JPH0866384A - Nuclear magnetic resonance imaging device - Google Patents

Abstract

PURPOSE: To restrain the adverse effect of a DC component existing in a receiving system on an image. CONSTITUTION: A sequence controller 52 controls the phase of an RF oscillation signal from an RF oscillation circuit 34, thereby setting a different offset in the phases of a carrier signal for RF pulses generated from an RF coil 12 and a reference signal from a phase detection circuit 43 at every slice encoding. Also, the controller 52 changes a phase offset amount, depending on a slice encoding amount, so that a difference in the phase offset amount becomes 180 degrees as a whole. The phases of carrier signals are reversed from each other at every adjacent phase encoding at each slice encoding process.

Description

Detailed Description of the Invention

[0001]

This invention relates to the nuclear magnetic resonance phenomenon (M
The present invention relates to a nuclear magnetic resonance imaging apparatus that performs imaging using the R phenomenon), and particularly to a nuclear magnetic resonance imaging apparatus that performs three-dimensional imaging.

[0002]

2. Description of the Related Art In a nuclear magnetic resonance imaging apparatus, an object is irradiated with an RF pulse to excite the same, and an NMR signal is generated from this. Then, the NMR signal generated in the subject is received by the RF coil, phase-detected, and then sampled at a predetermined rate by the A / D converter and converted into digital data. By the way, in such a receiving system, a DC offset, a DC drift, and the like are inevitable, and therefore, these DC components are mixed in the collected data, and an artifact is generated in the reconstructed image.

Therefore, conventionally, the phase of the excitation pulse is
A phase shift of 180 ° for each excitation and a corresponding signal phase shift of 180 ° has been used to suppress DC artifacts. That is, if the phase of the excitation pulse is shifted by 180 ° and the received signal is shifted by 180 °, the signal component will be the same as when it was not shifted, but the DC component is added after reception. 1
The phase is shifted by 80 degrees. Therefore, if the data collected for each excitation is adjacently arranged as a signal of each line in the raw data space and two-dimensional Fourier transform is performed,
Since the DC component becomes the signal with the highest phase difference, it shifts to both ends in the phase encoding direction of the image, and the image artifact due to the DC component can be made inconspicuous.

[0004]

However, there is a problem that the above-mentioned conventional method for operating the phase of the excitation pulse is not very useful for three-dimensional imaging. That is, in a two-dimensional image, the image artifact due to the DC component can be driven to both ends in the phase encoding direction, but in the three-dimensionally reconstructed image, the image artifact due to the DC component occurs in the central slice. I will end up.

In view of the above, it is an object of the present invention to provide a nuclear magnetic resonance imaging apparatus capable of performing three-dimensional imaging while suppressing adverse effects on a reconstructed image of a DC component inevitable in a receiving system.

[0006]

To achieve the above object, in a nuclear magnetic resonance imaging apparatus according to the present invention, a main magnetic field generating means for generating a static magnetic field and a means for generating a gradient magnetic field pulse for slice encoding are provided. A means for generating a gradient magnetic field pulse for reading, a means for generating a gradient magnetic field pulse for phase encoding, a means for irradiating a subject placed in a magnetic field with an RF pulse whose amplitude is modulated by an RF carrier signal, Receiving means for receiving the NMR signal from the specimen, phase-detecting the carrier signal as a reference signal and then sampling to obtain data, and processing means for arranging the data in a raw data space and performing a multidimensional Fourier transform process, Each slice encode has a different offset in the phase of the carrier signal and the phase is turned off as a whole. The phase offset amount is changed according to the slice encode amount so that the difference between the bit amounts becomes 180 °, and the phase of the carrier signal is inverted for each adjacent phase encode in each slice encode. It is characterized by including a phase control means.

[0007]

When performing three-dimensional imaging in a nuclear magnetic resonance imaging apparatus, a method of performing phase encoding in the case of two-dimensional imaging and coding a position in one direction in the slice plane as a phase is applied to the slice thickness direction. By applying it and changing the magnitude of the gradient magnetic field for slice encoding according to the number of slices, the position in the slice thickness direction is coded. Each slice encode has a different offset in the phase of the carrier signal of the RF pulse used as the reference signal for phase detection, and the phase offset is adjusted according to the slice encode amount so that the difference in the phase offset amount is 180 ° as a whole. Change the amount. Then, in the data obtained by the phase detection, the DC component is just 18 in the slice encode direction.
It appears with a phase change of 0 °, and DC artifacts can be transferred to the slices at both ends by Fourier transform in the slice encoding direction. Further, in the same slice encoding amount, the phases of the carrier signals are mutually inverted for each adjacent phase encoding, so that the DC artifacts are driven to both ends in the direction of the phase encoding gradient magnetic field in the reconstructed image of each slice. .

[0008]

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT A preferred embodiment of the present invention will now be described in detail with reference to the drawings. A nuclear magnetic resonance imaging apparatus according to an embodiment of the present invention is configured as shown in FIG. In FIG. 1, the main magnetic field magnet 10 is for generating a static magnetic field. Normal,
It consists of a superconducting magnet. A gradient magnetic field coil 11 is provided to generate a gradient magnetic field that is superimposed on this static magnetic field. The gradient magnetic field coil 11 is composed of three sets of coils,
By each of them, three gradient magnetic fields Gx, Gy, Gz whose magnetic field strengths are respectively inclined in the three axis directions of X, Y, Z are generated. These three-axis gradient magnetic fields Gx, Gy, G
By selecting one of z or combining them, a gradient magnetic field Gs for slice encoding in an arbitrary direction (S direction), a gradient magnetic field Gr for reading (and frequency encoding) in an arbitrary direction (R direction), arbitrary Direction (P
Direction), a gradient magnetic field Gp for phase encoding can be created.

A subject (not shown) is placed in the space to which the static magnetic field and the gradient magnetic field are applied. For this subject,
An RF coil 12 is attached for irradiating the subject with RF pulses and for receiving the NMR signal generated in the subject.

Each set of gradient magnetic field coil 11 includes:
Currents of predetermined waveforms are respectively supplied from the gradient magnetic field power amplifier 23. Under the control of the sequence controller 52, the digital gradient magnetic field waveform generator 21 outputs Gs, Gr, Gp.
Each digital pulse waveform of is generated, and this is D / A
It is converted into an analog signal by the converter 22 and input to the gradient magnetic field power amplifier 23. An amplified version of this analog signal is supplied to each set of coils of the gradient magnetic field coil 11, so that Gs, Gr, Gp pulses corresponding to the desired waveform output from the digital gradient magnetic field waveform generator 21 are generated. It will be. This makes it possible to generate Gs, Gr, and Gp pulses having pulse waveforms required for pulse sequences such as the spin echo method and the gradient echo method.

The RF pulse is applied to the subject from the RF coil 12. Therefore, the RF from the RF oscillation circuit 34
The carrier signal is amplitude-modulated by the amplitude modulation circuit 33, the modulated output is amplified by the RF power amplifier 35, and then supplied to the RF coil 12. The RF oscillation circuit 34 is controlled by the sequence controller 52 and generates an RF carrier signal having a frequency corresponding to the resonance frequency of the subject. The phase of the RF carrier signal is also controlled by the sequence controller 52, and the phase is inverted by 180 ° for each line and the phase offset amount is changed for each slice encoding as will be described later in detail. The amplitude modulation signal is a digital RF pulse waveform generated from the digital RF waveform generator 31 under the control of the sequence controller 52, and the D / A converter 32.
It is obtained by converting into an analog signal at.

When excited by such an RF pulse, an NMR signal is generated in the subject and the N
The MR signal is received by the RF coil 12, passed through the preamplifier 41 and the anti-aliasing filter 42, and is sent to the phase detection circuit 43 for phase detection. As the reference signal for the phase detection, the RF signal having a predetermined phase is sent from the RF oscillation circuit 34. The signal obtained by the phase detection is sent to the A / D converter 44, sequentially sampled at a predetermined sampling rate, and converted into a series of digital data. The data obtained from the A / D converter 44 is stored in the host computer 51.
3D Fourier transform processing for reconstructing an image. Also, this host computer 5
1 controls the sequence controller 52 according to the pulse sequence which comprises various imaging scans.

Here, the case of adopting a gradient echo sequence as a pulse sequence for imaging and performing three-dimensional imaging will be described. As shown in FIG. 2, an RF pulse having a predetermined flip angle is a Gs pulse Gs.
1 is given together with 1 to selectively excite a predetermined range in the slice thickness direction. After that, the Gr pulse is inverted to refocus the spin phase to generate an echo signal. Before this signal is generated, a Gp pulse is added to perform phase encoding. The magnitude of this Gp pulse is changed for each repetition. Here, it is assumed that the phase encoding amount is changed in order from the larger one to the smaller one. That is, FIG.
, Gp1, Gp2, ..., And an echo signal is generated for each of them to obtain data.

Gradient magnetic field pulse Gp for phase encoding
, Gp2, ... are all applied, then the slice encoding gradient magnetic field Gs is changed to give a slightly reduced pulse Gs2, and the pulses Gp1, Gp are similarly given.
2, ... Are sequentially given and an echo signal is generated for each of them to obtain data. When this is finished, a further reduced pulse Gs3 is given, and similarly Gp1, Gp2,
Get data for each. By repeating this for the number of slice encodes (the number of slices), data of many lines is collected.

The data of the large number of lines are three-dimensionally arranged in the raw data space as shown in FIG. That is, the data of one line is a data array at each sampling point in the time axis direction of one echo signal, but the data of each line is obtained by changing the Gp pulse for Gs1, that is, Gp1, Gp2 ,. , Respectively, and also for each of Gs2, Gs3, ... For each of Gp1, Gp2 ,. The data of each line is Gp1, Gp2, Gp3, Gp4, Gp5, ... For Gs1, Gs2, Gs3 ,.
3 in the vertical direction as shown in FIG.
Gs2, Gs3, ... Are arranged in this order in the depth direction.

By thus performing the three-dimensional Fourier transform on the three-dimensionally arranged data in the raw data space, the positions of the gradient magnetic field Gr, the gradient magnetic field Gp and the gradient magnetic field Gs are reproduced. To be done. That is,
A three-dimensional image can be reconstructed.

Here, the carrier signal of the RF pulse has a phase offset φ1 for the first slice encoding amount (Gs1) and then a slice encoding amount (Gs2).
Regarding the phase offset φ2 and the third slice encoding amount (Gs3) have different phase offsets such as the phase offset φ3. Further, the carrier signal of the RF pulse is 180 ° out of phase with respect to the same Gs pulse every Gp pulse. That is, for example, for Gs1, Gp1, Gp
The pulse sequence is performed in the order of 2, Gp3, Gp4, Gp5, ...
1 and Gp2, phase offset φ1 + 180 °, G
φ1 for p3, φ1 + 180 ° for Gp4, φ1 for Gp5,
... like this.

Then, the phase offset for each slice encode is set to be 180 ° in total. When the number of slice encodes is n, the difference between φ1 and φn is 180 °.

For convenience of explanation, it is assumed that the number of slices is 7 and the number of slice encodes is 7 as shown in FIG. At this time, the gradient magnetic field pulse for slice encoding is Gs1.
~ Gs7. Slice # 4 located at the center of the seven slices # 1 to # 7 and a gradient magnetic field in the slice thickness direction (S direction), that is, a slice encoding gradient magnetic field G.
When the center of s is coincident with the center of the magnetic field (not the direction of the magnetic field), φ1 = 0 °, φ7
= 180 °, φ2 to φ6 are shifted from each other by 180 ° / 7.

Then, since the received signal is detected based on the reference signal having the phase offsets of φ1, φ2, ..., The DC components contained in the received signal before detection are those phase offsets φ1, φ2. Appears as having a phase corresponding to, .... Therefore, in each data in the slice encode direction (depth direction in FIG. 3), the DC component appears as a phase component of φ1, φ2, ... It shifts to # 1 and # 7.

On the other hand, as shown in FIG. 5, slices # 4 and G located at the center of the seven slices # 1 to # 7.
In the case of eccentricity imaging in which the center of s is different, the carrier signal offsets are φ1 = 0 ° and φ7 = 180 as described above.
And φ2 to φ6 are each shifted by 180 ° / 7, the result of Fourier transform is -90 ° from the center phase.
Since the phase component in the range of + 90 ° is obtained and the positional relationship is reproduced by this, the slices # 1 and # 2 in FIG. 5 are reproduced as being at the positions shown by the dotted lines. Here, the position of slice # 6 is assumed to be the center of Gs as shown in FIG. Therefore, when the offset of the carrier signal is determined as described above, the DC component shifts to slice # 3 and slice # 2 (dotted line). After the Fourier transform, this positional relationship is returned from the dotted line position to the solid line position to normalize the positional relationship of each slice in the slice thickness direction. Then, the adjacent slices # 2 and # 3 near the center are processed. DC artifacts will appear in and.

Therefore, in such eccentric imaging, φ6 =
0 °, φ5 = 180 °, φ7, φ1, φ2, φ3,
φ4 is to be shifted from each other by 180 ° / 7. As a result, seven slices # 3, # 4, # 5, # 6, # 7 and # 1 are obtained as phase components in the range of -90 ° to + 90 ° by Fourier transform in the slice encoding direction.
When (dotted line) and # 2 (dotted line) are obtained, the DC artifacts appear as the fifth slice # 7 and the sixth slice # 1 (dotted line), so these slices # 1 and # 2 are represented by the solid line. The DC artifacts are transferred to the slices # 1 and # 7 at both ends when the DC artifact is returned to the position.

An image of each slice is obtained in this way. Since the phase of the carrier signal is inverted by 180 ° for each line of raw data in the phase encoding direction (vertical direction in FIG. 3), this carrier signal is By performing phase detection as the reference signal, the DC component contained in the reception signal before detection becomes a component having a 180 ° phase difference for each line. Therefore, in the image of each slice, the DC artifact appears at both ends of the gradient magnetic field Gp in the gradient direction and can be inconspicuous.

Although the gradient echo sequence is used in the above embodiment, it is needless to say that another pulse sequence such as a spin echo sequence can be used. Also, the order of phase encoding is changed in order from one with a larger positive phase encoding amount to one with a smaller positive phase encoding amount (and one with a larger negative amount). The order is not limited to this. The same applies to slice encoding, and in the above, the slice encoding amount is changed in order from the larger positive one to the smaller one (and the larger negative one), but a signal with each slice encoding amount can be obtained. Then, the temporal order can be different. Furthermore, the sequence encoding is changed for one slice encoding amount and the sequence is repeated. After all the phase encoding is completed, the next slice encoding amount is set and the phase encoding amount is not changed. The combination order of the phase encoding amount and the phase encoding amount may be random. Further, the functions of the host computer 51 and the sequence controller 52 may be separated.
Further, the specific configuration can be variously changed by inserting the anti-aliasing filter 42 after the phase detection circuit 43 in the receiving system.

[0025]

As described in the above embodiments, according to the nuclear magnetic resonance imaging apparatus of the present invention, even if the direct current component existing in the receiving system is unavoidable in three-dimensional imaging, its adverse effect is suppressed as much as possible. In addition, it is possible to obtain a three-dimensional image with excellent image quality.

[Brief description of drawings]

FIG. 1 is a block diagram of a nuclear magnetic resonance imaging apparatus according to an embodiment of the present invention.

FIG. 2 is a time chart showing a pulse sequence in the embodiment.

FIG. 3 is a diagram showing a three-dimensional raw data space.

FIG. 4 is a schematic diagram illustrating the positional relationship between the position of each slice and the slice encoding gradient magnetic field Gs.

FIG. 5 is a schematic diagram illustrating another positional relationship between the position of each slice and the slice encoding gradient magnetic field Gs.

[Explanation of symbols]

 10 Main Magnetic Field Magnet 11 Gradient Magnetic Field Coil 12 RF Coil 21 Digital Gradient Magnetic Field Waveform Generator 22, 32 D / A Converter 23 Gradient Magnetic Field Power Amplifier 31 Digital RF Waveform Generator 33 Amplitude Modulation Circuit 34 RF Oscillation Circuit 35 RF Power Amplifier 41 Preamplifier 42 Filter 43 Phase detection circuit 44 A / D converter 51 Host computer 52 Sequence controller

Claims (1)

[Claims] 1. A main magnetic field generating means for generating a static magnetic field, and a means for generating a gradient magnetic field pulse for slice encoding,
A unit for generating a gradient magnetic field pulse for reading, a unit for generating a gradient magnetic field pulse for phase encoding, and an RF in which an RF carrier signal is amplitude-modulated on a subject placed in a magnetic field.
A means for irradiating a pulse, a receiving means for receiving the NMR signal from the subject and performing phase detection using the carrier signal as a reference signal and then sampling to obtain data,
A processing means for arranging the data in the raw data space and performing a multi-dimensional Fourier transform process, and a different offset in the phase of the carrier signal for each slice encode, and the difference in the phase offset amount is 180 ° as a whole. A phase control means for changing the phase offset amount according to the slice encode amount, and inverting the phase of the carrier signal for each adjacent phase encode in each slice encode. Magnetic resonance imaging system.