JP7645739B2 - Magnetic resonance imaging apparatus and control method thereof - Google Patents

Description

The present invention relates to a magnetic resonance imaging device (hereinafter abbreviated as an MRI device), and in particular to a technology for obtaining images with reduced artifacts caused by FID (free induction decay) signals.

A typical imaging method for MRI devices is the spin-echo pulse sequence. In the spin-echo pulse sequence, an excitation RF pulse (90-degree pulse) is applied, followed by an inversion RF pulse (180-degree pulse) to diffuse the spins, and the NMR signals from the refocused spins are collected as echo signals. In such a spin-echo pulse sequence, in addition to the echo signals from the spins excited by the 90-degree pulse, the spins are also excited by the 180-degree pulse, and a signal (FID signal) is generated due to the free induction decay. The spin-echo signal is encoded by the phase encoding gradient magnetic field applied after the 90-degree pulse is applied, but since the FID signal is not encoded, it is collected as a zero-encoded signal superimposed on the spin echo. For this reason, in an image reconstructed from k-space data obtained by the spin-echo sequence, a zipper-shaped artifact (an artifact caused by the FID signal: hereinafter referred to as the FID artifact) appears in the center of the image parallel to the frequency encoding direction.

In imaging using spin echo pulse sequences, the phase of a 180-degree pulse is usually inverted each time the sequence is repeated in order to remove this FID artifact. This allows the FID artifact to be moved to both sides of the image (for example, Non-Patent Document 1).

Michael N. Hoff et al. “Artifacts in Magnetic Resonance Imaging”, Capter 9, pp165-190, Image Principles, Neck, and the Brain, Research Gate (2016)

As a high-speed MRI imaging method, a method (parallel imaging) has become common in which measurements are performed with a reduced number of measurement data, unmeasured data is estimated during reconstruction using sensitivity distribution information of the receiving coil, aliasing that occurs in the image due to the reduced measurement data is unfolded, and the image is reconstructed. When this parallel imaging is applied to a spin echo pulse sequence, there is a problem that the conventional FID artifact avoidance technique described above does not work. This is because when parallel imaging is applied to imaging using a spin echo pulse sequence, even if FID artifacts are shifted to both sides of the image before the aliasing is unfolded, FID artifacts appear in the image after unfolding, not at the edge of the image, but at a position overlapping with the subject image. Moreover, the higher the speed multiplication number, the more FID artifacts there are, so it has been difficult to shorten the measurement time by increasing the speed multiplication number.

The objective of the present invention is to provide an MRI device that can obtain images with artifacts removed even when parallel imaging is applied to imaging using a spin echo pulse sequence.

To solve the above problem, the MRI apparatus of the present invention adds an additional inversion RF pulse to a spin echo system pulse sequence including an excitation RF pulse and an inversion RF pulse, and measures the FID signal generated by this inversion RF pulse. At this time, the spin echo mixed with the FID signal generated by the first inversion RF pulse and the FID signal generated by the additional inversion RF pulse are phase encoded. In this way, an image of only the spin echo with the effects of FID eliminated is obtained by performing a difference calculation between k-space data or between reconstructed images.

That is, the MRI apparatus of the present invention includes a static magnetic field generating magnet that generates a static magnetic field, a radio frequency transmitting unit that irradiates an RF pulse to a subject placed in a static magnetic field space, a receiving unit that receives a nuclear magnetic resonance signal generated from the subject, a gradient magnetic field generating unit that generates a gradient magnetic field that encodes the nuclear magnetic resonance signal, a control unit that controls a pulse sequence that operates the radio frequency transmitting unit, the receiving unit, and the gradient magnetic field generating unit, and a calculation unit that generates an image using the nuclear magnetic resonance signal. In imaging using a spin echo system pulse sequence including an excitation RF pulse and an inversion RF pulse, the control unit uses a pulse sequence that applies an inversion RF pulse twice after application of an excitation RF pulse and applies a phase encoding gradient magnetic field pulse every two inversion RF pulses, and collects two sets of phase encoded nuclear magnetic resonance signals.

According to the present invention, by adding an inversion RF pulse and collecting phase-encoded echo signals after each inversion RF pulse, data in which the spin echo contains an FID signal and data consisting of only the FID signal are obtained, and the difference between these data provides data consisting of only the spin echo with the FID component removed. This makes it possible to suppress the occurrence of artifacts caused by the FID signal.

In addition, since the subtracted data does not contain FID components, even if the data contains spatially overlapping signals due to parallel imaging or the FID position changes due to the speed multiplication rate, there is no need to separate the FID components, and image reconstruction without artifacts caused by FID is possible using normal parallel imaging calculations.

The present invention can obtain data that does not contain FID components, so it can be applied to all imaging methods other than parallel imaging, so long as they use spin echo pulse sequences.

FIG. 1 is a functional block diagram showing an overall outline of an MRI apparatus. FIG. 2 is a diagram showing details of an imaging unit of the MRI apparatus in FIG. 1 . FIG. 4 is a flowchart showing the procedure of imaging according to the first embodiment. FIG. 1 shows a typical spin echo sequence. FIG. 2 is a diagram showing a pulse sequence according to the first embodiment. FIG. 13A and 13B are diagrams showing examples of slice selection according to the second embodiment.

Below, an embodiment of the MRI device of the present invention will be described with reference to the drawings.

As shown in FIG. 1, the MRI device 1 mainly comprises an imaging unit 10 including a static magnetic field generating unit 11, a high frequency transmitting unit 13, a receiving unit 14, and a gradient magnetic field generating unit 12, a control unit 30 that controls the imaging unit 10 according to a predetermined pulse sequence, and a calculation unit 20 that reconstructs an image using echo signals collected by the imaging unit 10.

The imaging unit 10 includes a static magnetic field generating unit 11, a gradient magnetic field generating unit 12, a transmitting unit 13, a receiving unit 14, and a signal processing unit 15. In addition, although not shown in FIG. 1, it includes a sequencer that operates these units in a certain procedure (pulse sequence). The details of the imaging unit 10 will be described with reference to FIG. 2.

The static magnetic field generating unit 11 is composed of a permanent magnet type, resistive conductive type, or superconductive type static magnetic field generating magnet (not shown). Depending on the direction of the static magnetic field generated by the static magnetic field generating magnet, there are different types such as vertical magnetic field type and horizontal magnetic field type. In the vertical magnetic field type, a uniform static magnetic field is generated in the space around the subject 2 in a direction perpendicular to the body axis, while in the horizontal magnetic field type, a uniform static magnetic field is generated in the direction of the body axis.

The gradient magnetic field generating unit 12 consists of gradient magnetic field coils 121 wound in the three directions of X, Y, and Z, which are the coordinate system (static coordinate system) of the MRI device, and gradient magnetic field power supplies 122 that drive each gradient magnetic field coil. By driving the gradient magnetic field power supplies 122 of each coil according to commands from the sequencer 16, gradient magnetic fields Gx, Gy, and Gz in the three directions of X, Y, and Z are applied to the static magnetic field.

The transmitter 13 irradiates the subject 2 with a radio frequency magnetic field pulse (RF pulse) to induce nuclear magnetic resonance in the nuclear spins of atoms that constitute the biological tissue of the subject 2, and is equipped with a radio frequency oscillator 131, a modulator 132, a radio frequency amplifier 133, and a radio frequency coil (transmitting coil) 134 on the transmitting side. The RF pulse output from the radio frequency oscillator 131 is amplitude modulated by the modulator 132 at a timing according to a command from the sequencer 16, and this amplitude modulated radio frequency pulse is amplified by the radio frequency amplifier 133 and then supplied to the transmitting coil 134 arranged close to the subject 2, thereby irradiating the RF pulse to the subject 2.

The intensity and phase of the RF pulse are controlled by the modulator 132, allowing 90-degree and 180-degree pulses of different intensities to be output, and their phases to be controlled.

The receiver 14 detects echo signals (NMR signals) emitted by nuclear magnetic resonance of atomic spins that constitute the biological tissue of the subject 2, and includes a receiver-side high-frequency coil (receiving coil) 141, a signal amplifier 142, a quadrature phase detector 143, and an A/D converter 144. The NMR signal of the response of the subject 2 induced by the electromagnetic waves irradiated from the transmitting coil 134 is detected by the receiver coil 141 arranged close to the subject 2, amplified by the signal amplifier 142, and then split into two orthogonal signals by the quadrature phase detector 143 at a timing according to a command from the sequencer 16, each of which is converted into a digital quantity by the A/D converter 144 and sent to the signal processing unit 15. When a coil combining multiple subcoils is used as the receiver coil 141, each of the subcoils is connected to the signal amplifier 142, the quadrature phase detector 143, and the A/D converter 144, and the signal processing unit 15 collects signals for each subcoil.

The signal processing unit 15 is equipped with a storage device 150 such as ROM or RAM, an external storage device 151 such as an optical disk or a magnetic disk, a display 152, and an operation unit 153 consisting of input devices such as a trackball, a mouse, and a keyboard. The signal processing unit 15 executes processes such as signal processing and image reconstruction, and displays the resulting tomographic image of the subject 2 on the display 152 and records it on the magnetic disk or the like of the external storage device 151. The operation unit 153 is disposed close to the display 152, and the operator interactively controls various processes of the MRI apparatus through the operation unit 153 while looking at the display 152.

In the example shown in FIG. 2, the MRI apparatus further includes a computer 50 equipped with a CPU and memory, and the computer 50 realizes part of the functions of the signal processing unit 15 described above, as well as the functions of the calculation unit 20 and the control unit 30 as shown in FIG. 1.

The control unit 30 includes, in addition to the function of controlling the entire device, a reception unit 31 that receives imaging conditions and the like input via the operation unit 153, an imaging control unit 32 that sets the pulse sequence used for imaging in the sequencer 16 and controls imaging via the sequencer 16, and a display control unit (not shown) that controls the display of images and GUI of the imaging on the display. The calculation unit 20 includes an image reconstruction unit 21 that reconstructs an image using the nuclear magnetic resonance signals collected by the receiving unit 14, and a difference calculation unit 22 that performs calculations such as the difference between data. Details of these functions will be described later.

Part or all of the functions of the calculation unit 20 are included in the signal processing unit 15 and are executed by the CPU reading a specific program stored in the storage device. Some of the functions performed by the calculator 50 can also be realized by hardware such as an ASIC (Application Specific Integrated Circuit) or an FPGA (Field Programmable Gate Array).

The sequencer 16 is a control means that repeatedly applies the above-mentioned RF pulses and gradient magnetic field pulses in a predetermined pulse sequence, and operates under the control of the control unit 30, sending various commands necessary for collecting data on tomographic images of the subject 2 to the radio frequency transmission unit 13, the gradient magnetic field generation unit 12, and the reception unit 14.

There are various pulse sequences depending on the imaging method, and they are stored in advance in the storage device 150 (151). The user selects a desired pulse sequence via the operation unit 153 and determines the imaging sequence by setting imaging parameters such as echo time (TE), repetition time (TR), imaging field of view (FOV), and speed ratio for parallel imaging. In this embodiment, the sequencer 16 executes a spin echo type pulse sequence.

Next, we will explain the operation of the MRI device of this embodiment, mainly the imaging control by the control unit 30 and the image generation by the calculation unit 20.

<Embodiment 1>
In this embodiment, a pulse sequence obtained by modifying a normal spin echo system pulse sequence is executed, and processing is performed to remove an FID signal caused by an inversion RF pulse during image reconstruction.

The operation of the MRI apparatus of this embodiment will be explained below with reference to Figure 3.

First, the control unit 30 receives imaging conditions such as a pulse sequence and imaging parameters through the reception unit 31 (S301). If the imaging is a spin echo type pulse sequence, the pulse sequence is calculated based on the imaging parameters and set in the sequencer 16 (S302). If the imaging parameters include a speed ratio for parallel imaging, a measurement matrix size according to the speed ratio is set.

The pulse sequence used in this embodiment is characterized by the use of two inversion RF pulses and the phase encoding of the echo signals acquired after the inversion RF pulses. To explain the difference between the pulse sequence of this embodiment and conventional pulse sequences, we will first briefly explain conventional spin echo type pulse sequences.

As shown in FIG. 4, a typical spin echo pulse sequence 400 applies an excitation RF pulse (also called a 90-degree pulse) 411 together with a slice gradient magnetic field 421 for selecting a predetermined slice of the subject 2 to excite spins in the predetermined slice. Next, a phase encoding gradient magnetic field 431 is applied, and an inversion RF pulse (also called a 180-degree pulse) 412 is applied together with a slice gradient magnetic field 422 at 1/2 (TE/2) of the echo time (the gradient magnetic field 420 is a gradient magnetic field that returns the phase of the spins). This generates an echo signal (NMR signal) 451 that peaks at the echo time. This echo signal is collected for a predetermined sampling time while applying a frequency encoding gradient magnetic field 441. After a predetermined repetition time TR has elapsed, the above-mentioned process from the excitation of the spins to the collection of the echo signal is repeated while changing the strength of the phase encoding gradient magnetic field 431, and a set number of echo signals are collected. At this time, as shown by the dotted line in the figure, an FID signal 461 is generated by the inversion RF pulse 412 and mixed into the spin echo 451. This causes FID artifacts in the image.

The spin echo pulse sequence of this embodiment generates such an FID signal 461 as a phase-encoded signal along with the spin echo, making it possible to remove it from the image during image reconstruction.

5 shows an example of a spin echo system pulse sequence 500 used in this embodiment. The same elements as those in the pulse sequence 400 shown in FIG. 4 are denoted by the same reference numerals.
5, in this pulse sequence 500, a phase encoding gradient magnetic field 531 is applied after the inversion RF pulse 412. That is, in the pulse sequence 400 in FIG. 4, after the application of the excitation RF 411, the phase encoding gradient magnetic field 431 that was applied before the application of the inversion RF pulse 412 is changed to after the inversion RF pulse 412. As a result, both the spin echo generated after the TE by the inversion RF pulse 412 and the FID of the inversion RF pulse 412 are phase encoded and collected as one echo signal 551.

In order to eliminate the influence of spin rotation in directions other than the phase encoding direction due to the change in the application timing of the phase encoding gradient magnetic field 531, a pair of crusher gradient magnetic fields 523, 543 are added to axes Gs, Gr other than the phase encoding axis Gp. After that, a readout gradient magnetic field 441 is applied to collect an echo signal 551, similar to the pulse sequence 400 in FIG. 4, but as described above, the FID 461 shown in FIG. 4 is phase encoded and measured in this echo signal 551.

Furthermore, the pulse sequence 500 applies a second inversion RF pulse 512 after measuring the echo signal 451. The application phase of the second RF pulse 512 is the same as that of the first RF pulse 412. When applying the second inversion RF pulse, a pair of crusher gradient magnetic fields 524, 544 are added to the axes Gs and Gr other than the phase encoding axis Gp, as in the first application, so that spin echoes are not generated. Next, after applying the phase encoding gradient magnetic field 532, a readout gradient magnetic field 542 is applied to collect an echo signal 552. This echo signal does not include spin echoes from spins excited by the RF excitation pulse 411, but is an FID in which spins are generated by the second inversion RF pulse 512, and is a phase-encoded signal. Finally, crusher gradient magnetic fields 525, 535 are applied to return the spins to their original state. In FIG. 5, crusher gradient magnetic fields are applied to the slice axis Gs and the phase encoding axis Gp, but they may be applied to three axes.

The above sequence is repeated in TR to collect a number of signals 551, 552 corresponding to a preset measurement matrix. As with conventional spin echo pulse sequences, the phase of the two RF inversion pulses may be alternated between 0 and π each time TR is repeated.

This pulse sequence 500 is completely different from conventional fast spin FSE, which repeatedly applies multiple inversion RF pulses in succession to measure multiple spin echoes, in that the echo signal 552 generated by the second inversion RF pulse is not generated as a spin echo, the FID signal generated by the inversion RF pulse is phase encoded, and the inversion RF pulse is limited to two times.

Note that while Figure 5 shows a two-dimensional imaging sequence in which phase encoding gradient magnetic fields 531 and 532 are applied in one axial direction (Gp), a three-dimensional imaging sequence in which they are applied in two axial directions may also be used.

The sequencer 16 controls the imaging unit 10 according to this pulse sequence 500, which causes the receiver 14 to collect two sets of signals 551, 552, each of which is converted into k-space data (FIG. 3: S303). If the receiver coil has multiple channels, two sets of k-space data are obtained for each channel. The calculation unit 20 uses each k-space data to perform image reconstruction (S304).

The processing in the calculation unit 20 includes image reconstruction processing (processing by the image reconstruction unit 21) and difference processing (processing by the difference calculation unit 22). The difference processing may be processing in the measurement space, i.e., processing before image reconstruction, or processing in the image space (difference after image reconstruction), and the order differs depending on the processing.

When performing subtraction in the measurement space, as shown in S3041 and S3042 in FIG. 3, two sets of k-space data (complex data), i.e., the k-space data of signal 452 acquired after the second inversion RF pulse is subtracted from the k-space data of signal 451 acquired after the first inversion RF pulse (S3041).

In the case of parallel imaging, each k-space data contains signals that are not spatially separated, but the FID signals, which are phase-encoded like the spin echo, are distributed in the same way in the two sets of data, so by subtracting the two, data from which the FID signal components have been removed, that is, k-space data containing only the spin echo, can be obtained.

The k-space data after subtraction is subjected to, for example, a fast Fourier transform to generate an image, as in normal image reconstruction methods. In the case of the PI method, a parallel imaging calculation is performed using sensitivity distribution information of each receiving coil (corresponding to the channels that make up the receiving coil) that has been calculated in advance to obtain an image without aliasing (S3042). Known parallel imaging calculation methods include the SENSE method and the GRAPPA method, and their explanations will be omitted here. Image reconstruction methods that include L1 norm minimization, such as compressed sensing, may also be applied.

On the other hand, when subtraction is performed in image space, as shown in S3043 and S3044 in FIG. 3, first, two sets of k-space data are reconstructed into image data (S3043). The image reconstruction method is the same as in step S3042. Complex subtraction is performed on the obtained image data. That is, the image reconstructed using the second signal 552 is subtracted from the image reconstructed using the first signal 551 to obtain an image from which artifacts caused by the FID signal have been removed (S3044).

Figure 6 shows images obtained using the S3043 and S3044 techniques. In Figure 6, the image on the left is created from the first signal (including spin echo and FID), the center is an image created from the second signal, and the left is the subtracted image. Although the signal intensity of the center image is significantly lower than the images on either side, a slight FID image can be seen, and it can be seen that by subtracting this image, an image of only the spin echo can be obtained.

As described above, according to this embodiment, as a spin echo system pulse sequence, the FID signal is phase encoded together with the echo signal, and a second inversion RF pulse is added to generate only the phase-encoded FID signal in a similar manner. This makes it possible to generate a signal that eliminates the influence of the FID signal, i.e., an image consisting only of the echo signal (spin echo).

<Embodiment 2>
In the first embodiment, an image of a region selected by the first and second inversion RF pulses is obtained by performing an operation between two sets of data collected after the application of these RF pulses, but in this embodiment, a slice selected by the second inversion RF pulse is made different from the slice selected when the excitation RF pulse and the first inversion RF pulse are applied, and echo signals are collected from different slices. This sequence is repeated while changing the slice position to be selected, and data of a plurality of slices is collected.

The shape of the pulse sequence in this embodiment is almost the same as that shown in FIG. 5, so the differences will be explained with reference to FIG. 5. In this embodiment, phase encoding is also applied to the FID signal generated by the inversion RF pulse 412, and after collecting the spin echo 551, the inversion RF pulse 512 is applied to phase encode the FID signal of the inversion RF pulse 512 to collect the echo signal 552, just like in embodiment 1. In this embodiment, the amount of the slice selection gradient magnetic field 522 applied simultaneously with the inversion RF pulse 512 is made different from the slice selection gradient magnetic field 422 to obtain a signal 552 in a cross section different from the cross section in which the spin echo 551 was obtained.

The first inversion RF pulse 412 and the second inversion RF pulse 512 select different slices, so in principle the nuclear magnetic resonance signal from the slice selected in the first inversion does not get into the second echo signal 552, so the crusher gradient magnetic fields 524 and 544 before and after the inversion RF pulse can be omitted, but here we use the crusher gradient magnetic fields 524 and 544 to reliably block the influence of the first echo.

The order of slices to be changed between the first spin echo 551 and the next echo signal 552 is not particularly limited. For example, as shown in FIG. 7(A), imaging 1 in which slice S1 is selected first and slice S2 is selected second, and imaging 2 in which slice S2 is selected first and slice S1 is selected third, may be repeated one or more sets to perform multi-slice imaging. Alternatively, as shown in FIG. 7(B), imaging in which slice S(n) is selected first and slice S(m) (n and m are slice numbers, where n ≠ m) are selected as one set may be performed sequentially while changing the slice position. In either case, the slice selected first and the slice selected next may be adjacent slices or may be separated.

In image reconstruction, image reconstruction including difference calculation is performed using data from the same slices, as in embodiment 1. That is, in the example of FIG. 7A, the spin echo 251 data (data from slice S1) obtained in imaging 1 and the echo signal 252 data (data from slice S1) obtained in imaging 2 are used to perform complex difference calculation in the measurement space before image reconstruction, or the image after image reconstruction is subjected to complex difference calculation, and an image of only the spin echo is obtained for slice S1. Similarly, an image is obtained for slice S2 using the data from that slice.

In this embodiment, by selecting a different slice for the first inversion RF pulse from the slice for the second inversion RF pulse, the influence of the first signal can be reliably removed from the echo collected after the second inversion RF pulse. As a result, a highly accurate difference image can be obtained.

Two embodiments of the MRI apparatus of the present invention have been described above, but when setting the imaging conditions, either embodiment 1 or embodiment 2 may be automatically selected depending on whether or not multi-slice imaging is performed, or the user may select via the operation unit (input device) 153.

11: static magnetic field generator, 12: gradient magnetic field generator, 13: transmitter (high frequency transmitter), 14: receiver, 15: signal processor, 16: sequencer, 20: calculator, 21: image reconstruction unit, 22: difference calculator, 30: controller, 31: receiver, 32: imaging controller, 50: computer

Claims (11)

a static magnetic field generating magnet that generates a static magnetic field, a radio frequency transmitting unit that irradiates an RF pulse to a subject placed in a static magnetic field space, a receiving unit that receives a nuclear magnetic resonance signal generated from the subject, a gradient magnetic field generating unit that generates a gradient magnetic field that encodes the nuclear magnetic resonance signal, a control unit that controls a pulse sequence that operates the radio frequency transmitting unit, the receiving unit, and the gradient magnetic field generating unit, and a calculation unit that generates an image using the nuclear magnetic resonance signal,
the control unit, in imaging using a spin echo system pulse sequence including application of an excitation RF pulse for exciting nuclear spins and application of an inversion RF pulse for inverting the excited nuclear spins, uses, as the pulse sequence, a pulse sequence in which, after application of the excitation RF pulse, the inversion RF pulse is applied twice and a phase encoding gradient magnetic field pulse is applied for each of the two inversion RF pulses, and two sets of phase encoded nuclear magnetic resonance signals are respectively acquired ;
The magnetic resonance imaging apparatus according to claim 1, wherein the calculation unit obtains an image from which FID has been removed by performing a difference calculation in a measurement space or an image space using the two sets of nuclear magnetic resonance signals . 2. The magnetic resonance imaging apparatus according to claim 1,
The magnetic resonance imaging apparatus according to claim 1, wherein the calculation unit performs complex difference between the two sets of nuclear magnetic resonance signals and performs image reconstruction using the data after the difference. 3. The magnetic resonance imaging apparatus according to claim 2,
the pulse sequence is a pulse sequence for a parallel imaging method in which a measurement matrix size determined by an encoding number is smaller than a matrix size of an image,
The magnetic resonance imaging apparatus according to claim 1, wherein the calculation unit performs a parallel imaging calculation on the subtracted data to reconstruct an image. 2. The magnetic resonance imaging apparatus according to claim 1,
The calculation unit reconstructs an image from each of the two sets of nuclear magnetic resonance signals, and generates an image reconstructed from only spin echoes of the nuclear magnetic resonance signals by calculating the complex difference between the two sets of reconstructed images. 5. The magnetic resonance imaging apparatus according to claim 4,
the pulse sequence is a pulse sequence for a parallel imaging method in which a measurement matrix size determined by an encoding number is smaller than a matrix size of an image,
The magnetic resonance imaging apparatus according to claim 1, wherein the calculation unit performs a parallel imaging calculation on each of the two sets of nuclear magnetic resonance signals to reconstruct an image. 2. The magnetic resonance imaging apparatus according to claim 1,
The control unit repeats the pulse sequence with a repetition time TR, and the phase of two inversion RF pulses within the same repetition time is the same, and the phase is changed for each repetition time. 2. The magnetic resonance imaging apparatus according to claim 1,
4. A magnetic resonance imaging apparatus, comprising: a magnetic resonance imaging unit for diffusing spins, the magnetic resonance imaging unit comprising: a first inversion RF pulse for inverting a first region of the magnetic field; 2. The magnetic resonance imaging apparatus according to claim 1,
The pulse sequence includes application of an excitation RF pulse, application of two inversion RF pulses, and application of a slice selection gradient magnetic field pulse for selecting a cross section of the subject at the same time;
A magnetic resonance imaging apparatus characterized in that a cross section selected by the slice selection gradient magnetic field pulse when a second inversion RF pulse is applied is different from a cross section selected when an excitation RF pulse and a first inversion RF pulse are applied. 9. The magnetic resonance imaging apparatus according to claim 8,
The magnetic resonance imaging apparatus according to claim 1, wherein the calculation unit generates an image using a combination of nuclear magnetic resonance signals having the same cross section as the two sets of nuclear magnetic resonance signals. 2. The magnetic resonance imaging apparatus according to claim 1,
The control unit has a reception unit that receives imaging conditions or user specifications, and determines a cross section for generating the two sets of nuclear magnetic resonance signals in accordance with the imaging conditions or user specifications received by the reception unit. In imaging using a spin echo system pulse sequence including application of an excitation RF pulse for exciting nuclear spins and application of an inversion RF pulse for inverting the excited nuclear spins, the pulse sequence is executed by applying the excitation RF pulse, applying the inversion RF pulse twice, applying a phase encoding gradient magnetic field pulse for each of the two inversion RF pulses, and acquiring two sets of phase encoded nuclear magnetic resonance signals,
A method for controlling a magnetic resonance imaging apparatus , comprising controlling the operation of the magnetic resonance imaging apparatus so as to obtain an image from which FID has been eliminated by differential calculation in measurement space or image space using the two sets of nuclear magnetic resonance signals .

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