Magnetic resonance imaging device, image correction method, and static magnetic field inhomogeneity correction method

Abstract

The present invention provides an MRI apparatus and correction method capable of correcting static magnetic field inhomogeneities caused by the application of a gradient magnetic field through simple calculations. Using the pulse sequence shape used in imaging, a simple calculation is performed to estimate image quality degradation caused by static magnetic field inhomogeneities caused by the application of a gradient magnetic field. Based on the estimated static magnetic field inhomogeneities and the pulse sequence shape, the amount of distortion in the image obtained by executing the pulse sequence is estimated, and the distortion of the reconstructed image is corrected. Alternatively, a compensation magnetic field output is calculated to eliminate the estimated static magnetic field inhomogeneities, superimposed on the compensation current of the active shimming, and supplied to the shimming coils.

Description

Technical Field

[0001] The present invention relates to a magnetic resonance imaging apparatus (hereinafter referred to as an MRI apparatus), and more particularly to a technique for reducing image quality degradation caused by non-uniform static magnetic field. Background Art

[0002] An MRI apparatus applies a high-frequency magnetic field to a subject placed in a uniform static magnetic field, causing nuclear magnetic resonance (NMR) of atomic nuclei (protons) in arbitrary regions of the subject. The resulting NMR signals (or echo signals) produce a tomographic image of that region. To selectively excite specific regions, a gradient magnetic field is applied along with the high-frequency magnetic field. Furthermore, the applied magnetic field and intensity of the gradient magnetic field are controlled to impart spatial positional information to the measured echo signals.

[0003] When a gradient magnetic field is applied by an MRI apparatus, the Z-axis is generated depending on the applied axis and applied intensity of the gradient magnetic field. 2 Static magnetic field inhomogeneities caused by Maxwell terms such as XY, ZY, and XZ lead to image quality degradation. Therefore, various methods for correcting static magnetic field inhomogeneities caused by Maxwell terms have been studied. For example, the technology described in Patent Document 1 proposes a method for calculating the time average of a gradient magnetic field in a given time unit according to a pulse sequence, and controlling the output of a shim coil that adjusts for static magnetic field inhomogeneities based on the calculated time average. This method allows correction of static magnetic field inhomogeneities generated when a gradient magnetic field is applied using a low-capacity power supply. Furthermore, Patent Document 2 proposes a method for pre-measuring the error magnetic field generated in the phase direction during the execution of an Echo Planar Imaging (EPI) sequence, and correcting the image based on this error magnetic field.

[0004] Prior art literature

[0005] Patent Literature

[0006] Patent Document 1: Japanese Patent Application Laid-Open No. 2017-86736

[0007] Patent Document 2: Japanese Patent Application Laid-Open No. 2007-117765

[0008] The technique in Patent Document 1 requires switching the compensation current (shimming current) flowing through the shim coils. Therefore, if the shim coils respond slowly, the corrected magnetic field output may not be controlled in real time, rendering it impractical. Furthermore, the relatively short switching of the shim currents can create new challenges, such as the generation of eddy currents. On the other hand, the technique in Patent Document 2 uses pre-measurement to determine the error magnetic field (static magnetic field inhomogeneity). Therefore, the imaging conditions and pulse sequence used in actual imaging must be consistent with the pre-measured conditions, resulting in the need to restrict the imaging conditions and pulse sequence. Summary of the Invention

[0009] Problems to be solved by the invention

[0010] An object of the present invention is to provide an MRI apparatus capable of simply correcting static magnetic field inhomogeneity caused by application of a gradient magnetic field without requiring prior measurement or output control of shim coils in accordance with a gradient magnetic field control signal.

[0011] Means for solving problems

[0012] To address the aforementioned issues, the MRI apparatus of the present invention estimates the effect of static magnetic field inhomogeneities on images caused by the application of a gradient magnetic field, based on the shape of the pulse sequence used during imaging. This estimation result is used to correct the captured images. Alternatively, based on the estimation result, a compensation current is pre-energized through the shim coils to suppress static magnetic field inhomogeneities caused by the application of a gradient magnetic field during imaging.

[0013] Specifically, the MRI apparatus of the present invention includes: a transmitter for irradiating a subject placed in a static magnetic field space with high-frequency pulses; a receiver for receiving nuclear magnetic resonance signals generated from the subject by the irradiation with the high-frequency pulses; a gradient magnetic field generator for generating a gradient magnetic field in the static magnetic field space; a controller for controlling the transmitter, the gradient magnetic field generator, and the receiver according to a predetermined pulse sequence; and an image generator for reconstructing an image of the subject based on the nuclear magnetic resonance signals. The apparatus further includes: a calculation unit for estimating static magnetic field inhomogeneities resulting from the application of the gradient magnetic field according to the predetermined pulse sequence. The calculation unit simply estimates static magnetic field inhomogeneities using the application intensity or average application intensity of the gradient magnetic field pulses determined by the predetermined pulse sequence.

[0014] According to one embodiment of the present invention, the calculation unit calculates distortion occurring in the image based on the estimated static magnetic field inhomogeneity, and the image generation unit corrects the reconstructed image based on the distortion calculated by the calculation unit.

[0015] According to another aspect of the present invention, the calculation unit calculates the compensation current for generating the correction magnetic field based on the estimated static magnetic field inhomogeneity, and the control unit performs control to pre-circulate the calculated compensation current.

[0016] Effects of the Invention

[0017] According to the present invention, image quality degradation caused by static magnetic field inhomogeneity resulting from the application of a gradient magnetic field can be corrected without real-time control of the correction magnetic field output. Furthermore, the error magnetic field generated secondary to the real-time variation of the correction magnetic field output can be eliminated. This provides significant results, particularly in low- and medium-field MRI apparatuses, which are significantly affected by static magnetic field inhomogeneity resulting from the application of a gradient magnetic field, and in apparatuses with slow shim coil response. Furthermore, since static magnetic field inhomogeneity is estimated based on the pulse sequence used during actual imaging, rather than preliminary measurements, there is no need to restrict imaging conditions or pulse sequences, thereby increasing the degree of freedom in imaging. BRIEF DESCRIPTION OF THE DRAWINGS

[0018] Figure 1 This is a block diagram showing the overall configuration of the MRI apparatus according to the present invention.

[0019] Figure 2 This is a functional block diagram of the arithmetic processing unit according to the first embodiment.

[0020] Figure 3 This is a flowchart showing static magnetic field inhomogeneity estimation and image correction processing according to the first embodiment.

[0021] Figure 4 This is a diagram showing an example of a sequence diagram of SE EPI.

[0022] Figure 5 This is a diagram showing the intensity of the gradient magnetic field applied in the phase encoding direction and the frequency encoding direction of the image.

[0023] Figure 6 : is a diagram showing an example of absolute coordinates of a perpendicular magnetic field device.

[0024] Figure 7 This is a diagram showing an example of estimating the inhomogeneity of the static magnetic field generated in the XZ plane.

[0025] Figure 8 This diagram displays the amount of image distortion due to static magnetic field inhomogeneity Bc as a vector.

[0026] Figure 9 is a diagram showing image distortion caused by static magnetic field inhomogeneity Bc, Figure 9 (a) shows the original image without distortion, Figure 9 (b) shows an image distorted by Bc.

[0027] Figure 10 is a diagram showing an example of correction of image distortion, Figure 10 (a) shows an image distorted by Bc. Figure 10 (b) shows the image after the distortion is corrected.

[0028] Figure 11 is a diagram showing an example of a pulse sequence used in a modification of the first embodiment. Figure 11 (A) shows a 2D SE sequence, Figure 11 (B) shows the 3D GE sequence.

[0029] Figure 12 This is a functional block diagram of the arithmetic processing unit according to the second embodiment.

[0030] Figure 13 This is a flowchart showing the process of estimating the adjustment value of static magnetic field inhomogeneity according to the second embodiment.

[0031] Figure 14 This is a diagram explaining the non-uniformity of the static magnetic field depending on the slice position.

[0032] Figure 15 This is a functional block diagram of the arithmetic processing unit according to the third embodiment.

[0033] Figure 16 This is a flowchart showing the process of estimating static magnetic field inhomogeneity and correcting static magnetic field inhomogeneity according to the third embodiment.

[0034] Figure 17 This is a diagram showing a flow of calculating adjustment values ​​of shim coil outputs according to the fourth embodiment.

[0035] Description of Reference Numerals

[0036] 21: Image generation unit, 22: Bc estimation unit, 23: Distortion calculation unit, 24: Adjustment value calculation unit, 25: Compensation output calculation unit, 100: MRI apparatus, 102: Static magnetic field generator, 103: Gradient magnetic field coil, 104: RF transmitting coil, 105: RF receiving coil, 107: Signal processing unit (receiving unit), 109: Shim coil (compensation magnetic field generator), 110: Shim power supply (compensation magnetic field generator), 111: Measurement control unit, 112: Overall control unit, 200: Processing unit (CPU). DETAILED DESCRIPTION

[0037] First, refer to Figure 1 The MRI apparatus according to the present invention will be described. Figure 1 This is a block diagram showing the overall configuration of an embodiment of an MRI apparatus according to the present invention.

[0038] This MRI apparatus 100 utilizes the NMR phenomenon to obtain tomographic images of a subject 101. It includes a static magnetic field generating magnet 102, a gradient magnetic field coil 103, and a gradient magnetic field power supply 108 (gradient magnetic field generating unit), an RF transmitting coil 104 and an RF transmitting unit 106 (transmitting unit), an RF receiving coil 105 and a signal processing unit 107 (receiving unit), a measurement control unit 111, an overall control unit (control unit) 112, a display and operation unit 118, and a bed 120 for allowing a tabletop on which the subject 101 is mounted to move in and out of the static magnetic field generating magnet 102. The MRI apparatus 100 may further include shim coils 109 and a shim power supply 110 (compensation magnetic field generating unit) for correcting static magnetic field inhomogeneities.

[0039] The static magnetic field generating magnet 102 is a magnet that generates a uniform static magnetic field in a direction perpendicular to the body axis of the object 101 if it is a vertical magnetic field method, and generates a uniform static magnetic field in the body axis direction if it is a horizontal magnetic field method, and a static magnetic field generating source of a permanent magnet method, a normal conductor method or a superconducting method is arranged around the object 101.

[0040] The gradient magnetic field coils 103 are coils wound in the X, Y, and Z axes of the MRI apparatus's actual spatial coordinate system (stationary coordinate system). Each gradient magnetic field coil is connected to a gradient magnetic field power supply 108 that drives it and supplies current. Specifically, the gradient magnetic field power supply 108 for each gradient magnetic field coil is driven according to commands from the measurement control unit 111, described later, to supply current to each gradient magnetic field coil. This generates gradient magnetic fields Gx, Gy, and Gz in the X, Y, and Z axes. The gradient magnetic field coils 103 and the gradient magnetic field power supply 108 are collectively referred to as a gradient magnetic field generating unit. When imaging a two-dimensional slice plane, a slice gradient magnetic field pulse (Gs) is applied in a direction orthogonal to the slice plane (imaging section) to set the slice plane for the subject 101. Phase encoding gradient magnetic field pulses (Gp) and frequency encoding (readout) gradient magnetic field pulses (Gf) are applied in the remaining two directions, which are orthogonal to the slice plane and mutually orthogonal to each other. Position information in each direction is encoded in the nuclear magnetic resonance signal (echo signal).

[0041] The shim coil 109 and the shim power supply 110 function as a compensation magnetic field generator. 2The coils for generating the spatially distributed magnetic field of spherical harmonic functions of order zero or higher, such as XY, ZY, and XZ, are connected to a shim power supply 110. To improve the uniformity of the static magnetic field, a magnetic field is generated using compensation currents supplied from the shim power supply 110. Alternatively, the gradient magnetic field coils 103 may also function as compensation magnetic field generators for the primary gradient magnetic field. In this case, the gradient magnetic field power supply 108 supplies the gradient magnetic field coils 103 with a compensation current for generating a correction magnetic field superimposed on the aforementioned gradient magnetic field current used to provide positional information.

[0042] The RF transmitting coil 104 is a coil that irradiates the subject 101 with RF pulses. It is connected to the RF transmitting unit 106 and supplied with a high-frequency pulse current. This induces the NMR phenomenon in the spins of atoms constituting the biological tissue of the subject 101. Specifically, the RF transmitting unit 106 is driven in accordance with instructions from the measurement control unit 111, described later, to amplitude-modulate the high-frequency pulses, amplify them, and then supply them to the RF transmitting coil 104, which is positioned close to the subject 101. This irradiates the subject 101 with RF pulses. The RF transmitting coil 104 and the RF transmitting unit 106 are collectively referred to as a transmitting unit.

[0043] The RF receiving coil 105 receives echo signals emitted by the NMR phenomenon of spins in the biological tissue constituting the subject 101, is connected to the signal processing unit 107, and transmits the received echo signals to the signal processing unit 107. The RF receiving coil 105 and the signal processing unit 107 are collectively referred to as a receiving unit.

[0044] The signal processing unit 107 detects and processes the echo signals received by the RF receiving coil 105. Specifically, in accordance with instructions from the measurement control unit 111 (described later), the signal processing unit 107 amplifies the received echo signals, divides them into two orthogonal signal systems using quadrature phase detection, samples a predetermined number of signals (e.g., 128, 256, 512, etc.), and performs A / D conversion on each sampled signal to convert it into a digital quantity. This yields the echo signals as time-series digital data (hereinafter referred to as echo data) consisting of the predetermined number of sampled data. The signal processing unit 107 then performs various processing on the echo data and transmits the processed echo data to the measurement control unit 111.

[0045] The measurement control unit 111 primarily sends various commands to the gradient magnetic field power supply 108, the RF transmitter 106, and the signal processing unit 107 for collecting echo data required for reconstructing tomographic images of the subject 101, thereby controlling these components. Specifically, the measurement control unit 111 operates under the control of the overall control unit 112, described later. Based on control data for a given pulse sequence, the control unit 111 controls the gradient magnetic field power supply 108, the RF transmitter 106, and the signal processing unit 107. The control unit repeatedly irradiates the subject 101 with RF pulses and applies gradient magnetic field pulses, as well as detects echo signals from the subject 101. This control unit controls the collection of echo data required for reconstructing images of the imaging region of the subject 101. During these repetitions, the application amount of the phase encoding gradient magnetic field is varied for two-dimensional imaging, and the application amount of the slice encoding gradient magnetic field is also varied for three-dimensional imaging. The number of phase codes per image can generally be selected as 128, 256, 512, etc., and the number of slice codes can generally be selected as 16, 32, 64, etc. Through these controls, echo data from the signal processing unit 107 is output to the overall control unit 112 .

[0046] The overall control unit 112 controls the measurement control unit 111, as well as various data processing functions and displays and saves processing results. It can be implemented on a computer or workstation equipped with a processing unit (CPU) 200, a memory 210, and an internal storage unit 220 such as a disk. However, some of its functions may be implemented by hardware such as an ASIC or FPGA. The overall control unit 112 composed of such a computer or workstation may also include a network interface 230 for connecting to an external network and may be connected to an external storage unit 240 such as an optical disk.

[0047] Specifically, the overall control unit 112 causes the measurement control unit 111 to collect echo data by executing an imaging sequence. Upon receiving the echo data from the measurement control unit 111, the CPU 200 stores the data in an area corresponding to k-space within the memory 210 based on the encoding information applied to the echo data. Storing the echo data in an area corresponding to k-space within the memory 210 is hereinafter referred to as "arranging the echo data in k-space." Furthermore, a set of echo data stored in an area corresponding to k-space within the memory 210 is also referred to as k-space data.

[0048] The processing unit (CPU) 200 performs signal processing, image reconstruction based on Fourier transform, and other processing on the k-space data, so that the image of the subject 101 as the result is displayed on the display and operation unit 118, recorded in the internal storage unit 210, the external storage unit 240, or transmitted to an external device via the network IF230.

[0049] The display and operation unit 118 includes a display unit for displaying reconstructed images of the subject 101, and an operation unit such as a trackball, mouse, and keyboard for inputting various control information for the MRI apparatus and control information for processing performed by the overall control unit 112. This operation unit is located close to the display unit, and the operator observes the display and interactively controls various processes of the MRI apparatus through the operation unit.

[0050] In the MRI apparatus of this embodiment, the processing unit 200, in addition to functioning as the image generator that performs the aforementioned image reconstruction, also functions as a static magnetic field inhomogeneity (Bc) estimation unit. This Bc estimation unit estimates static magnetic field inhomogeneity resulting from the application of a gradient magnetic field based on the pulse sequence used in imaging, as will be described in detail later. Based on the static magnetic field inhomogeneity estimated by the Bc estimation unit, image correction and the generation of a compensation current to correct the static magnetic field inhomogeneity are performed. An embodiment of an MRI apparatus equipped with static magnetic field inhomogeneity estimation processing functionality is described below.

[0051] <Implementation Method 1>

[0052] The MRI apparatus of this embodiment estimates static magnetic field inhomogeneity Bc caused by application of a gradient magnetic field during imaging based on the pulse sequence used for imaging, calculates a correction value based on the estimated static magnetic field inhomogeneity to correct image quality degradation caused by the static magnetic field inhomogeneity, and performs correction.

[0053] like Figure 2 As shown, the arithmetic processing unit 200 of this embodiment includes an image generating unit 21 for reconstructing an image using k-space data; a Bc estimating unit 22 for estimating static magnetic field inhomogeneity based on a pulse sequence; and a distortion calculating unit 23 for calculating distortion generated in the image based on the static magnetic field inhomogeneity estimated by the Bc estimating unit 22.

[0054] Below, refer to Figure 3 The processing of the MRI apparatus including the above-described arithmetic processing unit 200 will be described.

[0055] First, when the pulse sequence and imaging parameters are set, the measurement control unit 111 generates a pulse sequence used for imaging (hereinafter referred to as an imaging sequence) (S301). Here, as an example, the imaging sequence is set to an SE EPI sequence. Figure 4 An example of a SE EPI sequence is shown in FIG.

[0056] In the SE EPI sequence, such as Figure 4As shown, an excitation RF pulse (90-degree pulse) 401 is applied together with a slice selection gradient magnetic field pulse 403, and then an inversion RF pulse (180-degree pulse) 402 is applied together with a slice selection gradient magnetic field pulse 404 at a timing of TE (echo time) / 2. Then, the polarity of the frequency encoding (readout) gradient magnetic field pulse (Gf) 407 is converted at high speed, and a Blip pulse 406 is applied to the phase encoding gradient magnetic field, and a large number of echo signals 408 are acquired with a single excitation (shot). The phase encoding gradient magnetic field pulse 405 is a dephasing pulse. Figure 4 In the embodiment, although Blip pulses are used in the phase encoding direction, there are cases where a gradient magnetic field is continuously applied while echo signals are collected.

[0057] Once an imaging sequence is determined, the Bc estimation unit 22 uses the applied intensity or average applied intensity of the gradient magnetic field pulses included in the imaging sequence to simply estimate static magnetic field inhomogeneity (S302). The static magnetic field inhomogeneity caused by the gradient magnetic field pulses varies over time depending on the gradient magnetic field output. Therefore, rigorously determining this requires complex calculations, making it difficult to determine in real time. In this embodiment, static magnetic field inhomogeneity is simply estimated by removing the time factor.

[0058] That is, it is widely known that static magnetic field inhomogeneity Bc due to Maxwell terms generated when a gradient magnetic field is applied can be calculated by the following equation (1).

[0059] [Formula 1]

[0060]

[0061] x, y, z: real-space position [m]

[0062] t: time [ms]

[0063] G x , G y , G z : Inclined magnetic field strength of each axis [mT / m] (1)

[0064] According to formula (1), Bc is a function of the gradient magnetic field strength of each axis that changes with time. In the EPI sequence used in this embodiment, if the direction is expressed in real space (image space), the gradient magnetic field strength applied in the frequency encoding direction and the phase encoding direction can be expressed as follows: Figure 5 The magnitude of the inclination is represented by the inclination shown in (a) and (b). In practice, the intensity of the inclination magnetic field is switched and the echo signal is obtained, but it can be virtually regarded as applying the inclination magnetic field. Figure 5The Bc estimating unit 22 of this embodiment assumes that a fixed gradient magnetic field strength is applied, and uses the equation (2) which removes the time element instead of the above equation (1). This allows for simple calculation of static magnetic field inhomogeneity.

[0065] [Formula 2]

[0066]

[0067] x, y, z: real-space position [m]

[0068] G x , G y , G z : Inclined magnetic field strength of each axis [mT / m] (2)

[0069] For example, in Figure 6 In the vertical magnetic field MRI apparatus shown in FIG, when the slice direction is set to the Y axis, the phase encoding direction is set to the Z axis, and the frequency encoding direction is set to the X axis, Figure 4 In the SE EPI sequence, no gradient magnetic field is applied in the slice direction when the echo signal is acquired, so Gy is 0 mT / m, Gx is the frequency encoding gradient magnetic field intensity, and Gz is the phase encoding gradient magnetic field (Blip) intensity. Figure 4 The applied intensity of one frequency-encoding gradient magnetic field pulse 408 is the frequency-encoding gradient magnetic field intensity. The phase-encoding gradient magnetic field intensity can be set to the applied intensity of one Blip. However, since Blip pulses 406 are usually triangular rather than trapezoidal, the average applied intensity of one phase-encoding gradient magnetic field pulse 406 is used.

[0070] By substituting Gx, Gy, and Gz into equation (2), we can find Bc(x, y, z). Figure 7 An example of the static magnetic field inhomogeneity Bc calculated in this manner is shown. Figure 7 This is a diagram that estimates the static magnetic field inhomogeneity Bc generated in the xz plane at the center position (y=0) in the slice direction. Figure 5 As shown in (a) and (b), the intensity of the frequency encoding gradient magnetic field Gx is greater than the intensity of the phase encoding gradient magnetic field Gy. The four terms (z 2 、(x 2 +y 2 ) / 4, xz, yz) in the first term (Gx 2 z 2 ) is dominant, becoming Figure 7 As shown, the static magnetic field in the z direction is uneven and has a distribution that is a quadratic function of z.

[0071] Next, the distortion calculation unit 23 substitutes Bc estimated by the Bc estimation unit 22 and Gx and Gz substituted into the equation (2) into the next equation (3), thereby estimating the amount of image distortion (image quality degradation estimation process S303).

[0072] [Formula 3]

[0073]

[0074]

[0075] exist Figure 8 The example of the distortion amount when imaging with a FOV of 300 mm at the center of the magnetic field is shown for the slice position y=0 vector. Figure 9 Shown in accordance with Figure 8 An example of how the amount of distortion makes an image distorted. Figure 9 (a) is an undistorted image, and (b) is a distorted image of (a). The vertical direction of the paper is the phase encoding direction (Z direction). In this example, it can be seen that significant distortion occurs in the phase direction (Z axis).

[0076] If the imaging sequence is set for the measurement control unit 111 in the above-mentioned pulse sequence generation process S301, imaging is started under the control of the measurement control unit 111 to collect k-space data (S304), and the image generation unit 21 performs reconstruction processing such as Fourier transform on the acquired k-space data to acquire an image (S305). Distortion caused by the gradient magnetic field is generated in the image, and it can be estimated that the amount of distortion is roughly the amount of distortion calculated in S303. Therefore, the image generation unit 21 uses the estimated distortion amount to deform the reconstructed image in the opposite direction (S306). That is, Figure 8 The inverse vector of the distortion vector shown is used as a matrix, and the image is transformed by common image processing such as affine transformation to correct the distortion. Figure 10 The distorted image shown in (a) is converted to the corrected image shown in (b).

[0077] Through the above steps S301 to S306 , the pulse sequence actually used for imaging can be used to easily correct the distortion of the image caused by the static magnetic field inhomogeneity generated by the gradient magnetic field.

[0078] In addition, Figure 6In the example, the coordinate axes (X, Y, Z) of the device are consistent with the slice axis, frequency encoding axis, and phase encoding axis. However, the pulse sequence is usually mostly produced in the relative coordinate system of the slice axis, phase encoding axis, and frequency encoding axis. There are cases where it is impossible to obtain information of the absolute coordinate system that takes into account the oblique (rotation) and the like. In this case, the following formula (4) can be used to convert the relative coordinate system (x', y', z') to the absolute coordinate system (x, y, z). In formula (4), Rx, Ry, and Rz are matrices representing the oblique. For the gradient magnetic field strength, it can also be calculated based on the expansion in the absolute coordinate system.

[0079] [Formula 4]

[0080]

[0081]

[0082]

[0083]

[0084] θ x ,θ y θ z : Rotation angle relative to each axis (4)

[0085] As described above, according to this embodiment, static magnetic field inhomogeneity is simply estimated based on the gradient magnetic field strength determined by the imaging sequence. This eliminates the need for prior measurement of static magnetic field inhomogeneity associated with gradient magnetic field application and increases the flexibility in setting imaging conditions. Furthermore, since correction can be focused on components that affect the image, the process is simplified and significantly improves image quality.

[0086] <Variation of Embodiment 1>

[0087] In the first embodiment, an example of using the SE EPI sequence as an imaging sequence is described, but the present invention is also applicable to other sequences. Here, a case of using a 2D SE sequence as another sequence is described.

[0088] In a typical 2D SE sequence, such as Figure 11As shown in FIG. 5A , after applying an excitation RF pulse 501 along with a slice-selective gradient magnetic field pulse 503, an inversion RF pulse 502 is applied along with a slice-selective gradient magnetic field pulse 504 to generate an echo signal 508 with a peak at the echo time (TE). A readout gradient magnetic field (frequency-encoding gradient magnetic field pulse) 507 is then applied to collect the echo signal 508. At this time, a dephasing pulse 506 of the frequency-encoding gradient magnetic field pulse 507 is applied between the excitation RF pulse 501 and the inversion RF pulse 502, and a phase-encoding gradient magnetic field pulse 505 is applied to impart phase encoding to the generated echo signal 508. This sequence is repeated during the TR to collect k-space data.

[0089] For such an SE sequence, similarly to the first embodiment, the static magnetic field inhomogeneity Bc during execution of the pulse sequence can be estimated using equation (2) obtained by removing the time element from the gradient magnetic field intensity.

[0090] As an example, for Figure 6 In an MRI system with a vertical magnetic field, if the slice direction is set to the Y axis, the phase encoding direction is set to the Z axis, and the frequency encoding direction is set to the X axis, then when the echo signal is acquired, no gradient magnetic field is applied in the slice direction, so Gy in formula (2) is 0mT / m, and Gx is the intensity (pulse height) of the frequency encoding gradient magnetic field 507. The intensity of the phase encoding gradient magnetic field 505 is different at each repetition, so as Gz, the average value of all repeated phase encoding gradient magnetic fields is substituted. Usually, phase encoding applies the same intensity in positive and negative directions, so the actual Gz is 0mT / m. By substituting them into formula (2), Bc(x, y, z) can be estimated. For the case of tilted imaging, the tilted coordinate system can also be expanded into an absolute coordinate system by using formula (4), so that it can be estimated in the same way as in embodiment 1.

[0091] Similar to the first embodiment, the static magnetic field inhomogeneity estimated in this way is used to calculate the amount of distortion generated in the image by the equation (3), and the reconstructed image is corrected using the calculated amount of distortion.

[0092] The above describes a modification of the SE sequence applied to 2D, but the same is true for 3D. Figure 11 (B) shows an example of a 3D GE sequence. Figure 11 In (B), Figure 11Identical elements to those in (A) are denoted by the same reference numerals. In this GE sequence, readout is performed without using an inversion RF pulse. Instead, the inversion of gradient magnetic fields 506 and 507 generates and measures echo signals 508. Furthermore, as a significant difference from a 2D sequence, a slice encoding gradient magnetic field 513 is applied. Furthermore, gradient magnetic field pulses 514, 516, and 517 are rewind pulses for each axis, respectively.

[0093] When calculating static magnetic field inhomogeneity for such a 3D sequence, the slice gradient magnetic field intensity value in equation (2) can be substituted with the average value of all repeated slice encoding gradient magnetic field intensities, similar to the case of phase encoding gradient magnetic field intensity in an SE sequence. For example, when the slice direction is set to the Y axis, the average value of the slice encoding gradient magnetic field intensity is substituted as Gy. Typically, the encoding amount is the same in both positive and negative directions, so if averaged, it becomes 0 mT / m. For the other axes, the same is true for 2D sequences.

[0094] <Implementation Method 2>

[0095] In the first embodiment, static magnetic field inhomogeneity and image distortion were estimated by simple calculations based on the shape of the pulse sequence, thereby correcting image quality degradation. However, in the present embodiment, the accuracy of the distortion calculated by the distortion calculation unit is improved by utilizing prior measurement results using a phantom or the like.

[0096] The static magnetic field inhomogeneity estimated in Implementation 1 is due to the static magnetic field inhomogeneity imposed by the gradient magnetic field during imaging. However, although it is only slight, there is also inhomogeneity in the original static magnetic field strength. Furthermore, although the gradient magnetic field is designed to be centered around the static magnetic field center, in reality, there is a slight offset from the static magnetic field center. In this implementation, such inhomogeneity and offset are calculated as adjustment values ​​to adjust the distortion amount. Specifically, by adding an adjustment value to eliminate the error between the estimated value and the measured value of the static magnetic field inhomogeneity, equation (2) for calculating the static magnetic field inhomogeneity Bc can be expressed as the following equation (5).

[0097] [Formula 5]

[0098]

[0099] x, y, z: real-space position [m]

[0100] G x , G y , G z : Inclined magnetic field strength of each axis [mT / m]

[0101] C Z2 , CXγ , C xZ , C Yz : Correction cumulative

[0102] Δx, Δy, Δz: position correction values ​​(5)

[0103] In formula (5), C Z2 、C XY 、C XZ 、C YZ is a coefficient (correction coefficient) for adjusting the static magnetic field unevenness, and Δx, Δy, and Δz are correction values ​​(position correction values) for adjusting the magnetic field center position. In this embodiment, the distortion amount adjusted to include these correction coefficients and position correction values ​​is calculated.

[0104] Hereinafter, this embodiment will be described focusing on the points that are different from the first embodiment.

[0105] exist Figure 12 The functions of the arithmetic processing unit 200 of this embodiment are shown. Figure 12 In, with Figure 2 The same elements are denoted by the same reference numerals and repeated descriptions are omitted. As shown in the figure, the operation processing unit 200 includes, in addition to the image generation unit 21, the Bc estimation unit 22 for estimating static magnetic field inhomogeneity, and the distortion calculation unit 23, an adjustment value calculation unit 24 for calculating an adjustment value for adjusting the distortion amount calculated by the distortion calculation unit 23. The distortion amount can be compared with the embodiment 1 ( Figure 3 :S301~S303) are calculated in the same way.

[0106] Below, refer to Figure 13 An example of the processing performed by the adjustment value calculation unit 24 will be described.

[0107] First, before taking an image of the subject, two preliminary images are taken using a phantom or the like (S601, S602). One is to take a reference image (S601) under conditions that are not easily affected by static magnetic field inhomogeneity Bc. For example, a gradient echo sequence or a spin echo sequence can be used for imaging under conditions that are not easily affected by static magnetic field inhomogeneity. If it is an EPI sequence, in order to reduce the influence of distortion, the intensity of the tilted magnetic field of the Blip pulse can also be increased by setting it to multi-excitation to reduce the amount of distortion. In another imaging, an adjustment image is taken under conditions that generate a lot of distortion (S602). For example, a single-excitation EPI sequence is used. Any of these imaging can be performed first or later, but it is preferred that the actual imaging ( Figure 3 For example, it may be performed during adjustment of the MRI apparatus and the adjustment value obtained thereafter may be registered.

[0108] Next, the reference image and the adjustment image are used to perform repeated operations on the adjustment value for calculating the distortion amount (S603 to S606). First, the initial value of the adjustment value for adjusting the static magnetic field inhomogeneity is set (S603). As the initial value of the adjustment value (correction coefficient and position correction value), a value that can be taken as the adjustment value, for example, a value selected from a range that can be predicted empirically, can be used. Next, the distortion of the adjustment image is corrected using the adjustment value (S604). This process is similar to the process of embodiment 1 (S606), except that formula (5) is used instead of formula (2) as the calculation formula for estimating the static magnetic field inhomogeneity for calculating the distortion. Figure 3 : S302 to S304) are the same. That is, based on the pulse sequence used to capture the adjustment image, the gradient magnetic field intensity and the adjustment value set using the initial value are substituted into equation (5) to estimate the static magnetic field inhomogeneity Bc, and the distortion amount is calculated using equation (3). The calculated distortion amount is used to correct the adjustment image.

[0109] Next, the similarity between the corrected adjustment image and the reference image is calculated (S605). The similarity calculation can use common image processing techniques, such as normalized cross-correlation, etc. Then, S603 to S605 are repeated with different adjustment values.

[0110] The repetitive calculation may be repeated so that the range of values ​​that can be taken as the adjustment value is comprehensively changed, but Figure 13 In the example shown, a search range for adjustment values ​​is predetermined. After performing steps S603 to S605 using adjustment values ​​within the given search range, the search range is updated (S606) and the calculation is repeated. For example, a range update method can first distribute the adjustment values ​​over a wide range, mark adjustment values ​​indicating high similarity, and then, in the next step, narrow the search range to the vicinity of these adjustment values ​​and perform the search. By narrowing the range and repeating this search, the optimal adjustment value can be efficiently derived.

[0111] Finally, the adjustment value with the highest similarity is registered in the storage device 220 of the MRI apparatus (S607). The registered adjustment value is used when the Bc estimating unit 22 calculates static magnetic field inhomogeneity.

[0112] After registering the adjustment value through such preliminary measurement, the subject is imaged. The imaging process is the same as that of the first embodiment and its modified example. Figure 3 The illustrated process (S301 to S306) performs static magnetic field inhomogeneity estimation, imaging, image reconstruction, and image correction. However, in this embodiment, the Bc estimation unit 22 substitutes the adjustment value registered by prior measurement and the gradient magnetic field intensity acquired from the imaging sequence into equation (5) to perform static magnetic field inhomogeneity estimation processing (S302).

[0113] According to this embodiment, the error of each device in static magnetic field nonuniformity caused by application of a gradient magnetic field, the static magnetic field nonuniformity adjusted for the offset of the gradient magnetic field from the static magnetic field center, and the image distortion amount can be obtained, thereby enabling more accurate correction of image quality degradation.

[0114] <Variation 1 of Implementation Example 2>

[0115] In the second embodiment, as an example of the process of calculating the adjustment value by the adjustment value calculation unit 24, an example of calculating the adjustment value by repeated calculation is described ( Figure 13 ), but as another example, a method of performing inverse calculations on the adjustment value may be used without using the property of inhomogeneous magnetic field to perform repeated calculations.

[0116] like Figure 7 As shown, when the phase encoding direction is set to the Z-axis direction, the quadratic function distribution that takes the magnetic field center (z = 0) as the minimum value in the magnetic field inhomogeneity is dominant. In other words, the error between the position where the image distortion is least and the position of the image center (magnetic field center) is Δz. In addition, the distribution in the Z-axis direction is dominant, so by making the coefficient C of the first term on the right side of the formula (5) Z 2Optimization to improve calibration accuracy is obvious.

[0117] In this modified example, this fact is exploited and the error Δz in the image center position is used to simply calculate the correction coefficient for this term based on the first term on the right side of equation (5). The error Δz in the image center position can be calculated as the offset between the center position of the reference image and the center of the adjustment image whose distortion has been corrected with Δz = 0. The magnetic field inhomogeneity Bc is calculated by multiplying this error Δz by the gradient magnetic field strength of Gz (equation (3)). Next, in equation (5), C is replaced by Z2 The correction coefficients other than 0 are set, and the magnetic field inhomogeneity calculated as above is substituted into formula (5) to obtain C Z2 Perform the inverse operation.

[0118] According to this method, the adjustment value of the Z-axis direction with the greatest influence can be easily obtained without repeated calculations. In addition, the adjustment value obtained by this method can be set as the initial value and the adjustment value can be adjusted. Figure 13 Repeated operations to search for the optimal value.

[0119] <Variation 2 of Implementation Example 2>

[0120] In the second embodiment, the adjustment value (correction coefficient) for static magnetic field inhomogeneity is calculated when the slice center is the magnetic field center. However, different values ​​may be used depending on the spatial position.

[0121] Normally, the image center (slice center) is set as the magnetic field center and the image is taken. However, when taking a relatively large image along the slice direction, the slice position may be offset from the magnetic field center. The dominant magnetic field inhomogeneity in the Z-axis direction is as follows: Figure 7 As shown in , it becomes a quadratic function distribution, but if the position in the slice direction (Y direction) changes, it becomes Figure 14 As schematically shown, there is a possibility that the distribution may change due to the influence of errors in static magnetic field and gradient magnetic field output.

[0122] In this modification, when the adjustment value is obtained by pre-shooting ( Figure 13 (Steps S601 and S602) measure magnetic field inhomogeneities over a wide range of the magnetic field space by capturing images at slice positions offset from the magnetic field center. Adjustment values ​​specific to each slice position are calculated and registered. By applying adjustment values ​​specific to each slice position, the positional dependency of the correction effect can be reduced.

[0123] Furthermore, as another variation, conditions such as the reception bandwidth (BW) used when capturing reference and adjustment images can be changed, the distortion amount can be calculated at different gradient magnetic field intensities, and the adjustment value can be set as a function corresponding to the gradient magnetic field intensity. This approach eliminates the effects of localized inhomogeneities in the static magnetic field space and nonlinearities in the gradient magnetic field output, achieving high correction accuracy.

[0124] <Implementation Method 3>

[0125] In embodiments 1 and 2, image distortion is corrected as post-processing using estimated static magnetic field inhomogeneities. However, in this embodiment, the output of a shim coil (compensation magnetic field generator) that corrects the estimated static magnetic field inhomogeneities is calculated, and a compensation magnetic field is output from the shim coil in advance, thereby reducing image quality degradation.

[0126] The static magnetic field inhomogeneity correction associated with gradient magnetic field application in this embodiment does not control the compensation magnetic field in real time as in conventional techniques, but rather outputs a constant compensation magnetic field in advance. Therefore, while the versatility and accuracy of image quality degradation are reduced compared to real-time correction, image distortion can be easily corrected, and the gradient magnetic field coil does not require high responsiveness. This makes it possible to easily achieve excellent distortion reduction in MRI systems, such as low-magnetic-field machines, where static magnetic field inhomogeneity caused by gradient magnetic field application significantly affects image quality.

[0127] exist Figure 15 The functions of the arithmetic processing unit 200 in this embodiment are shown. Figure 15 In, with Figure 2、 Figure 12 The same elements are denoted by the same reference numerals, and repeated descriptions are omitted. Figure 15 As shown, the calculation processing unit 200 includes a compensation output calculation unit 25 in addition to the image generation unit 21 and the Bc estimation unit 22. The compensation output calculation unit 25 calculates the output value of a constant compensation current flowing through the shim coil 109 based on the static magnetic field inhomogeneity estimated by the Bc estimation unit 22. The measurement control unit 111 ( Figure 1 ) controls the supply of a compensation current corresponding to the calculated output value to the shim power supply 110.

[0128] exist Figure 16 The processing flow of this embodiment is shown in FIG. Figure 16 In, with Figure 3 Identical processes are denoted by the same reference numerals, and repeated descriptions are omitted. Here, it is assumed that conventional active shimming is used to achieve static magnetic field uniformity. Active shimming is a method of improving static magnetic field uniformity by measuring the uniformity of the static magnetic field in advance and applying a compensation current to the shim coil 109. However, the simultaneous use of active shimming is not essential.

[0129] like Figure 16 As shown in the flowchart of FIG. 3 , in processes S301 and S302 , similarly to the first and second embodiments, static magnetic field inhomogeneity Bc is estimated based on the pulse sequence shape.

[0130] Next, in S313, the compensation output calculation unit 25 calculates the output of the compensation current for generating the compensation magnetic field that cancels the static magnetic field inhomogeneity Bc. The compensation magnetic field that cancels the static magnetic field inhomogeneity Bc is, for example, positive and negative. Figure 7 The output of the quadratic function gradient magnetic field with the opposite quadratic function distribution can be calculated based on the relationship between the magnetic field generated by the shim coil and the output. The calculated output (Bc compensation amount) is added to the compensation current calculated by active shimming.

[0131] As a result, static magnetic field inhomogeneity Bc generated by gradient magnetic field application during echo signal acquisition is corrected, and thus a distortion-free image can be obtained.

[0132] This embodiment corrects static magnetic field uniformity, including static magnetic field inhomogeneity Bc caused by the application of a gradient magnetic field during imaging, which cannot be addressed by conventional active shimming methods that rely on prior measurements. This provides a high degree of image degradation prevention for imaging, where static magnetic field inhomogeneity caused by the application of a gradient magnetic field cannot be ignored. Furthermore, the output for eliminating static magnetic field inhomogeneity is applied as a fixed-value compensation current, eliminating the need for switching shim coils and the requirement for high responsiveness. Furthermore, there is no concern about eddy current generation (error magnetic field generation) associated with switching.

[0133] Furthermore, in the MRI apparatus of this embodiment, a Bc compensation amount is added in addition to the original compensation current. Therefore, at times when frequency-encoding gradient magnetic fields or phase-encoding gradient magnetic fields are not applied, such as during slice selection gradient magnetic field application, there is a possibility that static magnetic field uniformity may be adversely affected. However, if the compensation current can be changed at a timing that allows the shim coil to respond in time, this problem can be avoided by changing the compensation current amount during pulse sequence execution. For example, when using frequency-selective excitation for fat suppression, the Bc amount may not be applied, and energization may begin after the fat suppression pulse is applied.

[0134] <Implementation Method 4>

[0135] In the third embodiment, the compensation current to be additionally energized to the shim coil is determined based on the static magnetic field inhomogeneity Bc estimated from the pulse sequence shape. However, in the present embodiment, similar to the second embodiment, equation (5) is used instead of equation (2) as the equation for estimating the static magnetic field inhomogeneity Bc, and its adjustment values ​​(correction coefficients and position correction values) are calculated in advance using a phantom, thereby improving the accuracy of image quality improvement.

[0136] In addition to the arithmetic processing unit 200 of this embodiment Figure 15 The calculation processing unit 200 of the second embodiment shown in FIG. 1 is provided with an adjustment value calculation unit. Figure 15 The structure is the same as that of FIG, and the illustration is omitted. In addition, even in this embodiment, the outline of the processing flow is the same as that of FIG. Figure 16 Same, therefore, the following, according to Figure 17 The flowchart shown in FIG. 1 illustrates an example of calculation of the adjustment value. In the following description, the flowchart used in the description of the second and third embodiments is appropriately cited ( Figure 13 as well as Figure 15 ) for explanation.

[0137] First, in S701, with Figure 13Similarly to S601, a reference image without distortion is captured using a phantom. Next, repeated operations are performed to search for appropriate adjustment values ​​(S702 to S705). To this end, an initial value for the adjustment value is first set (S702), and the static magnetic field inhomogeneity Bc caused by the gradient magnetic field pulse is calculated using this adjustment value and formula (5). The gradient magnetic field intensities Gx, Gy, and Gz used in formula (5) can be, for example, the gradient magnetic field intensities used in the imaging sequence used in the following S703.

[0138] The output values ​​of the shim coils used to correct the calculated static magnetic field inhomogeneity Bc are calculated, the compensation currents of the shim coils are changed, and an adjustment image is captured using the same phantom as the reference image (S703). The similarity between the obtained adjustment image and the reference image acquired in S701 is calculated (S704). The method for calculating the similarity is the same as S605 in Implementation 2. If the similarity is high, the adjustment value set in S702 is appropriate, and the search range for the adjustment value is changed (S705). S702 to S704 are repeated until a high similarity is achieved. The method for determining the adjustment value and setting the search range is the same as in Implementation 2, and a method of narrowing the search range from a large range can be appropriately adopted.

[0139] Finally, the adjustment value indicating the highest similarity is registered (S706). These processes can be performed independently of the actual imaging. Figure 16 Each process (S301 to S305) is performed as described in the flow.

[0140] In this embodiment, as in Embodiment 2, the adjustment value may be set to a different value depending on the spatial position (slice position). Furthermore, during the acquisition of the reference image and the adjustment value image, the compensation output may be calculated using different gradient magnetic field intensities by changing conditions such as the reception bandwidth BW, with the adjustment value being a function corresponding to the gradient magnetic field intensity.

[0141] According to this embodiment, the accuracy of the compensation output can be further improved, and an image can be obtained in which image distortion caused by the gradient magnetic field is eliminated.

[0142] While several embodiments and variations of the MRI apparatus of the present invention have been described above, the present invention is not limited to these embodiments and variations. These embodiments and variations may be combined as long as they do not conflict technically. Furthermore, known correction units, etc., may be added. Furthermore, some elements of the structures described in the embodiments and variations may be omitted.

Claims

1. A magnetic resonance imaging device, characterized in that have: A transmitting unit irradiates a subject placed in a static magnetic field space with a high-frequency pulse; a receiving unit receives a nuclear magnetic resonance signal generated from the subject by the irradiation of the high-frequency pulse; a gradient magnetic field generating unit generates a gradient magnetic field in the static magnetic field space; a control unit controls the transmitting unit, the gradient magnetic field generating unit, and the receiving unit according to a given pulse sequence; an image generating unit reconstructs an image of the subject based on the nuclear magnetic resonance signal; and a static magnetic field inhomogeneity estimating unit estimates the static magnetic field inhomogeneity generated by the application of the gradient magnetic field according to the given pulse sequence. The static magnetic field inhomogeneity estimation unit uses the applied intensity or average applied intensity of the gradient magnetic field pulse determined by the given pulse sequence to simply estimate the static magnetic field inhomogeneity by removing the time factor, calculates the error generated in the simple estimation of the static magnetic field inhomogeneity by prior imaging, registers it as an adjustment value, and uses the adjustment value to improve the estimation accuracy of the static magnetic field inhomogeneity.

2. The magnetic resonance imaging apparatus according to claim 1, wherein The gradient magnetic field includes gradient magnetic fields in the slice selection direction, the phase encoding direction, and the readout direction. The static magnetic field inhomogeneity estimating unit estimates the static magnetic field inhomogeneity using the applied intensity or average applied intensity of the phase encoding gradient magnetic field pulse and the readout gradient magnetic field.

3. The magnetic resonance imaging apparatus according to claim 2, wherein: The static magnetic field inhomogeneity estimating unit estimates the static magnetic field inhomogeneity based on a spatial position.

4. The magnetic resonance imaging apparatus according to claim 1, wherein The magnetic resonance imaging apparatus further includes a storage unit in which an adjustment value for correcting the static magnetic field inhomogeneity simply estimated by the static magnetic field inhomogeneity estimating unit is pre-registered. The static magnetic field inhomogeneity estimating unit adjusts the estimated static magnetic field inhomogeneity based on the adjustment value registered in the storage unit.

5. The magnetic resonance imaging apparatus according to claim 4, wherein The adjustment value includes a correction coefficient for adjusting the intensity of the static magnetic field non-uniformity and a position correction value for adjusting the center position of the static magnetic field.

6. The magnetic resonance imaging apparatus according to claim 4, wherein: The magnetic resonance imaging apparatus further includes an adjustment value calculation unit that calculates the adjustment value by repeatedly performing operations to maximize a similarity between an image obtained by correcting a previously measured image including an influence of the static magnetic field inhomogeneity using the static magnetic field inhomogeneity simply estimated by the static magnetic field inhomogeneity estimation unit and a reference image not including an influence of the static magnetic field inhomogeneity.

7. The magnetic resonance imaging apparatus according to claim 1, wherein The magnetic resonance imaging apparatus further includes a distortion calculation unit configured to calculate distortion generated in the image based on the static magnetic field inhomogeneity estimated by the static magnetic field inhomogeneity estimation unit. The image generation unit corrects and reconstructs the image based on the distortion calculated by the distortion calculation unit.

8. The magnetic resonance imaging apparatus according to claim 1, wherein The magnetic resonance imaging apparatus further comprises: a compensation magnetic field generating unit that generates a magnetic field having a spatial distribution that complies with a spherical harmonic function of order zero or higher; and a compensation output calculation unit that calculates a compensation current to be supplied to the compensation magnetic field generation unit based on the static magnetic field inhomogeneity estimated by the static magnetic field inhomogeneity estimation unit, The control unit supplies the calculated compensation current to the compensation magnetic field generating unit in advance.

9. The magnetic resonance imaging apparatus according to claim 8, wherein The control unit finely adjusts the compensation current supplied to the compensation magnetic field generating unit using an adjustment value calculated based on a distortion amount of an image measured in advance.

10. An image correction method for correcting a reconstructed image captured by a magnetic resonance imaging device, the image correction method comprising: A step of simply estimating static magnetic field inhomogeneity by removing the time factor based on a pulse sequence used for imaging using the application intensity or average application intensity of the gradient magnetic field pulse determined by the pulse sequence; a step of calculating an amount of distortion generated in a reconstructed image based on the estimated static magnetic field inhomogeneity; a step of correcting the reconstructed image using the calculated distortion amount; a step of calculating an error generated in simply estimating the static magnetic field inhomogeneity by taking an image in advance, and registering the error as an adjustment value; as well as The step of estimating the static magnetic field inhomogeneity includes using the adjustment value to improve the estimation accuracy of the static magnetic field inhomogeneity.

11. A static magnetic field inhomogeneity correction method, using shim coils to correct static magnetic field inhomogeneity in a magnetic resonance imaging device, the static magnetic field inhomogeneity correction method comprising: A step of simply estimating the static magnetic field inhomogeneity by removing the time factor based on a pulse sequence used for imaging using the application intensity or average application intensity of the gradient magnetic field pulse determined by the pulse sequence; The step of calculating a compensation current value supplied to the shim coil according to the estimated non-uniform static magnetic field; The step of energizing the shim coil with a current having the calculated compensation current value; a step of calculating an error generated in simply estimating the static magnetic field inhomogeneity by taking an image in advance, and registering the error as an adjustment value; as well as The step of estimating the static magnetic field inhomogeneity includes using the adjustment value to improve the estimation accuracy of the static magnetic field inhomogeneity.

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