JP4763874B2 - Magnetic resonance imaging system - Google Patents

Description

[0001]
BACKGROUND OF THE INVENTION
The present invention relates to a magnetic resonance imaging apparatus that reduces noise based on a change in pulse sequence.
[0002]
[Prior art]
In general, a magnetic resonance imaging apparatus main body for obtaining a magnetic resonance image is provided with a gantry for forming a magnetic resonance diagnosis space and inserting a subject into the diagnosis space. The gantry is equipped with various units, especially a static magnetic field magnet such as a superconducting magnet that generates a static magnetic field in the diagnostic space, a gradient coil that generates a linear gradient magnetic field superimposed on the static magnetic field, and a high-frequency signal transmission / reception An RF coil or the like is an essential element. When imaging a magnetic resonance image, a static magnetic field magnet, a gradient magnetic field coil, an RF coil, and the like are driven according to a pulse sequence, and magnetic resonance signals necessary for imaging are collected. A magnetic resonance image representing a cross section of the subject is generated based on the collected magnetic resonance signals.
[0003]
In the conventional magnetic resonance imaging apparatus, there is a problem that electromagnetic force is generated by driving the gradient coil according to the pulse sequence, which causes mechanical distortion (vibration) in the gradient coil and generates noise. . Needless to say, noise causes discomfort and fear to the subject, and it is necessary to reduce this as much as possible.
[0004]
In view of this, several approaches have been tried to reduce noise, such as considering the gradient magnetic field coil as a noise generation source and accommodating the gradient magnetic field coil in a vacuum vessel in order to suppress the propagation of the vibration of the gradient magnetic field coil to the outside.
[0005]
[Problems to be solved by the invention]
However, the conventional approaches have been mainly aimed at improving the mechanical structure (mechanical) of the apparatus, such as housing the gradient coil in a vacuum vessel as described above.
[0006]
The present invention has been made in view of the above circumstances, and an object of the present invention is to realize a new noise reduction different from that due to mechanical structural improvement in a magnetic resonance imaging apparatus.
[0007]
[Means for Solving the Problems]
In order to solve the above problems and achieve the object, the present invention is configured as follows.
[0008]
The magnetic resonance imaging apparatus according to the present invention uses the first pulse sequence in which the slew rate of the gradient magnetic field is the maximum slew rate regardless of the strength of the gradient magnetic field, and the rising time of the gradient magnetic field as the gradient magnetic field strength. Regardless of whether the second pulse sequence is set to the time when the gradient magnetic field rises to the maximum gradient magnetic field intensity at the maximum slew rate, the first pulse sequence or the second pulse can be switched. According to a sequence, a gradient magnetic field and an RF pulse are applied to a subject arranged in a static magnetic field, a magnetic resonance signal is obtained from the subject, and the magnetic resonance signal is reconstructed and the subject is reconstructed. A magnetic resonance imaging apparatus.
[0012]
DETAILED DESCRIPTION OF THE INVENTION
Hereinafter, embodiments of the present invention will be described with reference to the drawings.
[0013]
FIG. 1 is a block diagram showing a schematic configuration of a magnetic resonance imaging apparatus according to an embodiment of the present invention. In the gantry 20, a static magnetic field magnet 1, an X-axis / Y-axis / Z-axis gradient magnetic field coil 2, and a transmission / reception coil 3 are provided. The transmission / reception coil 3 is embedded in the top plate of the bed 13. Alternatively, the transmission / reception coil 3 may be directly attached to the subject. In addition, separate coils dedicated for transmission and reception may be used instead of such a coil for both transmission and reception. The static magnetic field magnet 1 as a static magnetic field generator is configured using, for example, a superconducting coil or a normal conducting coil. The X axis / Y axis / Z axis gradient magnetic field coil 2 is a coil for generating an X axis gradient magnetic field Gx, a Y axis gradient magnetic field Gy, and a Z axis gradient magnetic field Gz. The transmission / reception coil 3 generates a radio frequency (RF) pulse as a selective excitation pulse for selecting a slice, and is used to detect a magnetic resonance signal (MR signal) generated by magnetic resonance. The subject P placed on the top plate of the bed 13 is inserted into an imageable region (a spherical region in which an imaging magnetic field is formed and can be diagnosed only in this region) in the gantry 20. The
[0014]
The static magnetic field magnet 1 is driven by a static magnetic field control device 4. The transmitter / receiver coil 3 is driven by a transmitter 5 during RF transmission, and is coupled to a receiver 6 when a magnetic resonance signal is detected. The X axis / Y axis / Z axis gradient magnetic field coil 2 is driven by an X axis gradient magnetic field power source 7, a Y axis gradient magnetic field power source 8, and a Z axis gradient magnetic field power source 9.
[0015]
The X-axis gradient magnetic field power supply 7, the Y-axis gradient magnetic field power supply 8, the Z-axis gradient magnetic field power supply 9, and the transmitter 5 are driven by the sequencer 10 according to a predetermined sequence, and the X-axis gradient magnetic field Gx, Y-axis gradient magnetic field Gy, Z-axis gradient magnetic field A magnetic field Gz and a radio frequency (RF) pulse are generated according to a predetermined pulse sequence described later. In this case, the X-axis gradient magnetic field Gx, the Y-axis gradient magnetic field Gy, and the Z-axis gradient magnetic field Gz are mainly used as, for example, a phase encode gradient magnetic field Ge, a read gradient magnetic field Gr, and a slice gradient magnetic field Gs. The computer system 11 drives and controls the sequencer 10 and takes a magnetic resonance signal received by the receiver 6 and performs predetermined signal processing to generate a tomographic image of the subject and display it on the display unit 12.
[0016]
In this embodiment, the X-axis gradient magnetic field Gx, the Y-axis gradient magnetic field Gy, and the Z-axis gradient magnetic field Gz are collectively referred to as “gradient magnetic field”, and the gradient magnetic field unit applied to the gradient magnetic field in accordance with the pulse sequence is described. The amount of change per time is referred to as “slew rate (dB / dt)”. In particular, the case where the slew rate is constant is referred to as “constant slew rate (CSR)”. The case where the rising time of the gradient magnetic field is constant is referred to as “constant rise time (CRT)”.
[0017]
If attention is paid to the symmetry of the trapezoidal gradient magnetic field pulse waveform, if the gradient magnetic field pulse rises at a certain rate, the pulse falls at the same rate. If the gradient magnetic field pulse rises at a certain time, the time at which the pulse falls also becomes constant at the same time.
[0018]
Hereinafter, some embodiments relating to the noise reduction function focusing on the waveform characteristics of the gradient magnetic field on the pulse sequence will be described. Note that the embodiments described below can be implemented in combination with other embodiments as appropriate.
[0019]
(First embodiment)
FIG. 2 is a diagram illustrating a waveform of a gradient magnetic field pulse according to the first embodiment. In the figure, P1 shows the waveform of the gradient magnetic field pulse when the constant rise time (CRT) is applied, and P2 shows the gradient magnetic field pulse when the constant slew rate (CSR) is applied. The waveform is shown. The waveforms of these gradient magnetic field pulses P1 and P2 are both trapezoidal waves.
[0020]
In the present embodiment, it is assumed that the gradient magnetic field pulses P1 and P2 shown in FIG. 2 do not have the maximum intensity. That is, the maximum value of the current supplied to the gradient coil that is the source of these pulses does not become the maximum value MAX shown in the graph.
[0021]
The gradient magnetic field pulse P2 to which the constant slew rate is applied has a waveform characteristic that always rises at a constant slew rate. On the other hand, the gradient magnetic field P1 to which the constant rise time is applied has a waveform characteristic that it always rises with a constant rise time. In the pulse waveform diagram, the areas of the regions surrounded by the waveforms of the gradient magnetic field pulses P1 and P2 are the same.
[0022]
For the case where the strength of the gradient magnetic field does not become the maximum, both the gradient magnetic field pulses P1 and P2 are compared. As can be seen from FIG. 2, the rising gradient of the pulse is smaller for the constant rise time gradient magnetic field pulse P1 than for the constant slew rate gradient magnetic field pulse P2. When the rising gradient is smaller (gradual), the intensity of generation of electromagnetic force that causes mechanical distortion becomes smaller. Therefore, the noise at the constant rise time is smaller than that at the constant slew rate only when the intensity of the gradient magnetic field does not become maximum. When the gradient magnetic field strength is raised to the maximum, there is no difference in the pulse waveform between the constant slew rate and the constant rise time, that is, the gradient magnetic field pulses P2 and P1 have the same waveform. However, in the pulse sequence related to magnetic resonance imaging, the maximum gradient magnetic field strength is not always produced, so the constant rise time is less noise than the constant slew rate when viewed throughout the pulse sequence. Become. More specifically, the noise reduction effect is about several decibels to tens of decibels [dB].
[0023]
FIG. 3 is a graph showing an example of noise measurement results for both constant slew rate (CSR) / constant rise time (CRT). According to the graph, first, it can be seen that CRT has lower noise than CSR. In CSR, noise increases monotonically with increasing CSR values (steeper rises).
[0024]
The apparatus of the present embodiment is configured to be able to appropriately switch between a constant slew rate pulse sequence and a constant rise time pulse sequence for the same magnetic resonance imaging.
[0025]
More specifically, for the same (specifically, for example, the same spin echo sequence) magnetic resonance imaging, a pulse sequence for normal imaging to which a constant slew rate is applied and a constant rise time are applied. And a pulse sequence for noise reduction imaging. These pulse sequences are stored in a storage unit (not shown) connected to the computer system 11. When a predetermined instruction is input from the operator, a read command corresponding to the instruction input is issued from the computer system 11, and a pulse sequence is selectively read from the storage unit and supplied to the sequencer 10. Thus, by changing the pulse sequence supplied to the sequencer 10, switching between normal imaging (CSR) and noise reduction type (CRT) imaging can be performed.
As described above, according to the present embodiment, a new noise reduction function without mechanical structural change based on switching between a constant slew rate pulse sequence and a constant rise time pulse sequence is provided. A magnetic resonance imaging apparatus can be provided.
[0026]
(Second Embodiment)
The second embodiment relates to noise reduction through optimization of constant rise time (CRT) and constant slew rate (CSR). In the second embodiment, the pulse sequence is optimized for each of the CRT and CSR in accordance with imaging. The second embodiment can be applied to each of a CRT type magnetic resonance imaging apparatus or a CSR type magnetic resonance imaging apparatus, and can be switched between a CRT mode and a CSR mode as in the first embodiment. It can also be applied to magnetic resonance imaging apparatuses.
[0027]
FIG. 4 is a graph showing the relationship between the CRT value and the sound pressure [Pa] of noise in a CRT magnetic resonance imaging apparatus. In this graph, the horizontal axis represents the 1 / CRT value (rise time).
[0028]
In the present embodiment, in consideration of the fact that the maximum gradient magnetic field rise is not always required in normal imaging, the minimum CRT value required for the imaging is set.
[0029]
For example, as shown in the graph of FIG. 4, when the current CRT value is Vr1 and the minimum CRT value necessary for the imaging is Vr2, the CRT value is changed from Vr1 to Vr2. As a result, the sound pressure of noise can be reduced from Pr1 to Pr2.
[0030]
In extending the CRT value, a high noise reduction effect is obtained until a certain CRT value is reached, but the reduction effect tends to decrease after the CRT value. Specifically, it is not effective to extend the CRT value beyond 1.2 to 1.5 [ms].
[0031]
On the other hand, FIG. 5 is a graph showing the relationship between the CSR value (slew rate) and the sound pressure [Pa] of noise in the CSR magnetic resonance imaging apparatus. As already described with reference to FIG. 3, in CSR, noise increases monotonically with increasing CSR value (steeper rise). In turn, reducing the CSR value also reduces noise.
[0032]
As in the case of CRT, in consideration of the fact that the maximum gradient magnetic field rise is not always required in normal imaging, the minimum CSR value required for the imaging is set.
[0033]
For example, as shown in the graph of FIG. 5, when the current CSR value is Vs1 and the minimum CSR value required for the photographing is Vs2, the CSR value is changed from Vs1 to Vs2. As a result, the sound pressure of noise can be reduced from Ps1 to Ps2.
[0034]
As for the switching of CRT value or CSR value, there are two realization methods: automatic switching and manual switching. The automatic switching is configured such that a minimum CRT value or CSR value necessary for the photographing is obtained in advance for each photographing, and the value is automatically switched according to the photographing selection. One implementation relies on selective switching between preset pulse sequences corresponding to CRT values or CSR values.
[0035]
On the other hand, in manual switching, the operator can manually switch the CRT value or CSR value to a desired value via the input device.
[0036]
According to the second embodiment, finer noise drive control can be performed in each of the CRT or CSR.
[0037]
(Third embodiment)
In the second embodiment, noise reduction is achieved by switching the CSR value to a value that provides the minimum necessary shooting conditions. On the other hand, in the present embodiment, the CSR value is obtained in reverse from the allowable noise level, and the limit photographing condition is determined from the CSR value. In particular, this embodiment is effective when noise reduction is given the highest priority in order to obtain a quiet photographing environment such as photographing children.
[0038]
This embodiment will be described with reference to FIG. When the allowable noise level Ps is set, an allowable CSR value Vs is obtained. As can be seen from the figure, the sound pressure [Pa] of the noise is proportional to the CSR value. Next, the limit of photographing conditions is obtained from Vs. The imaging conditions here are a minimum FOV (field of view), a minimum slice thickness, a maximum number of images to be captured, and the like.
[0039]
As a modification of the present embodiment, for each allowable noise level, a plurality of pulse sequences whose noise is lower than that is associated with each other, and the corresponding plurality of pulse sequences are displayed in a list according to the selection of the allowable noise level. The operator may select a desired sequence.
[0040]
In the case of the present embodiment, photographing exceeding the limit of the photographing condition is not performed, and instead, noise does not exceed the allowable level Ps, and desired silent photographing can be realized. Needless to say, this embodiment can be applied not only to CSR but also to a CRT magnetic resonance imaging apparatus. Further, the present invention is not limited to the above-described embodiment, and can be implemented with various modifications.
[0041]
【The invention's effect】
As described above, according to the present invention, it is possible to provide a magnetic resonance imaging apparatus that realizes a new noise reduction different from that due to mechanical structural improvements.
[Brief description of the drawings]
FIG. 1 is a block diagram showing a schematic configuration of a magnetic resonance imaging apparatus according to an embodiment of the present invention. FIG. 2 is a diagram showing a waveform of a gradient magnetic field pulse according to the first embodiment. Graph showing an example of noise measurement results for both (CSR) / constant rise time (CRT). Fig. 4 shows the relationship between CRT value and noise sound pressure [Pa] in a CRT magnetic resonance imaging system. Graph [Fig. 5] Graph showing the relationship between CSR value (slew rate) and noise sound pressure [Pa] in CSR magnetic resonance imaging system [Fig. 6] The relationship between allowable noise level and slew rate value Graph to show 【Explanation of symbols】
DESCRIPTION OF SYMBOLS 1 ... Static magnetic field magnet 2 ... X-axis / Y-axis / Z-axis gradient magnetic field coil 3 ... Transmission / reception coil 4 ... Static magnetic field control apparatus 5 ... Transmitter 6 ... Receiver 7 ... X-axis gradient magnetic field amplifier 8 ... Y-axis gradient magnetic field amplifier 9 ... Z-axis gradient magnetic field amplifier 10 ... Sequencer 11 ... Computer system 12 ... Display unit

Claims (2)

The first pulse sequence in which the slew rate of the gradient magnetic field is the maximum regardless of the strength of the gradient magnetic field and the rise time of the gradient magnetic field is the maximum slew rate regardless of the strength of the gradient magnetic field. Means capable of switching between the second pulse sequence, which is the time when the gradient magnetic field rises up to the maximum gradient magnetic field intensity at the rate,
In accordance with the first pulse sequence or the second pulse sequence, a gradient magnetic field and an RF pulse are applied to a subject placed in a static magnetic field to obtain a magnetic resonance signal from the subject, and the magnetic Means for obtaining a magnetic resonance image of the subject by reconstructing the resonance signal;
A magnetic resonance imaging apparatus comprising: It said switchable means is the same spin echo sequence, a magnetic resonance imaging apparatus according to claim 1, characterized in that switching between, said second pulse sequence and said first pulse sequence.

Images