JP5683984B2 - Magnetic resonance imaging apparatus and nonlinear distortion correction method - Google Patents

Description

  The present invention relates to a magnetic resonance imaging (hereinafter referred to as “MRI”) apparatus, and more particularly, to a technique for correcting image distortion caused by gradient nonlinearity.

  An MRI apparatus measures an NMR (nuclear magnetic resonance) signal generated by a nuclear spin constituting a subject, particularly a human tissue, and two-dimensionally or three-dimensionally describes the form and function of the head, abdomen, limbs, and the like. It is a device that automatically images. In imaging, the NMR signal is given different phase encoding depending on the gradient magnetic field and is frequency-encoded to be measured as time series data. The measured NMR signal is reconstructed into an image by two-dimensional or three-dimensional Fourier transform.

  In order to obtain a correct reconstructed image, the gradient magnetic field distribution that imparts phase encoding to the NMR signal must be linear. However, in the current MRI apparatus, the linearity of the gradient magnetic field distribution is sacrificed in order to shorten the bore of the magnet or improve the speed of the gradient magnetic field. Such nonlinearity of the gradient magnetic field distribution causes distortion in the reconstructed image. Therefore, a correct image cannot be obtained. For this reason, a technique has been proposed in which the distortion amount of the reconstructed image due to the nonlinearity of the gradient magnetic field distribution is estimated from the distortion of the gradient magnetic field distribution, and the image distortion is corrected (see, for example, Patent Document 1).

JP 2009-226199 A

  The distortion correction method described in Patent Document 1 is a technique in which a distortion correction table for a three-dimensional space is stored in advance and correction is performed using the distortion correction table. The distortion correction table stores a plurality of two-dimensional distortion amounts obtained by a magnetic field measuring instrument or analytically and corresponds to a three-dimensional space. When a correction table is stored in advance, the accuracy of the correction table affects the accuracy of distortion correction. Further, when the correction table and the imaging position do not exactly match, a correction table corresponding to the imaging position must be calculated from nearby values.

  By the way, since the MRI apparatus can perform multi-slice imaging of a subject having a three-dimensional volume, it is possible to perform a detailed examination of the state inside the subject. In addition, the MRI apparatus is characterized by being capable of imaging a cross section having an arbitrary inclination. Specifically, in the MRI apparatus, three sets of orthogonal gradient magnetic fields are used and assigned to the slice direction, the phase encoding direction, and the frequency encoding direction, respectively, and imaging can be performed from various angles depending on the combination. In addition to the axial section, a sagittal section and a coronal section perpendicular to the axial section, and an oblique section whose direction perpendicular to the plane is not perpendicular to the gradient magnetic field axis can be taken.

  However, if the imaging position of the oblique section inclined at an arbitrary angle and the correction table do not match, a correction table corresponding to the imaging position must be calculated from nearby values. However, if the deviation from the correction table is large, it may be difficult to perform distortion correction with high accuracy by making the correction table correspond.

  Therefore, unless sufficient distortion correction due to nonuniform gradient magnetic field distribution is performed, the MR image may be distorted, blurred, or missing a signal.

  Therefore, an object of the present invention is to calculate the distortion amount of the gradient magnetic field in the cross section at the time when the imaging cross section is determined, from the shape of the gradient magnetic field coil using the Bio-Savart law, and to perform arbitrary imaging. An object of the present invention is to provide an MRI apparatus that corrects a distortion amount of a gradient magnetic field in a reconstructed image caused by nonlinearity of a gradient magnetic field distribution in a cross section.

  In order to achieve the above object, the outline of a representative one of the magnetic resonance imaging apparatuses of the present invention will be briefly described as follows.

  Static magnetic field generating means for generating a uniform static magnetic field in a space in which the subject is accommodated, gradient magnetic field generating means for generating a gradient magnetic field superimposed on the static magnetic field, and signal detection for detecting a magnetic resonance signal generated from the subject And a processing unit for generating a reconstructed image in the space of the subject based on the magnetic resonance signal, the magnetic resonance imaging apparatus comprising: a coil pattern constituting a gradient magnetic field generating unit in the processing unit The gradient magnetic field strength in an arbitrary imaging cross section in the space is calculated based on the image, the gradient magnetic field distortion amount in the imaging cross section is calculated using the calculated gradient magnetic field strength, and the reconstructed image is calculated based on the gradient magnetic field distortion amount. By correcting the amount of distortion, a reconstructed image from which the gradient magnetic field strength distortion has been removed is obtained.

  The outline of a representative one of the gradient magnetic field nonlinearity correction methods of the magnetic resonance imaging apparatus of the present invention will be briefly described as follows.

  Static magnetic field generating means for generating a uniform static magnetic field in a space in which the subject is accommodated, gradient magnetic field generating means for generating a gradient magnetic field superimposed on the static magnetic field, and signal detection for detecting a magnetic resonance signal generated from the subject A method for correcting non-linearity of a gradient magnetic field in a magnetic resonance imaging apparatus comprising: means for generating a reconstructed image in a space of a subject based on a magnetic resonance signal, and determining an imaging cross section of the subject Then, a first step of acquiring a parameter for specifying the imaging section, and a second step of calculating the distortion amount of the gradient magnetic field at each position on the imaging section acquired in the first step using the formula (A). The maximum or average value of the gradient amount of the gradient magnetic field calculated in the step 2 is compared with a preset reference value, and either the maximum value or the average value of the gradient amount of the gradient magnetic field is the reference value. If it does not exceed, either the third step waiting until the next new imaging cross-sectional area is determined and the maximum value or average value of the gradient amount of the gradient magnetic field calculated in the second step is the reference value. 4th step and 4th step of calculating the distortion amount of the gradient magnetic field for the other region of the imaging cross-sectional region that has not been subjected to the calculation process, using the formula (A) that is the calculation formula After completing the fifth step of calculating the oblique unit vector using the imaging cross-section parameters acquired in step S4, and the reconstruction processing of the captured image of each imaging cross-section, the values calculated in the fourth step and the fifth step are calculated. And a sixth step of performing distortion amount correction processing on the captured image.

    (Where r is the position vector from the origin of the space to the position for calculating the amount of strain of the gradient magnetic field, δr is the amount of strain of the gradient magnetic field at position r, Gr is the strength of the gradient magnetic field, and B (r) is the distance r from the origin. (It shows the magnetic flux density at a position that is far away.)

  According to the present invention, the amount of distortion of the reconstructed image due to the non-linearity of the gradient magnetic field distribution in an arbitrary imaging section is calculated from the shape of the gradient magnetic field coil, and correction is performed. Can be obtained.

It is a figure for demonstrating the whole structure of this invention. It is a figure explaining the formula which calculates | requires magnetic flux density. It is a figure which shows virtual magnetic field strength distribution and actual magnetic field strength distribution. It is the figure which defined the vector of the oblique image. It is a flowchart explaining the flow of a process of this invention.

  Hereinafter, preferred embodiments of the MRI apparatus of the present invention will be described in detail with reference to the accompanying drawings. Note that components having the same function are denoted by the same reference symbols throughout the drawings for describing the embodiments of the invention, and the repetitive description thereof is omitted.

  First, an overall outline of an example of an MRI apparatus according to the present invention will be described with reference to FIG. FIG. 1 is a block diagram showing the overall configuration of an embodiment of an MRI apparatus according to the present invention. This MRI apparatus obtains a tomographic image of a subject using the NMR phenomenon. As shown in FIG. 1, the MRI apparatus includes a static magnetic field generation system 2, a gradient magnetic field generation system 3, a transmission system 5, The receiving system 6, the signal processing system 7, the sequencer 4, and a central processing unit (CPU) 8 are provided.

  The static magnetic field generation system 2 generates a uniform static magnetic field in the direction perpendicular to the body axis in the space around the subject 1 if the vertical magnetic field method is used, and in the body axis direction if the horizontal magnetic field method is used. Thus, a permanent magnet type, normal conduction type or superconducting type static magnetic field generation source is arranged around the subject 1.

  The gradient magnetic field generation system 3 includes a gradient magnetic field coil 9 that applies gradient magnetic fields in the three-axis directions of X, Y, and Z, which are coordinate systems (stationary coordinate systems) of the MRI apparatus, and gradient magnetic fields that drive the respective gradient magnetic field coils. A gradient power supply Gx, Gy, Gz is applied in the X, Y, and Z triaxial directions by driving the gradient magnetic field power supply 10 of each coil in accordance with a command from the sequencer 4 described later. . At the time of imaging, a slice direction gradient magnetic field pulse (Gs) is applied in a direction orthogonal to the slice plane (imaging cross section) to set a slice plane for the subject 1, and the remaining two orthogonal to the slice plane and orthogonal to each other. A phase encoding direction gradient magnetic field pulse (Gp) and a frequency encoding direction gradient magnetic field pulse (Gf) are applied in one direction, and position information in each direction is encoded in the echo signal.

  The sequencer 4 is a control unit that repeatedly applies a high-frequency magnetic field pulse (hereinafter referred to as “RF pulse”) and a gradient magnetic field pulse in a predetermined pulse sequence. The sequencer 4 operates under the control of the CPU 8 and collects tomographic image data of the subject 1. Various commands necessary for the transmission are sent to the transmission system 5, the gradient magnetic field generation system 3, and the reception system 6.

  The transmission system 5 irradiates the subject 1 with an RF pulse in order to cause nuclear magnetic resonance to occur in the nuclear spins of the atoms constituting the living tissue of the subject 1, and includes a high-frequency oscillator 11, a modulator 12, and a high-frequency amplifier. 13 and a high frequency coil (transmission coil) 14a on the transmission side. The RF pulse output from the high-frequency oscillator 11 is amplitude-modulated by the modulator 12 at a timing according to a command from the sequencer 4, and after the amplitude-modulated RF pulse is amplified by the high-frequency amplifier 13, the RF pulse is arranged close to the subject 1. By supplying to the high frequency coil 14a, the subject 1 is irradiated with the RF pulse.

  The receiving system 6 detects an echo signal (NMR signal) emitted by nuclear magnetic resonance of nuclear spins constituting the living tissue of the subject 1. The receiving system 6 receives a high-frequency coil (receiving coil) 14 b on the receiving side and a signal amplifier 15. And a quadrature phase detector 16 and an A / D converter 17. After the NMR signal of the response of the subject 1 induced by the electromagnetic wave irradiated from the high frequency coil 14a on the transmission side is detected by the high frequency coil 14b arranged close to the subject 1 and amplified by the signal amplifier 15, The quadrature phase detector 16 divides the signal into two orthogonal signals at the timing according to the command from the sequencer 4, and each signal is converted into a digital quantity by the A / D converter 17 and sent to the signal processing system 7.

  The signal processing system 7 performs various data processing and display and storage of processing results, and includes an external storage device such as an optical disk 19 and a magnetic disk 18 and a display 20 including a CRT or the like. When data from the reception system 6 is input to the CPU 8, the CPU 8 executes processing such as signal processing and image reconstruction, and displays the tomographic image of the subject 1 as a result on the display 20 and an external storage device. On the magnetic disk 18 or the like.

  The operation unit 25 inputs various control information of the MRI apparatus and control information of processing performed by the signal processing system 7 and includes a trackball or mouse 23 and a keyboard 24. The operation unit 25 is arranged in the vicinity of the display 20, and the operator controls various processes of the MRI apparatus interactively through the operation unit 25 while looking at the display 20.

  In FIG. 1, the high-frequency coil 14a and the gradient magnetic field coil 9 on the transmission side are opposed to the subject 1 in the static magnetic field space of the static magnetic field generation system 2 into which the subject 1 is inserted. If the horizontal magnetic field method is used, the subject 1 is installed so as to surround it. The high-frequency coil 14b on the receiving side is installed so as to face or surround the subject 1.

  At present, the radionuclide to be imaged by the MRI apparatus is a hydrogen nucleus (proton) which is a main constituent material of the subject as widely used clinically. Information on the spatial distribution of the proton density and the spatial distribution of the relaxation time of the excited state is imaged, thereby imaging the form or function of the human head, abdomen, limbs, etc. two-dimensionally or three-dimensionally.

  Next, embodiments of the present invention will be described with reference to the drawings.

  FIG. 2A shows a space in which a static magnetic field is formed by the static magnetic field magnet 31, that is, an imaging region (FOV: Field Of View) 33. A gradient magnetic field coil 9a is provided above the imaging region 33, and a static magnetic field magnet 31 is provided above the gradient magnetic field coil 9a. In addition, the gradient magnetic field coil 9a shown here has shown only one of a pair of Y-axis gradient magnetic field coils. The coil constitutes a closed circuit (C) 32.

  In order to represent the position of the imaging region 33, X, Y and Z axes are given as shown in the figure. A plane parallel to the plane determined by each of the two axes is a reference plane (axial plane, sagittal plane, coronal plane, etc.), the origin of the X, Y, and Z axes is indicated by symbol O, and is separated from the origin O by a distance r. The position on the oblique section 34 is P.

  Here, the oblique section 34 is a plane inclined at an angle of ψ (psi) with respect to a predetermined inclination from the reference plane, for example, a plane determined by the X axis and the Z axis. Even a plane determined by the Y-axis is a plane inclined at an angle of φ (phi). In the present embodiment, the following description will be made on the assumption that the cross section of the subject is imaged on the plane.

  Needless to say, the present embodiment can be applied even when ψ (psy) is zero, φ (phi) is zero, or both are zero.

The magnetic flux density at the position P at this time is obtained below.
When the current I flows through the closed circuit (C) 32 shown in FIG. 2A according to Bio-Savart's law, the magnetic flux density B (r) at the position P away from the origin O by the distance r is

It is represented by Here, ds represents a minute region (line segment) on the closed circuit (C) 32, μ represents magnetic permeability, vector s represents a vector starting from the origin O and ending at the minute region ds.
Further, the integration of Expression (1) indicates that line integration is performed along the closed circuit (C) 32. Note that the distance from the minute region ds to the position P is expressed as an rs vector.

Incidentally, the MR imaging is executed according to a pulse sequence of an imaging method such as an SE (Spin Echo) method or an FE (Field Echo) method under the control of the sequencer 4. A k-space database is formed by using digital raw data output from the sequencer 4 as k-space.
Image reconstruction processing such as two-dimensional or three-dimensional Fourier transform processing or maximum value projection processing is performed on the acquired k-space data, and a reconstruction image region (FOV) in a non-linear actual Z-axis gradient magnetic field Is generated.

  FIG. 2B shows a virtual magnetic field strength distribution and an actual magnetic field strength distribution of the Z-axis gradient magnetic field formed from the origin O of the imaging region 33 shown in FIG. 2A toward the outer end of the imaging region 33.

The virtual magnetic field strength distribution extends from the origin O to the outer end of the imaging region 33 while maintaining linearity.
On the other hand, the actual magnetic field strength distribution extends while maintaining linearity from the origin O to a certain region, but the linearity cannot be maintained from a certain region, and the magnetic field strength distribution shows non-linearity. The non-linearity shown in the figure is only an example. The nonlinear characteristic varies depending on the MRI apparatus to be handled.

If the reconstruction FOV is in a narrow range, that is, the reconstruction FOV (1) shown in the figure, the virtual and actual magnetic field strengths at the right end of the reconstruction FOV (1) are Z _V (1 ) And Z _R (1), and almost no difference in magnetic field strength is observed.

On the other hand, when the reconstruction FOV is in a wide range, that is, in the reconstruction FOV (2) shown in the figure, the virtual and actual magnetic field strengths at the right end of the reconstruction FOV (2) are Z _V (2 ) And Z _R (2), and the difference in the magnetic field strength appears remarkably.

  Therefore, since the actual intensity distribution of the gradient magnetic field is non-linear, the magnetic field intensity at the position P deviates from the virtual magnetic field intensity distribution having linearity and is distorted.

Next, the amount of distortion is determined using the above equation (1).
The amount of distortion of the gradient magnetic field at the position P can be obtained by the following equation using the gradient magnetic field strength Gr.

The coordinate of r is set to (x, y, z) in the coordinate system of the stationary coordinate system, and the x component, y component, and z component of the gradient magnetic field strength at the position P represented by the coordinate r are Gx, Gy, Gz, respectively. Then, the above formula is

And can be decomposed into each axis component.

  Actually, as described above, in many cases, the imaging section is tilted with respect to the reference plane, that is, the imaging section is not perpendicular or parallel to the stationary coordinate system (when it is oblique).

  Hereinafter, a calculation method when there is an oblique is described below. If there is no oblique, one component of a unit vector described later may be set to zero.

The horizontal vector and vertical vector unit vector for the captured image shown in FIG. 3 are represented by R (r _x , r _y , r _z ) and C (c _x , c _y , c _y ), respectively. Assuming that the distortion amount δR of the gradient magnetic field in the horizontal direction is the distortion amount δC of the gradient magnetic field in the vertical direction, the distortion amount of each gradient magnetic field is the unit vector and the distortion amount (reference value) of the gradient magnetic field when not oblique. Using equation (3),

It is expressed. As described above, at all points on the image, the distortion of the broken image is corrected by moving the pixels by δR and δC in the horizontal and vertical directions of the image, respectively.

  The above description relates to the calculation method of the magnetic field strength at the position P and the calculation of the distortion amount of the gradient magnetic field at the position P. By repeating the same method over the entire oblique section 34, the oblique section 34 is obtained. The magnetic field strength distribution over the entire surface is obtained. The distortion amount of the gradient magnetic field over the entire surface can be calculated from the calculated magnetic field strength distribution over the entire surface and the virtual magnetic field strength distribution. With this correction method, it is possible to obtain a captured image with reduced defects such as blurring and artifacts in the oblique section 34.

  Since the calculation of the gradient magnetic field distortion amount can be executed at the stage when the imaging cross section is determined before the image reconstruction is completed, the gradient magnetic field distortion amount is calculated and the image is acquired when the position information is determined. Can be divided into two stages so as to execute distortion correction according to the distortion amount of the gradient magnetic field.

  As shown in FIG. 2B, the distortion amount of the gradient magnetic field increases as the distance from the magnetic field center (here, the origin O is the magnetic field center) increases. Therefore, the distortion amount of the gradient magnetic field at the position farthest from the center of the magnetic field in the imaging cross section is calculated first, and if the distortion amount of the gradient magnetic field is below the allowable range, the calculation of the gradient magnetic field distortion amount and the distortion correction are performed. It is also possible not to execute.

Next, the flow of processing will be described with reference to FIG.
After the imaging section is determined, a parameter for specifying the imaging section is acquired (step 401).
An imaging cross-sectional area farthest from the magnetic field center (outside edge of the reconstructed FOV area) is extracted from the imaging cross-section acquired in step 401, and each of the imaging cross-sectional areas on the imaging cross-sectional area is expressed by Expression (2) and Expression (3) The amount of gradient magnetic field distortion at the position is calculated (step 402).
When either the maximum value or the average value of the distortion amount of the gradient magnetic field calculated at step 402 exceeds the reference value, step 404 is executed, and when it does not exceed, the next new imaging section is determined. (Step 403).
In step 404, the distortion amount of the gradient magnetic field is calculated for the remaining uncalculated cross section using the same calculation formula as in step 402. An oblique unit vector is calculated using the imaging section parameters acquired in step 401 (step 405).
Step 401 to step 405 can be executed immediately after the imaging section is determined. After the reconstruction processing of the image of each imaging section is completed, the distortion correction processing is performed on the image using the values calculated in Step 404 and Step 405 (Step 406).

  As described above, according to the present embodiment, it is possible to correct distortion of an oblique image. As in the past, a correction data table at a specific position is prepared in advance, and when there is no corresponding position coordinate of the image, this correction is compared to a method using a neighboring value or a method obtained by an interpolation method as appropriate. Can correct the distortion amount of the gradient magnetic field with high accuracy. In particular, this is effective in correcting the amount of distortion of the gradient magnetic field in an oblique image. The reason is as described above.

1: subject, 2: static magnetic field generation system, 3: gradient magnetic field generation system, 4: sequencer, 5: transmission system, 6: reception system, 7: signal processing system, 8: central processing unit (CPU), 9: Gradient magnetic field coil, 9a: Y axis gradient magnetic field coil, 10: Gradient magnetic field power supply, 11: High frequency transmitter, 12: Modulator, 13: High frequency amplifier, 14a: High frequency coil (transmitting coil), 14b: High frequency coil (receiving coil) ), 15: signal amplifier, 16: quadrature detector, 17: A / D converter, 18: magnetic disk, 19: optical disk, 20: display, 21: ROM, 22: RAM, 23: trackball or mouse, 24: keyboard, 31: static magnetic field magnet, 32: closed circuit (C), 33: imaging region, 34: oblique section.

Claims (6)

A static magnetic field generating means for generating a uniform static magnetic field in a space for accommodating a subject;
A gradient magnetic field generating means for generating a gradient magnetic field superimposed on the static magnetic field;
Signal detection means for detecting a magnetic resonance signal generated from the subject;
Processing means for generating a reconstructed image in the space of the subject based on the magnetic resonance signal, and a magnetic resonance imaging apparatus comprising:
In the processing means,
Calculate the gradient magnetic field strength at any imaging cross section in the space based on the pattern of the coils constituting the gradient magnetic field generating means,
Calculate the amount of distortion of the gradient magnetic field in the imaging cross section using the calculated gradient magnetic field strength,
Obtaining a reconstructed image in which the distortion of the gradient magnetic field strength is removed by correcting the distortion amount of the reconstructed image based on the distortion amount of the gradient magnetic field;
The calculation of the gradient amount of the gradient magnetic field in the imaging section compares the maximum or average value of the gradient amount of the gradient magnetic field calculated using the gradient magnetic field strength with a reference value set based on the virtual gradient magnetic field strength. The magnetic resonance imaging apparatus is executed when one of the values exceeds the reference value. The magnetic resonance imaging apparatus according to claim 1.
The calculation of the amount of distortion of the gradient magnetic field in the imaging section starts from the outer end of the space toward the center of the gradient magnetic field. The magnetic resonance imaging apparatus according to claim 1.
Processing for calculating the distortion amount of the gradient magnetic field in the imaging section, and detecting a magnetic resonance signal generated from the subject using the signal detecting means, and reconstructing the subject based on the magnetic resonance signal The magnetic resonance imaging apparatus is characterized in that the process of generating is performed in parallel in time series. A static magnetic field generating means for generating a uniform static magnetic field in a space accommodating the subject; a gradient magnetic field generating means for generating a gradient magnetic field superimposed on the static magnetic field; and a magnetic resonance signal generated from the subject A gradient magnetic field non-linearity correction method in a magnetic resonance imaging apparatus comprising: a signal detection unit; and a processing unit that generates a reconstructed image of the subject in the space based on the magnetic resonance signal,
A first step of determining an imaging cross section of the subject and then obtaining a parameter identifying the imaging cross section;
A second step of calculating the distortion amount of the gradient magnetic field at each position on the imaging cross section acquired in the first step using the formula (A);
The maximum value or average value of the strain amount of the gradient magnetic field calculated in the second step is compared with a preset reference value, and either the maximum value or the average value of the strain amount of the gradient magnetic field is the reference value. If not, the third step of waiting until the next new imaging cross-sectional area is determined;
When either the maximum value or the average value of the strain amount of the gradient magnetic field calculated in the second step exceeds the reference value, the calculation process is still performed using the equation (A). A fourth step of calculating the amount of distortion of the gradient magnetic field for another region of the imaging cross-sectional region that has not been performed;
A fifth step of calculating an oblique unit vector using the imaging section parameters acquired in the first step;
A sixth step of performing a distortion amount correction process on the captured image using the values calculated in the fourth step and the fifth step after the reconstruction process of the captured image of each imaging section is completed; A method for correcting non-linear distortion of a gradient magnetic field, comprising:

(Where r is a position vector from the origin in the space to the position for calculating the amount of strain of the gradient magnetic field, δr is the amount of strain of the gradient magnetic field at position r, Gr is the strength of the gradient magnetic field, and B (r) is the origin. (The magnetic flux density at a position away from the center by a distance r is shown.) The gradient magnetic field non-linearity correction method according to claim 4 ,
In the second step, an imaging cross-sectional area located in contact with the outer edge of the space is extracted, and the distortion amount of the gradient magnetic field at each position on the imaging cross-sectional area is calculated using the equation (A). A gradient magnetic field non-linearity correction distortion method characterized by: The gradient magnetic field non-linearity correction method according to claim 4 ,
A process of calculating a distortion amount of the gradient magnetic field in the imaging section in the second step, a magnetic resonance signal generated from the subject using the signal detection means, and detecting the magnetic resonance signal based on the magnetic resonance signal. A gradient magnetic field nonlinear distortion correction method, characterized in that processing for generating a reconstructed image of a specimen is performed in parallel in time series.

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