JP7412787B2 - Nuclear magnetic resonance imaging device, nuclear magnetic resonance imaging method, and program - Google Patents

Description

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 the process of applying specific gradient magnetic fields to a subject in a static magnetic field and irradiating specific RF (Radio Frequency) pulses (excitation pulses) to cause specific atoms within the subject to undergo nuclear magnetic resonance. 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 (i.e., a cross-sectional image), a three-dimensional image) is generated from this signal. An NMR signal to which position information is added by applying a gradient magnetic field or the like measured by MRI is also called an MRI signal. Further, a set of RF pulses and gradient magnetic fields that are applied with specific intensity and timing to obtain an NMR signal is called a pulse sequence.

The aforementioned gradient magnetic field connects the real space where the subject is located and the subject's spatial frequency space (k-space) through a Fourier transform relationship, and the MRI signal mentioned above reflects the information of the subject's k-space. are doing. Therefore, in MRI, k-space information of a subject is collected discretely from an MRI signal, and an image of the subject in real space is reconstructed by performing discrete inverse Fourier transform on the obtained discrete data.

As a method for collecting k-space data from MRI signals, three-axis gradient magnetic fields (X, Y, and Z axes) are controlled by three-axis gradient magnetic field coils, and the lattice of the Cartesian coordinate system of the object's k-space is A method of extracting point data line by line (line scan), a method of sequentially extracting data on the polar coordinate system of the subject's k space along multiple radial straight lines or spiral curves passing through the origin (radial scan) /Spiral Scan) etc. have been put into practical use.

An MRI signal has a convex portion with high signal intensity, which is called an echo, which is generated by manipulating a gradient magnetic field with a designed pulse sequence. The echo and signal changes before and after the echo contain spatial information of the subject. Echoes resulting from the application of a gradient magnetic field are called gradient echoes. Echoes resulting from continuous application of a high frequency magnetic field (eg, application of an excitation pulse followed by application of an inverted pulse) are called spin echoes. Regardless of the cause of the echo, the time from the application of the excitation pulse to the occurrence of the echo is called the echo time (TE).

The distribution, density, and relaxation time of the NMR atomic nuclei of the target within the object differ depending on the situation, so research and development of MRI signal integration methods (particularly pulse sequences) suitable for the target are progressing rapidly. There is.

For example, a partial echo method is known as one such method. In the partial echo method, the gradient echo (GRE) method shortens the TE and acquires the MRI signal until half of the signal acquisition time _ta has elapsed from the TE, thereby obtaining data corresponding to approximately half of the k-space. MRI Generate an image. FIG. 5 shows a pulse sequence when applying the partial echo method in the GRE method. For example, an example in which correction is performed using 0 fill is disclosed in Patent Document 1.

US Patent No. 6,166,545

In conventional MRI, a cross-sectional image of a living body is generated based on the presence or absence of ¹ H (protons), which are abundantly present in a subject such as a living body. However, with the aim of improving diagnostic accuracy and expanding the scope of research, it is desired to develop an MRI device that targets other NMR nuclei. For example, attempts to develop a ²³ Na-MRI device have begun in hopes of visualizing sodium in living bodies and early detection of pathologies, particularly those related to the brain and kidneys.

However, it is known that the NMR signal of ²³ Na has lower sensitivity 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 1/20,000 of the NMR signal of ¹ H.

Furthermore, in many subjects including living bodies, the abundance of NMR nuclei other than ¹ H is smaller than that of ¹ H, making it difficult to obtain MRI signals with signal strength necessary for image acquisition.

In order to obtain an MRI signal with a practical signal strength in MRI targeting atoms other than ^1H , it is simply necessary to obtain an integrated signal obtained by acquiring and integrating MRI signals corresponding to the same image multiple times. It is conceivable to generate an image from ²³ Na, but when the signal strength is low, such as with 23 Na, it is necessary to select a magnet with a high static magnetic field strength to increase the NMR signal strength, and furthermore, it is necessary to greatly increase the number of signal collection repetitions. However, this may exceed the measurement time that the patient can tolerate. Therefore, in order to increase the number of repetitions and keep the measurement time within the allowable limit, we need to reduce the repetition time (TR) required to acquire a single MRI signal, especially the echo time required from the application of an excitation pulse to the generation of an 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 constructed from half of the imaging data, there is a risk that false image artifacts may often be mixed in. In particular, in the case of MRI imaging using the gradient echo method, Hermitian symmetry is incomplete due to spatial inhomogeneity of the static magnetic field, and the problem of false image artifacts is significant.

Further, in order to put ²³ Na-MRI into practical use, it has been desired to further improve the S/N ratio of ²³ Na NMR signals. ²³ Na Sodium is an NMR nucleus with a spin quantum number of 3/2. In addition, in the case of ^1H (proton), it often forms practically large molecules or complexes due to hydrogen bonding with _H2O , but ^23Na exists almost as a single ion even in living organisms. It is thought that there are. For this reason, ²³ Na has a short T2 relaxation time ranging from several ms (milliseconds) to about 30 ms. Due to these two characteristics, in the case of ²³ Na-MRI, ^1H (protons) (spin quantum number 1/2, main T2 relaxation time is tens of ms to hundreds of ms) contained in biological water and lipids. It has been difficult to simply apply the methodology of image contrast generation in the imaging theory of ms).

In view of the above, an object of the present invention is to provide a nuclear magnetic resonance imaging apparatus, a nuclear magnetic resonance imaging method, and a program that can suitably take MRI images.

The nuclear magnetic resonance imaging apparatus according to the first aspect of the present invention includes:
a static magnetic field forming section that forms a static magnetic field;
an object holding unit that holds the object within the static magnetic field;
applying a pulse sequence comprising an excitation pulse, a phase-encoding gradient magnetic field, and a frequency-encoding gradient magnetic field to the object in the static magnetic field; application of the excitation pulse produces an NMR signal from the object; and the frequency-encoding gradient magnetic field a pulse application unit that dephases and then rephases the NMR signal by applying;
a detecting unit that detects each of the NMR signals phase-encoded by the phase-encoding gradient magnetic fields of different amplitudes while the pulse applying unit is applying the frequency-encoding gradient magnetic field to rephase the NMR signal; ,
an image generation unit that generates an image from the NMR signal detected by the detection unit;
Equipped with
The pulse application unit applies to the object the pulse sequence such that an echo peak occurs before half of the signal acquisition time from application of the frequency encoded gradient magnetic field for rephasing the NMR signal,
The detection unit detects the NMR signal over the entire range from the application of the frequency encode gradient magnetic field for rephasing the NMR signal until the signal acquisition time elapses,
The image generation unit generates the image from the entire NMR signal detected over the entire area,
It is characterized by

The nuclear magnetic resonance imaging method according to the second aspect of the present invention includes:
applying a pulse sequence comprising an excitation pulse, a phase-encoding gradient magnetic field, and a frequency-encoding gradient magnetic field to an object in a static magnetic field, the application of the excitation pulse producing an NMR signal from the object, and the application of the frequency-encoding gradient magnetic field; a pulse application step of dephasing and then rephasing the NMR signal;
a detection step of detecting each of the NMR signals phase-encoded by the phase-encoding gradient magnetic fields of different amplitudes while the frequency-encoding gradient magnetic field is applied in the pulse application step to rephase the NMR signal; ,
an image generation step of generating an image from the NMR signal detected in the detection step;
Equipped with
In the pulse application step, the pulse sequence is applied to the object such that the echo peak occurs before half of the signal acquisition time from the application of the frequency encoded gradient magnetic field for rephasing the NMR signal,
In the detection step, the NMR signal is detected over the entire range from the application of the frequency encode gradient magnetic field for rephasing the NMR signal until the signal acquisition time elapses;
In the image generation step, the image is generated from the entire NMR signal detected over the entire area.
It is characterized by

The program according to the third aspect of the present invention is
computer,
The pulse application unit applies a pulse sequence including an excitation pulse, a phase encoded gradient magnetic field, and a frequency encoded gradient magnetic field to an object in a static magnetic field, generates an NMR signal from the object by application of the excitation pulse, and generates an NMR signal from the object by applying the excitation pulse, and A pulse applying means for dephasing and then rephasing the NMR signal by applying an encoding gradient magnetic field, the pulse applying means dephasing and then rephasing the NMR signal, wherein an echo peak is generated before half of the signal acquisition time from application of the frequency encoding gradient magnetic field for rephasing the NMR signal. pulse application means for applying the pulse sequence to the object such that
A detection unit detects each of the NMR signals phase-encoded by the phase-encoded gradient magnetic field having different amplitudes while the pulse application unit is applying the frequency-encoded gradient magnetic field to rephase the NMR signal. detection means for detecting the NMR signal over the entire range from application of the frequency encoded gradient magnetic field for rephasing the NMR signal to elapse of the signal acquisition time;
Image generation means for causing an image generation unit to generate an image from the NMR signal detected by the detection unit, the image generation unit causing the image generation unit to generate the image from the entire NMR signal detected over the entire area;
function as

According to the present invention, MRI images can be suitably captured.

FIG. 1 is a block diagram showing the configuration of an MRI apparatus according to an embodiment of the present invention. FIG. 1 is a schematic diagram showing a pulse sequence according to an embodiment of the present invention. 3 is a flowchart of image generation processing according to an embodiment of the present invention. FIG. 2 is a schematic diagram showing a pulse sequence of a conventional GRE method. FIG. 2 is a schematic diagram showing a pulse sequence of a partial echo method. FIG. 2 is a schematic diagram showing a quantum pulse. 2 is a graph showing the intensity of ¹ H-NMR signal and ²³ Na-NMR signal of saline solution versus repetition time. ²³ is a cross-sectional image of the vicinity of a mouse kidney obtained by Na-MRI. (a) shows cross-sectional images of the surgical group and (b) shows the control group. It is a graph showing the intensity of the ¹ H-NMR signal and ²³ Na-NMR signal of saline measured at the same measurement time versus the repetition time. ^1H / ^23Na -This is a photograph of the measurement table of the 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, an MRI apparatus 100 according to an embodiment of the present invention includes a static magnetic field coil 10, a gradient magnetic field generation section 20, an RF pulse application section 30, a reception section 40, a control device 50, It includes a display section 60 and an operation section 70.

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 same axis (Z axis), and , is provided in a casing (not shown).

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

The static magnetic field coil 10 forms a uniform static magnetic field within the bore 3 at several ppm level (for example, 1.5 Tesla to 21 Tesla). For example, the static magnetic field coil 10 may be formed such that the central axis of the static magnetic field coil 10 or the central axis of magnetic flux passes through the bore 3 (in particular, the central axis of the bore 3). The static magnetic field that is formed is a horizontal magnetic field that is generally 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 a control device 50 via a static magnetic field coil drive section (not shown). Note that 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 a superconducting coil or a 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 include a shim coil group (not shown) for correcting static magnetic field uniformity, and the static magnetic field can be made more uniform by such a shim coil group. Note that any static magnetic field generating unit can be used instead of the static magnetic field coil 10 as long as it can form a uniform static magnetic field.

The gradient magnetic field generation unit 20 generates independent three-axis gradient magnetic fields necessary for imaging in the bore 3, and includes a gradient magnetic field coil 21, a gradient magnetic field coil drive unit 22 that drives the gradient magnetic field coil 21, and a gradient magnetic field coil drive unit 22 that drives the gradient magnetic field coil 21. , has.

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 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 magnetic field 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. Note that a unidirectional gradient magnetic field may play two or more roles within the same pulse sequence.

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

In FIG. 1, the horizontal direction of the paper is the X-axis direction, the vertical direction of the paper is the Y-axis direction, and the normal direction of the paper is the Z-axis direction, and the gradient axes of the gradient magnetic field described above are parallel to these X, Y, and Z axes. Although it is depicted as being applied, the gradient axes of the three gradient magnetic fields (e.g. slice axis, phase axis, frequency axis) can be limited to the above-mentioned X, Y, and Z axes as long as they maintain orthogonality to each other. It is not necessary to match some or all of these.

The RF pulse application unit 30 is for applying an RF pulse to the object 1 to generate nuclear magnetic resonance, and includes an RF coil 31 and an RF coil drive unit 32 that drives the RF coil 31.

The RF coil 31 forms a high frequency magnetic field in the static magnetic field space for exciting the nuclear spin of the NMR atomic nucleus targeted in the object 1 . Forming a high frequency magnetic field in this manner is also referred to as applying or transmitting an RF pulse. The RF coil 31 has the function of transmitting RF pulses and also the function of receiving nuclear magnetic resonance (NMR) signals, which are electromagnetic waves generated by excited nuclear spins. Note that an NMR signal to which position information is added by applying a gradient magnetic field or the like measured by MRI is referred to as an MRI signal. Note that instead of the RF coil 31, an RF pulse transmitting coil and an MRI signal receiving coil may be configured separately.

Further, it is desirable that the RF coil 31 or the RF pulse transmitting coil be large, and that the MRI signal receiving coil be small, and it is desirable that the size be the minimum necessary size. The smaller the RF coil 31 or the RF pulse transmission coil is, the lower the excitation power can be suppressed (specific absorption rate (SAR) of the RF pulse, such as when the RF pulse is applied continuously in a relatively short time by increasing the number of repetitions). Rate) can be lowered). The smaller the RF coil 31 or the RF pulse receiving coil is, the more it can avoid receiving MRI signals from unnecessary parts (for example, parts of the subject 1 that are not imaged). In particular, it is desirable to narrow the receiving area of these receiving coils so that the repetition time and echo time can be shortened, thereby eliminating the need for selective excitation pulses.

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 drive section 32 causes the RF coil 31 to generate an RF pulse corresponding to the Larmor frequency determined by the type of target atom and the magnetic field strength 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 it can shorten the echo time (TE). Further, in order to generate a so-called T1 image contrast, a 180° pulse (commonly known as an Inversion Recovery pulse) for inverting nuclear magnetization may be prefixed, and several ms of IR time may be further arranged as a waiting time. This makes it possible to reduce nuclear magnetization signals having a specific T1 time (not shown).

The receiving section (detecting section) 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 signals necessary for image reconstruction be removed by an analog or digital frequency filter in the receiving section 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 section 51 and a storage section 52, and drives, for example, a computer that controls the overall operation of the MRI apparatus 100, an RF coil drive section 32, and a gradient magnetic field coil drive section 22 using a pulse sequence. 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 operate the MRI apparatus. The operation of each part of 100 is controlled.

The storage unit 52 is composed of a ROM (Read Only Memory), a RAM (Random Access Memory), etc., and stores a CPU and various operating programs (a program PG for executing image generation processing to be described later) as necessary. (including) data, table data to be described later, etc. are stored in advance. The RAM of the storage unit 52 temporarily stores data representing various calculation results, data representing discrimination results, etc., and performs calculations. The storage unit 52 can also be placed on the Internet.

The control section 51 includes a pulse control section 51a and an image generation section 51b as functional sections.

The pulse control unit 51a controls the gradient magnetic field generation unit 20 and the RF pulse application unit 30 based on data of a pulse sequence 300 described below indicating a pulse sequence of the RF pulse, the phase encoding gradient magnetic fields Gp1 and Gp2, and the frequency encoding gradient magnetic field Gr. The drive control is performed. The data of the pulse sequence 300 may be stored in advance in the storage unit 52, or may be input or reset by a user operation via the operation unit 70, which will be 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 encode gradient magnetic fields Gp1 and Gp2 and the frequency encode gradient magnetic field Gr. The image generation unit 51b generates a three-dimensional image of the object 1 by performing three-dimensional inverse Fourier transform on this k-space data.

The display unit 60 displays a waveform showing the intensity of the MRI signal over time, which allows confirmation of the echo peak position while scanning the k-space, an 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), an organic electroluminescent display (OELD), or the like. Note that when the MRI apparatus 100 is configured to allow remote work, a personal computer (for example, a notebook computer) or other portable electronic terminal may be used as the display section 60 and the operation section 70 described below.

The operation unit 70 accepts an operation by a user and supplies an operation signal to the control device 50 in accordance with the accepted operation. The operation unit 70 includes, for example, a keyboard with a pointing device, a mouse, 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 by operating the operation unit 70 .

(Image generation processing)
The image generation process will be explained with reference to FIG. 3. The image generation process is executed by the control unit 51 in response to a start operation from the operation unit 70 by the user, for example. It is assumed that the object 1 placed on the holding section 2 is set within the bore 3 before the start of the image generation process.

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 within the bore 3 (step S11). Note that 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 the object 1. Therefore, a group of shim coils (not shown) for correcting the static magnetic field uniformity may be provided in the static magnetic field coil 10 to make the static magnetic field even 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 pulse sequence 300 below (pulse application process), and receives the k-space data of the object 1 via the reception unit 40. The information is acquired (detection step) and stored in the storage unit 52 (total step S12).

(Pulse sequence 300)
FIG. 2 shows a pulse sequence 300 used in this embodiment. The pulse sequence 300 is similar to the pulse sequence of the gradient echo (GRE) method, except for the position of the TE (echo time). In FIG. 2, "RF" is the sequence of α° pulses 311 that are excitation pulses, "Gp1" is the first phase encoding gradient magnetic field, "Gp2" is the second phase encoding gradient magnetic field, and "Gr" is the frequency encoding gradient. Magnetic field (readout gradient field), "S" indicates MRI signal. The first phase encoding gradient magnetic field, the second phase encoding gradient magnetic field, and the frequency encoding gradient magnetic field are orthogonal to each other, eg, parallel to the X, Y, and Z axes, respectively. The signal acquisition time t _a is a time width for selecting data used for image reconstruction. The signal acquisition time t _a is the same as or shorter than the application time of the frequency encoding gradient magnetic field 342 .

First, the α° pulse 311 causes α° excitation of spins. The flip angle α° is 90° or less.

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

For phase encoding, first and second phase encoding gradient magnetic fields 321 and 331 are applied while sequentially varying the amplitude each time an α° pulse 311 is applied (if signal integration is performed, a series of α° pulses 311 is applied). This is done by For example, the first phase encoding gradient magnetic field 321 is set to N1 stages (for example, 128 stages), the second phase encoding gradient magnetic field 331 is set to N2 stages (for example, 32 stages), and the number of repetitions (number of signal integrations) is set to N times. (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 will be repeated N1 × N2 times, If one repetition time (TR) is counted as one measurement, the measurement will be repeated N×N1×N2 times. The signal integration may be performed in an order in which N1×N2 measurements are repeated N times.

In parallel with each phase encode, the macroscopic nuclear magnetization is dephased (flyback) by applying a frequency encoding gradient magnetic field 341, and then the spins are rephased (refocused) by applying a frequency encoding gradient magnetic field 342. to generate a gradient echo MRI signal 352. The signal intensity reaches its maximum at the time after TE from α° excitation.

The integral value of the frequency encode gradient magnetic field 341 for dephasing (the stippled area of the frequency encode gradient magnetic field 341 in FIG. 2) is the integral value from the application of the frequency encode gradient magnetic field 342 to TE (the stippled area of the frequency encode gradient magnetic field 341 in FIG. 2). is equal to the stippled area). Therefore, by changing the integral value of the frequency encoding gradient magnetic field 341 for dephasing (for example, by changing the application period, intensity, and/or waveform) without changing the frequency encoding gradient magnetic field 342, it is possible to change the timing at which TE occurs. can be adjusted. Conventionally, as shown in FIG. 4, the frequency encode gradient magnetic field 343 is adjusted so that TE is at the center of the signal acquisition time _ta of the frequency encode gradient magnetic field 342, but in this embodiment, as shown in FIG. , the TE (i.e., the echo peak) is greater than half the signal acquisition time t _a from application of the frequency-encoded gradient magnetic field 342, and specifically between one-tenth and one-third of the start of the signal acquisition time t _a . In between, the frequency encoding gradient magnetic field 341 is adjusted. In FIG. 2, the start time of the application of the frequency encode 342 and the start time of the signal acquisition time _ta are the same, but the timing is intended to be the time when the current value of the frequency encode 342 rises or falls 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 to 50 μs to 500 μs in many cases. Ideally, it is expected that there will be no delay, so the description will be omitted. In addition, in FIGS. 2 and 4, the application end time of the frequency encode 342 is the same as the signal acquisition time _{t a} determined based on the width of frequency encoding applied to the target 1, but the period from the start of application of the frequency encode 342 to the end of application is The application time is arbitrary as long as it is longer than the signal collection time _ta . The frequency encode 342 applied _after the signal acquisition time ta has the effect of promoting relaxation recovery of the nuclear magnetization signal of the subject.

The MRI signal 352 thus obtained is digitized via the receiving section 40. The control unit 51 acquires data at M points (for example, 128 points) from the digitized MRI signal 352 at each repetition time over the entire signal acquisition time t _a of the frequency encoded gradient magnetic field 342, and stores the data in the storage unit 52. to be memorized. After that, the control unit 51 selects the amplitudes of the first and second phase encoding gradient magnetic fields 321 and 331 and the signal collection time t _a of the frequency encoding gradient magnetic field 342 corresponding to the same position of the target 1 among the data. (where the combination of positions is the same) is integrated, and 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 object 1 obtained as described above (for example, the k-space data of the object 1 consisting of N1×N2×M pieces of data). (Step S13). Specifically, a three-dimensional discrete inverse Fourier transform is applied to this k-space data to reconstruct a three-dimensional image of the object 1. With the recent development of information technology, MRI image reconstruction is sometimes performed using artificial intelligence AI or deep learning instead of inverse Fourier transform. These may be applied to reconstructing a three-dimensional image of object 1.

(Effects of this embodiment)
In the conventional GRE method, the peak of the MRI signal is at the center of the signal acquisition time _ta , as shown in Figure 4, but in MRI targeting low-sensitivity NMR nuclei (e.g., ²³ Na), the peak of the MRI signal is The intensity was reduced by T2 and T2 ^* decay, easily confused with noise, and difficult to detect.

On the other hand, the MRI apparatus 100 according to the present embodiment can capture signals with priority given to echo peaks having rough image contour information. In particular, when ²³ Na is the target of MRI, the signal component of ²³ Na Among them, really fast signal components with a T2 decay time of 2 to 3 ms or less can be used for image generation. Therefore, according to this embodiment, an MRI image can be suitably generated even in MRI targeting atoms with low sensitivity.

Furthermore, as shown in FIG. 5, although the partial echo method can shorten the TE and measurement time, there is a problem in that false image artifacts are introduced due to the configuration of the k-space. Furthermore, when collecting signals using the GRE method, phase variations occur due to T2 ^* attenuation, making it difficult to apply the partial echo method. Furthermore, the partial echo method cannot be combined with other MRI measurement methods (eg, compressed sensing method) due to fundamental problems.

On the other hand, in _this embodiment, as shown in _FIG . Since there is a low frequency component whose signal strength is higher than the high frequency component in the latter half of the signal acquisition time _ta where the signal strength attenuation is large, the pixel bandwidth is narrowed to reduce noise while giving priority to image generation. It has been achieved. Since homodyne reconstruction with phase encoding removed is not performed, the occurrence of false image artifacts is reduced. Therefore, according to this embodiment, an MRI image can be generated more favorably than the partial echo method. Furthermore, the image generation processing method according to this embodiment can be combined with various MRI measurement methods, as shown in the following modified examples.

(Modified example)
The present invention is not limited to the above embodiments and drawings. It is possible to make changes (including deletion of constituent elements) as appropriate without changing the gist of the present invention. Furthermore, in this specification, descriptions of known technical matters have been omitted as appropriate in order to facilitate understanding of the present invention.

(Modification 1)
In the embodiments described above, the GRE method is used, but a pre-RF pulse may also be provided. Thereby, it is possible to obtain an image having a contrast depending on the type of pre-RF pulse. For NMR nuclei with short relaxation times (eg ²³ Na), it is desirable to introduce a pre-RF pulse after each excitation pulse. Also, in the case of NMR nuclei with long relaxation times (e.g. ¹ H), a pre-RF pulse may be laid for every excitation pulse or for every multiple excitation pulses (e.g. when integrating the signal N times). (before the first of N excitation pulses).

For example, when using ²³ Na as the NMR nucleus, a density-enhanced image can be obtained if the signal is acquired within the relaxation time, but by applying a pre-RF pulse, it is possible to obtain an image with contrast other than density emphasis. It will be done. When using ²³ Na as the NMR nucleus, as explained in Example 1, unlike when using ¹ H as the NMR nucleus, it is not possible to obtain contrast other than density enhancement by simply changing the repetition time. , the provision of contrast by the introduction of a pre-RF pulse is advantageous.

Further, at this time, a quantum pulse may be applied as a pre-RF pulse as shown in FIG. 5. With such quantum pulses, it is expected that tissue-specific contrast can be obtained due to spin energy transitions. With conventional ¹ H-MRI, the relaxation time is long, ranging from several tens of ms to several hundred ms, so it is possible to discriminate between tissues using differences in relaxation time or chemical shift excitation. However, in the case of NMR nuclei with short relaxation times (for example, in the case of ²³ Na, which has a short relaxation time of 20 ms or less), such discrimination is not possible. According to quantum pulses, even in NMR nuclei with short relaxation times, it is possible to emphasize tissue differences before the signal attenuates due to spin energy transitions.

In the quantum pulse, after a 90° pulse (+X direction) is applied, a 180° pulse is applied after a time τ has elapsed, and then a 180° pulse is continuously applied every 2τ. At this time, the continuous 180° pulse is called 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 (Carr-Purcell-Meiboom-Gill) may also be used.

Note that 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. In addition, position information is not added to the NMR signal generated by the pre-RF pulse by phase encoding and frequency encoding before application of the excitation pulse, and the pulse sequence according to the above-mentioned pulse sequence 300 or other variations is used. After applying the excitation pulse, position information is added by phase encoding and frequency encoding.

(Modification 2)
In the embodiment described above, a 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 for each slice, an MRI image is converted by a two-dimensional Fourier transform. may be generated. Moreover, the reconstruction of this MRI image may be performed by artificial intelligence AI or deep learning instead of inverse Fourier transform.

(Modification 3)
If the echo peak is too close to the start time of the signal acquisition time _ta of the frequency encoder 342, the signal intensities of the MRI signals measured at the _start and end points of the signal acquisition time ta will differ greatly, and the same waveform will continue periodically. There is also a risk that the premise of the discrete inverse Fourier transform that the In this case, for example, false image artifacts such as cast ringing artifacts (multiple edges on edges) occur in the MRI image. Therefore, it is preferable that the echo peak is far enough away from the start time of the signal acquisition time t _a of the frequency encoding gradient magnetic field 342 to an extent that such image artifacts do not occur. The presence or absence of false image artifacts can be confirmed by human visual inspection of the completed MRI image, and the position of the echo peak can be compared to the MRI image generated according to the above embodiment for the same part of the same object 1 in the same way as in the conventional method. The correlation coefficient is calculated by comparing the reference image centered at the signal acquisition time t _a with the reference image using a correlation coefficient (for example, a normalized cross-correlation coefficient obtained by ZNCC (Zero-mean Normalized Cross-Correlation)). It can also be determined whether it exceeds a certain value (for example, 0.7 in the case of a normalized cross-correlation coefficient determined by ZNCC). In addition, based on the results of a preliminary experiment using object 1 or a reference object of the same type as object 1, it is determined that the position of the echo peak in step S12 of the image generation process is such that no false image artifact is visually confirmed or there is no correlation as described above. It is preferable that the number is adjusted to be equal to or greater than a certain value. The false images may be removed using artificial intelligence or deep learning.

(Modification 4)
If target 1 is a sample that is uniform in the thickness direction or a sample with high uniformity in the thickness direction (for example, if target 1 is a thin film material), frequency encoding is performed in the thickness direction (for example, a direction perpendicular to the thin film). A two-dimensional MRI image of a cross section in the thickness direction of the sample is obtained by applying first and second phase encoding gradient magnetic fields in directions perpendicular to the thickness direction and perpendicular to each other (for example, in a direction parallel to the thin film). In this case, the echo peak position may be generated at the same time or before the start of signal collection. When applying a frequency encoded gradient magnetic field in the thickness direction of such a sample (i.e., setting the signal readout axis in the thickness direction), the MR signal draws a very gentle mountain shape around the peak, _so The importance of the echo peak position is relatively low.

(Modification 5)
In the embodiments described above, in order to suppress signal attenuation, it is desirable that the repetition time be equal to or less than the T2 relaxation time. 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.

In addition, in the above-described embodiment, the position of the echo peak is determined from the application of the excitation _pulse to the elapse of the T2 relaxation time, which is the main component that can be observed, regardless of the length of the signal collection time t a (multiple T2 relaxation times If the component consists of an excitation pulse, it is preferable that the excitation pulse occurs between the application of the excitation pulse and the elapse of the longest T2 relaxation time, the shortest T2 relaxation time, or a T2 relaxation time between these. may occur until one-half, one-third, or one-tenth of the T2 relaxation time (if the T2 relaxation time consists of multiple components, any of those T2 relaxation times) has elapsed since the application of preferable. When the lifetime of a signal with a short relaxation time component is intentionally extended by applying a pre-excitation pulse, the apparently longer T2 relaxation time is referred to.

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

(Modification 7)
Any other MRI techniques can be combined with the embodiments described above. For example, a compressed sensing method or a multi-quantum coherence method may be employed. In particular, the compressed sensing method selects the phase encoding in a Gaussian-random manner, so it cannot be applied to the partial echo method.

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

(Modification 9)
Although it is assumed that the program PG for executing the image generation process is stored in advance in the storage unit 52, it may be distributed/provided by a removable 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 telecommunications network or the like.

(Modification 10)
In order to obtain signals derived from two or more types of atoms (for example, ^{a 1} H-NMR signal and a ²³ Na-NMR signal), for example, in order to enable simultaneous imaging of two types of atoms, the RF coil 31 is used. Alternatively, a coil capable of transmitting and receiving two or more types of RF pulses of different frequencies corresponding to two or more types of atoms may be employed. For example, such coils may include a single transmitter/receiver coil tuned to the Larmor frequency of two or more atoms, such as a double-tuned birdcage coil, a double-tuned surface coil, a double-tuned saddle coil, a double-tuned butterfly coil, etc. A type coil, a double-tuned solenoid coil, etc. may be used, or two or more transmitting/receiving coils individually tuned to the Larmor frequencies of two or more types of atoms may be used.

Furthermore, when configuring an RF pulse transmitting coil and an MRI signal receiving coil separately instead of the RF coil 31, a single coil tuned to the Larmor frequency of two or more types of atoms may be used as the RF pulse transmitting coil. A transmitting coil or two or more transmitting coils individually tuned to the Larmor frequencies of two or more types of atoms may be used, and as an MRI signal receiving coil, a coil tuned to the Larmor frequencies of two or more types of atoms may be used. A single receiving coil or two or more receiving coils individually tuned to the Larmor frequencies of two or more types of atoms may be used.

When using the above-mentioned coil as the RF coil 31 or the MRI signal receiving coil, interference between RF receiving cables for different atoms is suppressed and a high S/N ratio can be obtained within a compact space. Coils and associated resonant circuits and RF receiving cables (coaxial cables) may be arranged. For example, when there are two types of atoms to be measured, the resonant circuit (excluding the coil) and RF receiving cable for one atom are the same as the resonant circuit (excluding the coil) and RF receiving cable for the other atom. from a single coil or two coils of different inductance placed in close proximity (e.g., stacked or coaxial) in opposite directions as shown in FIG. Preferably, they are arranged so as to extend.

(Example)
(Example 1: Comparison of sensitivity between ¹ H-NMR signal and ²³ Na-NMR signal)
For the comparison of NMR signal intensity, we used ¹ H (9.4 Tesla) or ²³ Na (9.4 Tesla) RF coils (inner diameter φ15 mm, gap 10 mm, 2 turns x 2 Helmholtz type) with similar diameter and shape. A plastic test tube containing 2 ml of saline solution (0.9 w/v%) was placed inside the coil, an excitation pulse was applied to the saline solution, and FID signals were measured. 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 from a separately prepared high-frequency transmitter to the preamplifier as a pseudo signal, and the intensity is adjusted to the intensity of the NMR signal collected in advance. By adjusting the output of the high frequency transmitter, the signal strength (dBm) of ¹ H and ²³ Na NMR signals after the RF coil was measured. The measurement results are shown in FIG. The intensity of the ¹ H-NMR signal derived from this sample was approximately 20,000 times the intensity 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 Figure 7 (TR < 50 ms) is depressed due to T1 saturation due to fast repeated TR, but ²³ Na exists as a single atom and does not interact with other atoms or molecules. Since it is not bonded, the exchange of relaxation energy with the surroundings does not occur as frequently as with ^1H (protons), so what is occurring on the left side of Figure 7 is not T1 emphasis, but an overall signal drop. is reasonable. Therefore, the ²³ Na-NMR signal obtained with a pulse sequence with TR<50 ms (for example, TR of about 20 ms) is considered to reflect 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 T2 relaxation time of various saline solutions were measured. The T1 relaxation times of saline solutions (2 ml) at various concentrations (0.9 w/v % to 26.4 w/v %) measured by the saturation recovery method were 100 ms or less. Most of the T2 relaxation times of various concentrations measured by fitting the attenuation curve using a 90-degree pulse were about 20 ms, and very fast relaxation time components of about 2 to 3 ms were also observed.

(Example 3: MRI image capture with mouse)
An MRI image of a mouse was taken using a ²³ Na-MRI apparatus equipped with a holding part suitable for holding a mouse and having the configuration shown in FIG. 1. The configuration of this MRI device is similar to the conventional ¹ H-MRI device, except that it uses a transmitting/receiving RF coil that can obtain NMR signals from ²³ Na and transmit excitation RF pulses, and performs the image generation process described above. be. For details of such a device, please refer to Japanese Patent Application Laid-Open No. 2015-145853. As a pulse sequence, a three-dimensional gradient echo method was used, and phase encoding and frequency encoding were performed with a pixel count of 64 pixels x 64 pixels for an imaging area of 420 μm x 420 μm, the repetition time was 20 ms, and the signal acquisition time was _ta. was 5.12 ms, the number of integrations was 20, and the imaging time was 7 minutes. At this time, the frequency encoding gradient magnetic field for dephasing was adjusted so that the echo peak occurred within 2 ms from the start of application of the frequency encoding gradient magnetic field for rephasing. In addition, the mice were C57BL/6, and those that underwent renal ischemia reperfusion surgery (surgery group) and those that did not undergo surgery (control group) were used. Figure 8 shows cross-sectional images of the vicinity of the kidneys in the surgical group and the control group.

(Example 4: Relationship between repetition time and integrated signal value at the same measurement time in ²³ Na-NMR signal measurement)
²³ Na-NMR signals and ^{1 H} of various saline samples using a ²³ Na-NMR device and a 1 H-NMR device in Example 1, with a measurement time of 100 seconds and various repetition times. - The integrated signal value of the NMR signal was measured. The results are shown in FIG. Note that 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 x 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 although the signal strength increases in proportion to the number of integrations N, noise has redundancy and decreases only by √N. Furthermore, since TR is an important element that determines 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 (the top graph in FIG. 9) decreases approximately to the -0.5 power of the measurement time, while the noise (the bottom graph in FIG. 9) also decreases. , decreases by approximately -0.5 power of measurement time. Therefore, the S/N ratio of ¹ H-NMR signal to noise remains almost constant even if the measurement time changes. This means that even if the number of signal integrations N is increased by decreasing the repetition time TR, the S/N ratio will not improve if the measurement time is constant.

On the other hand, ^{the 23} Na-NMR signal (the graph between the graph of the ¹ H-NMR signal and the noise graft in FIG. 9) decreases at a rate of about -0.7 to about -0.9 power of the measurement time. At this time, the S/N ratio of the ²³ Na-NMR signal to noise decreases by the -0.2 to -0.4 power with respect to the measurement time (x ^-0.7 ÷ x ^-0.5 = x ^{- 0.2} ~ 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 decreasing the repetition time TR and increasing the number of signal integrations N.

In addition, in the ²³ Na-NMR signal, there is a negative correlation between the multiplier of the exponential regression equation and the Na concentration. On the other hand, the signal intensity is measured by varying the repetition time with the same measurement time, and the multiplier of the exponential regression equation for each sample is determined, and the relational expression (for example, a linear function) between the Na concentration and the multiplier is determined. Then, from now on, measure the signal intensity of a sample with unknown Na concentration by varying the repetition time with the same measurement time, find the multiplier of the exponential regression equation for the sample, and apply the multiplier to the above relational equation. By applying this, the Na concentration in the sample can be measured. Further, this concentration measurement method is considered to be applicable to atomic nuclei other than Na (for example, atomic nuclei with a spin quantum number of 3/2).

(Example 5: Observation of Na dynamics in mouse body using ¹ H/ ²³ Na-dual MRI device)
^{The 1} H/ ²³ Na-dual MRI system uses a 400 MHz super-electromagnetic magnet (manufactured by JASTEC, Narrow bore series, bore diameter 54 mm) and a measurement device placed inside the bore of this super-electromagnetic magnet as shown in Figure 10. Includes stand. The measurement stand has a cylindrical shape with one side open as shown in FIG. 10(a), and a bottom plate to which a transmitting/receiving surface coil for ¹ H/ ²³ Na measurement is attached is fitted into the open part. As shown in Fig. 10(a), a surface coil for ^1H and a surface coil for ^23Na are placed on top of each other on the surface of the bottom plate, and during measurement, a mouse is placed between these coils and the bottom plate. The coil is placed around the kidneys on the dorsal side of the mouse. In addition, on the back of the bottom plate, as shown in Figure 10(b), there is a section of the resonant circuit for the ^1H surface coil excluding the coil, and a portion of the resonant circuit of the RF cable and the ^23Na surface coil. The removed portion and the RF cable are arranged to extend in opposite directions. The rest of the configuration is the same as the conventional polyatomic MRI apparatus, except that signal acquisition is performed for ²³ Na based on the above-described embodiment.

Using the above ¹ H/ ²³ Na-dual MRI device, Na dynamics in the mouse body after furosemide treatment (1 mmcc injected into the kidney with a syringe) was observed. By repeating the measurement of ¹ H/ ²³ Na alternately on the same cross section, ¹ H images and ²³ Na images were obtained during the same period. Throughout the observation period, no changes were observed in the ¹ H images due to furosemide treatment. On the other hand, in the ^23Na image, the Na concentration in the kidney increased (signal intensity increased) until about 2 minutes after furosemide treatment, but after that, the Na concentration in the kidney decreased (signal intensity decreased). ), instead, Na concentration increased in the large intestine. This result was consistent with the conventional findings based on furosemide treatment, and showed that this device is suitable for observing Na dynamics over time in the mouse body.

This invention is capable of various embodiments and modifications without departing from the broad spirit and scope of this invention. Further, the embodiments described above are for explaining the present invention, and do not limit the scope of the present invention. That is, the scope of the invention is indicated by the claims rather than the embodiments. Various modifications made within the scope of the claims and the meaning of the invention equivalent thereto are considered to be within the scope of this invention.

This application claims priority 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 this specification. , the claims, and the drawings are hereby incorporated by reference in their entirety.

DESCRIPTION OF SYMBOLS 100... MRI apparatus 1... Target 2... Holding part 3... Bore 10... Static magnetic field coil 20... Gradient magnetic field generation part, 21... Gradient magnetic field coil, 22... Gradient magnetic field coil drive part 30... RF pulse application part, 31... RF coil , 32...RF coil drive unit 40...Reception 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 encode gradient magnetic field 331...Second phase encode gradient magnetic field 341, 342...Frequency encode gradient magnetic field

Claims (8)

a static magnetic field forming section that forms a static magnetic field;
an object holding unit that holds the object within the static magnetic field;
applying a pulse sequence comprising an excitation pulse, a phase-encoding gradient magnetic field, and a frequency-encoding gradient magnetic field to the object in the static magnetic field; application of the excitation pulse produces an NMR signal from the object; and the frequency-encoding gradient magnetic field a pulse application unit that dephases and then rephases the NMR signal by applying;
a detecting unit that detects each of the NMR signals phase-encoded by the phase-encoding gradient magnetic fields of different amplitudes while the pulse applying unit is applying the frequency-encoding gradient magnetic field to rephase the NMR signal; ,
an image generation unit that generates an image from the NMR signal detected by the detection unit;
Equipped with
The pulse application unit is configured to apply a signal at least half of the signal acquisition time from the application of the frequency encoded gradient magnetic field for rephasing the NMR signal, and from among the plurality of T2 relaxation times of the atoms to be measured from the application of the excitation pulse. applying to the subject the pulse sequence such that the echo peak occurs before a time equal to the longest one has elapsed;
The detection unit detects the NMR signal over the entire range from the application of the frequency encode gradient magnetic field for rephasing the NMR signal until the signal acquisition time elapses,
The image generation unit generates the image from the entire NMR signal detected over the entire area,
generating a contrast image by introducing into the pulse sequence a pre-RF pulse comprising successive 180° pulses of alternating polarity;
A nuclear magnetic resonance imaging device characterized by: The nuclear magnetic resonance imaging apparatus according to claim 1, wherein the atom to be measured is an atomic nucleus with 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 atom to be measured is ²³ Na,
the repetition time is less than 30ms,
the image is a density distribution image;
The nuclear magnetic resonance imaging apparatus according to any one of claims 1 to 3. The pulse application unit applies the pulse sequence having the same phase encode to the target multiple times,
The image generation unit integrates the NMR signals obtained from the plurality of pulse sequences having the same phase encoding, and generates the image based on the integrated NMR signals.
The nuclear magnetic resonance imaging apparatus according to any one of claims 1 to 4. The pulse application unit is capable of applying 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 5. applying a pulse sequence comprising an excitation pulse, a phase-encoding gradient magnetic field, and a frequency-encoding gradient magnetic field to an object in a static magnetic field, the application of the excitation pulse producing an NMR signal from the object, and the application of the frequency-encoding gradient magnetic field; a pulse application step of dephasing and then rephasing the NMR signal;
a detection step of detecting each of the NMR signals phase-encoded by the phase-encoding gradient magnetic fields of different amplitudes while the frequency-encoding gradient magnetic field is applied in the pulse application step to rephase the NMR signal; ,
an image generation step of generating an image from the NMR signal detected in the detection step;
Equipped with
In the pulse application step, the frequency encoded gradient magnetic field for rephasing the NMR signal is applied before half of the signal acquisition time, and from the application of the excitation pulse to one of the plurality of T2 relaxation times of the atoms to be measured. applying to the subject the pulse sequence such that the echo peak occurs before a time equal to the longest one has elapsed;
In the detection step, the NMR signal is detected over the entire range from the application of the frequency encode gradient magnetic field for rephasing the NMR signal until the signal acquisition time elapses;
In the image generation step, the image is generated from the entire NMR signal detected over the entire area,
generating a contrast image by introducing into the pulse sequence a pre-RF pulse comprising successive 180° pulses of alternating polarity;
A nuclear magnetic resonance imaging method characterized by: computer,
The pulse application unit applies a pulse sequence including an excitation pulse, a phase encoded gradient magnetic field, and a frequency encoded gradient magnetic field to an object in a static magnetic field, generates an NMR signal from the object by application of the excitation pulse, and generates an NMR signal from the object by applying the excitation pulse, and A pulse application means for dephasing and then rephasing the NMR signal by applying an encoding gradient magnetic field, the pulse application means being earlier than half of the signal acquisition time from the application of the frequency encoding gradient magnetic field for rephasing the NMR signal, and Applying to the object the pulse sequence such that the echo peak occurs before a time equal to the non-longest T2 relaxation time of the plurality of T2 relaxation times of the atoms of the measurement object elapses after application of the excitation pulse, and the pulse sequence includes: pulsing means for producing a contrast image by laying down a pre-RF pulse comprising successive 180° pulses of alternating polarity;
A detection unit detects each of the NMR signals phase-encoded by the phase-encoded gradient magnetic field having different amplitudes while the pulse application unit is applying the frequency-encoded gradient magnetic field to rephase the NMR signal. detection means for detecting the NMR signal over the entire range from application of the frequency encoded gradient magnetic field for rephasing the NMR signal to elapse of the signal acquisition time;
Image generation means for causing an image generation unit to generate an image from the NMR signal detected by the detection unit, the image generation unit causing the image generation unit to generate the image from the entire NMR signal detected over the entire area;
A program that functions as

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