations (3) and (4) to determine the phase gradient (θ) in the readout direction (hereinafter referred to as “PG in the readout direction ( A phase gradient (θ) contrast image for each voxel using the PG signal was displayed on the display device 16 as shown in FIG. 8C.

  Here, in FIG. 8C, when the phase gradient (θ) tends to increase from the left side to the right side, bright display (that is, display close to white) is performed, and the phase gradient (θ) is changed from the left side to the right side. If the phase gradient (θ) does not change, a neutral color is displayed (that is, a display close to gray).

  Further, in order to enable comparison with the display result of FIG. 8, FIG. 9 shows a gray level display of the average intensity of the phase obtained under the same experimental conditions (the higher the intensity, the brighter the display). doing).

  According to the comparison between the display result of FIG. 8C and the display result of FIG. 9, the display result of the phase gradient (θ) of FIG. 8C can reproduce the change of the phase intensity shown in FIG. It is considered reasonable.

  As described above, the magnetic resonance imaging method of the present embodiment can derive the phase gradient (θ) of the proton magnetic moment in the readout direction in the voxel that is a unit of the three-dimensional image analysis region, and as a result. The contrast image for each voxel of the phase gradient (θ) in the readout direction can be appropriately displayed on the computer 6. Therefore, the visualization of the contrast image of such a phase gradient (θ) can be expected to lead to the development of functional research of the human brain, and thus breakthrough advances in diagnostic technology in brain medicine.

  The present invention is not limited to the first and second embodiments and Examples 1 and 2 described above. In addition, the specific components of each part and each process are not limited to the first and second embodiments and Examples 1 and 2, and various modifications can be made without departing from the spirit of the present invention. is there.

  According to the magnetic resonance imaging method of the present invention, it is possible to derive a phase gradient of the magnetic moment of protons in which nuclear magnetic resonance has occurred in a predetermined direction in the voxel.

  Therefore, the present invention can be used as an imaging method of magnetic resonance imaging using nuclear magnetic resonance of protons, and can be applied to, for example, diagnostic imaging techniques in brain medicine.

1 is a block diagram illustrating a configuration example of a magnetic resonance imaging apparatus (MRI apparatus) according to a first embodiment of the present invention. It is the figure which showed an example of the pulse sequence used for the magnetic resonance imaging method of 1st Embodiment of this invention. The horizontal axis represents the phase difference of the proton magnetic moment, and the vertical axis represents the signal intensity of the echo signal Echo, depicting the relationship S = S ₀ Sinc (X). It is the figure which showed an example of the experimental result by the phantom of Example 1 as a photograph. It is the figure which showed an example of the pulse sequence used for the magnetic resonance imaging method of 2nd Embodiment of this invention. It is a schematic diagram explaining the construction method of the image by the magnetic resonance imaging method of 2nd Embodiment of this invention. It is the figure which showed an example of k space by the phantom of Example 2 as a photograph. It is the figure which showed an example of the experimental result by the phantom of Example 2 as a photograph. It is the figure which showed the example of a display of the phase intensity by the phantom of Example 2 as a photograph.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 MRI apparatus 2 Static magnetic field generating magnet 3 High frequency transmission coil 4 High frequency reception coil 5 Gradient magnetic field coil 6 Computer (image processing apparatus)
7 Sequencer 8 High-frequency generator 9 Modulator 10, 12 Amplifier 11 Gradient magnetic field generator 13 Phase detector 14 AD converter 15 Storage device 16 Display device

Claims (5)

Magnetic resonance imaging using nuclear magnetic resonance of protons,
The following formula (1) formulated with the phase difference between the center of the voxel and the end of the voxel as an independent variable (X) and the signal intensity correlated with the density of the proton as a dependent variable (S) Can be used to derive a gradient of the phase of the magnetic moment of the proton in a given direction in the voxel ,
The gradient of the phase at the end of the voxel relative to the center of the voxel is the variable term of the independent variable (X);
Gradient magnetic field pulses are applied in the predetermined direction, the phase at the end of the voxel in the predetermined direction is shifted with reference to the center of the voxel, and the independent variable (X) includes the shifted phase. Magnetic resonance imaging capable of enhancing the difference of the dependent variable (S) corresponding to the change of the variable (X) .
S = S ₀ Sinc (X) (1)
However,
S ₀ : Signal strength when there is no phase shift The dependent variable (S) when the phase at the end in the predetermined direction of the voxel is linearly shifted by π / 2 with respect to the center of the voxel, and the independent variable (X) includes the shifted phase. Is the first signal strength,
The dependent variable (S) when the phase at the end in the predetermined direction of the voxel is linearly shifted by −π / 2 with respect to the center of the voxel, and the independent variable (X) includes the shifted phase. ) As the second signal strength,
The difference between the first signal strength and the second signal strength and the sum between the first signal strength and the second signal strength are taken, and the difference is divided between the sum and the sum. ,
The signal intensity ratio obtained by dividing, using the relationship between the phase of the Ziegler stringent, magnetic resonance imaging according to claim 1, wherein directing the Ziegler stringent of the phase. The gradient magnetic field pulse is applied to the slice selection direction, and the phase at the end of the voxel in the slice selection direction with respect to the center of the voxel is linearly shifted by π / 2,
Wherein also applied gradient magnetic field pulses in the opposite direction slice selection direction, shifted with respect to the center of the voxel the phase at the end of the slice selection direction of the voxels - [pi] / 2, the linear, claim 2 Magnetic resonance imaging as described. In the readout direction to which the gradient magnetic field pulse is applied, k-space is acquired by an imaging scan,
An image corresponding to the first signal intensity and an image corresponding to the second signal intensity are obtained by performing a Fourier transform on each half of the k space divided with respect to the center of the k space,
The magnetic resonance imaging method according to claim 2 , wherein an image corresponding to the gradient of the phase is acquired using each of the acquired images. A static magnetic field generating means capable of generating a uniform static magnetic field in a space where the subject is placed;
A gradient magnetic field generating means capable of generating a gradient magnetic field in three orthogonal directions in the space;
High-frequency pulse generating means capable of irradiating a high-frequency pulse causing nuclear magnetic resonance to protons in a static magnetic field constituting the subject's living tissue;
Detection means for detecting a signal depending on the magnetic moment state of the proton in which the nuclear magnetic resonance has occurred;
5. A magnetic resonance imaging apparatus comprising: an image processing device that derives a phase gradient according to any one of claims 1 to 4 based on the signal and displays a contrast image of the phase gradient for each of the voxels. .

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