JP5449903B2 - Magnetic resonance imaging system - Google Patents

Description

  The present invention relates to a magnetic resonance imaging technique. In particular, the present invention relates to a method for estimating an object parameter by calculation.

  A magnetic resonance imaging (MRI) apparatus is a medical diagnostic imaging apparatus that causes nuclear magnetic resonance to occur in hydrogen nuclei in an arbitrary plane that crosses a subject, and that takes a tomographic image in the plane from the generated nuclear magnetic resonance signal. . In general, a magnetic resonance signal (echo) is generated when an excitation pulse is applied to excite the magnetization in the plane at the same time as a slice gradient magnetic field that specifies the imaging plane is applied, and the excited magnetization converges. Get. In order to give positional information to the magnetization, a phase encoding gradient magnetic field and a readout gradient magnetic field in directions perpendicular to each other in the tomographic plane are applied between the excitation and the echo. The measured echoes are arranged in k-space with the horizontal axis being kx and the vertical axis being ky, and image reconstruction is performed by inverse Fourier transform.

  The high-frequency magnetic field pulse and the gradient magnetic field pulse for generating the echo are applied based on a preset imaging sequence. Various imaging sequences are known depending on the purpose. For example, the spin echo (SE) method and the gradient echo (GE) method measure the echo by repeating the imaging sequence in a short time. By measuring one echo per excitation of magnetization and changing the intensity of the phase encoding gradient magnetic field pulse for each excitation, this method measures the number of echoes necessary to obtain one tomographic image. is there.

  The pixel value of the MRI image is determined by this imaging sequence, imaging parameters including the field of view and repetition time of the imaging sequence, and object parameters such as the magnetization density and relaxation time of the object.

For example, the pixel value of FLASH, which is one of the GE systems (here, the transverse magnetization intensity Mxy) is expressed by the following equation as a solution of Bloch's equation, which is a basic equation of the magnetic resonance phenomenon.
(Formula 1)

Here, TR and α are imaging parameters, which are repetition time and flip angle, respectively. T1 is the longitudinal relaxation time of the subject parameter. Further, in the SE method, when TR is sufficiently longer than T1, it is expressed by the following formula.
(Formula 2)

Here, TE is the echo time of the imaging parameter, and T2 is the longitudinal relaxation time of the subject parameter.

As one of the MRI imaging methods, there is a method in which a plurality of images are acquired using different imaging parameters and subject parameters are calculated for each pixel. An image having the obtained object parameter as a pixel value is called a calculation image or a map.
For example, when obtaining T2 of the subject parameter using SE, the following is performed. Two images, images A and B, are taken with an SE with TR sufficiently longer than T1 and TE set to 20 ms and 40 ms, respectively. For each pixel, the value of image A and the value of image B are fitted to Equation 2 using the least square method or the like to obtain T2.

  In addition to this, the pixel value of the MRI image depends on apparatus parameters such as a static magnetic field distribution, a high-frequency magnetic field (B1) distribution, and a sensitivity distribution of the receiving coil. For example, if the high-frequency magnetic field distribution is not uniform, the flip angle α is not spatially constant. Therefore, in FLASH, the pixel value is also spatially changed from Equation 1. This distribution is called a B1 map, and is generally obtained by using a DAM (double angle method, Non-Patent Document 1) that calculates from two images taken with the flip angle doubled.

  On the other hand, there is a method of creating an image by numerical simulation (Patent Document 1; Patent 3404191). An object model in which spins are arranged on lattice points, an imaging sequence, imaging parameters, and apparatus parameters are input, and Bloch's equation, which is a basic equation of the magnetic resonance phenomenon, is solved and a magnetic resonance signal is output. An image under a given condition can be obtained by reconstructing an image of the magnetic resonance signal.

Patent 3340419 Cunningham CH, Pauly JM, Nayak KS, Saturated Double-Angle Method for Rapid B1 + Mapping. Magn. Reson. In Med. 2006; 55: 1326-1333.

  In the above method, since the relationship (signal function) between the imaging parameter and the subject parameter or the device parameter and the pixel value needs to be analytically determined, the applicable imaging sequence is limited, and the imaging time and image quality are limited. There are big restrictions. The image obtained by the above-described method is an image in which the subject parameter itself is a pixel value and is suitable for use in quantitative analysis. However, an image used for diagnosis by a doctor is not necessarily suitable because the contrast is different from an image used for normal diagnosis.

  The present invention has been made in view of the above circumstances, and it is suitable for a technique for obtaining a map without restriction even in an imaging sequence in which a signal function is not analytically obtained, and the obtained subject parameters are suitable for diagnosis. Providing technology for image contrast.

According to the present invention, even in an imaging sequence in which a signal function is not analytically obtained, the object parameter and the apparatus parameter are estimated by obtaining the relationship by numerical simulation. Further, an image is created by performing a numerical simulation on the obtained object parameter or apparatus parameter using an arbitrary imaging sequence and imaging parameters.

In particular,
An imaging means for applying a high-frequency magnetic field and a gradient magnetic field to a subject placed in a static magnetic field to detect a magnetic resonance signal generated from the subject, and an operation for processing the magnetic resonance signal detected by the imaging means A magnetic resonance imaging apparatus comprising: means; and control means for controlling the imaging means and the computing means,
The shooting means takes multiple images with different shooting parameters,
The computing means estimates subject parameters or device parameters from the plurality of images,
The calculation means provides a magnetic resonance imaging apparatus characterized in that an image is created by numerical simulation using the estimated object parameter or apparatus parameter, imaging sequence, and imaging parameter as inputs.

According to the present invention, since there is no restriction on the imaging sequence that can be used, a high-speed imaging method can be used (imaging time can be shortened).
In addition, since an imaging sequence and imaging parameters for obtaining an optimal image contrast for diagnosis can be selected at the time of diagnosis rather than at the time of imaging, an improvement in imaging throughput and an improvement in diagnostic ability can be expected.

<< First Embodiment >>
Hereinafter, a first embodiment to which the present invention is applied will be described. Hereinafter, in all the drawings for explaining the embodiments of the present invention, those having the same function are denoted by the same reference numerals, and repeated explanation thereof is omitted.

First, the MRI apparatus of this embodiment will be described. FIG. 1 is a block diagram showing a schematic configuration of the MRI apparatus 100 of the present embodiment. The MRI apparatus 100 irradiates a high-frequency magnetic field and a nuclear magnetism with a magnet 101 that generates a static magnetic field, a gradient coil 102 that generates a gradient magnetic field, a sequencer 104, a gradient magnetic field power source 105, a high-frequency magnetic field generator 106, and the like. A transmission / reception coil 107 that detects a resonance signal, a receiver 108, a calculator 109, a display 110, and a storage medium 111 are provided. Although a single transmission / reception coil 107 is shown in the figure, a transmission coil and a reception coil may be provided separately.
A subject (for example, a living body) 103 is placed on a bed (table) in a static magnetic field space generated by a magnet 101. The sequencer 104 sends commands to the gradient magnetic field power source 105 and the high frequency magnetic field generator 106 to generate a gradient magnetic field and a high frequency magnetic field, respectively. The high frequency magnetic field is applied to the subject 103 through the transmission / reception coil 107. A nuclear magnetic resonance signal generated from the subject 103 is received by the transmission / reception coil 107 and detected by the receiver 108. The sequencer 104 sets a nuclear magnetic resonance frequency (detection reference frequency f0) as a reference for detection. The detected signal is sent to the computer 109, where signal processing such as image reconstruction is performed. The result is displayed on the display 110. The detected signal and measurement conditions can be stored in the storage medium 111 as necessary.

  The sequencer 104 normally performs control so that each device operates at a timing and intensity programmed in advance. Among the programs, a program that particularly describes a high-frequency magnetic field, a gradient magnetic field, and the timing and intensity of signal reception is called a pulse sequence (imaging sequence). In the MRI apparatus 100 of the present embodiment, any pulse sequence can be used. In the present embodiment, a case where a GE type RF-spoiled GRASS sequence shown in FIG. 1 is provided as an example will be described.

  The computer 109 instructs the sequencer 104 to measure a nuclear magnetic resonance signal (echo) according to the RF-SPOILED GRASS sequence, and an echo measurement unit that arranges the measured echo in the k space, and an echo arranged in the k space. An image reconstruction unit that reconstructs an image from the image, and a parameter estimation unit that estimates an object parameter or a device parameter from the reconstructed image and a signal function. Further, prior to measurement, a signal function calculation unit is provided that performs numerical simulation according to the RF-SPOILED GRASS sequence to create an image and calculates a signal function. Each of these functions is realized when the CPU of the computer 109 loads the program stored in the storage medium 111 to the memory and executes it. Alternatively, the program of the signal function calculation unit may be executed by a computer other than the computer 109, and the obtained signal function may be stored in the storage medium 111.

  First, a pulse sequence (hereinafter referred to as an RF-SPOILED GRASS sequence) used when the echo measurement unit performs RF-SPOILED GRASS imaging will be described. FIG. 2 is a sequence diagram of the RF-SPOILED GRASS sequence. In the figure, RF, Gs, Gp, and Gr represent a high-frequency magnetic field, a slice gradient magnetic field, a phase encoding gradient magnetic field, and a readout gradient magnetic field, respectively.

  Along with the application of the slice gradient magnetic field pulse 201, a high frequency magnetic field (RF) pulse 202 is irradiated to excite magnetization of a slice in the target object. Next, after applying a slice rephase gradient magnetic field pulse 203, a phase encode gradient magnetic field pulse 204 for adding position information in the phase encode direction to the phase of magnetization, and a dephase readout gradient magnetic field 205, position information in the readout direction A magnetic resonance signal (echo) 207 is measured while applying a readout gradient magnetic field pulse 206 for adding. Finally, a phase encoding gradient magnetic field pulse 209 for dephase is applied. While changing the intensity of the phase encoding gradient magnetic field pulses 204 and 209 (phase encoding amount kp) and changing the increment value of the RF pulse phase by 117 degrees (the phase of the nth RF pulse is θ (n ) = θ (n-1) + 117n) Repeatedly at a repetition time TR, and an echo necessary for obtaining one image is measured. Each echo is arranged in the k space as shown in (b), and an image is reconstructed by two-dimensional inverse Fourier transform. This pulse sequence has a feature that an image in which T1 (longitudinal relaxation time) is emphasized can be obtained.

The flow of processing is shown in FIG.
First, a signal function 602 is created 601 by numerical simulation. The shooting parameters that can be changed with RF-spoiled GRASS are FA (flip angle), TR (repetition time), TE (echo time), and θ (RF phase increment value). Of these, the RF phase increment value is fast. Generally, it is fixed at 117 degrees so that an image contrast with less T2 dependency equivalent to FLASH, which is one of the photographing methods, can be obtained. When this θ is changed, the T2 dependence of the image contrast changes greatly.
The signal function fs of RF-SPOILED GRASS is expressed as follows.
(Formula 3)

Here, T1, T2, Cs, and ρ are the longitudinal relaxation time, the lateral relaxation time, the chemical shift, and the spin density of the subject parameter, respectively. B0, B1, and Sc are the apparatus parameter magnetic field intensity, transmission coil irradiation intensity, and reception coil sensitivity, respectively. Since B1 is a factor of FA at the time of shooting, it is in the form of a product with FA.

  The imaging parameters are comprehensively changed with respect to arbitrary values of the subject parameters T1, T2, and Cs, a signal is created by numerical simulation, and a signal function is created by interpolation. The spin density ρ, B1, and Sc to be imaged are set constant (for example, 1). In addition, B0 is set to be the same as the magnetic field strength (for example, 3T) of the apparatus used for photographing.

In numerical simulation, an object model in which spins are arranged on lattice points, an imaging sequence, imaging parameters, and apparatus parameters are input, and Bloch's equation, which is a basic equation of the magnetic resonance phenomenon, is solved and a magnetic resonance signal is output. The object model is given as the spatial distribution of spins (γ, M0, T1, T2, Cs). Here, γ is the gyromagnetic ratio, M0 is the thermal equilibrium magnetization (spin density), and T1 and T2 are the longitudinal relaxation time and the transverse relaxation time, respectively. An image under a given condition can be obtained by reconstructing an image of the magnetic resonance signal.
Bloch's equation is a first-order linear ordinary differential equation and is expressed by the following equation.
(Formula 4)

  Here, (x, y, z) represents a three-dimensional orthogonal coordinate system, and z is equal to the direction of the static magnetic field (intensity is B0). Further, (Mx, My, Mz) is spin, Gx, Gy, Gz are gradient magnetic field strengths in the subscript direction, H1 is high-frequency magnetic field strength, and f0 is the frequency of the rotating coordinate system.

Next, a plurality of images are captured by changing the imaging parameters FA, TR, TE, and θ 603, and the object parameter and device parameter 605 are estimated 604 by fitting the signal value I for each pixel to fs 604. However, since it is difficult to control the device parameters freely at the time of shooting, and it cannot be separated only by changing the shooting parameters except for B1, fitting is performed for f with variable conversion as follows: . Thus, the subject parameters T1, T2 and the device parameter B1, and Δf0 and a, which are products of the subject parameter and the device parameter, are estimated. For the function fitting, for example, a least square method can be used.
(Formula 5)

  Finally, an image is created 607 using the subject parameter and the device parameter. A signal function fs is used to create an image. The object parameter is usually left as it is, and the apparatus parameter is changed to a target value to obtain a signal from fs. For example, when an image with a uniform B1 is created, fs is calculated with B1 = 1. Thereby, an image without B1 nonuniformity can be obtained. When it is desired to create images of different magnetic field strengths or TR and TE, signals are calculated by changing B0, TR and TE, respectively. Thereby, an image with an optimal image contrast can be obtained.

  However, at this time, if fs for the given device parameter is not created in 601, it is necessary to perform numerical simulation for that parameter and add fs.

  As described above, since the signal function is configured by numerical simulation, it is possible to estimate the subject parameter and the apparatus parameter even for the imaging sequence in which the signal function is not analytically obtained. In addition, since an image with apparatus parameters different from the measurement can be created using a signal function, a high-quality and high-contrast image can be obtained, and the diagnostic ability can be improved.

As an example, we show the results of actually estimating the subject parameter T1 and the instrument parameter B1 by the above-mentioned method using a 1.5T MRI apparatus using RF-SPOILED GRASS. First, the image used to construct the signal function Create using numerical simulation. As shown in Fig. 4, the analyte model is a model in which the spin density and chemical shift distributed in T1 and T2 are both uniform (ρ = 1, Cs = 0 ppm). The distribution ranges of T1 and T2 are approximately 50--3000 ms and 30--1500 ms, respectively, which are roughly equivalent to humans. Since the distribution range is wide, the change rate with respect to the position is made exponential to reduce the rate of change for small T1 and T2. As a result, a decrease in the accuracy of the signal can be suppressed. The Hanning window is applied to the uniform spin density distribution. Thereby, truncation artifact can be suppressed. In order to ensure sufficient calculation accuracy, 256 spins / pixel are arranged in the lead-out direction (800 spins / pixel when the FA is 10 degrees or less), and 4 spins / pixel are arranged in the phase encoding direction. If the number of spins is less than this, the calculation accuracy is insufficient, and striped artifacts that are not actually seen in the actual machine occur.

Using this object model as a subject of imaging, a total of 45 images are created by changing FA and TR as follows, and a signal function is constructed by cubic function interpolation using the signal values of these 45 images. .
FA 1, 3, 5, 10, 15, 30, 40, 50, 60 degrees
TR 10, 15, 20, 25, 30 ms
As for other imaging parameters, θ is fixed at 117 degrees, TE is fixed at 5 ms, the number of pixels is 64x64, and RF Prep is 500. Here, RF Prep is the number of repetitions of executing the imaging sequence without measuring the signal before signal measurement, and is necessary for making the spin almost in a steady state. In RF-SPOILED GRASS, if this value is changed, the image contrast also changes, so it is necessary to keep the same value when creating the signal function and when shooting the image for parameter estimation. Artifacts occur when the length of the object is long. If the object model is 500, no artifacts will occur.

The apparatus parameters are such that the magnetic field strength is uniform and B0 = 1.5 T.
Signal function created as described above (FIG. 10)

FIG. 5 shows the strength of the. For the five types of T1 and T2 combinations, the horizontal and vertical axes are displayed as FA and TR, respectively.
The object parameter and the apparatus parameter are estimated using this signal function. Here, an estimation result when an image is captured using a phantom whose T1 and T2 are known is shown. Phantoms are five types of nickel chloride aqueous solutions with different concentrations as shown in Fig. 6a. The respective T1, T2 values are shown in FIG. 6b.

The shooting parameters are as follows. A total of 12 images were shot with varying FA and TR.

FA 15, 25, 35, 45 degrees
TR 10, 20, 30 ms
θ 117 degrees, TE 5 ms, Number of pixels 128x128, RF Prep 500

The phantom T1, B1, and a are estimated by function fitting using this image. In this case, the fitting function is obtained from Equation 5 (FIG. 11-1).

Thus, function fitting is performed by the least square method of relative error expressed by the following equation (FIG. 11-2).

Here, χ is the sum of the residuals of the signal function and the phantom pixel value, and I is the pixel value in FA and TR.

  FIG. 7b shows the estimation result on one line passing through the vicinity of the center of the phantom (FIG. 7a). In FIG. 7b, the horizontal axis is the position, and the vertical axis is the value of each parameter. FIG. 7c shows the average value of the estimated values of T1 and T2 together with the phantom T1 and T2 values, with the estimated value on the left and the phantom on the right for each nickel chloride concentration.

  From FIG. 7b, the value of a tends to decrease the center of the visual field (15 mM). a is the product of spin density and receiver coil sensitivity. This is because the spin densities of the solutions of each concentration are almost equal, and the sensitivity of the receiving coil tends to become smaller at the center. In contrast to Ba, B1 has a center of field of view of almost 1 and is smaller in the periphery. This is qualitatively consistent with the characteristics of the transmission coil. From Fig. 7b and Fig. 7c, it can be seen that the estimated value of T1 is almost similar to the T1 value of the phantom.

  In this example, since the estimated value of T2 is also calculated, FIG. 7bc also shows the result. From FIG. 7bc, an almost constant estimated value of T2 is obtained regardless of the concentration. This indicates that the RF-SPOILED GRASS having a constant θ of 117 degrees is not correctly estimated because the pixel value of the image hardly depends on T2. If the value of θ is also changed in the same way as TR and FA, the pixel value becomes dependent on T2, so that T2 can be estimated correctly at the same time.

Finally, an image when B1 is uniform is created using T1 and T2 in FIG. 7b. Here, the case where FA is 45 degrees and TR is 30 ms is created. In this case, the pixel value of the image to be created can be obtained from the following equation (FIG. 11-3) in which each parameter is substituted into Equation 3.

  Fig. 8 shows (a) an image when the result of Fig. 7b is put into a, B1, T1, T2 of this equation, (b) an image when B1 is uniform (B1 = 1), (c) both (D) shows the original photographed image. FIG. 8a is almost the same image as FIG. 8d, indicating that the estimation of this method is correct. Further, in FIG. 8b, compared to FIG. 8a, the signal on the short left phantom of T1 is strong, and the signal on the right phantom with long T1 is weak. This reflects that B1 at both ends of the field of view in FIG. 8b is larger than in FIG. 8a, and the FA is accordingly increased. Thus, it can be seen that an image obtained by removing the influence of nonuniformity on B1 can be easily obtained.

  Next, a method for handling a diffusion coefficient as an object parameter will be described. Also in this case, as described above, a signal function is constructed by numerical simulation, a plurality of images are photographed by changing the b value which is a photographing parameter for determining the diffusion enhancement degree, and a diffusion coefficient is estimated by function fitting. Images with different b values are created from the signal function using the estimated diffusion coefficient.

  Here, a DWEPI sequence will be described as an example of the diffusion weighted imaging method. FIG. 9 is a sequence diagram of the DWEPI sequence. In the figure, RF, Gs, Gp, and Gr represent axes of a high-frequency magnetic field, a slice gradient magnetic field, a phase encoding gradient magnetic field, and a readout gradient magnetic field, respectively.

  First, a high-frequency magnetic field (RF) pulse 202 having a proton resonance frequency fh is irradiated along with the application of a slice direction gradient magnetic field pulse 201 in the z direction to excite protons in a predetermined slice in the target object. Then, after applying the slice rephase gradient magnetic field pulse 203 and the dephasing phase encoding gradient magnetic field pulse 204 for adding the position information in the phase encoding direction (y direction) to the magnetization, the 180-degree pulse 208 is irradiated and read. In order to add position information in the out direction (x direction), a plurality of magnetic resonance signals (echoes) 207 are measured while applying a dephasing readout gradient magnetic field 205 and a positive and negative alternating readout gradient magnetic field pulse 206. At this time, in order to add position information in the phase encoding direction (y direction), a blip-shaped gradient magnetic field 210 is applied every time the echo 207 is measured.

  The MPG pulse 211 is applied before and after the 180 degree pulse 208. Here, as an example, the case of applying in the lead-out direction (x direction) is shown. The MPG pulse may be applied in any of the slice direction, the readout direction, and the phase encoding direction. Here, in order to perform sufficient diffusion weighting, an MPG pulse having a large intensity and a long application time is required.

It is known that when shooting parameters other than the MPG intensity and time are fixed, the DWEPI signal I and the signal function fs are expressed as follows (FIG. 11-4).

  Here, I0 is a signal in the absence of MPG, D is a diffusion coefficient, Δ is an application time of one MPG pulse, Δ is an interval between MPG pulses, and G is an intensity of the MPG pulse. b is a value indicating the degree of diffusion emphasis. When G is increased and b is increased, the diffusion enhancement degree is increased, but the signal is decreased and the SN ratio is decreased.

  Next, a plurality of images are shot by changing the shooting parameter b, and D and I0 are estimated by fitting the pixel value I to fs. For the function fitting, for example, a least square method can be used.

  Finally, an image is created using the subject parameters D and I0. A signal function fs is used to create an image. The object parameter is usually kept as it is, and the apparatus parameter b is changed to a target value to obtain a signal from fs.

  According to this method, for example, an image corresponding to a large b can be created even when the SN ratio is insufficient. That is, a plurality of images can be taken with a small b to obtain D and I0, and an image with a large b can be created from fs. In this case, the SN ratio can be expected to be better than an image taken with b increased. This is because an image having a large b value is created based on a small b value image having a high SN ratio.

A more general signal function fs when the imaging parameters are not fixed is expressed as follows (FIG. 11-5), with the diffusion coefficient D as the subject parameter and b as the imaging parameter in Equation 3.

An imaging parameter is comprehensively changed with respect to an arbitrary value of an object parameter, a signal is created by numerical simulation, and a signal function is created by interpolation. The spin density ρ, B1, and Sc to be imaged are set constant (for example, 1). In addition, B0 is set to be the same as the magnetic field strength (for example, 3T) of the apparatus used for photographing.

The basic equation of magnetic resonance including the diffusion coefficient is the Bloch-Torrey equation, which is expressed by the following equation (FIG. 11-6).

  By solving this equation, a numerical simulation can be performed to obtain a signal function.

Next, a plurality of images are captured by changing the imaging parameters FA, TR, TE, θ, and b 603, and the object parameter and the apparatus parameter 605 are estimated 604 by fitting the signal value I for each pixel to fs. However, since it is difficult to control the device parameters freely at the time of shooting, and it cannot be separated only by changing the shooting parameters except for B1, fitting is performed for f with variable conversion as follows: . Thus, the subject parameters T1, T2, and D and the device parameter B1, and Δf0 and a that are products of the subject parameter and the device parameter are estimated. For function fitting, for example, the least square method can be used (FIG. 11-7).

  Finally, an image is created using the subject parameter and the apparatus parameter. A signal function fs is used to create an image. The object parameter is usually left as it is, and the apparatus parameter is changed to a target value to obtain a signal from fs. For example, when it is desired to create an image with a magnetic field strength different from that at the time of shooting, f is calculated by changing B0. For example, based on an image taken at 3T, it is possible to create an image contrast equivalent to 7T with a higher magnetic field. When the magnetic field is increased, the frequency shift due to Cs is large, so that there is an advantage that the frequency resolution of spectroscopy is increased. When it is desired to create images of different FA, TR, and TE, signals are calculated by changing FA, TR, and TE, respectively. Thereby, an image with an optimal image contrast can be obtained.

It is a block diagram which shows schematic structure of the MRI apparatus of embodiment of this invention. It is a figure explaining the sequence diagram and k space of RF-SPOILED GRASS in the embodiment of the present invention. It is a processing flow in the embodiment of the present invention. It is a figure which shows the subject model utilized for the numerical simulation for constructing | assembling the signal function in embodiment of this invention. It is a figure which shows a part of signal function in embodiment of this invention. It is the phantom of the object parameter estimation object and its relaxation time in the embodiment of the present invention. It is an estimation result of an object parameter by an embodiment of the present invention. It is the image produced by changing the apparatus parameter by embodiment of this invention. It is a sequence diagram of DWEPI in the embodiment of the present invention. It is a figure which shows the numerical formula as described in the specification of this invention. It is a figure which shows the numerical formula as described in the specification of this invention.

101: Magnet for generating a static magnetic field, 102: Gradient magnetic field coil, 103: Subject, 104: Sequencer, 105: Gradient magnetic field power source, 106: High-frequency magnetic field generator, 107: Probe, 108: Receiver, 109: Computer, 110: Display, 111: Storage medium

Claims (7)

An imaging means for applying a high-frequency magnetic field and a gradient magnetic field to a subject placed in a static magnetic field to detect a magnetic resonance signal generated from the subject, and an operation for processing the magnetic resonance signal detected by the imaging means A magnetic resonance imaging apparatus comprising: means; and control means for controlling the imaging means and the computing means,
The photographing means photographs a plurality of images with different photographing parameters,
The calculation means estimates, from the plurality of images, a subject parameter and a device parameter including any one of magnetic field strength, irradiation coil sensitivity, or reception coil sensitivity by function fitting to a signal function,
The signal function is a function that returns a signal when an imaging sequence, the imaging parameters, the apparatus parameters, and the subject parameters are given,
The magnetic resonance imaging apparatus characterized in that the calculation means generates an image for a parameter in which at least a part of the estimated apparatus parameter is different.   The said calculating means produces | generates an image by the numerical simulation which inputs the parameter from which the said estimated object parameter or apparatus parameter, the said imaging | photography sequence, and the said some imaging | photography parameter differ. Magnetic resonance imaging device.   The said calculating means produces | generates an image by using as a signal the value of the signal function with respect to the parameter from which the estimated said object parameter or apparatus parameter, imaging | photography sequence, and a part of said imaging parameters differ. Magnetic resonance imaging device. The signal function is obtained by interpolation from signal values of a plurality of images obtained by numerical simulation,
The magnetic resonance imaging apparatus according to claim 1, wherein the numerical simulation receives the imaging parameter, the apparatus parameter, and an object parameter, and outputs the magnetic resonance signal.   The magnetic resonance imaging apparatus according to claim 1, wherein the numerical simulation solves Bloch's equation.   The magnetic resonance imaging apparatus according to claim 1, wherein the imaging parameter includes any one of a repetition time, a high-frequency magnetic field intensity, and a high-frequency magnetic field phase.   The magnetic resonance imaging apparatus according to claim 1, wherein the subject parameter includes any one of a longitudinal relaxation time, a transverse relaxation time, a spin density, a resonance frequency, and a diffusion coefficient.