JPH06285038A - Magnetic resonance video system - Google Patents

Abstract

PURPOSE:To provide the magnetic resonance video system which can measure T1 with high accuracy in a short observation time. CONSTITUTION:This system is constituted of a means for collecting and reconstituting data by applying high-frequency magnetic field pulses composed of a 180 degree pulse first and a 90 degree pulse secondly, a means for collecting and reconstituting data by applying high-frequency magnetic field pulses composed of a 90 degree pulse first and a 90 degree pulse secondly, a means for generating data by synthesizing both the data, and a means for executing a non-linear least square method by using this data train deriving magnetization MO of a thermal equilibrium state and a spin-nucleon relaxation time T1.

Description

Detailed Description of the Invention

[0001]

FIELD OF THE INVENTION The present invention relates to a magnetic resonance imaging apparatus. Particularly, in the subject, the spin-nucleon relaxation time T1
And a technique for measuring the spin-spin relaxation time T2. Further, the present invention relates to a method of correcting signal mixture due to Gibbs ringing.

[0002]

2. Description of the Related Art Magnetic resonance imaging is a technique in which a group of nuclei with unique magnetic moments is placed in a uniform static magnetic field.
This is a phenomenon in which the energy of a high-frequency magnetic field rotating at a specific frequency is absorbed resonantly, and this phenomenon is used to visualize the chemical and physical phenomena of a substance. Among these physical phenomena, spin-nucleon relaxation and spin-spin relaxation are important phenomena, and this phenomenon is used in many diagnostics. This is a spin-nucleon relaxation time T ₁ and spin-description of this relaxation process in the normal part and the lesion part.
This is because the spin relaxation times T ₂ are different. Apart from this, it is necessary to measure T ₁ and T ₂ also from the viewpoint of quantification. This is because, normally, under the conditions of a pulse sequence for imaging, the magnetic resonance signal is affected by T1 and T _2, for the quantification because shall correction of T ₁ and T ₂ is there.

When measuring T ₁ , normally IR (Inver
sion recovery) method is used. The pulse sequence of the IR method is shown in FIG. This method is pre-pulse 18
This is a method in which magnetization in a thermal equilibrium state is inverted in the −Z direction by a 0 degree pulse, and a 90 degree pulse which is an observation pulse is applied after t _1i time for observation. By changing the time t _1i between the pre-pulse and the observation pulse, the magnetization recovery process as shown in FIG. 4 is obtained, and M ₀ , T ₁ is obtained by curve fitting.
Is the method of asking for.

However, in this method, it is necessary to wait until the magnetization M ₀ in the thermal equilibrium state is restored after the 90-degree pulse is applied. Therefore, the 90-degree pulse and the second 180
The time pulse interval T _D had to be set to 5 times T ₁ . Therefore, the observation time Tobs is given by the following equation, and there is a problem that it takes much time. In addition,
In Expression 1, n is the number of t _1i .

[0005]

[Equation 1]

Therefore, in order to solve the above problem, FI
The R (Fast Inversion Recovery) method was developed. In this method, in order to shorten the observation time, T _{D is set} to T ₁ which is 2 to 3 times, and in the same manner as the IR method, the time t 1 _i of the prepulse and the observation pulse is changed to obtain the magnetization recovery process. is there. Fitting is performed on the recovery process obtained in this way to obtain M ₀ and T ₁ . If T _D = 2T ₁ , Tobs is given by the following equation.

[0007]

[Equation 2]

However, even this FIR method requires an observation time corresponding to T _D. On the other hand, in the SR (Saturation Recovery) method in which the prepulse is 90 degrees, T _D can be set to almost 0, and therefore the observation time is given by the following equation. The pulse sequence of this SR method is shown in FIG.
Shown in.

[0009]

[Equation 3]

As is clear from the above equations 2 and 3,
In the SR method, the observation time can be shortened as compared with other methods. However, as shown in FIG. 6, the change in magnetization is 0 to M ₀ , which is about 1/2 that of the FIR method.
Become. For this reason, the SR method has a problem that the accuracy is deteriorated.

On the other hand, there are several methods for measuring T ₂ as follows. That is, 1 of T ₂ measurement method
One is a method in which the echo time is changed and measured, and some obtained signals are obtained by processing such as curve fitting. The pulse sequence used at this time is shown in FIG.
4 shows. The change of the echo signal can be described by T ₂ , and the echo signal train is exp
It can be represented by (-TE / T2). However,
This method has a problem that the observation time is long because data must be collected every time the echo time is changed. Further, if TE is lengthened, it is likely to be affected by spin diffusion and deviates from the curve of exp (-TE / T2), resulting in a problem that measurement accuracy is deteriorated.

As a solution to this problem, Call and Pu
There is known a method called a CP method in which a 90-degree pulse is connected to a 180-degree pulse, such as 90-180 to 180 -... proposed by rcell. This pulse sequence is shown in FIG. 15, and the state of curve fitting is shown in FIG.
Shown in. In this pulse sequence, the spin has 9
Since it is affected only between 0 and 180 degrees or between 180 and 180 degrees, the influence of diffusion can be reduced by shortening these intervals. However, this method has a problem that it is susceptible to the imperfections of the 180-degree pulse.

Meiboom and Gill are 90 degrees X'-180 degrees y'-18 that change the phase between 90 degrees and 180 degrees by 90 degrees.
CPMG (Call-P which is a pulse sequence of 0 degree y '...
urcell-MeibooM-Gill) method was devised. This pulse sequence is shown in FIG. In this method, the odd-numbered echoes are affected by the imperfections of the high-frequency magnetic field, while the even-numbered echoes are not affected. Therefore, if curve fitting is performed using only even-numbered echo signal intensities and T ₂ is obtained, there is no effect due to incompleteness of the high-frequency magnetic field. This state is shown in FIG. However, there is a problem that the odd-numbered echo signals cannot be used for curve fitting and the data of the odd-numbered echoes are wasted.

On the other hand, the T ₂ distribution can be measured by a method that combines the CPMG method with phase encoding and frequency encoding. However, this imaging method also has a problem that the odd-numbered echo signals are wasted as in the T ₂ measurement method.

Although the quantification is mentioned above, the Gibbs ringing phenomenon is also a problem in the quantification. The Gibbs ringing phenomenon is a phenomenon in which a signal is mixed into another voxel because the voxel size is finite. This Gibbs ringing phenomenon can be explained by band limitation in the frequency space. The density distribution of the substance is shown in Fig. 25 (a).
Consider the case where it is expressed as. When this is Fourier transformed, it becomes as shown in FIG. However, since the voxel size is finite, the band is limited in the k-space as shown in (c) of the figure, and the data actually collected in the k-space becomes the data as shown in (d) of the figure. . The same figure (b) is obtained by Fourier transforming the same figure (e), and the same figure (c) is obtained by Fourier transforming the same figure (f). The same figure (g) in which and are convolved is obtained. As a result, Gibbs ringing occurs, and mixing of signals into other voxels occurs. This effect becomes very problematic when the high signal region and the low signal region are close to each other. For example, head imaging of the 31P metabolite creatine phosphate or adenosine triphosphate.
In muscle, the content of creatine phosphate and adenosine triphosphate is about 7 to 8 times that of brain. For this reason, the accuracy of measurement in the brain deteriorates due to the mixing of muscle signals into the brain,
It was a problem.

[0016]

As described above, the conventional FIR method allows accurate measurement, but has a drawback that the observation time is long. On the other hand, in the SR method, the observation time can be shortened, but on the other hand, the accuracy is deteriorated.

On the other hand, in the conventional T2 measuring method, it is useless because the collected odd-numbered data cannot be used,
There was a problem that the measurement accuracy decreased. Further, in the conventional magnetic resonance imaging apparatus, when the high signal region and the low signal region are close to each other, there is a problem that an image cannot be accurately imaged in the low signal region due to mixing of signals by Gibbs ringing.

[0018]

In order to solve the above-mentioned conventional problems, the present invention generates a high-frequency magnetic field and a gradient magnetic field to a subject placed in a uniform static magnetic field to generate the subject. In a magnetic resonance imaging apparatus for collecting magnetic resonance signals to obtain a magnetic resonance image and applying a high-frequency magnetic field pulse composed of a first pulse of 180 degree pulse and a second pulse of 90 degree pulse A second applying means for applying a high-frequency magnetic field pulse in which the first pulse and the second pulse are 90-degree pulses, and the first or second
Means for collecting magnetic resonance signals generated from the subject on the basis of the applying means for obtaining first or second reconstruction data, and the first reconstruction data and the second data obtained by this means. Of the magnetic resonance imaging apparatus and means for obtaining the synthetic data by synthesizing the reconstructed data of 1) and the means for obtaining the magnetization or spin-nucleon relaxation time in the thermal equilibrium state based on the synthetic data obtained by this means. Constitute.

Further, according to the present invention, by applying a high frequency magnetic field and a gradient magnetic field to a subject placed in a uniform static magnetic field, magnetic resonance signals generated from the subject are collected to obtain a magnetic resonance image. In a magnetic resonance imaging apparatus, a first high frequency magnetic field pulse having a pulse angle of 90 degrees is applied 90
Degree high frequency magnetic field pulse applying means and a 180 degree high frequency magnetic field for applying a plurality of second high frequency magnetic field pulses having a phase difference of 90 degrees to the first high frequency magnetic field pulse and a pulse angle of about 180 degrees. Pulse applying means, means for collecting magnetic resonance signals generated from the subject corresponding to each second high-frequency magnetic field pulse, and odd-numbered application of the 180-degree high-frequency magnetic field pulse applying means A magnetic resonance imaging apparatus is constituted by means for obtaining the spin-spin relaxation time based on each of the corresponding magnetic resonance signals and the magnetic resonance signals corresponding to the even-numbered application.

Further, according to the present invention, by applying a high frequency magnetic field and a gradient magnetic field to a subject placed in a uniform static magnetic field, magnetic resonance signals generated from the subject are collected to obtain a magnetic resonance image. In the magnetic resonance imaging apparatus, among a plurality of voxels forming the magnetic resonance image, a means for calculating the mixing amount of the image signal from one voxel to another voxel, and based on the mixing amount obtained by this means A magnetic resonance imaging apparatus is constituted by means for obtaining a corrected image from the image signal of the one voxel.

[0021]

According to the invention of claim 1, it is possible to obtain the magnetization reversed in the -Z direction by acquiring by the FIR method for the portion where the time between the prepulse and the observation pulse is short,
It is possible to observe the magnetization in the Z direction by acquiring by the SR method the portion where the time between the pre-pulse and the observation pulse is long. In other words, since the change in the magnetization to be recovered can be increased to M _0, which is almost twice as large as in the FIR method, the accuracy is F
It can be made equivalent to the IR method. Further, since the FIR method that requires the waiting time T _D after the observation pulse is used only in the portion where the pre-pulse and the observation pulse time are short and the SR method that can make T _D almost 0 is used in the long portion, the FIR method is used. It becomes possible to measure in a short observation time. That is, with the above configuration, it is possible to accurately determine M ₀ and T ₁ .

According to the second aspect of the present invention, the nonlinear least squares method can be performed by using all of the large number of echoes formed by the high frequency magnetic field pulse.
The measurement accuracy of 2 is improved.

Furthermore, according to the third aspect of the present invention, it is possible to calculate the mixing amount of the signal from the high signal voxel to the low signal voxel, and use this to correct the signal mixing due to the Gibbs ringing phenomenon. Therefore, it is possible to obtain an accurate image.

[0024]

Embodiments of the present invention will be described below with reference to the drawings. FIG. 2 is a block diagram showing the configuration of the magnetic resonance imaging apparatus according to the embodiment of the present invention. In the same figure, a static magnetic field magnet 1 and a gradient coil 2 and a shim coil 4 provided inside the static magnetic field magnet 1 make a static magnetic field uniform on a subject (not shown) and linear in three directions x, y, z orthogonal to each other in the same direction. A gradient magnetic field having a gradient magnetic field distribution is applied. The gradient coil 2 is driven by a gradient coil power supply 5, and the shim coil 4 is driven by a shim coil power supply 6. Gradient coil 2
The probe 3 provided on the inside applies a high-frequency magnetic field to the subject by being supplied with the high-frequency signal from the transmitter 7, and receives a magnetic resonance signal from the subject. The probe 3 may be used for both transmission and reception, or may be separately provided for transmission and reception. The magnetic resonance signal received by the probe 3 is detected by the receiver 8 and then transferred to the data collector 9 where it is A / D converted and then sent to the computer system 10 for data processing.

The above gradient coil power supply 5, shim coil power supply 6, receiving unit 8 and data collecting unit 9 are all controlled by the pulse sequence control unit 12, and the pulse sequence control unit 12 is controlled by the computer system 10. The computer system 10 is controlled by a command from the console 11. From the data collection unit 9 to the computer system 1
The magnetic resonance signal input to 0 is subjected to Fourier transform or the like, and image data of the density distribution of desired nuclei in the subject is reconstructed based on the Fourier transform. This image data is sent to the image display 13 and displayed as an image.

Next, a method for obtaining M ₀ and T ₁ will be described. As the pulse sequence, the pulse sequence of the FIR method shown in FIG. 7 and the pulse sequence of the SR method shown in FIG. 5 are used. In the FIR method, the signal magnitude Mobs follows Equation 4, and the relationship between Mobs and t1i is as shown in FIG.

[0027]

[Equation 4] In the SR method, the magnitude of the signal follows Mobs and t _1i according to the following equation.
The relationship between and is as shown in FIG.

[0028]

[Equation 5] In the present invention, in order to increase the change in magnetization, a short t _1i
For, the data is collected by the pulse sequence of the FIR method, and for long t _1i , the pulse sequence of the SR method is used. Figure 1 shows the magnetization recovery process at this time.
Is. As a result, the observation time is given by
It can be shorter than the FIR method.

[0029]

[Equation 6] Here, the model formula is set as the following formula.

[0030]

[Equation 7] Using this model formula, fitting is performed by the nonlinear least squares method. With this method, the amount of change in the magnitude of magnetization can be approximately doubled as in the FIR method, so that the same accuracy as in the FIR method can be obtained. Next, a case where the flip angle of the high frequency magnetic field is not 90 degrees or 180 degrees will be described. In this case, Equation 8 is used in the FIR method and Equation 9 is used in the SR method.
Follow

[0031]

[Equation 8]

[0032]

[Equation 9] As in Equation 8 or Equation 9, h regarding the flip angle is added to the parameter. In this case, equation 10 is used instead of the model equation shown in equation 7 above.

[0033]

[Equation 10] Fitting is performed using this model formula to determine M ₀ and T ₁ .

In order to image the T ₁ distribution by this method, the pulse sequence of FIG. 9 or 11 is used for the FIR method, and the pulse sequence of FIG. 10 or 12 is used for the SR method. A sequence may be used.

Next, a method of measuring T ₂ will be described. FIG. 17 is a diagram showing a pulse sequence of the CPMG method, and this pulse sequence is used in this embodiment.
When the pulse is deviated from 180 degrees, the spin exhibits the following behavior in the CPMG method. Where 180
It is assumed that the deviation from the degree is α and the pulse is a (180−α) degree pulse.

First, the 90-degree pulse causes the spin to fall on the y'axis of the rotating coordinate system (FIG. 19 (a)). Next, by the interval τ between the 90-degree pulse and the (180-α) -degree pulse, it spreads as shown in FIG. 19B due to the nonuniformity of the magnetic field. Here, the (180-α) degree pulse applied in the y ′ direction reverses the spin where it floats from the x′y ′ plane ((c) of the same figure). After that, the spins move in the y′-axis direction and gather at a position floating from the y′-axis as shown in FIG. This is the first echo. In the time of τ until the next (180-α) degree pulse, it moves as shown in FIG.
A pulse of −α) is applied ((f) in the same figure). And
The spins move again in the y'axis direction, and after τ time, the spins gather on the y'axis ((g) in the same figure). This is the second echo. In this way, the odd-numbered echoes are gathered at a position floating from the y'axis, while the even-numbered echoes are gathered on y '. Therefore, the echo signal train becomes as shown in FIG. 18, and the even-numbered echo signal Meven (TE) becomes as shown in the following equation.

[0037]

[Equation 11]

On the other hand, the odd-numbered echoes gather at a position slightly floating from the y'axis, but the odd-numbered echo signals Moddd
(TE) is also attenuated at T ₂ , and can be expressed by the following equation, where k is a proportional constant.

[0039]

[Equation 12]

The data is arranged as shown in FIG. 13 in order to use both the even-numbered echo and the odd-numbered echo. In this data array, the model formula of Expression 11 is used in the area of FIG. 11A, and the model expression of Expression 12 is used in the area of FIG. Using this model formula, M
Non-linear least squares method with ₀ , k, T ₂ as parameters is performed. Thereby, both the even-numbered echo and the odd-numbered echo can be used, and the T ₂ measurement accuracy is improved.

The method of measuring T ₂ has been described above. On the other hand, T
To obtain the ₂ distribution, the pulse sequence shown in FIG. 20 or 21 is used. With this pulse sequence,
An echo signal train as shown in FIG. 18 is obtained for each pixel. These signal sequences are converted into data sequences as shown in FIG. After that, curve fitting is performed for each pixel using the model formulas shown in Formula 11 and Formula 12. With this method, the T ₂ distribution can be obtained.

Finally, the correction method in metabolite imaging will be described with reference to FIG. First, a case where a high signal region and a low signal region are known in advance in metabolite imaging will be described. For example, in imaging of ³¹ P metabolites, creatine phosphate and adenosine triphosphate correspond to this. It is known in advance that the difference between these metabolites in the muscle and the brain is about 7 to 8 times, and the influence of Gibbs ringing is large. In this case, the position of the muscle is confirmed by the 1 H image, and the amount of the signal mixed into the brain may be calculated.

First, the distribution within the voxel of the metabolite image is obtained based on the 1 H image (step 1). However, the voxel for which the distribution is obtained may be only the voxel in which the brain and the muscle are mixed as shown in FIG. 23A or the voxel in which the muscle exists only in a part of the voxel as shown in FIG. 23B. For these voxels, the voxel interior is divided into M × M × M to obtain the distribution. FIG. 24 shows the correspondence between the voxel coordinates and the coordinates after subdivision. However, only the one-dimensional direction is shown. Next, in step 2, the amount of mixture is calculated. here,
The coordinates of the high signal voxel are (p, q, r), and the voxel of the mixing destination is (n, l, m). The concentration distribution in the high signal voxel necessary for the calculation of the amount of mixture can be expressed by Expression 13. Since the voxel signal can be considered as a muscle signal based on the difference in concentration between the muscle and the brain, the muscle signal of each pixel is the voxel signal value divided by the number of pixels including muscle.

[0044]

[Equation 13]

[0045]

[Equation 14] The following formula is used to calculate the amount of contamination.

[0046]

[Equation 15] In Expression 15, M is an even number here.

Next, a correction image is obtained by performing correction based on the obtained mixture amount (step 3). The correction method is obtained by subtracting the mixing amount from the signal value for low signal voxels. For high-signal voxels, the signal amount is added to the low-signal voxels. With that, the correction is completed.

Next, another embodiment will be described. First, in the metabolite image, a voxel in which the image signal ratio of adjacent voxels is a or more is searched for, and a high signal voxel is found. a is set in advance. Next, the high signal area on the 1H image corresponding to this high signal voxel is confirmed. In the correction method described above with reference to FIG. 22, this high signal area corresponds to the muscle signal. That is, the processing of FIG. 22 performed on the previous muscle signal is performed on this high signal area. As described above, the corrected image can be obtained.

In the method described above, the distribution of the metabolite image in the high-signal voxel is considered to be uniform, but it can also be calculated by separately metabolizing the metabolite image with high resolution. . With this method, the distribution within the voxel can be obtained, and the mixing amount can be calculated and corrected by the method of FIG.

[0050]

As described above, according to the present invention, it is possible to accurately obtain T ₁ in a short observation time. Further, in the invention relating to the T ₂ measuring method, T ₂ can be measured with high accuracy, and T ₂ distribution can be measured with high accuracy. Further, according to the invention related to Gibbs ringing, the amount of mixing from the high signal region to the low signal region can be obtained, and an accurate image can be obtained.

[Brief description of drawings]

FIG. 1 is a diagram showing a magnetization recovery process when the present method is used.

FIG. 2 is a block diagram of a magnetic resonance imaging apparatus.

FIG. 3 is a diagram showing a pulse sequence of the IR method.

FIG. 4 is a diagram showing a magnetization recovery process in the IR method.

FIG. 5 is a diagram showing a pulse sequence of the SR method.

FIG. 6 is a diagram showing a magnetization recovery process in the SR method.

FIG. 7 is a diagram showing a pulse sequence of the IR method.

FIG. 8 is a diagram showing a magnetization recovery process in the FIR method.

FIG. 9 is a diagram showing a pulse sequence for obtaining a T ₁ distribution which is an embodiment of the present invention.

FIG. 10 is a diagram showing a pulse sequence for obtaining a T ₁ distribution which is an embodiment of the present invention.

FIG. 11 is a diagram showing a pulse sequence for obtaining a T ₁ distribution which is an embodiment of the present invention.

FIG. 12 is a diagram showing a pulse sequence for obtaining a T ₁ distribution which is an embodiment of the present invention.

FIG. 13 is a view for explaining an example in which an echo signal train collected by the CPMG method is divided into even-numbered echoes and odd-numbered echoes and data is arrayed.

FIG. 14 is a diagram showing a sequence of high-frequency magnetic field pulses for spin echo.

FIG. 15 is a diagram showing a pulse sequence for performing the CP method.

FIG. 16 is a diagram showing a state in which an echo signal train by the CP method is curve-fitted.

FIG. 17 is a diagram showing a pulse sequence for performing the CPMG method.

FIG. 18 is a diagram showing a state in which an echo signal train by the CPMG method is curve-fitted.

FIG. 19 is a diagram for explaining the behavior of spins in the CPMG method.

FIG. 20 is a diagram showing a pulse sequence for obtaining a T ₂ distribution which is an embodiment of the present invention.

FIG. 21 is a diagram showing a pulse sequence for obtaining a T ₂ distribution which is an embodiment of the present invention.

FIG. 22 is a flowchart showing a method for correcting the influence of Gibbs ringing.

FIG. 23 is a diagram showing the distribution of voxels in a metabolite image.

FIG. 24 is a diagram showing coordinates of voxels and subdivided voxels.

FIG. 25 is a diagram for explaining the Gibbs ringing phenomenon.

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 ... Static magnetic field magnet 2 ... Gradient coil 3 ... Shim coil 4 ... Probe 5 ... Gradient coil 6 ... Shim coil power supply 7 ... Transmitting unit 8 ... Receiving unit 9 ... Data collecting unit 10 ... Computer system 11 ... Console 12 ... Pulse sequence control unit 13 … Image display

Claims (5)

[Claims] 1. A magnetic resonance imaging apparatus for obtaining a magnetic resonance image by collecting a magnetic resonance signal generated from the subject by applying a high frequency magnetic field and a gradient magnetic field to the subject placed in a uniform static magnetic field. In, the first pulse is a 180-degree pulse and the second pulse is a 90-degree pulse, and the first applying means applies a high-frequency magnetic field pulse, and the first pulse and the second pulse are 90-degree pulses. Second applying means for applying a high-frequency magnetic field pulse; means for collecting magnetic resonance signals generated from the subject based on the first or second applying means to obtain first or second reconstruction data A means for synthesizing the first reconstructed data and the second reconstructed data obtained by this means to obtain synthetic data, and a thermal equilibrium state based on the synthesized data obtained by this means. Magnetization or spin - magnetic resonance imaging apparatus characterized by comprising the nucleons relaxation time or means for obtaining both. 2. The first applying means applies when the time interval between the first pulse and the second pulse is a predetermined time or less, and the second applying means applies the first pulse and the second pulse. The magnetic resonance imaging apparatus according to claim 1, wherein the magnetic field is applied when the time interval with the second pulse is a predetermined time or more. 3. A magnetic resonance imaging apparatus for obtaining a magnetic resonance image by collecting a magnetic resonance signal generated from the subject by applying a high frequency magnetic field and a gradient magnetic field to the subject placed in a uniform static magnetic field. A 90 degree high frequency magnetic field pulse applying means for applying a first high frequency magnetic field pulse having a pulse angle of 90 degrees; and a phase difference of 90 degrees with respect to the first high frequency magnetic field pulse, and the pulse angle is substantially 180 degree high frequency magnetic field pulse applying means for applying a second high frequency magnetic field pulse of 180 degrees a plurality of times, and means for collecting magnetic resonance signals generated from the subject corresponding to each of the second high frequency magnetic field pulses Of the application by the 180-degree high-frequency magnetic field pulse applying means, a magnetic resonance signal corresponding to an odd-numbered application and a magnetic resonance signal corresponding to an even-numbered application are respectively applied. A magnetic resonance imaging apparatus comprising: a means for determining a pin-spin relaxation time. 4. A magnetic resonance imaging apparatus for obtaining a magnetic resonance image by collecting a magnetic resonance signal generated from the subject by applying a high frequency magnetic field and a gradient magnetic field to the subject placed in a uniform static magnetic field. In the plurality of voxels constituting the magnetic resonance image, means for calculating the mixing amount of the image signal from one voxel to another voxel, and the one voxel based on the mixing amount obtained by this means And a means for obtaining a corrected image from the image signal of the magnetic resonance imaging apparatus. 5. A magnetic resonance imaging apparatus for obtaining a magnetic resonance image by collecting a magnetic resonance signal generated from the subject by applying a high frequency magnetic field and a gradient magnetic field to the subject placed in a uniform static magnetic field. A means for blocking a high resolution image level region expressed at a predetermined resolution into a low resolution image level region, and a region in which a predetermined body tissue is partially included in the low resolution image level region, Means for extracting from the blocked area; means for equally dividing the signal from the area of the low resolution image level by the number of pixels of the high resolution image level of the body tissue contained in the blocked area; Means for calculating the amount of reduction of the image signal due to Gibbs ringing and the amount of mixing into other regions using the signal distribution in the region, and a low decomposition based on the amount of reduction and the amount of mixing. Performing correction of the signal distribution from the area of the image level, a magnetic resonance imaging apparatus characterized by comprising a means for reconstructing a magnetic resonance image of the low resolution image level.