WO2020235505A1

WO2020235505A1 - Nuclear magnetic resonance imaging device, nuclear magnetic resonance imaging method, and program

Info

Publication number: WO2020235505A1
Application number: PCT/JP2020/019524
Authority: WO
Inventors: 智之拝師; 亮平忰田; 佐々木　進
Original assignee: 国立大学法人新潟大学
Priority date: 2019-05-17
Filing date: 2020-05-15
Publication date: 2020-11-26
Also published as: JPWO2020235505A1; JP7412787B2

Abstract

A nuclear magnetic resonance imaging device (100) comprises: a magnetostatic coil (10); a holding unit (2); a pulse application unit (30); a reception unit (40); and an image generation unit (51b). The pulse application unit (30) applies a target with a pulse sequence of which echo peak appears before a half of a signal acquisition time from application of a frequency encoding gradient magnetic field for rephasing an NMR signal. The reception unit (40) detects an NMR signal over the entire range from the application of the frequency encoding gradient magnetic field for rephasing the NMR signal to the lapse of the signal acquisition time. The image generation unit (51b) generates an image from the entire NMR signal detected over the above-mentioned entire range.

Description

Nuclear magnetic resonance imaging equipment, nuclear magnetic resonance imaging methods, and programs

The present invention relates to a nuclear magnetic resonance imaging apparatus, a nuclear magnetic resonance imaging method, and a program.

Nuclear magnetic resonance imaging (MRI) is to irradiate a specific RF (Radio Frequency) pulse (excitation pulse) while applying a specific gradient magnetic field to a subject in a static magnetic field to cause nuclear magnetic resonance of a specific atom in the subject. In this method, the induced current generated in the receiving coil is acquired as a nuclear magnetic resonance (NMR) signal, and an image of the subject (for example, a two-dimensional image (that is, a cross-sectional image) or a three-dimensional image) is generated from this signal. An NMR signal to which position information is added by applying a gradient magnetic field measured by MRI is also particularly called an MRI signal. Also, the set of RF pulses and gradient magnetic fields applied at a particular intensity and timing to acquire an NMR signal is called a pulse sequence.

The above-mentioned gradient magnetic field connects the real space in which the subject is placed and the spatial frequency space (k-space) of the subject by the relationship of Fourier transform, and the above-mentioned MRI signal reflects the information in the k-space of the subject. doing. Therefore, in MRI, information in the k-space of the subject is discretely collected from the MRI signal, and the obtained discrete data is subjected to a discrete inverse Fourier transform to reconstruct the image of the subject in the real space.

As a method of collecting k-space data from the MRI signal, a triaxial gradient magnetic field (X-axis, Y-axis, Z-axis) is controlled by a triaxial gradient magnetic field coil to control a lattice of the Cartesian coordinate system of the subject in k-space. A method of extracting data on points one by one (line scan), and a method of sequentially extracting data on the polar coordinate system of the subject's k space along a plurality of radial straight lines or spiral curves passing through the origin (radial scan). / Spiral scan) etc. have been put into practical use.

The MRI signal has a convex part with strong signal strength generated by manipulating the gradient magnetic field by the designed pulse sequence, and this is called an echo. The echo and the signal changes before and after it include the spatial information of the subject. Echoes caused by the application of a gradient magnetic field are called gradient echoes. An echo caused by continuous application of a high-frequency magnetic field (for example, application of an inversion pulse after application of an excitation pulse) is called a spin echo. Regardless of the cause of the echo, the time from the application of the excitation pulse to the generation of the echo is called the echo time (TE).

Since the distribution, density, and relaxation time of the target NMR nuclei in the subject differ depending on the situation, research and development of MRI signal integration methods (particularly pulse sequences) suitable for the target are progressing day by day. There is.

For example, as one of these methods, the partial echo method is known. The partial echo method, gradient echo (GRE) method, to shorten the TE, the data corresponding to about half of k-space by obtains the MRI signal until the elapse of the time of half the signal acquisition time t _a from the TE Is collected, and the rest of the data in k-space is appropriately corrected (for example, 0 fill (zero fill), complementation by duplication based on Elmeet symmetry), and then the inverse Fourier transform is performed to perform MRI. Generate an image. FIG. 5 shows a pulse sequence when the partial echo method is applied in the GRE method. For example, Patent Document 1 discloses an example of performing correction with 0 fill.

U.S. Pat. No. 6,166,545

In conventional MRI, to generate a cross-sectional image of the subject based on the presence or absence of abundant in subjects such as bio ^{1 H} (proton). However, with the aim of improving diagnostic accuracy and expanding research subjects, it is desired to develop an MRI apparatus that targets other NMR nuclei. For example, with the expectation that sodium in the living body will be visualized and pathological conditions related to the brain and kidneys will be detected at an early stage, attempts to develop a ²³ Na-MRI apparatus have begun.

^However, it is known that the NMR signals of 23 Na is less sensitive than the NMR signal of ¹ H (proton). For example, as shown in FIG. 7, when physiological saline (0.9% NaCl aqueous solution, 2 ml) was measured, the NMR signal of ²³ Na was about 1/20000 of the NMR signal of ¹ H.

Moreover, since the abundance of many NMR nuclei other than ^{1 H} in the object containing the living body is smaller than the ^{1 H,} it is difficult to obtain the MRI signal having a signal strength required to capture the image.

^{1 In} order to obtain an MRI signal with a practical signal strength in an MRI targeting an atom other than H, the integrated signal obtained by simply acquiring and integrating the MRI signals corresponding to the same image multiple times. However, when the signal strength is low such as ²³ Na, it is necessary to select a magnet with a high static magnetic field strength to increase the NMR signal strength and greatly increase the number of repetitions of signal collection. Yes, this can exceed the measurement time that the patient can tolerate. Therefore, in order to keep the measurement time within the permissible limit while increasing the number of repetitions, the repetition time (TR) required for acquiring one MRI signal, particularly the echo time required from the application of the excitation pulse to the generation of the echo of the MRI signal. It is necessary to make (TE) as short as possible.

However, in the partial echo method as in Patent Document 1, since the k-space is composed of half of the captured data, there is a risk that false image artifacts are often mixed. In particular, in the case of MRI imaging by the gradient echo method, the Hermitian symmetry becomes incomplete due to the inhomogeneity of the spatial static magnetic field and the like, and the problem of false image artifacts is remarkable.

Further, in order to put ²³ Na-MRI to practical use, further improvement of the S / N ratio in the NMR signal of ²³ Na has been desired. ²³ Na sodium is an NMR nucleus with a spin quantum number of 3/2. Moreover, in the case of ¹ H (proton), it often becomes a large molecule or complex in practice due to the hydrogen bond of H ₂ O, but ²³ Na exists as almost a single ion in the living body. It is thought that there is. For this reason, T2 relaxation time of ²³ Na is as short as several ms (milliseconds) to about 30 ms. Due to these two characteristics, in the case of ²³ Na-MRI, ¹ H (proton) (spin quantum number 1/2, main T2 relaxation time is several tens of ms to several hundreds) contained in the water and lipid of the living body. It was difficult to simply apply the methodology of image contrast generation in the imaging theory of ms).

In view of the above, it is an object of the present invention to provide a nuclear magnetic resonance imaging apparatus, a nuclear magnetic resonance imaging method, and a program capable of suitably capturing an MRI image.

The nuclear magnetic resonance imaging apparatus according to the first aspect of the present invention is
A static magnetic field forming part that forms a static magnetic field,
An object holding unit that holds an object in the static magnetic field,
A pulse sequence including an excitation pulse, a phase-encoded gradient magnetic field, and a frequency-encoded gradient magnetic field is applied to the object in the static magnetic field, and an NMR signal is generated from the object by applying the excitation pulse, and the frequency-encoded gradient magnetic field is generated. A pulse application unit that dephases the NMR signal and then rephases it by applying
A detection unit that detects each of the NMR signals phase-encoded by the phase-encoded gradient magnetic field of different amplitudes while the frequency-encoded gradient magnetic field is applied by the pulse application unit to rephase the NMR signal. ,
An image generation unit that generates an image from the NMR signal detected by the detection unit,
With
The pulse application unit applies the pulse sequence to the target so that the echo peak comes before half of the signal acquisition time from the application of the frequency-encoded gradient magnetic field for rephase of the NMR signal.
The detection unit detects the NMR signal over the entire range from the application of the frequency-encoded gradient magnetic field for rephase of the NMR signal to the elapse of the signal acquisition time.
The image generation unit generates the image from the entire NMR signal detected over the entire area.
It is characterized by that.

The nuclear magnetic resonance imaging method according to the second aspect of the present invention is
A pulse sequence including an excitation pulse, a phase-encoded gradient magnetic field, and a frequency-encoded gradient magnetic field is applied to an object in a static magnetic field, an NMR signal is generated from the object by applying the excitation pulse, and the frequency-encoded gradient magnetic field is applied. A pulse application step in which the NMR signal is dephased and then rephased by
A detection step of detecting each of the NMR signals phase-encoded by the phase-encoded gradient magnetic field having different amplitudes while the NMR signal is being rephased by applying the frequency-encoded gradient magnetic field by the pulse application step. ,
An image generation step of generating an image from the NMR signal detected in the detection step, and
With
In the pulse application step, the pulse sequence is applied to the target so that the echo peak comes before half of the signal acquisition time from the application of the frequency-encoded gradient magnetic field for rephase of the NMR signal.
In the detection step, the NMR signal is detected over the entire range from the application of the frequency-encoded gradient magnetic field for rephase of the NMR signal to the elapse of the signal acquisition time.
In the image generation step, the image is generated from the entire NMR signal detected over the entire area.
It is characterized by that.

The program according to the third aspect of the present invention
Computer,
A pulse sequence including an excitation pulse, a phase-encoded gradient magnetic field, and a frequency-encoded gradient magnetic field is applied to a target in a static magnetic field to a pulse application unit, and an NMR signal is generated from the target by applying the excitation pulse to generate an NMR signal at the frequency. A pulse application means that dephases the NMR signal by applying an encode gradient magnetic field and then rephases the NMR signal, and echo peaks before half of the signal acquisition time from the application of the frequency encode gradient magnetic field for rephase the NMR signal. A pulse application means for applying the pulse sequence to the target.
Each of the NMR signals phase-encoded by the phase-encoded gradient magnetic field of different amplitude is detected in the detection unit while the frequency-encoded gradient magnetic field is applied by the pulse application unit to rephase the NMR signal. A detection means for detecting the NMR signal over the entire range from the application of the frequency-encoded gradient magnetic field for rephase the NMR signal to the elapse of the signal acquisition time.
An image generation means for causing an image generation unit to generate an image from the NMR signal detected by the detection unit, and for generating the image from the entire NMR signal detected over the entire area.
To function as.

According to the present invention, an MRI image can be preferably taken.

It is a block diagram which shows the structure of the MRI apparatus which concerns on one Embodiment of this invention. It is a schematic diagram which shows the pulse sequence which concerns on one Embodiment of this invention. It is a flowchart of image generation processing which concerns on one Embodiment of this invention. It is a schematic diagram which shows the pulse sequence of the conventional GRE method. It is a schematic diagram which shows the pulse sequence of a partial echo method. It is a schematic diagram which shows the quantum pulse. It is a graph which shows the intensity of ¹ H-NMR signal of saline solution and ²³ Na-NMR signal with respect to the repetition time. ²³ It is a cross-sectional image of the vicinity of the kidney of a mouse by Na-MRI. (A) shows a cross-sectional image of the surgery group, and (b) shows a cross-sectional image of the control group. It is a graph which shows the intensity of ¹ H-NMR signal and ²³ Na-NMR signal of the saline solution measured at the same measurement time with respect to the repetition time. It is a photograph of the measuring table of the ¹ H / ²³ Na-dual MRI apparatus. (A) is a photograph of the measuring table taken from above, and (b) is a photograph of the bottom plate removed from the measuring table and taken from the back side.

An embodiment of the present invention will be described with reference to the drawings.

(Configuration of MRI apparatus 100)
As shown in FIG. 1, the MRI apparatus 100 according to the embodiment of the present invention includes a static magnetic field coil 10, a gradient magnetic field generating unit 20, an RF pulse applying unit 30, a receiving unit 40, and a control device 50. A display unit 60 and an operation unit 70 are provided.

The static magnetic field coil 10, the gradient magnetic field coil 21 included in the gradient magnetic field generation unit 20, and the RF (Radio Frequency) coil 31 included in the RF pulse application unit 30 are arranged around the coaxial (Z axis), for example. , It is provided in a housing (not shown).

The object 1 to be photographed is held in the bore 3 (inspection space) in the housing by the holding unit 2. For example, if the subject 1 is a human, the holding unit 2 may be a sleeper. The holding unit 2 moves, for example, the object 1 in the bore 3 according to the imaging site (for example, horizontally moves, vertically moves, or rotates), or moves the object 1 from outside the bore 3 into the bore 3. It may be provided with a transport means for causing. Further, the subject 1 may be a part instead of the whole, and may be not only the whole body of a human being but also a part thereof (for example, the head or the abdomen). It is preferable that the shape of the MRI apparatus 100, particularly the structure around the bore 3, for example, the gradient magnetic field coil 21 and the RF (Radio Frequency) coil 31 is optimized for the shape of the object 1.

The static magnetic field coil 10 forms a uniform static magnetic field (for example, 1.5 tesla to 21 tesla) at a level of several ppm in the bore 3. For example, the static magnetic field coil 10 may be formed so that the central axis of the static magnetic field coil 10 or the central axis of the magnetic flux passes through the bore 3 (particularly, the central axis of the bore 3). The static magnetic field formed is a horizontal magnetic field substantially parallel to the Z direction. The static magnetic field coil 10 is composed of, for example, a superconducting coil or a normal conducting coil, and is driven under the control of the control device 50 via a static magnetic field coil driving unit (not shown). The static magnetic field coil 10 may be driven and controlled by a control system independent of the control device 50. Further, the configuration for generating the static magnetic field is not limited to the superconducting coil and the normal conducting coil, and for example, a permanent magnet (for example, 1 tesla or less) may be used. The static magnetic field coil 10 may have a shim coil group (not shown) for correcting the static magnetic field uniformity, and the static magnetic field can be further made uniform by such a shim coil group. If a uniform static magnetic field can be formed, an arbitrary static magnetic field forming portion can be adopted instead of the static magnetic field coil 10.

The gradient magnetic field generating unit 20 generates an independent three-axis gradient magnetic field necessary for imaging in the bore 3, and includes a gradient magnetic field coil 21 and a gradient magnetic field coil driving unit 22 for driving the gradient magnetic field coil 21. , Have.

The gradient magnetic field coil 21 generates a gradient magnetic field that gives a gradient to the strength of the static magnetic field formed by the static magnetic field coil 10 in three axial directions orthogonal to each other. Therefore, the gradient magnetic field coil 21 has three systems of coils. The gradient magnetic fields in the three axes orthogonal to each other are, for example, a slice gradient magnetic field in the slice axis direction, a phase encode gradient magnetic field in the phase axis direction, or a frequency encode gradient in the frequency axis direction, depending on the pulse sequence used for imaging. Used as a magnetic field. The slice gradient magnetic field is a gradient magnetic field for slice selection. The phase-encoded gradient magnetic field and the frequency-encoded gradient magnetic field are gradient magnetic fields for measuring the spatial distribution of resonant elements. In some cases, the gradient magnetic field in one direction plays two or more roles in the same pulse sequence.

The gradient magnetic field coil drive unit 22 supplies a drive signal to the gradient magnetic field coil 21 to generate a gradient magnetic field under the control of the control device 50. The gradient magnetic field coil drive unit 22 has three systems of drive circuits (not shown) corresponding to the three systems of coils included in the gradient magnetic field coil 21.

In FIG. 1, the left-right direction of the paper surface is the X-axis direction, the vertical direction of the paper surface is the Y-axis direction, and the normal direction of the paper surface is the Z-axis direction, and the gradient axes of the above-mentioned gradient magnetic field are parallel to these X, Y, and Z axes. Although drawn as being applied, the gradient axes of the three gradient magnetic fields (for example, the slice axis, the phase axis, and the frequency axis) are limited to the above-mentioned X, Y, and Z axes as long as they are orthogonal to each other. It does not have to match some or all of these.

The RF pulse application unit 30 is for applying an RF pulse for generating nuclear magnetic resonance to the object 1, and has an RF coil 31 and an RF coil drive unit 32 for driving the RF coil 31.

The RF coil 31 forms a high-frequency magnetic field in the static magnetic field space for exciting the nuclear spins of the target NMR nucleus in the target 1. Forming a high-frequency magnetic field in this way is also referred to as application or transmission of RF pulses. The RF coil 31 has a function of transmitting an RF pulse and a function of receiving a nuclear magnetic resonance (NMR) signal which is an electromagnetic wave generated by an excited nuclear spin. An NMR signal to which position information is added by applying a gradient magnetic field measured by MRI is called an MRI signal. Instead of the RF coil 31, the RF pulse transmission coil and the MRI signal reception coil may be separately configured.

Further, it is desirable that the RF coil 31 or the RF pulse transmission coil is large, and the MRI signal reception coil is small, and it is desirable that the size is the minimum necessary. The smaller the RF coil 31 or the RF pulse transmission coil, the smaller the excitation power (SAR: Specific Absorption), such as when the number of repetitions is increased and the coil is continuously applied in a relatively short time. Rate) can be lowered). The smaller the RF coil 31 or the RF pulse receiving coil is, the more it is possible to prevent receiving an MRI signal from an unnecessary portion (for example, a portion that is not imaged in the object 1). In particular, it is desirable to narrow the receiving region of these receiving coils to eliminate the need for selective excitation pulses so that the repetition time and echo time can be shortened.

The RF coil drive unit 32 supplies a drive signal to the RF coil 31 to drive the RF coil 31 under the control of the control device 50. Specifically, the RF coil driving unit 32 generates an RF pulse corresponding to the Larmor frequency determined by the type of target atom and the magnetic field strength in the RF coil 31 as an excitation pulse. The excitation pulse may be a hard pulse, a Gaussian pulse, or an adiabatic pulse, but it is particularly desirable to employ a hard pulse because the echo time (TE) can be shortened. Further, in order to generate so-called T1 image contrast, a 180 ° pulse (commonly known as an Inversion Recovery pulse) that inverts the nuclear magnetization may be preceded, and a few ms of IR time may be arranged as a waiting time. This makes it possible to reduce the signal of nuclear magnetization having a specific T1 time (not shown).

The receiving unit (detecting unit) 40 is connected to the RF coil 31 and detects the MRI signal received by the RF coil 31. The receiving unit 40 digitally converts the detected MRI signal and transmits it to the control device 50. It is desirable that unnecessary signals outside the MRI signal necessary for image reconstruction are removed by an analog or digital frequency filter in the receiving unit 40 or the control device 50. For example, a steep bandpass filter may be used as the analog frequency filter.

The control device 50 includes a control unit 51 and a storage unit 52, and for example, drives a computer that controls the overall operation of the MRI device 100, an RF coil drive unit 32, and a gradient coil drive unit 22 by a pulse sequence. It consists of a sequencer.

The control unit 51 is composed of a CPU (Central Processing Unit), an FPGA (Field-Programmable Gate Array), an ASIC (Application Specific Integrated Circuit), etc., and executes an operation program stored in the storage unit 52 to execute an MRI apparatus. Control the operation of each part of 100.

The storage unit 52 is composed of a ROM (Read Only Memory), a RAM (Random Access Memory), etc., holds a CPU as needed, and provides various operation programs (program PG for executing image generation processing described later). (Including) data, table data described later, etc. are stored in advance. The RAM of the storage unit 52 temporarily stores data indicating various calculation results, data indicating discrimination results, and the like, and performs calculations. The storage unit 52 can also be arranged on the Internet.

The control unit 51 includes a pulse control unit 51a and an image generation unit 51b as functional units.

The pulse control unit 51a includes the gradient magnetic field generation unit 20 and the RF pulse application unit 30 based on the data of the pulse sequence 300 described later showing the pulse sequences of the RF pulse, the phase-encoded gradient magnetic fields Gp1 and Gp2, and the frequency-encoded gradient magnetic field Gr. Drive control of. The data of the pulse sequence 300 may be stored in the storage unit 52 in advance, or may be input or reset by a user operation via the operation unit 70 described later.

The image generation unit 51b stores the digitally converted MR signal data collected by the reception unit 40 in the storage unit 52. The data constitutes a three-dimensional Fourier space (k space) by the gradients of the phase-encoded gradient magnetic fields Gp1 and Gp2 and the frequency-encoded gradient magnetic field Gr. The image generation unit 51b generates a three-dimensional image of the target 1 by performing a three-dimensional inverse Fourier transform on the data in the k-space.

The display unit 60 displays a waveform showing the intensity of the MRI signal over time that can confirm the echo peak position while scanning the k space, the MRI image generated by the image generation unit 51b, and various other information. The display is composed of, for example, a liquid crystal display (LCD: Liquid Crystal Display), an organic EL display (OELD: Organic Electro Luminescent Display), and the like. When the MRI apparatus 100 is configured to enable remote work, a personal computer (for example, a notebook computer) or another portable electronic terminal may be used as the display unit 60 and the operation unit 70 described later.

The operation unit 70 receives an operation by the user and supplies an operation signal corresponding to the received operation to the control device 50. The operation unit 70 is composed of, for example, a keyboard and a mouse provided with a pointing device, a touch panel integrated with the display unit 60, and the like. For example, the MRI apparatus 100 is configured so that the user can input data of the pulse sequence 300 or the like by an operation from the operation unit 70.

(Image generation processing)
The image generation process will be described with reference to FIG. The image generation process is executed by, for example, the control unit 51 in response to a start operation from the operation unit 70 by the user. Before the start of the image generation process, it is assumed that the target 1 mounted on the holding unit 2 is set in the bore 3.

When the image generation process is started, first, the control unit 51 drives the static magnetic field coil 10 via a static magnetic field coil drive unit (not shown) to form a uniform static magnetic field in the bore 3 (step S11). A static magnetic field may be formed in the bore 3 in advance before the image generation process. The uniformity of the static magnetic field is affected by the composition and shape of subject 1. Therefore, the static magnetic field coil 10 may be provided with a shim coil group (not shown) for correcting the uniformity of the static magnetic field to further make the static magnetic field more uniform.

Next, the control unit 51 drives and controls the gradient magnetic field generation unit 20 and the RF pulse application unit 30 according to the following pulse sequence 300 (pulse application step), and inputs the data in the k-space of the target 1 via the reception unit 40. It is acquired (detection step) and stored in the storage unit 52 (in total, step S12).

(Pulse Sequence 300)
The pulse sequence 300 used in this embodiment is shown in FIG. The pulse sequence 300 is the same as the pulse sequence of the gradient echo (GRE: Gradient Echo) method except for the position of TE (echo time). In FIG. 2, “RF” is a sequence of α ° pulses 311 which are excitation pulses, “Gp1” is a first phase-encoded gradient magnetic field, “Gp2” is a second phase-encoded gradient magnetic field, and “Gr” is a frequency-encoded gradient. Magnetic field (lead-out gradient magnetic field), "S" indicates an MRI signal. The first phase-encoded gradient magnetic field, the second phase-encoded gradient magnetic field, and the frequency-encoded gradient magnetic field are orthogonal to each other and are, for example, parallel to the X-axis, Y-axis, and Z-axis, respectively. Signal acquisition time t _a is the time width of selecting the data to be used for image reconstruction. Signal acquisition time t _a is equal to or application time of the frequency encoding gradient magnetic field 342, a short time is used.

First, α ° excitation of spin is performed by α ° pulse 311. The flip angle α ° is 90 ° or less.

After α ° excitation, phase encoding and frequency encoding are performed while the free induction decay (FID) signal 351 is generated.

In the phase encoding, the first and second phase encoding gradient

magnetic fields

321 and 331 are applied while sequentially changing the amplitude for each application of the α ° pulse 311 (application of a series of α ° pulses 311 if signal integration is performed). It is done by. For example, the first phase-encoded gradient magnetic field 321 is set to N1 steps (for example, 128 steps), the second phase-encoded gradient magnetic field 331 is set to N2 steps (for example, 32 steps), and the number of repetitions (number of signal integrations) is N times. When (N is 1 or more), a series of pulse sequences in which α ° pulses are applied N times for signal integration and the same phase encoding and frequency encoding are performed for each application is repeated N1 × N2 times. If the repetition time (TR) is counted as one measurement, the measurement is repeated N × N1 × N2 times. The signal integration may be in the order of repeating the measurement of N1 × N2 N times.

In parallel with each phase encoding, a frequency-encoded gradient magnetic field 341 is applied to dephase (flyback) the macroscopic nuclear magnetization, and then a frequency-encoded gradient magnetic field 342 is applied to rephase (refocus) the spin. Then, the MRI signal 352 of the gradient gradient echo is generated. The signal strength becomes maximum after TE from α ° excitation.

The integrated value of the frequency-encoded gradient magnetic field 341 for dephase (the stippled area of the frequency-encoded gradient magnetic field 341 in FIG. 2) is the integrated value from the application of the frequency-encoded gradient magnetic field 342 to TE (the integrated value of the frequency-encoded gradient magnetic field 342 in FIG. 2). It is equal to the stippling area). Therefore, by changing the integral value of the frequency-encoded gradient magnetic field 341 for dephasé (for example, depending on the application period, intensity, and / or waveform) without changing the frequency-encoded gradient magnetic field 342, the TE generation time is generated. Can be adjusted. Conventionally, as shown in FIG. 4, but TE frequency encoding gradient magnetic field 343 such that the center of the signal acquisition time t _a frequency encode gradient magnetic field 342 is adjusted, in the present embodiment, as shown in FIG. 2 a, TE (i.e., echo peak), before the half from the application of the signal acquisition time t _a frequency encoding gradient magnetic field 342, in particular, from the tenth of the start of the signal acquisition time t _a up to one third The frequency-encoded gradient magnetic field 341 is adjusted so that it is in between. Although the start time and the signal acquisition starting time of the time t _a the application of Figure 2 the frequency encoding 342 is in the same, intended time timing the current value of the frequency encoding 342 falls upstanding or stand up to the set value ing. This rise (fall) delay is due to the characteristics of the gradient magnetic field coil drive unit 22, and is fixedly adjusted from 50 μs to 500 μs in many cases. Ideally, it is expected that there will be no delay, so the description is omitted. In FIG. 2 and 4, application end time of the frequency encoding 342, the object 1 has been the same as the signal acquisition time t _a which is determined based on the width of performing frequency encoding, to application end from the start of application of the frequency encoding 342 application time is arbitrary as long as the signal acquisition time t _a or more. Frequency encoding 342 is applied after the signal acquisition time t _a has the effect of promoting the relaxation recovery of the nuclear magnetization signal of an object.

The MRI signal 352 thus obtained is digitized via the receiving unit 40. Controller 51 over the MRI signal 352 is digitized for each repetition time to the full width of the signal acquisition time _{t a} frequency encoding gradient magnetic field 342 M points (e.g., 128 points) to get the data of the storage unit 52 Remember in. Thereafter, the control unit 51 corresponds to the same position of the object 1 among the data (e.g., first and second phase-encoding gradient

magnetic field

321 and 331 amplitude and frequency encoding gradient magnetic field 342 signal acquisition time in t _a of (The combination of the positions is the same) N signal strengths are integrated, the integrated signal strength at the position is calculated, and stored in the storage unit 52.

Finally, the control unit 51 generates an image from the k-space data of the target 1 obtained as described above (for example, the k-space data of the target 1 consisting of N1 × N2 × M data). (Step S13). Specifically, the data in the k-space is subjected to a three-dimensional discrete inverse Fourier transform to reconstruct the three-dimensional image of the object 1. Due to the recent development of information technology, MRI image reconstruction may be carried out by artificial intelligence AI or deep learning instead of inverse Fourier transform. These may be applied to the reconstruction of the three-dimensional image of the object 1.

(Effect of this embodiment)
In conventional GRE method, as shown in FIG. 4, but in the center of the peak signal acquisition time _{t a} of the MRI signal, the low-sensitivity NMR nuclei ^(e.g., 23 Na) in MRI targeted to, the MRI signal The intensity was reduced by T2 and T2 ^* attenuation, easily mixed with noise, and difficult to detect.

On the other hand, in the MRI apparatus 100 according to the present embodiment, the signal can be captured with priority given to the echo peak having rough image contour information. In particular, when ²³ Na is targeted for MRI, the signal component of ²³ Na Among them, a really fast signal component having a T2 attenuation time of 2 to 3 ms or less can be used for image generation. Therefore, according to the present embodiment, it is possible to suitably generate an MRI image even in an MRI targeting a low-sensitivity atom.

Further, in the partial echo method, as shown in FIG. 5, although the TE and the measurement time can be shortened, there is a problem of mixing false image artifacts due to the configuration of the k space. Further, when the signal is collected by the GRE method, the partial echo method is difficult to apply because the phase variation occurs due to T2 ^* attenuation. In addition, the partial echo method could not be combined with other MRI measurement methods (eg, compressed sensing methods) due to principle problems.

On the other hand, in this embodiment, present in the first half of the peak of the MRI signal is the signal acquisition time t _a (in particular, between one-tenth of the signal acquisition time t _a third of) as shown in FIG. 2, MRI since the high low-frequency component as compared with the signal intensity is high-frequency components in the second half of the attenuation is large signal acquisition time t _a signal strength is present, while giving priority to image generation, a narrow and low-noise pixel bandwidth It has been realized. Since the homodyne reconstruction with the phase encoding removed is not performed, the occurrence of false image artifacts is reduced. Therefore, according to the present embodiment, an MRI image can be preferably generated as compared with the partial echo method. Further, the image generation processing method according to the present embodiment can be combined with various MRI measurement methods as shown in the following modified examples.

(Modification example)
The present invention is not limited to the above embodiments and drawings. Changes (including deletion of components) can be made as appropriate without changing the gist of the present invention. Further, in the present specification, in order to facilitate the understanding of the present invention, the description of known technical matters is appropriately omitted.

(Modification example 1)
In the above-described embodiment, the GRE method is adopted, but a pre-RF pulse may be further laid. This makes it possible to obtain an image having contrast according to the type of pre-RF pulse. In the case of NMR nuclei with a short relaxation time (for example, ²³ Na), it is desirable to lay a pre-RF pulse for each excitation pulse. Further, in the case of an NMR nucleus having a long relaxation time (for example, ¹ H), the pre-RF pulse may be laid for each excitation pulse, or for each of a plurality of excitation pulses (for example, when the signal is integrated N times). If so, it may be laid (before the first one of N excitation pulses).

For example, when ²³ Na is used as an NMR nucleus, if a signal is acquired within the relaxation time, a density-weighted image can be obtained, but by laying a pre-RF pulse, an image with contrast other than density-weighted can be obtained. Be done. ^{When 23} Na is used as an NMR nucleus, as described in Example 1, a contrast other than density enhancement cannot be obtained by simply changing the repetition time as in the case where ¹ H is used as an NMR nucleus. It is advantageous to give contrast by laying a pre-RF pulse.

Further, at this time, a quantum pulse may be applied as a pre-RF pulse as shown in FIG. According to such a quantum pulse, it is expected that a tissue-specific contrast can be obtained by the transition of spin energy. If the conventional ^{1 H-MRI,} because the relaxation time of several hundred ms and long from a few tens of ms, was possible the discrimination of using the differences and chemical shift excitation of relaxation time organization. However, in the case of NMR nuclei having a short relaxation time (for example, in the case of ²³ Na having a short relaxation time of 20 ms or less), such discrimination is impossible. According to the quantum pulse, even in such an NMR nucleus with a short relaxation time, the transition of spin energy makes it possible to emphasize the tissue difference before the signal is attenuated.

In the quantum pulse, after applying a 90 ° pulse (+ X direction), a 180 ° pulse is applied after the lapse of time τ, and then a 180 ° pulse is continuously applied every 2τ of time. At this time, the continuous 180 ° pulse is referred to as a pulse train (APCP (Alternating Polarity Carr Purcell)) in which 180 ° pulses of opposite polarity (−X direction) are continuously applied from the second pulse, as shown in FIG. S.Watanabe and S.Sasaki.J.Jpn.Appl.Phys.Express Lett. (2003)) or a pulse train ((CPMG) in which 180 ° pulses (+ Y direction) of the same polarity are continuously applied. (Pulse train called Carr-Purcell-Meiboom-Gill) may be used.

For MRI signal generation, the NMR signal generated by the pre-RF pulse and the subsequent excitation pulse may be a free induction decay (FID) signal or a spin echo signal. Further, the NMR signal generated by the pre-RF pulse does not have position information added by phase encoding and frequency encoding before the excitation pulse is applied, and the pulse sequence according to the above-mentioned pulse sequence 300 or other modifications is provided. Therefore, after the excitation pulse is applied, the position information is added by phase encoding and frequency encoding.

(Modification 2)
In the above-described embodiment, the three-dimensional Fourier transform is performed, but the phase encoding in the Y direction is omitted, and instead, slice selection is performed along the Y direction, and the MRI image is obtained by the two-dimensional Fourier transform for each slice. It may be generated. Further, the reconstruction of the MRI image may be performed by artificial intelligence AI or deep learning instead of the inverse Fourier transform.

(Modification 3)
When the echo peak is too close to the start time of the signal acquisition time t _a frequency encoding 342, the signal strength of the MRI signal measured at the starting point and the end point of the signal acquisition time t _a discrepancy greater, the same waveform is periodically continuously There is a possibility that the premise of the discrete inverse Fourier transform that it is In this case, for example, a false image artifact such as a cast ringing false image (multiple edges can be formed on the edge) occurs in the MRI image. Therefore, it echoes peaks are preferably separated from the start time of the signal acquisition time of the frequency encoding gradient magnetic field 342 t _a to the extent that such a false image artifacts do not occur. The presence or absence of the false image artifact can be confirmed by visually observing the completed MRI image, and the MRI image generated according to the above-described embodiment and the position of the echo peak are the same as in the conventional method for the same part of the same object 1. in the reference image with a focus of the signal acquisition time t _a, compared with a correlation coefficient (e.g., ZNCC (Zero-mean normalized cross -correlation) by the obtained normalized cross-correlation coefficient), the correlation coefficient It can also be determined whether or not it exceeds a certain value (for example, 0.7 for the normalized cross-correlation coefficient obtained by ZNCC). Further, based on the result of the preliminary experiment using the object 1 or the reference object of the same type as the object 1, the position of the echo peak in step S12 of the image generation process is not visually confirmed as a false image artifact or the above-mentioned correlation. It is preferable to adjust the number so that it becomes a certain value or more. False images may be removed by artificial intelligence AI or deep learning.

(Modification example 4)
When the object 1 is a sample that is uniform in the thickness direction or a sample that is highly uniform in the thickness direction (for example, when the object 1 is a thin film material), frequency encoding is performed in the thickness direction (for example, the direction perpendicular to the thin film). A two-dimensional MRI image of the cross section of the sample in the thickness direction is obtained by applying the first and second phase-encoded gradient magnetic fields in the direction perpendicular to the gradient magnetic direction and the thickness direction and perpendicular to each other (for example, the direction parallel to the thin film). It may be generated, and in this case, the echo peak position may be at the same time as the start of signal collection or before the start of signal collection. In this sample, applying a frequency encoding gradient magnetic field in the thickness direction (i.e., to set the signal reading-axis thickness direction) case, the peak around the MR signal draws a very loose chevron, for the signal acquisition time t _a The importance of the echo peak position is relatively low.

(Modification 5)
In the above-described embodiment, the repetition time is preferably T2 relaxation time or less in order to suppress signal attenuation. For example, when the NMR nucleus is ²³ Na, the repetition time is preferably 30 ms, 20 ms, 10 ms, or 6 ms or less.

Further, in the embodiment described above, regardless of the length of the signal acquisition time t _a, the position of the echo peaks, is between (T2 relaxation time to lapse T2 relaxation time which is a main component that can be observed from the application of the excitation pulse a plurality If it is composed of components, it preferably occurs between the application of the excitation pulse and the longest T2 relaxation time, the shortest T2 relaxation time, or the elapse of the T2 relaxation time between them), and in particular, the excitation pulse. It may occur within one-half, one-third, or one-tenth of the T2 relaxation time (or any of those T2 relaxation times if the T2 relaxation time consists of multiple components) from the application of. preferable. If the application of the pre-excitation pulse intentionally extends the signal lifetime of the short relaxation time component, refer to the apparently longer T2 relaxation time.

(Modification 6)
In the above-described embodiment, the NMR nucleus is arbitrary, and may be a nucleus having a spin quantum number of 3/2 or ²³ Na.

(Modification 7)
Any other MRI technique can be combined with the above embodiments. For example, a compressed sensing method or a multi-quantum coherence method may be adopted. In particular, the compressed sensing method does not apply to the partial echo method because it selects the phase encoding as Gaussian-random.

(Modification 8)
If the pulse sequence 300 described above or the pulse sequence according to the modification described above can be applied to the object 1, the configuration of the MRI apparatus 100 is arbitrary, and various configurations can be adopted.

(Modification 9)
The program PG that executes the image generation process is stored in the storage unit 52 in advance, but may be distributed and provided by a detachable recording medium. Further, the program PG may be downloaded from another device connected to the MRI apparatus 100. Further, the MRI apparatus 100 may execute each process according to the program PG by exchanging various data with other devices via a telecommunication network or the like.

(Modification example 10)
RF coil 31 to obtain signals derived from two or more types of atoms (eg, ¹ H-NMR signal and ²³ Na-NMR signal), for example, to enable simultaneous imaging of two types of atoms. As a result, a coil capable of transmitting and receiving two or more types of RF pulses having different frequencies corresponding to two or more types of atoms may be adopted. For example, such coils include a single transmit / receive coil tuned to the Lamore frequency of two or more atoms, such as a double tuned bird cage coil, a double tuned surface coil, a double tuned saddle coil, a double tuned butterfly. A mold coil, a double tuning solenoid coil, or the like may be used, or two or more transmission / reception coils individually tuned to the Lamore frequency of two or more kinds of atoms may be used.

Further, when the RF pulse transmission coil and the MRI signal reception coil are separately configured instead of the RF coil 31, the RF pulse transmission coil is a single coil tuned to the Lamore frequency of two or more kinds of atoms. A transmission coil of the above or two or more transmission coils individually tuned to the Lamore frequency of two or more types of atoms may be used, and the coil for receiving MRI signals is tuned to the Lamore frequency of two or more types of atoms. A single receiving coil or two or more receiving coils individually tuned to the Lamore frequency of two or more atoms may be used.

When the above-mentioned coil is used as the RF coil 31 or the MRI signal receiving coil, different atoms are used so that interference is suppressed between RF receiving cables for different atoms and a high S / N ratio can be obtained in a compact space. The coil and the accompanying resonance circuit and RF receiving cable (coaxial cable) may be arranged. For example, when there are two types of atoms to be measured, the resonance circuit for one atom (however, the part excluding the coil) and the RF receiving cable are the resonance circuit for the other atom (however, the part excluding the coil) and RF. From a single coil or two coils with different inductances arranged in close proximity (eg, overlapping or coaxially), arranged so that they do not overlap the receiving cable, in opposite directions, as shown in FIG. It is preferably arranged so as to extend.

(Example)
(Example 1: Comparison of sensitivity between ^{1 1} H-NMR signal and ²³ Na-NMR signal)
A comparison of the NMR signal intensity, similar diameter and ^{shape, 1} for H (9.4 tesla) or ²³ for Na (9.4 tesla) RF coil (inner diameter 15 mm, the Helmholtz-type gap 10 mm, 2 vol x2) A plastic test tube containing 2 ml of saline solution (0.9 w / v%) was placed in the coil using an NMR apparatus equipped with the above, and an excitation pulse was applied to the saline solution to measure FID signals. In order to measure the absolute intensity of the NMR signal, a continuous sine wave with an intensity of -120 dBm to 0 dBm is input to the preamplifier as a pseudo signal from a separately prepared high-frequency transmitter so that the intensity of the NMR signal collected in advance is obtained. By adjusting the output of the high-frequency transmitter, the signal strength (dBm) of the ¹ H and ²³ Na NMR signals after the RF coil was measured, respectively. The measurement results are shown in FIG. The intensity of the ¹ H-NMR signal derived from this sample was approximately 20,000 times that of the ²³ Na-NMR signal derived from the same sample.

It can be seen that the ²³ Na-NMR signal on the left side of FIG. 7 (TR <50 ms) is depressed due to T1 saturation due to fast repeating TR, but ²³ Na exists as a single atom and is also present with other atomic molecules. Since it is not coupled, the transfer of relaxation energy to the surroundings does not occur as frequently as ¹ H (proton), so it is considered that what is occurring on the left side of Fig. 7 is not T1 weighted but the signal drop as a whole. Is reasonable. Therefore, it is considered that the ²³ Na-NMR signal obtained by the pulse sequence of TR <50 ms (for example, TR is about 20 ms) reflects the Na density distribution.

(Example 2: T1 relaxation time and T2 relaxation time)
Using the ²³ Na-NMR apparatus of Example 1, the T1 relaxation time and the T2 relaxation time were measured for various saline solutions. The T1 relaxation time of various concentrations (0.9 w / v% to 26.4 w / v%) of saline solution (2 ml) measured by the saturation recovery method was 100 ms or less. Most of the T2 relaxation times of various concentrations measured by fitting the attenuation curve with a 90-degree pulse were about 20 ms, and a very fast relaxation time component of about 2 to 3 ms was also observed.

(Example 3: MRI image shooting with a mouse)
An MRI image was taken with a mouse using a ²³ Na-MRI apparatus provided with a holding portion suitable for holding the mouse and having the configuration shown in FIG. The configuration of the MRI ^apparatus, using both transmission and reception RF coils that can send obtained and excitation RF pulse NMR signals from 23 Na, except for performing the image generation process described above, the same as conventional ¹ H-MRI device is there. For details of such a device, refer to Japanese Patent Application Laid-Open No. 2015-145853. The pulse sequence, using the three-dimensional gradient echo method for imaging a region of 420 [mu] m × 420 [mu] m, performs phase encoding and frequency encoding in the number of pixels 64 pixels × 64 pixels, repetition time 20 ms, the signal acquisition time _{t a} Was 5.12 ms, the number of integrations was 20, and the imaging time was 7 minutes. At this time, the frequency-encoded gradient magnetic field for dephasé was adjusted so that the echo peak occurred within 2 ms from the start of application of the frequency-encoded gradient magnetic field for rephase. The mice were C57BL / 6, and those subjected to renal ischemia-reperfusion surgery (surgery group) and those without surgery (control group) were used. Cross-sectional images of the vicinity of the kidney in the surgery group and the control group are shown in FIG.

(Example 4: Relationship between repetition time and integrated signal value at the same measurement time in ²³ Na-NMR signal measurement)
Using ²³ Na-NMR apparatus and IH-NMR apparatus in Example 1, the measurement time is 100 seconds, when the various changing repetition time, ²³ Na-NMR signals and ¹ H at various brine sample -The integrated signal value of the NMR signal was measured. The result is shown in FIG. The exponential regression equation (y = a × x ^b ) shown in FIG. 9 is when the measurement time is x and the repetition time is y.

When the repetition time is TR and the number of integrations is N, the measurement time is t = TR × N. Generally, in the case of ¹ H-MRI, the signal-to-noise ratio (S / N ratio) is proportional to √t. It is known. This is because the signal strength increases in proportion to the number of integrations N, but the noise has redundancy and decreases only by √N. Further, since TR is an important factor for determining the image contrast of a subject in ¹ H-MRI, it is not possible to set it extremely long or short from the intended TR.

As shown in FIG. 9, the ¹ H-NMR signal (top graph in FIG. 9) decreases at about -0.5th power of the measurement time, while noise (bottom graph in FIG. 9) also decreases. , It decreases by about -0.5th power of the measurement time. Therefore, the S / N ratio of ^{1 1} H-NMR signal to noise is substantially constant even if the measurement time changes. This means that even if the repetition time TR is reduced and the signal integration number N is earned, the S / N ratio does not improve if the measurement time is constant.

On the other hand, the ²³ Na-NMR signal (graph in FIG. 9 between the graph of the ¹ H-NMR signal and the graph of the noise graft) decreases in the measurement time of about -0.7 to the -0.9 power. At this ^time, 23 S / N ratio of the Na-NMR signal-to-noise is reduced at -0.2 to -0.4 squared relative measurement time ^{^{(x -0.7 ÷ x -0.5 = x}} - ^0.2 to x- ^0.9 ÷ x- ^0.5 = x- ^0.4 ). This means that in the case of a ²³ Na-NMR signal, even if the measurement time is constant, the S / N ratio can be improved by reducing the repetition time TR and increasing the signal integration number N.

Further, in the ²³ Na-NMR signal, the multiplier of the exponential regression equation and the Na concentration have a negative correlation. Therefore, in advance, a sample containing various concentrations of Na (for example, saline solution having various concentrations) is used. On the other hand, the signal strength is measured by changing the repetition time variously with the same measurement time, the multiplier of the exponential regression equation for each sample is obtained, and the relational expression (for example, linear function) between the Na concentration and the multiplier is obtained. After that, in a sample with an unknown Na concentration, the signal strength is measured by variously changing the repetition time at the same measurement time, the multiplier of the exponential regression equation in the sample is obtained, and the multiplier is used as the relational expression. By applying to, the Na concentration in the sample can be measured. It is also considered that this concentration measurement method can be applied to nuclei other than Na (for example, nuclei having a spin quantum number of 3/2).

(Example 5: Observation of Na dynamics in a mouse by a ¹ H / ²³ Na-dual MRI apparatus)
^{The 1} H / ²³ Na-dual MRI apparatus is a 400 MHz super-electromagnetic magnet (manufactured by JASTEC, Narrow bore series, bore diameter 54 mm) and a measurement as shown in FIG. 10 arranged in the bore of this super-electromagnetic magnet. Including the stand. Measurement table is a cylindrical shape with open one side as shown in FIG. 10 (a), from the open portion, fitting the bottom plate transmission and reception surface coil is mounted for ^{1 H} / 23 ^Na measurements. On the surface of the bottom plate, as shown in FIG. 10 (a), are arranged to overlap ¹ and ^H for surface coils and ^{23 Na} for surface coils, the time of measurement, the mouse between the coils and the bottom plate It is placed and the coil hits around the kidney on the back side of the mouse. Further, on the back surface of the bottom plate, as shown in FIG. 10 (b), the coil of the resonance circuit portion and the RF cable and ^{23 Na} for surface coils except for the coils of the resonant circuit for of the ^{1 H} for surface coils The part to be removed and the RF cable are arranged so as to extend in opposite directions. Other configurations are the same as those of the conventional multi-atomic MRI apparatus except that signals are acquired for ²³ Na based on the above-described embodiment.

Using the above ¹ H / ²³ Na-dual MRI apparatus, Na dynamics in mice after furosemide treatment (1 mmcc was injected into the kidney with a syringe) was observed. ^By alternately repeating the measurement of ¹ H / ²³ Na in the same cross section, ¹ H images and ²³ Na in the same period were obtained. Throughout the observation period, the by ^{1 H} image, change by furosemide treatment was not observed. On the other hand, in the ²³ Na image, the Na concentration in the kidney increased (signal intensity increased) until about 2 minutes after the furosemide treatment, but after that, the Na concentration in the kidney decreased (signal intensity decreased). ), Instead, the Na concentration increased in the large intestine. This result is consistent with the conventional findings of furosemide treatment, and showed that this device is suitable for observing Na dynamics in the body of mice over time.

The present invention allows various embodiments and modifications without departing from the broad spirit and scope of the present invention. Moreover, the above-described embodiment is for explaining the present invention, and does not limit the scope of the present invention. That is, the scope of the present invention is indicated by the scope of claims, not by the embodiment. Then, various modifications made within the scope of the claims and the equivalent meaning of the invention are considered to be within the scope of the present invention.

Regarding the present application, priority is claimed based on Japanese Patent Application No. 2019-093569 filed on May 17, 2019, and the specification of Japanese Patent Application No. 2019-093569 is included in the present specification. , The scope of claims and the entire drawing shall be incorporated as a reference.

100 ... MRI device 1 ... Target 2 ... Holding unit 3 ... Bore 10 ... Static magnetic field coil 20 ... Gradient coil generator, 21 ... Gradient coil, 22 ... Gradient coil drive unit 30 ... RF pulse application unit, 31 ... RF coil , 32 ... RF coil drive unit 40 ... Receiver unit 50 ... Control device 51 ... Control unit, 51a ... Pulse control unit, 51b ... Image generation unit 52 ... Storage unit, PG ... Program 60 ... Display unit 70 ... Operation unit 300 ... Pulse Sequence 311 ... α ° pulse 351 ... signal 352 ... MRI signal 321 ... first phase-encoded gradient coil 331 ... second phase-encoded

gradient coil

341, 342 ... frequency-encoded gradient magnetic field

Claims

A static magnetic field forming part that forms a static magnetic field,
An object holding unit that holds an object in the static magnetic field,
A pulse sequence including an excitation pulse, a phase-encoded gradient magnetic field, and a frequency-encoded gradient magnetic field is applied to the object in the static magnetic field, and an NMR signal is generated from the object by applying the excitation pulse, and the frequency-encoded gradient magnetic field is generated. A pulse application unit that dephases the NMR signal and then rephases it by applying
A detection unit that detects each of the NMR signals phase-encoded by the phase-encoded gradient magnetic field of different amplitudes while the frequency-encoded gradient magnetic field is applied by the pulse application unit to rephase the NMR signal. ,
An image generation unit that generates an image from the NMR signal detected by the detection unit,
With
The pulse application unit applies the pulse sequence to the target so that the echo peak comes before half of the signal acquisition time from the application of the frequency-encoded gradient magnetic field for rephase of the NMR signal.
The detection unit detects the NMR signal over the entire range from the application of the frequency-encoded gradient magnetic field for rephase of the NMR signal to the elapse of the signal acquisition time.
The image generation unit generates the image from the entire NMR signal detected over the entire area.
A nuclear magnetic resonance imaging apparatus characterized by this.
The nuclear magnetic resonance imaging apparatus according to claim 1, wherein the atom to be measured is an atomic nucleus having a spin quantum number of 3/2.
The nuclear magnetic resonance imaging apparatus according to claim 2, wherein the atom to be measured is 23 Na.
The nuclear magnetic resonance imaging apparatus according to any one of claims 1 to 3, which generates a contrast image by laying a pre-RF pulse in the pulse sequence.
The atom to be measured is 23 Na,
The repetition time is less than 30 ms
The image is a density distribution image,
The nuclear magnetic resonance imaging apparatus according to any one of claims 1 to 4.
The pulse application unit applies the pulse sequence having the same phase encoding to the target a plurality of times.
The image generation unit integrates the NMR signals obtained from the pulse sequences having the same phase encoding a plurality of times, and generates the image based on the integrated NMR signals.
The nuclear magnetic resonance imaging apparatus according to any one of claims 1 to 5.
The pulse application unit can apply two or more types of excitation pulses having different frequencies corresponding to two types of atoms.
The nuclear magnetic resonance imaging apparatus according to any one of claims 1 to 6.
A pulse sequence including an excitation pulse, a phase-encoded gradient magnetic field, and a frequency-encoded gradient magnetic field is applied to an object in a static magnetic field, an NMR signal is generated from the object by applying the excitation pulse, and the frequency-encoded gradient magnetic field is applied. A pulse application step in which the NMR signal is dephased and then rephased by
A detection step of detecting each of the NMR signals phase-encoded by the phase-encoded gradient magnetic field having different amplitudes while the NMR signal is being rephased by applying the frequency-encoded gradient magnetic field by the pulse application step. ,
An image generation step of generating an image from the NMR signal detected in the detection step, and
With
In the pulse application step, the pulse sequence is applied to the target so that the echo peak comes before half of the signal acquisition time from the application of the frequency-encoded gradient magnetic field for rephase of the NMR signal.
In the detection step, the NMR signal is detected over the entire range from the application of the frequency-encoded gradient magnetic field for rephase of the NMR signal to the elapse of the signal acquisition time.
In the image generation step, the image is generated from the entire NMR signal detected over the entire area.
A nuclear magnetic resonance imaging method characterized by this.
Computer,
A pulse sequence including an excitation pulse, a phase-encoded gradient magnetic field, and a frequency-encoded gradient magnetic field is applied to a target in a static magnetic field to a pulse application unit, and an NMR signal is generated from the target by applying the excitation pulse to generate an NMR signal at the frequency. A pulse application means that dephases the NMR signal by applying an encode gradient magnetic field and then rephases the NMR signal, and echo peaks before half of the signal acquisition time from the application of the frequency encode gradient magnetic field for rephase the NMR signal. A pulse application means for applying the pulse sequence to the target.
Each of the NMR signals phase-encoded by the phase-encoded gradient magnetic field of different amplitude is detected in the detection unit while the frequency-encoded gradient magnetic field is applied by the pulse application unit to rephase the NMR signal. A detection means for detecting the NMR signal over the entire range from the application of the frequency-encoded gradient magnetic field for rephase the NMR signal to the elapse of the signal acquisition time.
An image generation means for causing an image generation unit to generate an image from the NMR signal detected by the detection unit, and for generating the image from the entire NMR signal detected over the entire area.
A program that functions as.