JPH08336505A - Magnetic resonance imaging system - Google Patents

Abstract

PURPOSE: To improve the spatial resolution of, especially the region of interest of an image without extending photographing time and without narrowing a photographing visual field. CONSTITUTION: Phase encode gradient magnetic fields 26 with different change steps are formed by a CPU 1 for control, and they are set on a CPU 2 for sequence. The phase encode gradient magnetic fields 26 are applied with a 90 deg. pulse, a 180 deg. pulse and other gradient magnetic fields according to a pulse sequence, and obtained plural echo signals 18 are divided into echo signal groups with different kinds of photographing visual fields and spatial resolution by the CPU 1 for control, and also, the images of them are reconfigured, and one image is obtained from respective images by sampling and combining the one with high spatial resolution to the region of interest area and the one with low spatial resolution to the regions other than that. In this way, the whole position relation of the image is recognized, and also, the spatial resolution in the interested area is improved, and moreover, the photographing time is shortened.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a magnetic resonance imaging (hereinafter referred to as MRI) apparatus for obtaining a tomographic image by utilizing an NMR phenomenon, and more particularly to an MRI apparatus suitable for improving spatial resolution near the center of an image. .

[0002]

2. Description of the Related Art An MRI apparatus utilizes an NMR phenomenon to make a nuclear spin (hereinafter, referred to as a nuclear spin) at a desired inspection site in an object.
(Referred to as spin) density distribution, relaxation time distribution, etc. are measured, an image is reconstructed from the measured data, and the cross section of the subject is displayed as an image. In this MRI apparatus, a subject 7 is placed in a static magnetic field generator 4 that generates a static magnetic field of about 0.02 to 2 Tesla as shown in FIG. At this time, the spins in the subject 7 perform a precession motion with the direction of the static magnetic field as the central axis at a frequency determined by the strength H0 of the static magnetic field. This frequency is called the Larmor frequency. The Larmor frequency ν0 can be expressed by ν0 = γ / 2π · H0 (1). Here, γ is a gyromagnetic ratio and has a unique value for each type of nucleus. When the angular velocity of the Larmor precession is ω0, there is a relationship of ω0 = 2πν0 (2), and therefore ω0 = γH0 (3).

Here, when a high frequency magnetic field (electromagnetic wave) having a frequency equal to the Larmor frequency ν0 of the atomic nucleus to be measured by the high frequency irradiation coil 11 is applied to the subject 7, spins are excited and a transition is made to a high energy state. When this high-frequency magnetic field is cut off, the spin returns to its original low energy state. The electromagnetic wave emitted at this time is received by the high-frequency receiving coil 14, amplified by the amplifier 15, separated into two kinds of signals by the quadrature detector 16, digitized by the A / D converter 17, and then outputted to the control CPU 1. send. Control CPU 1
Then, reconstruction calculation is performed based on this data, and the calculated data is displayed on the display 18 as a tomographic image of the subject 7. The high-frequency magnetic field is obtained by amplitude-modulating the output signal of the synthesizer 8 by the modulator 9 according to the signal sent by the sequence CPU 2 controlled by the control CPU 1 and by amplifying it by the high-frequency irradiation coil amplifier 10. Obtained by sending to 11.

Further, the MRI apparatus includes a static magnetic field generator 4
In addition to the high-frequency irradiation coil 11 and the high-frequency reception coil 14, a gradient magnetic field coil 13 for generating a gradient magnetic field for giving the detected signal positional information in space is provided. This gradient magnetic field coil 13 is used for the sequence CPU 2
An electric current is supplied from the gradient magnetic field coil power supply 12 which operates by the signal from to generate a gradient magnetic field. And the subject 7
It is possible to perform imaging by determining the imaging field of view and the like and applying these static magnetic field, high-frequency magnetic field, and gradient magnetic field to the subject 7.

In the MRI apparatus configured as described above, a two-dimensional Fourier imaging method is a commonly used imaging method at present. FIG. 9 shows a schematic pulse sequence of a typical spin echo method among the imaging methods. In this pulse sequence, first 90 ° pulse 2
3 and slice encode gradient magnetic field 25 are applied. Then, when the static magnetic field direction is the Z direction, each spin in the slice falls in the XY plane orthogonal to it and rotates in the XY plane. Next, a phase encode gradient magnetic field 29 and a frequency encode gradient magnetic field 27 are applied, and a 180 ° pulse 24 is applied to the slice encode gradient magnetic field 25 after a time TE / 2 when the echo time until the echo signal 28 is generated is TE. 'Add with Then, each spin is reversed. After that, the frequency encoding gradient magnetic field 2
Add 7 '. The frequency-encoding gradient magnetic field 27 'is applied so that the echo signal 28 generated by focusing the spins becomes maximum at time TE.

Here, in order to construct a tomographic image, it is necessary to obtain the spatial distribution of signals. For this purpose, a linear gradient magnetic field is used. A spatial magnetic field gradient can be created by superimposing a gradient magnetic field on a uniform static magnetic field. As described above, since the spin rotation frequency is proportional to the magnetic field strength, the spin rotation frequency is spatially different in the state in which a gradient magnetic field is applied. Therefore, the position of each spin can be known by examining the frequency of the spin in the state where the gradient magnetic field is applied and the phase of the spin after applying the gradient magnetic field for a predetermined time. For this purpose, a phase encoding gradient magnetic field 29, a frequency encoding gradient magnetic field 27 '
Is being applied.

Using the pulse sequence described above as a basic unit, the intensity of the phase-encoding gradient magnetic field is changed each time, and a predetermined number of times, for example 256 times, are repeated at a constant repetition time (TR). The spatial distribution of macroscopic magnetization is obtained by two-dimensional inverse Fourier transform of the measurement signal thus obtained, and a tomographic image of the subject 7 is obtained. The above-mentioned basic principles of MRI are detailed in "NMR Medicine" (Basic and Clinical) (edited by Nuclear Magnetic Resonance Medicine, Maruzen Co., Ltd., issued January 20, 1984).

In such an MRI apparatus, the magnetic field strength of the phase encode gradient magnetic field is changed every TR and applied repeatedly, and the increment amount Gpst of the gradient magnetic field is increased with each repetition.
ep was constant, that is, applied at equal intervals. This increase amount Gpstep is proportional to the reciprocal of the imaging field of view FOV, that is, if the increase amount Gpstep is large, the imaging field of view FOV is narrow, and if the increase amount Gpstep is small, the imaging field of view FOV.
Becomes wider. Gpstep = N / FOV (4) Here, N is a constant. The relationship between the phase encoding gradient magnetic field and the field of view FOV will be described with reference to FIG. In the imaging field of view FOV in the phase encode direction, the phase difference of spins selectively excited by application of a phase encode gradient magnetic field for one step is 2π (π to −) in the phase encode direction.
The range of π) is the field of view. Further, the phase difference of spins is proportional to the application amount of the gradient magnetic field, and if the application time is constant, the phase difference increases as the increase amount (slope) increases. Since the shooting field of view is the range of the phase difference of 2π,
The larger the amount of increase (tilt), the narrower the field of view for photography. For example, if there are two gradient magnetic fields of X and Y, and the increase amount (slope) of Y is twice the increase amount (slope) of X, and the imaging field of view is FOVx with the gradient magnetic field of X, the Y position is increased. Since the phase difference is proportional to the increase amount (gradient) of the gradient magnetic field, it becomes 4π in the same range as X. Since the imaging field of view has a phase difference of 2π, the imaging field of view FOVy is ½ of X. That is,
If the amount of increase (gradient) of the gradient magnetic field is doubled, the field of view for photography is halved, and if it is tripled, it is ⅓.

A pixel R for forming an image is represented by a size obtained by dividing the field of view FOV by the number of times PRJ of application of the phase encoding gradient magnetic field. R = FOV / PRJ (5) Furthermore, the imaging time T for one image is represented by the product of the time TR during the repetition of the high frequency magnetic field and the repetition number PRJ. T = TR.PRJ (6) Then, in the MRI apparatus having such a relationship, in order to increase the spatial resolution of the image, that is, to reduce the pixel R, the number of repetitions PRJ is increased from the equation (5), or
The field of view FOV may be reduced. However, equation (6)
If the number of repetitions PRJ is further increased, the shooting time T becomes longer, and if the shooting field of view FOV is made smaller, the field of view becomes narrower, making it difficult to grasp the overall positional relationship.

For example, in the field of view FOV of the whole image as shown in FIG. 3D, the phase encoding gradient magnetic fields a, b and c and the images obtained when the following imaging parameters are described. When the phase encode gradient magnetic field a is used as a reference, the number of repetitions PRJ and the imaging field of view FOV are expressed as follows. a; PRJa = 4 FOVa = FOV (7) b; PRJb = 4 FOVb = FOV / 2 = FOVa / 2 (8) c; PRJc = 8 FOVc = FOV = FOVa (9)

From the equations (7), (8) and (9), the imaging field of view is as shown in FIG. 3, the phase-encoding gradient magnetic fields a and c are the whole image, and the phase-encoding gradient magnetic field b is half of the whole image. . First, the increase amount Gpsstep of each phase encoding gradient magnetic field can be calculated from the equation (4) as follows: Gpstepa = N / FOVa (10) Gpstepb = N / FOVb = N / FOVa / 2 = 2N / FOVa = 2Gpstepa (11) Gpstepc = N / FOVc = N / FOVa = Gpstepa (12). Here, N is a constant. Further, the position of each phase encoding gradient magnetic field with respect to the number of repetitions is i
Then, it is expressed as follows. − (PRJ / 2) ≦ i <PRJ / 2 (13) where i = integer

From the equation (13), the position of the phase encode gradient magnetic field a takes values of ia = 1, 0, -1, -2, and the phase encode gradient magnetic field b is the phase encode gradient magnetic field a and the number of repetitions PRJ. Since it is the same, ib = 1,
It takes values of 0, -1, and -2, and the phase encoding gradient magnetic field c
Takes values of ic = 3, 2, 1, 0, -1, -2, -3, -4. Each increment G at such a position i
The pstep is represented by the equations (10), (11) and (12), and is shown in FIG. The increase Gpsstepb of the phase-encoding gradient magnetic field b due to such a relationship is 2 compared to the others.
Double. Next, the pixel R is expressed by the formula (5) as follows: Ra = FOVa / 4 (14) Rb = FOVb / 4 = FOVa / 2/4 = FOVa / 8 = Ra / 2 (15) Rc = FOVc / 8 = FOVa / 8 = Ra / 2 (16), and the number of pixels Ra of the phase encoding gradient magnetic field a is double that of the others. Here, since FIG. 4 shows an image on a pixel-by-pixel basis, there are four images of the phase encoding gradient magnetic fields a and b in the phase direction and eight images of the phase encoding gradient magnetic field c.
Divided into two. If this is shown in memory units for display, the phase-encoding gradient magnetic fields b and c are displayed as they are on the display or the like, but the phase-encoding gradient magnetic field a is a compressed image, so it is enlarged and displayed when displayed. . In order to simplify the description, the display memory is divided into eight, and one division of the memory is used for one repeat number PRJ. And the shooting time T
From equation (6), Ta = TR × 4 (16) Tb = TR × 4 = Ta (17) Tc = TR × 8 = 2Ta (18), and the imaging time Tc of the phase encode gradient magnetic field c is Doubled.

[0013]

As described above, when the photographic field of view is the entire image and the number of repetitions is 4, the pixel becomes large and the spatial resolution is lowered as shown in FIG. 3 (a). . Therefore, if you narrow the field of view to 1/2,
Improves spatial resolution without changing the shooting time.
However, as shown in FIG. 3 (b), only the vicinity of the center can be photographed, making it difficult to grasp the overall positional relationship. If the number of repetitions is doubled, the number of pixels can be reduced to 1/2 without changing the photographing field as shown in FIG.
The shooting time was doubled. Therefore, an object of the present invention is to improve the spatial resolution of an image, particularly a region of interest, without extending the imaging time and narrowing the imaging field of view.

[0014]

In order to achieve the above object, the present invention provides a high-frequency magnetic field for exciting nuclear spins in a subject and a slice encoding for applying an imaging slice position of the subject simultaneously with the high-frequency magnetic field. A gradient magnetic field, and a phase-encoding gradient magnetic field applied by changing the magnetic field strength stepwise to give phase information to the excited nuclear spins,
Sequence control means for applying a frequency-encoding gradient magnetic field that gives frequency information to the nuclear spins to which the phase information has been applied in order to read out an NMR signal according to a pulse sequence, and an NM obtained by this pulse sequence.
In the magnetic resonance imaging apparatus having a means for reconstructing an image of the R signal, the sequence control means executes the pulse sequence by changing the changing step of the magnetic field strength in the low encode region and the high encode region of the phase encode.

Further, the image reconstructing means divides and forms the NMR signals obtained by the pulse sequence into a plurality of NMR signal groups having different imaging fields of view and spatial resolutions during image reconstruction, and these NMR signal groups are formed. Image is reconstructed separately, a desired image region is extracted from each image obtained by the image reconstruction, and these extracted image regions are combined to obtain one image. In particular, the change step of the phase encode gradient magnetic field is made larger in the high encode area than in the low encode area, and a plurality of NMR signal groups obtained by division are obtained by a pulse sequence in which the change steps of the phase encode are at equal intervals. NM
Divide for each R signal.

Furthermore, a plurality of NMR images for image reconstruction are used.
The signal group is an NMR signal group formed only by NMR signals,
An image formed by combining a signal group in which 0 data is added to an NMR signal and obtained by combining is extracted from an image having a high spatial resolution in a region corresponding to the region of interest, and a space lower than the region of interest in the other regions. It is extracted from the resolution image and synthesized. Furthermore, an image obtained by reconstructing an image of a plurality of NMR signal groups or an NMR signal group obtained by adding an NMR signal and 0 data to an image other than the desired image area is set to 0.
An image to which data is added and 0 data is added is inversely converted into signal data, these signal data are combined, and images are reconstructed to obtain one image.

[0017]

The step of changing the magnetic field strength of the phase encode gradient magnetic field is applied so that it is different between the low encode region and the high encode region, and the NMR signal obtained by the pulse sequence including this phase encode gradient magnetic field is converted into a plurality of NMR signals. The image is reconstructed by dividing it into signal groups. The images obtained by this are images with different field of view and spatial resolution,
A desired image area is extracted from these images and combined to obtain one image. As a result, an image with high spatial resolution can be combined with the region of interest and an image with low spatial resolution can be combined with other regions, so that the spatial resolution can be increased and the overall positional relationship can be grasped. Further, since the spatial resolution is high only in the region of interest and the spatial resolution is sufficient to grasp the positional relationship in the other regions, the imaging time can be shortened as compared with the same spatial resolution and imaging field of view.

[0018]

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS An embodiment of the present invention will be described below with reference to FIGS. 1 is a block diagram showing the configuration of an MRI apparatus for carrying out the present embodiment, FIG. 2 is a diagram showing a pulse sequence of the present embodiment, FIG. 3 is a diagram showing a relationship between an imaging field of view, pixels, and the number of repetitions, FIG. Is a diagram showing a combination of phase encoding gradient magnetic fields, and FIG. 5 is a diagram showing image reconstruction of the present embodiment. The MRI apparatus of this embodiment is roughly classified into a control C
It comprises a PU 1, a sequence CPU 2, a transmission system 3, a static magnetic field generating magnet 4, a reception system 5, and a signal processing system 6.

The control CPU 1 has a type of pulse sequence previously input from an input device such as a keyboard 22, and the like.
A sequence CPU according to a program according to a photographing parameter including a photographing field of view FOV and a repetition number PRJ
2, each of the transmission system 3, the reception system 5, and the signal processing system 6 is controlled. The sequence CPU 2 operates based on a control command from the control CPU 1 and sends various commands necessary for collecting tomographic image data of the subject 7 to the transmission system 3 and the static magnetic field generating magnet 4.
It is output to the gradient magnetic field generating system 21 and the receiving system 5.

The transmission system 3 has a synthesizer 8, a modulator 9 and an irradiation coil 11 as a high frequency coil. The high frequency pulse from the synthesizer 8 is amplitude-modulated by the modulator 9 according to a command from the sequence CPU 2, and this amplitude modulation is performed. The generated high-frequency pulse is amplified through the high-frequency amplifier 10 and supplied to the irradiation coil 11 to irradiate the subject 7 with a predetermined pulsed electromagnetic wave.

The static magnetic field generating magnet 4 is for generating a uniform static magnetic field around the subject 7 in an arbitrary direction. Inside the static magnetic field generating magnet 4, the irradiation coil 11 is provided.
Besides, a gradient magnetic field coil 13 for generating a gradient magnetic field and a receiving coil 14 of the receiving system 5 are installed. The gradient magnetic field generation system 21 includes a gradient magnetic field coil 13 capable of independently applying a gradient magnetic field in Cartesian coordinate axis directions orthogonal to each other.
And a gradient magnetic field coil power supply 12 that supplies a current to the gradient magnetic field coil 13.

The receiving system 5 includes a receiving coil 14 as a high frequency coil and an amplifier 15 connected to the receiving coil 14.
And a quadrature detector 16 and an A / D converter 17, and when the receiving coil 14 detects an NMR signal from the subject 7, the signal is amplified by an amplifier 15, and the amplified signal is The quadrature phase detector 16 that uses the output signal of the synthesizer 8 as a reference signal is used as two systems of signals, and the two systems of signals are sampled by the A / D converter 17 and controlled as real part data and imaginary part data in Fourier transform. For CP
Output to U1.

The signal processing system 6 has an external storage device such as a magnetic disk 20 and an optical disk 19 and a display 18 such as a CRT. When the data from the receiving system 5 is input to the control CPU 1, control is performed. CPU1 for signal processing,
Processing such as image reconstruction is executed, and a desired sectional image of the subject 7 as a result is displayed on the display 18 and recorded on the magnetic disk 20 or the like of the external storage device. The hard disk of the magnetic disk 20 of the signal processing system 6 stores a plurality of pulse sequence programs.

Next, the pulse sequence of the spin echo method of this embodiment will be described with reference to FIG. First, keyboard 2
2, the pulse sequence of the spin echo method is selectively input, the field of view FOV, the number of repetitions PR
Input shooting parameters including J. As a result, the pulse sequence is read out from the magnetic disk 20 and input to the control CPU 1, and each photographing parameter is also input to the control CPU 1. The control CPU 1 outputs a command related to a pulse sequence to be executed to the sequence CPU 2 based on the input data. When the pulse sequence of FIG. 2 is started, the 90 ° pulse 23 and the slice encoding gradient magnetic field 25 are first applied. This excites nuclear spins in the imaging slice. Then, position information and phase diffusion are given to the nuclear spins excited by the application of the phase encode gradient magnetic field 26 and the frequency encode gradient magnetic field 27, and a 180 ° pulse 24 and a slice encode gradient magnetic field 25 ′ are applied. The slice encoding gradient magnetic field 25 ′ applies a gradient magnetic field having the same magnitude as the slice encoding gradient magnetic field 25. Also,
When the frequency encode gradient magnetic field 27 'that maximizes the echo signal 28 is applied at time TE, the echo signal 28 is generated from within the imaging slice and is measured. When the first echo signal 28 is obtained, the next 90 ° pulse 23 is applied after the repetition time TR has elapsed. This operation is repeated for the number of times PRJ, and the application amount (strength × application time) of the phase encoding gradient magnetic field 26 is changed and applied for each repetition. Then, a tomographic image is obtained by reconstructing an image based on the echo signal 28 for the number of repetitions PRJ.

Next, the phase encoding gradient magnetic field 26 of this embodiment will be described. First, a plurality of shooting parameters (see FIGS. 3 and 4) are stored in a memory (not shown) in the control CPU 1.
(A) of the imaging field of view FOVa, the increase amount Gpsstepa of the phase encoding gradient magnetic field a, and the number of repetitions PRJ
a, the field of view FOVb shown in (b) of FIGS. 3 and 4, the increase amount Gpsstepb of the phase encoding gradient magnetic field b, the number of repetitions PRJb, and the like) are stored. The field of view FOVd is shown in FIG.
The entire image as shown in (d) is set, and the number of repetitions PRJd is set. When each setting is made, control C
PU1 synthesizes shooting parameters. For example, when the field of view FOV is the whole image and the number of repetitions PRJd is 6, the field of view is the whole image from the memory.
The phase encode gradient magnetic field a shown in (a) is selected, and further the phase encode gradient magnetic field b shown in FIG. 4 (b) in which the pixel R near the center of the image is small is selected. The ratio of the increase amount Gpsstepa of the phase encode gradient magnetic field a to the increase amount Gpstepb of the phase encode gradient magnetic field b is 1: 2, and even if the number of repetitions PRJa and PRJb is the same, ia and ib are the same (1, 0 , -1, -2) when
The intensity of the phase encoding gradient magnetic field b is twice as great.

When these phase-encoding gradient magnetic fields a and b are combined, the number of repetitions PRJd becomes 6 as shown in FIG. 4D because the same gradient magnetic field exists. The phase encode gradient magnetic field d and the phase encode gradient magnetic fields a and b
As for the relationship with, id = 2 is ib = 1, and id = 1 is ia =
1 and id = 0, ia, ib = 0, and id = -1, ia =-
1, id = -2 is ia = -2 and ib = -1, id =-
3 corresponds to ib = -2, respectively. The respective increments are 1 for id = 2 to -2 and 2 for id = -2 to -3. That is, the synthesized phase encode gradient magnetic field d
Is in a form in which the phase encode gradient magnetic field a enters between the phase encode gradient magnetic fields b. Then, the phase-encoding gradient magnetic field d having such an increase amount and strength is set in the sequence CPU 2 and imaging is performed in a pulse sequence as shown in FIG.

Next, the image reconstruction of the echo signal obtained by the above pulse sequence will be described. First,
The obtained echo signal is divided into two. In this division,
The phase encoding gradient magnetic field a (ia = 1, 0, −
Echo signals (id = 1, 0, −) obtained by 1, −2)
1, -2) and the phase encode gradient magnetic field b (ib = 1,
Echo signals (id = 2, 0, -1, -2)
0, -2, -3). At this time, id = 0
Since the echo signal of 1 and the echo signal of id = -2 include both phase encode gradient magnetic fields a and b, two echo signals are created and distributed to the data group for image reconstruction. Then, these echo signals are image-reconstructed and stored in a display memory (not shown) in the control CPU.

When the echo signal group obtained by applying the phase-encoding gradient magnetic field a is image-reconstructed, in the case of displaying in memory units for display, as shown in FIG. 5C, the image of the whole image is 4 in the phase direction. The image is stored in the memory for the columns, that is, becomes a compressed image, and this is enlarged as shown in FIG. 5E to be an image of the entire memory for display (memory of 8 rows and 8 columns). Also,
When the echo signal group obtained by applying the phase encode gradient magnetic field b is image-reconstructed, the image near the center is stored in the memory for four columns in the phase direction as shown in FIG. 5D. Then, the images at both ends of the image of FIG. 5 (e) excluding the image of 8 rows and 4 columns near the center are extracted, and the extracted image and the image of the phase encoding gradient magnetic field b shown in FIG. 5 (d) ( An image of 8 rows and 4 columns near the center) is synthesized. Then, since the image of the phase encoding gradient magnetic field a has large pixels and the image of the phase encoding gradient magnetic field b has small pixels, the image near the center has small pixels, and the images at both ends have pixels. It will be big.

The field of view FOVd at this time is expressed by equation (7).
From (8), it is expressed as follows. FOVd = FOVa = 2FOVb The field of view FOVd shows the same overall image as the field of view FOVa. Next, since the number of repetitions PRJd of the pixel Rd is 6, the pixel Rd is as follows. Rd = FOVd / 6 However, the pixel Rd ′ of the image in the vicinity of the center and the pixel Rd ″ of the images at both ends are expressed by the equations (14) and (15), respectively, and Rd ′ = FOVd ′ / 4 = FOVb / 4 = FOVa / 2 / 4 = FOVa / 8 Rd ″ = FOVd ″ / 4 = FOVa / 4. The pixel Rd ′ of the image near the center has the property of the phase encoding gradient magnetic field b, and the pixel R of the image at both ends is
d ″ has a property due to the phase encode gradient magnetic field a,
Furthermore, since the number of repetitions PRJd is 6 for all pixels Rd, the phase encoding gradient magnetic field a is 1.5 times larger. Next, the shooting time Td is calculated by the equations (16), (17), (1)
From 8), it becomes as follows. Td = TR × 6 = 1.5 Ta = 1.5 Tb = Tc / 1.5 Since the number of repetitions is smaller than that of the conventional phase encoding gradient magnetic field c in which a pixel is small and an entire image can be obtained, 1.
You can shoot in 1/5.

That is, in this embodiment, a phase encode gradient magnetic field d, which is a combination of two conventional phase encode gradient magnetic fields a and b, is applied, and the obtained echo signal is divided corresponding to the phase encode gradient magnetic fields a and b. At the same time, the images are reconstructed, and the images near the center are the images by the phase encoding gradient magnetic field b, and the other images are the images by the phase encoding gradient magnetic field a, and these are combined to obtain one image. This makes it possible to grasp the overall positional relationship, increase the spatial resolution near the center of the image, and shorten the shooting time.

Further, in this embodiment, the image of the phase encoding gradient magnetic field a is enlarged, and the images on both sides of the enlarged image are combined with the image of the phase encoding gradient magnetic field b to form one image. However, as shown in FIG. 6, an image of the phase encoding gradient magnetic field a, that is, images on both sides of the compressed image,
One image may be created by combining the image of the phase encoding gradient magnetic field b. In this case, the obtained image has a normal size in the vicinity of the center and is compressed at both ends, but since the amount of information does not change at all, the same effect as described above can be obtained.

Next, a second embodiment will be described with reference to FIG. In this embodiment, the synthesis of the phase encode gradient magnetic fields a and b and the division of the echo signal are performed in the same manner as in the first embodiment. After the division into the echo signal groups of the phase encode gradient magnetic fields a and b, the echo signal group of the phase encode gradient magnetic field b is first reconstructed in the same manner as in the first embodiment. Then, before the image reconstruction of the echo signal group of the phase encoding gradient magnetic field a, the high frequency region (id = 3, id = 3,
2 and -3, -4) are provided, 0 data is added to this high frequency band, and a signal group to which 0 data is added is reconstructed. The images obtained by such image reconstruction are, as in the first embodiment, an image with large pixels in the whole image with the phase-encoding gradient magnetic field a and an image with small pixels near the center in the phase-encoding gradient magnetic field b. can get. Then, as in the first embodiment, the images at both ends of the image of the phase-encoding gradient magnetic field a, excluding the image of 8 rows and 4 columns near the center, are extracted, and the extracted images are the images of the phase-encoding gradient magnetic field b. (Image in 8 rows and 4 columns near the center) is combined to form one image. As described above, as in the first embodiment, the overall positional relationship can be grasped, the spatial resolution near the center of the image can be increased, and the photographing time can be shortened.

Next, a third embodiment will be described with reference to FIG. Also in this embodiment, the phase encode gradient magnetic fields a and b
Are combined and the echo signal is divided in the same manner as in the first embodiment, 0 data is added to the echo signal of the phase encoding gradient magnetic field a and image reconstruction, and image reconstruction of the echo signal of the phase encoding gradient magnetic field b is performed. The operations up to are performed in the same manner as in the second embodiment (Fig. 8 (a)-
(E)). Then, after reconstructing the image of the echo signal of the phase-encoding gradient magnetic field a, as shown in FIG. 8G, only both ends of this image are left and 0 data is added near the center. After the image of the echo signal of the phase encoding gradient magnetic field b is reconstructed, 0 data is added to both ends of this image as shown in FIG. An image to which 0 data is added near the center and an image to which 0 data is added to both ends are respectively subjected to inverse Fourier transform to obtain signal data (FIG. 8).
(H), (i)). Then, the respective signal data are combined into one image of signal data (FIG. 8 (j)), and the combined signal data is reconstructed into an image to obtain an image (FIG. 8 (k)). As described above, in order to add 0 data to the image of each divided echo signal obtained by the image reconstruction, inversely transform it back to the echo signal, synthesize the echo signal, and perform image reconstruction again, The connection of images can be smoothed as compared to composing images.

In each of the above embodiments, two phase encode gradient magnetic fields are combined, but three or more phase encode gradient magnetic fields may be combined, whereby an image near the center is obtained. The higher the spatial resolution, the lower the spatial resolution at both ends. Further, it is only applicable to control of the phase encoding gradient magnetic field and can be applied to pulse sequences other than the spin echo method.

The phase-encoding gradient magnetic field of this embodiment is obtained by combining phase-encoding gradient magnetic fields at equal intervals. The parameters are input from the keyboard 22, the position i and the increment Gpstep are obtained from the parameters, and the phase is calculated. The encoding gradient magnetic field may be determined. As an example, it may be obtained from the following equation. When | i | <β Gpstep = α × i / | i | | i | ≧ β Gpstep = n × α × i / | i | α, β, n are constants, α is an increase amount, β Indicates the position where the increase amount changes, and n indicates the change rate of the increase amount. For example, if α is 1, β is 2, and n is 2, then FIG.
The phase encoding gradient magnetic field d is similar to. That is, the increase amount Gpstepd of the phase encode gradient magnetic field d when the position i changes from −2 to −3 becomes −2, and the other increase amounts become 1 or −1. The value of β is 1
If the number is increased to 2 or more, the increase amount Gpstepd is 3
It will change more than once.

[0036]

According to the present invention, the step of changing the magnetic field strength of the phase encode gradient magnetic field is applied so that it is different between the low encode region and the high encode region, and a pulse sequence including this phase encode gradient magnetic field is obtained. The obtained NMR signal is divided into a plurality of NMR signal groups to reconstruct an image. The images obtained by this are images with different photographing fields of view and spatial resolutions, and desired image regions are taken out from these images and combined to obtain one image. As a result, the overall positional relationship of the image can be grasped and the spatial resolution of the region of interest can be improved. Further, since the spatial resolution is high only in the region of interest and the spatial resolution is sufficient to grasp the positional relationship in the other regions, the imaging time can be shortened as compared with the same spatial resolution and imaging field of view.

[Brief description of drawings]

FIG. 1 is a block diagram showing a configuration of an MRI apparatus including the present invention.

FIG. 2 is a diagram showing a pulse sequence of the present invention.

FIG. 3 is a diagram showing a relationship between an imaging visual field, pixels, and the number of repetitions.

FIG. 4 is a diagram showing a combination of phase encoding gradient magnetic fields.

FIG. 5 is a diagram showing image reconstruction of the first embodiment.

FIG. 6 is a diagram showing image reconstruction of the first embodiment.

FIG. 7 is a diagram showing image reconstruction of the second embodiment.

FIG. 8 is a diagram showing image reconstruction of the third embodiment.

FIG. 9 is a diagram showing a conventional pulse sequence.

[Explanation of symbols]

 1 Control CPU 2 Sequence CPU 23 90 ° pulse 24 180 ° pulse 25 Slice encoding gradient magnetic field 26 Phase encoding gradient magnetic field 27 Frequency encoding gradient magnetic field 28 Echo signal

Claims (7)

[Claims] 1. A high frequency magnetic field for exciting nuclear spins in a subject, a slice encode gradient magnetic field applied simultaneously with the high frequency magnetic field to select an imaging slice position of the subject, and a magnetic field strength applied stepwise. A phase encoding gradient magnetic field that provides phase information to the excited nuclear spins;
Sequence control means for applying a frequency-encoding gradient magnetic field that gives frequency information to the nuclear spins to which the phase information has been applied in order to read out an NMR signal according to a pulse sequence, and an NM obtained by this pulse sequence.
In a magnetic resonance imaging apparatus having a means for reconstructing an image of an R signal, the sequence control means executes a pulse sequence by changing the changing step of the magnetic field strength in a low encode region and a high encode region of phase encode. And a magnetic resonance imaging apparatus. 2. A high frequency magnetic field for exciting nuclear spins in a subject, a slice encoding gradient magnetic field applied simultaneously with the high frequency magnetic field to select an imaging slice position of the subject, and a magnetic field strength applied stepwisely. A phase encoding gradient magnetic field that provides phase information to the excited nuclear spins;
Sequence control means for applying a frequency-encoding gradient magnetic field that gives frequency information to the nuclear spins to which the phase information has been applied in order to read out an NMR signal according to a pulse sequence, and an NM obtained by this pulse sequence.
In a magnetic resonance imaging apparatus having a means for reconstructing an image of an R signal, the sequence control means executes a pulse sequence by changing the changing step of the magnetic field intensity in a low encode region and a high encode region of phase encode, The image reconstructing means divides and forms the NMR signals obtained by the pulse sequence into a plurality of NMR signal groups having different imaging fields of view and spatial resolutions during image reconstruction, and forms these NMR signal groups separately. A magnetic resonance imaging apparatus characterized by reconstructing and extracting a desired image region from each image obtained by image reconstruction, and combining these extracted image regions to obtain one image. . 3. The magnetic resonance imaging apparatus according to claim 1, wherein the step of changing the phase encode gradient magnetic field is larger in the high encode region than in the low encode region. 4. The magnetic resonance imaging apparatus according to claim 2, wherein the plurality of NMR signal groups obtained by the division are divided for each NMR signal obtained by a pulse sequence in which the change steps of phase encoding are at equal intervals. 5. A plurality of NMR signal groups used for image reconstruction include an NMR signal group formed only by NMR signals and an NMR signal group.
3. A signal group formed by adding 0 data to a signal.
The magnetic resonance imaging apparatus according to. 6. An image obtained by synthesizing is extracted from an image having a high spatial resolution in a region corresponding to the region of interest, and is extracted and synthesized from an image having a lower spatial resolution than the region of interest in the other regions. Item 2. A magnetic resonance imaging apparatus according to Item 2. 7. An image obtained by reconstructing an image of a plurality of NMR signal groups or an NMR signal group in which NMR signals and 0 data are added, and 0 data is added to areas other than a desired image area to obtain 0 data. 6. The magnetic resonance imaging apparatus according to claim 2, wherein the image added with is inversely converted into signal data, these signal data are combined, and the image is reconstructed to obtain one image.