JPH11262478A - Magnetic resonance imaging device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To reduce imaging time by reducing the rise time of a readout tilted magnetic field. SOLUTION: A tilted magnetic field applied at the same time as the irradiation of an amplitude modulated RF burst 1 to a subject for imaging is produced as composite magnetic fields by use of tilted magnetic field generators for two of three orthogonal axes, and then the titled magnetic fields of the tilted magnetic field generators for the two axes have their polarities inverted. A group of echoes is generated by the composite magnetic fields to reduce the intensity of the titled magnetic field produced by each of the tilted magnetic field generators of the two axes, to thereby reduce the rise time of each tilted magnetic field. In this case, the amplitude of the titled magnetic field of one of the titled magnetic field generators during the generation of the echoes is set to be greater than that of the other titled magnetic field applied during the irradiation of the RF burst, and the amplitude of the tilted magnetic field of the other titled magnetic field generator during the generation of the group of echoes is set to be smaller than the amplitude of the other tilted magnetic field applied during the irradiation of the RF burst.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a magnetic resonance imaging apparatus, and more particularly to a magnetic resonance imaging apparatus using an RF burst composed of a plurality of sub-pulses as an excitation RF pulse.

[0002]

2. Description of the Related Art A magnetic resonance imaging apparatus or a magnetic resonance imaging apparatus is an apparatus for measuring a tomographic image of a subject using nuclear magnetic resonance. That is, in order to excite the nuclei in the subject, a high-frequency magnetic field pulse (RF pulse) having a frequency proportional to the strength of the static magnetic field is applied to the subject placed in the static magnetic field, and the magnetic resonance generated from the subject accordingly. A signal is detected, and a tomographic image is reconstructed based on the detected magnetic resonance signal.

An imaging method using an RF burst in which the amplitude of a plurality of high frequency magnetic field sub-pulses discretely arranged on a time axis is modulated by a function such as a sinc function as an excitation RF pulse for exciting an atomic nucleus in a subject is known. (Japanese Patent Application No. 7-117015). For example, when a Fourier transform is performed on an RF burst on the time axis that is amplitude-modulated by the sinc function, a square periodic wave having a specific width is obtained on the frequency axis. Here, assuming that the interval between the sub-pulses constituting the RF burst on the time axis is u [second], the period of the square periodic wave on the frequency axis is 1 / u [Hz]. Also, assuming that the cycle of the sinc function obtained by amplitude-modulating the RF burst on the time axis is T [sec],
The width of the square periodic wave on the frequency axis is 1 / T [Hz].
In a magnetic resonance imaging apparatus, since a frequency band measured by a gradient magnetic field is allocated to coordinates in a real space,
The intensity of the gradient magnetic field applied simultaneously with the excitation RF pulse is R
In the case of no change during the F pulse application period, the waveform on the frequency axis represents the excitation profile of the nucleus in real space as it is, that is, the absolute value of the transverse magnetization. In particular, when an RF burst subjected to amplitude modulation with T = 2u is used as an excitation RF pulse, an excited region and an unexcited region appear alternately in the same volume. Therefore, the carrier frequencies of the amplitude-modulated RF burst with T = 2u are 1 /
When two RF bursts shifted (changed) by (2u) [Hz] are used, the excited region and the non-excited region are just interchanged, so that almost all of the nuclei in the imaging cross section of the subject are excited. Shooting can be performed. That is, the first of the two amplitude modulated RF bursts
Immediately after imaging by applying the amplitude-modulated RF burst to the subject, imaging is performed by applying the second amplitude-modulated RF burst to the subject, and the two are subjected to two-dimensional inverse Fourier transform and then combined to achieve resolution. It is known to improve the performance (Japanese Patent Application No. 8-74960).

FIG. 8 shows a conventional example of a photographing sequence using such an amplitude-modulated RF burst. 8, the horizontal axis represents time, the vertical axis represents the intensity of an amplitude-modulated RF burst or a gradient magnetic field, and the timing chart of each axis shows, in order from the top, a high-frequency magnetic field RF, a gradient magnetic field Gx in the x-axis direction, and a gradient magnetic field Gx in the y-axis direction. The gradient magnetic field Gy, the gradient magnetic field Gz in the z-axis direction, and the echo echo are shown. The excitation RF burst 1 has five RF sub-pulses amplitude-modulated by a sinc function. As shown, an excitation RF burst 1 and a gradient magnetic field 32 in the x direction are shown.
Applying -0 at the same time excites nuclei in the subject.
Next, when the subject is simultaneously applied with the 180-degree pulse 31 and the slice gradient magnetic field 4 in the z direction, a section having a specific width perpendicular to the direction in which the slice gradient magnetic field is applied (imaging section)
The magnetic moments of the nuclei at are reversed, and the magnetic moments of nuclei outside the cross section have different phases. Next, a spin echo 6-1 is generated by applying a readout gradient magnetic field 32-1-1 in the x-axis direction. After observing the echo 6-1 and inverting the polarity of the readout gradient magnetic field 32 and applying the gradient magnetic field 32-2-1, a field echo 6-2 is generated. As described above, the five echoes corresponding to the five sub-pulses are regarded as one set, and the readout gradient magnetic field inversion is repeated to generate n sets of echo groups, and the number of echoes required for image reconstruction is measured.
The gradient magnetic field Gy in the y-direction is repeatedly applied as a phase encoding gradient magnetic field 33 in reverse with the application and reversal of the readout gradient magnetic field 32.

[0005]

In the conventional photographing sequence shown in FIG. 8, an echo is measured (sampling).
What can be done is a portion where the intensity of the readout gradient magnetic field 32 in the x direction is constant (32-1-1, 32-2-1,..., 32).
-n-1), and the portions (32-1-2, 32-2-2,...) where the intensity of the readout gradient magnetic field 32 changes.
., 32-n-2) is wasted time in which echo cannot be measured. Here, if the rise time or fall time (hereinafter, collectively referred to as rise time) during which the intensity of the gradient magnetic field changes can be reduced, the photographing time can be reduced.

FIG. 8 shows that the rise time can be reduced by increasing the change speed of the gradient magnetic field at the time of reversal. That is, the readout gradient magnetic field 3
2 changing part (32-1-2,32-2-2, ...
., 32-n-2) may be increased. But,
In order to increase the rate of change related to the reversal of the gradient magnetic field, it is necessary to increase the capacity of the gradient magnetic field power supply, which increases the cost and is limited.

In FIG. 8, if the amplitude of the read-out gradient magnetic field 32 is reduced, the rise time is reduced even if the gradient magnetic field changes at the same speed. However, the field of view of the magnetic resonance imaging apparatus is inversely proportional to the product of the readout gradient magnetic field intensity and the sampling rate (time interval when sampling a signal). Therefore, the field of view increases when the readout gradient magnetic field intensity is reduced. As the field of view increases, the spatial resolution of the image deteriorates. On the other hand, if the intensity of the readout gradient magnetic field is reduced, the sampling rate is increased, and the product of the two is kept constant, the field of view does not change. Is a fixed part (32-1-1, 32-2-1, ...,
Since the time of (32-n-1) is lengthened, it does not meet the purpose of shortening the shooting time.

As described above, if the intensity of the read-out gradient magnetic field is reduced in order to reduce the rise time of the read-out gradient magnetic field, the field of view becomes large, so that the spatial resolution is deteriorated. If the capacity of the gradient magnetic field power supply is increased, the rise time of the readout gradient magnetic field can be reduced, but there is a disadvantage that the cost increases.

SUMMARY OF THE INVENTION It is an object of the present invention to solve the above problems and to reduce the rise time of the read-out gradient magnetic field and shorten the photographing time without increasing the cost.

[0010]

According to the magnetic resonance imaging apparatus of the present invention, in order to solve the above-mentioned problems, an object arranged in a predetermined magnetic field space is subjected to an R-modulation with amplitude modulation by a sinc function.
After applying an excitation RF pulse composed of an F burst and a gradient magnetic field in a predetermined direction almost simultaneously, an echo accompanying nuclear magnetic resonance generated in the subject is applied with a gradient magnetic field in substantially the same direction as the gradient magnetic field in the predetermined direction. In order to generate a tomographic image of the subject based on the detected echo group, a gradient magnetic field in a predetermined direction is applied at the same time as applying the excitation RF pulse to the subject. Is generated as a composite magnetic field by driving two gradient magnetic field generators that generate a gradient magnetic field of two axes out of three orthogonal axes, and the polarity of the gradient magnetic fields of the two gradient magnetic field generators is inverted to generate an echo group. Is generated.

The magnitude of the combined magnetic field of two gradient magnetic field generators having two orthogonal axes is √2 times the magnetic field intensity generated by one gradient magnetic field generator. If it is set so as to be strong, the magnetic field strength generated by one gradient magnetic field generator may be 1 / √2 times the conventional value. Accordingly, the rise time is also reduced by a factor of 1 / 、 2. Therefore, without increasing the cost of increasing the capacity of the gradient magnetic field generator, the rise time of the read-out gradient magnetic field is reduced and imaging is performed. Time can be reduced.

Here, the amplitude of the gradient magnetic field at the time of generation of the echo group generated by each of the two gradient magnetic field generators is
It can be almost the same.

The amplitude of the first axial gradient magnetic field generated by one of the two gradient magnetic field generators at the time of generation of the echo group is the first gradient magnetic field applied simultaneously with the application of the excitation RF pulse. The amplitude of the second gradient magnetic field generated by the other gradient magnetic field generator when the echo group is generated is greater than the amplitude of the gradient magnetic field in the axial direction
By setting the amplitude to be smaller than the amplitude of the gradient magnetic field applied in the second axial direction at the same time as the application of the pulse, it is possible to shift the phase in the k-space in the phase encoding direction. In this case, the time when the amplitude of each of the gradient magnetic fields generated from the two gradient magnetic field generators applied when the echo group is first generated becomes zero differs between the two. Also, the time when the output of the first axial gradient magnetic field applied when the echo group is first generated is zero,
This is set to be later than the time when the output of the gradient magnetic field in the second axial direction applied when the echo group is first generated becomes zero.

[0014]

Embodiments of the present invention will be described below with reference to the drawings. (First Embodiment) FIG. 1 shows an imaging sequence of a magnetic resonance imaging apparatus according to an embodiment of the present invention, and FIG. 2 shows a configuration example of the magnetic resonance imaging apparatus. As shown in FIG. 2, the magnetic resonance imaging apparatus includes a magnet 101 for generating a static magnetic field, and a gradient magnetic field generating coil 102 for generating a gradient magnetic field, and a subject 103 is placed in a magnetic field formed by these. In general, the gradient magnetic field generating coil 102 is composed of three-axis gradient magnetic field coils of x, y, and z. Then, the sequencer 104 sends a command to the gradient magnetic field power supply 105 and the RF pulse generator 106, generates a gradient magnetic field by the gradient magnetic field generating coil 102, and causes the probe 107 to generate an RF pulse.

Normally, the RF pulse is generated by the RF pulse generator 1
06 is amplified by the RF power amplifier 115 and applied to the subject 103 through the probe 107. The signal generated from the subject 103 is received by the probe 107 and detected by the receiver 108. The magnetic resonance frequency used as a reference for detection is set by the sequencer 104. The detected signal is sent to a computer 109, where signal processing such as image reconstruction is performed, and the result is displayed on a display 110. If necessary, the storage medium 1
11 can also store signals and measurement conditions.

When it is necessary to adjust the uniformity of the static magnetic field, the shim coil 112 is used. The shim coil 112 includes a plurality of channels, and a current is supplied from a shim power supply 113. When adjusting the uniformity of the static magnetic field, the current flowing through each coil is controlled by the sequencer 104. The sequencer 104 sends a command to the shim power supply 113 to generate an additional magnetic field from the coil 112 so as to correct the non-uniformity of the static magnetic field. The sequencer 104 normally controls each device to operate at a timing and intensity programmed in advance. Among these programs, the one that particularly describes the RF pulse, the gradient magnetic field, and the timing and intensity of signal reception is called an imaging sequence.

Here, the photographing sequence of the embodiment of the present invention shown in FIG. 1 will be described. The horizontal axis in FIG. 1 is time,
The vertical axis represents the amplitude (intensity) of the RF pulse, the gradient magnetic field, and the like,
RF is a high frequency magnetic field, Gx, Gy, and Gz are x,
The echo is a gradient magnetic field in the y- and z-axis directions. As shown, a gradient magnetic field 2-0 in the x direction and a gradient magnetic field 3-0 in the y direction
At the same time as 0, the excitation RF pulse 1 is applied to excite the nuclei in the subject. Here, the gradient magnetic field 2-0 and the gradient magnetic field 3-0
Are equal in amplitude. Next, when the subject is simultaneously applied with the 180-degree pulse 31 and the gradient magnetic field 4 in the z direction as a slice gradient magnetic field, the magnetic moment of a nucleus in a cross section having a specific width perpendicular to the z direction is reversed, and the magnetic moment moves out of the cross section. The magnetic moment of a certain nucleus is out of phase. next,
The echo 6-1 is generated by applying the readout gradient magnetic field 2-1-1 and the gradient magnetic field 3-1-1. Echo 6-
After observing 1, the polarity of the gradient magnetic field is reversed to change the gradient magnetic field 2.
When -2-1 and the gradient magnetic field 3-2-1 are applied, a field echo 6-2 is generated.

Here, as shown in FIG. 3, the excitation RF pulse 1 is obtained by amplitude-modulating five sub-pulses 11 of a high-frequency magnetic field with a sinc function 12. The five echoes corresponding to the sub-pulses 11 are regarded as one set, and n sets of echo groups 6 are generated by repeating the inversion of the readout gradient magnetic fields 2 and 3, and the number of echoes necessary for image reconstruction are obtained. Is measured. That is, a gradient magnetic field to be applied at the same time as irradiating the subject with the RF burst amplitude-modulated by the sinc function is generated as a combined magnetic field using a gradient magnetic field coil of two of the three axes x, y, and z. further,
An echo group is generated by reversing the polarity of the gradient magnetic field using the biaxial gradient magnetic field coil. Then, the echo group is measured to reconstruct an image.

FIG. 1 is compared with the imaging sequence of FIG. 8 which is a conventional example. The echo measurement time, that is, for example, the gradient magnetic field 2-1-1 or 3-1-1 in FIG.
And the application time of the gradient magnetic field 32-1-1 in FIG. 8 are the same. On the other hand, the rise time of the gradient magnetic field in FIG. 1 is about 71% of the rise time in FIG.
It has become. This is because the intensity of the composite magnetic field of the gradient magnetic fields in the x and y directions in FIG. 1 is set equal to the gradient magnetic field intensity in the x direction in FIG. That is, when two orthogonal vectors having the same magnitude are combined, the magnitude of the vector becomes √2 times, and the direction becomes a vector in a direction of 45 degrees with respect to the two vectors. Then, it can be understood that the sequence shown in FIG. 1 can shorten the photographing time as compared with the sequence shown in FIG. 8 by the reduced rise time. For example, the photographing time of one image can be reduced by about 15%.

Since the readout gradient magnetic field intensity and the sampling rate are the same in both FIG. 1 and FIG. 8, the field of view is the same, and when the same number of echoes are measured, the spatial resolution is the same. However, when the image obtained by performing the two-dimensional Fourier transform on the echo obtained in the sequence shown in FIG. 8 is the image shown in FIG. 6A, the echo obtained in the sequence shown in FIG. The image obtained by performing the two-dimensional Fourier transform is obtained by rotating FIG. 6A by 45 degrees as shown in FIG. 6B.

The photographing sequence of FIG. 1 will be described in more detail. As described above, in the sequence shown in FIG. 1, the gradient magnetic field applied simultaneously with irradiating the subject with the RF burst amplitude-modulated by the sinc function is x, y, z.
An echo group is generated by using a two-axis gradient magnetic field coil to generate a synthetic magnetic field and using the two-axis gradient magnetic field coil to invert the polarity of the gradient magnetic field. In FIG. 1, the gradient magnetic fields 52 and 53 shown by broken lines show the gradient magnetic fields 2-0 and 3-0 having the same amplitude as the gradient magnetic fields, respectively, and their polarities are simultaneously inverted. That is, the times when the outputs of the gradient magnetic fields 52 and 53 become zero are equal.

In FIG. 1, portions (2-1-1, 1-2-2-1,..., 2-2-1) of the gradient magnetic field 2 in the x direction when the echo group is generated.
It can be seen that the amplitude of (n-1) is set to be smaller than the amplitude of the gradient magnetic field in the x direction applied simultaneously with the RF burst irradiation. Conversely, the portions of the gradient group 3 in the y-direction when the echo group is generated (3-1-1, 3-2-1, ..., 3-
It can be seen that the amplitude of (n-1) is set to be larger than the amplitude of the gradient magnetic field in the y direction applied simultaneously with the RF burst irradiation. Furthermore, the time when the output of the gradient magnetic field 3-1-1 in the y-axis direction applied when the echo group is first generated becomes zero (the portion where the gradient magnetic field changes 3-2-2 crosses the time axis) The time during which the output of the gradient magnetic field 2-1-1 applied in the x-axis direction applied when the echo group is first generated becomes zero (the part where the gradient magnetic field changes is 2-2-2). It can be seen that the time is set later than the time (intersecting the axis). The reason for setting the amplitude of the gradient magnetic field in such a relationship is to provide spatial information corresponding to the scan position in the k-space.

By the way, in the imaging sequence of FIG. 1, the amplitude of the gradient magnetic field in the y-axis direction (3-1-1, 1
3-2-1,..., 3-n-1) is set to be larger than the amplitude of the gradient magnetic field 3-0 in the y direction applied simultaneously with the RF burst irradiation. If the change speed is the same, the rise time of the gradient magnetic field at the time of generation of the echo group is longer than the rise time of the gradient magnetic field 3-0. However, when the number of pixels of the image obtained by the imaging sequence of FIG. 1 is, for example, 128, the amplitude of the gradient magnetic field at the time of generation of the echo group is 3
It is only about 1% larger than the amplitude of −0. Therefore, the rise time of the gradient magnetic field at the time of generation of the echo group is
It is only about 1% longer than the rise time of the gradient magnetic field 3-0. Rather than the effect of increasing the rising time by about 1%, the rising magnetic field is applied as a composite magnetic field using a biaxial gradient magnetic field coil of x and y simultaneously with irradiating the RF burst to the subject. The effect of reducing the time by about 30% is much greater. This effect becomes more remarkable as the number of pixels of the image increases.

As described above, the gradient magnetic field applied at the same time as irradiating the subject with the RF burst amplitude-modulated by the sinc function is applied to the gradient magnetic field coils of two of the three axes x, y, and z. Generated as a synthetic magnetic field,
An echo group is generated by inverting the polarity of the gradient magnetic field using the two-axis gradient magnetic field coil, and a first echo generated from the gradient magnetic field coil of the first axis of any one of the two axes. The amplitude of the axial gradient magnetic field at the time of generation of the echo group is larger than the amplitude of the first axial gradient magnetic field applied simultaneously with the RF burst irradiation, while the second gradient magnetic field generated from the second magnetic field gradient coil is applied. The amplitude of the gradient magnetic field in the axial direction during generation of the echo group is set smaller than the amplitude of the second gradient magnetic field in the second axial direction applied simultaneously with the RF burst irradiation. Further, when the echo group is generated first, The time during which the output of the applied gradient magnetic field in the first axial direction becomes zero is the first time applied when the echo group was first generated.
By setting the time later than the time when the output of the gradient magnetic field in the axial direction becomes zero, the rise time of the gradient magnetic field can be reduced by about 3 times.
0% can be shortened, and as a result, there is an effect that the photographing time can be shortened as compared with the related art.

In FIG. 1, the readout gradient magnetic field application direction is selected in the x direction, the encode gradient magnetic field application direction is in the y direction, and the slice gradient magnetic field application direction is in the z direction. You can choose the direction. For example, the same cross section can be obtained by selecting the readout gradient magnetic field application direction in the y direction, the encode gradient magnetic field application direction in the x direction, and the slice gradient magnetic field application direction in the z direction. Alternatively, another section can be imaged by selecting the readout gradient magnetic field application direction in the z direction, the encode gradient magnetic field application direction in the x direction, and the slice gradient magnetic field application direction in the y direction.

By the way, as described in the prior art,
When the RF burst on the time axis subjected to amplitude modulation by the nc function is subjected to Fourier transform, a square periodic wave 13 having a specific width on the frequency axis is obtained as shown in FIG. Here, as shown in FIG. 3, when the interval between the sub-pulses 11 is u [seconds], the period of the square periodic wave 13 is 1 / u [Hz]. If the cycle of the sinc function is T [seconds], a square periodic wave 1
The width of 3 is 1 / T [Hz]. In a magnetic resonance imaging apparatus, a frequency band measured by a gradient magnetic field is allocated to coordinates in a real space. Therefore, when the intensity of the gradient magnetic field applied simultaneously with the excitation RF pulse is constant during the application period of the RF pulse, the waveform on the frequency axis directly represents the excitation profile in the real space of the nucleus, that is, the absolute value of the transverse magnetization. Will be. In particular, when an excitation RF pulse with T = 2u is used, an excited region and an unexcited region alternately appear in the same volume. Therefore,
As shown in FIG. 3, when two RF bursts in which the carrier frequency of the RF burst of T = 2u is shifted (changed) by 1 / (2u) [Hz] from each other, the excited region and the excited region are excited. The area that has not been replaced is just replaced
The imaging can be performed by exciting almost all of the nuclei in the imaging section of the subject. It is known that the resolution of an image is improved by combining the echo groups corresponding to the two RF bursts after performing the two-dimensional inverse Fourier transform. Note that the function for amplitude-modulating the RF burst does not need to be strictly a sinc function, but may be any function that becomes square when Fourier-transformed.

As a specific example, the above two amplitude modulations R
Immediately after the first amplitude-modulated RF burst of the F burst is applied to the subject and imaging is performed, immediately after the second amplitude-modulated RF burst is applied to the object and imaging is performed, and each is subjected to two-dimensional inverse Fourier transform, , There is a method of synthesis. The details of this method are described in Japanese Patent Application No. 8-74960. FIG. 4 shows the positional relationship between one pixel of the image obtained by this method and the excitation profile. For example, under the condition that the width of the square periodic wave 13 in the excitation profile when the first amplitude modulation RF burst is applied to the subject is 3 mm, and the pixel size of the captured image in the readout gradient magnetic field application direction is 6 mm. Photographing is performed according to the photographing sequence shown in FIG. 1, and a plurality of obtained echoes are subjected to two-dimensional inverse Fourier transform. At this time, an amplitude-modulated RF burst composed of an infinite number of sub-pulses is excited R
If used as an F pulse, the excitation profile will be a perfect square periodic wave. However, since the imaging sequence of FIG. 1 uses five sub-pulses, the excitation profile does not become a perfect square periodic wave, but here, the excitation profile of FIG. 4 can be regarded as a substantially square periodic wave. It will be described as. One pixel of this image has information of only a 3 mm area that is actually excited. A second amplitude-modulated RF burst whose carrier frequency is shifted by 1 / (2u) [Hz] is applied to a subject to be photographed, and subjected to two-dimensional inverse Fourier transform.
The pixel has information of only the 3 mm area that was not excited earlier. If two images are alternately arranged every other pixel and synthesized, an image with a spatial resolution of 3 mm in the readout gradient magnetic field application direction can be obtained. That is, the pixel size of the captured image in the readout gradient magnetic field application direction can be changed from 6 mm to 3 mm. In the application of the first amplitude modulation RF burst and the second amplitude modulation RF burst, the excitation regions are different from each other, so that there is no need to wait for the recovery of the magnetization between the two imagings. As mentioned above,
When five sub-pulses are used, the excitation profile does not become a perfect square periodic wave, so that a slight leakage occurs in the adjacent pixel, but this is not a practical problem. As described above, continuous imaging is performed by using the first amplitude modulation RF burst and the second amplitude modulation RF burst as the excitation RF pulse 1 of the imaging sequence in FIG. 1 to synthesize two images. Thereby, imaging can be performed by exciting almost all of the nuclei in the imaging section of the subject.

(Second Embodiment) FIG. 7 shows another embodiment of the photographing sequence according to the features of the present invention. In FIG. 7, each axis is the same as that in FIG. As shown in the drawing, the present embodiment differs from the imaging sequence of FIG. 1 in that the polarities of the gradient magnetic field 2 in the x-axis direction and the gradient magnetic field 3 in the y-axis direction at the time of readout are reversed. The gradient magnetic field strength in the x- and y-axis directions at the time of excitation is the same as in FIG. 1, but the gradient magnetic field strength in the x-axis direction at the time of readout is smaller than that at the time of excitation, and the gradient magnetic field strength in the y-axis direction is lower at the time of excitation. The difference is that it is set larger.

When FIG. 7 is compared with the conventional photographing sequence of FIG. 8, the time for measuring the echo, that is,
For example, the application time of the gradient magnetic fields 2 and 3 at the time of readout in FIG. 7 is the same as the application time of the gradient magnetic field 32-1-1 in FIG. However, as in the case of FIG. 1, the rise time of the gradient magnetic field in FIG. 7 is about 71% of the rise time in FIG. Accordingly, it can be seen that the sequence shown in FIG. 7 can reduce the photographing time by the amount corresponding to the shortened rise time, as compared with the sequence shown in FIG.

Regarding the photographing sequence of the present embodiment,
This will be described in more detail. The sequence shown in FIG.
A gradient magnetic field applied simultaneously with irradiating an RF burst 1 amplitude-modulated by the nc function to a subject is generated as a composite magnetic field using a gradient magnetic field coil of two of the three axes x, y, and z. . An echo group is generated by inverting the polarity of the gradient magnetic field using the biaxial gradient magnetic field coil. In FIG. 7, the gradient magnetic field 6 indicated by a broken line
Numerals 2 and 63 denote gradient magnetic fields having the same amplitude as the gradient magnetic fields 2-0 and 3-0, respectively, and their polarities are simultaneously inverted.
That is, the times when the outputs of the gradient magnetic fields 62 and 63 become zero are equal. From FIG. 7, the portion of the gradient magnetic field 2 in the x direction when the echo group is generated (2-1-1, 2-2-1,..., 2-n-1)
Is set to be larger than the amplitude of the gradient magnetic field 2-0 in the x direction applied simultaneously with the RF burst irradiation. Further, the amplitude of the portion (3-1-1, 3-2-1,..., 3-n-1) of the gradient magnetic field 3 in the y direction at the time of generation of the echo group is y applied simultaneously with the RF burst irradiation. It can be seen that the amplitude is set smaller than the amplitude of the gradient magnetic field 3-0 in the direction. Further, the time during which the output of the gradient magnetic field 2-1-1 in the x-axis direction applied when the echo group is first generated becomes zero (the portion where the gradient magnetic field changes 2-2-2 crosses the time axis) Is the time during which the output of the gradient magnetic field 3-1-1 in the y-axis direction applied when the echo group is first generated becomes zero (the time where the gradient magnetic field changes, 3-2-2 is the time). It can be seen that the time is set later than the time (intersecting the axis).

As described above, the gradient magnetic field applied simultaneously with irradiating the subject with the RF burst amplitude-modulated by the sinc function is expressed by a gradient magnetic field coil of two axes of three axes x, y and z. Generated as a synthetic magnetic field,
An echo group is generated by inverting the polarity of the gradient magnetic field using the two-axis gradient magnetic field coil, and a first echo generated from the gradient magnetic field coil of the first axis of any one of the two axes. The amplitude of the axial gradient magnetic field at the time of generation of the echo group is larger than the amplitude of the first axial gradient magnetic field applied simultaneously with the RF burst irradiation, while the second gradient magnetic field generated from the second magnetic field gradient coil is applied. The amplitude of the gradient magnetic field in the axial direction during generation of the echo group is set smaller than the amplitude of the second gradient magnetic field in the second axial direction applied simultaneously with the RF burst irradiation. Further, when the echo group is generated first, The time when the output of the applied first gradient magnetic field in the axial direction becomes zero is the second time applied when the echo group is first generated.
By setting the time later than the time when the output of the gradient magnetic field in the axial direction becomes zero, the rise time of the gradient magnetic field can be reduced by about 3 times.
0% can be shortened, and as a result, there is an effect that the photographing time can be shortened as compared with the related art.

Although the imaging sequence according to the present invention has been described above based on the embodiment, similarly to the imaging sequences other than the above, the gradient magnetic field applied simultaneously with the irradiation of the RF burst to the subject is represented by x, x, An echo group is generated by using a gradient magnetic field coil of two axes out of three axes of y and z to generate a synthetic magnetic field, and further inverting the polarity of the gradient magnetic field by using the gradient magnetic field coils of the two axes. ,
The amplitude at the time of generating an echo group of the gradient magnetic field in the first axial direction generated from the gradient coil of the first axis of one of the two axes is the first magnetic field applied simultaneously with the RF burst irradiation. The amplitude of the second axial gradient magnetic field generated from the second axis gradient magnetic field coil when the echo group is generated is larger than the axial gradient magnetic field amplitude, Is set smaller than the amplitude of the gradient magnetic field in the axial direction, and the time when the output of the gradient magnetic field in the first axial direction applied when the echo group is generated for the first time becomes zero, the echo group is first set. By setting the output time of the gradient magnetic field in the second axial direction applied at the time of generation to be zero, the rise time of the gradient magnetic field can be reduced by about 30%. Shortening compared to It is possible.

For example, the echo 6-1 of the first set is 1
The spin echo does not need to be generated by an 80-degree pulse, but may be a field echo generated by reversing the readout gradient magnetic field. Similarly, the echoes of the second and subsequent sets need not be field echoes generated by reversing the readout gradient magnetic field, but may be spin echoes generated by 180-degree pulses. Also, in three-dimensional imaging in which echoes are measured while applying a phase-encoding gradient magnetic field while changing the encoding amount in multiple steps in the z-direction and three-dimensional inverse Fourier transform is performed on the echoes to obtain a three-dimensional image, the present invention is also applicable to the present invention. By using the present invention, the photographing time can be reduced. The directions of applying the gradient magnetic fields of x, y, and z can be selected as desired depending on the purpose. For example, if the slice gradient magnetic field application direction is selected in the y direction, another cross section can be imaged.

[0034]

According to the present invention, the rise time of the gradient magnetic field can be shortened, and the photographing time can be shortened by the shortened time as compared with the related art.

[Brief description of the drawings]

FIG. 1 is a diagram showing an embodiment of a photographing sequence according to a feature of the present invention.

FIG. 2 is an overall configuration diagram of a magnetic resonance imaging apparatus to which the present invention is applied.

FIG. 3 shows the carrier frequency of two amplitude modulated RF bursts as 1
It is a figure explaining the relation on the frequency axis at the time of shifting / (2u) [Hz].

FIG. 4 is a diagram illustrating an excited state inside a subject.

FIG. 5 is a diagram illustrating an RF burst amplitude-modulated by a sinc function.

FIG. 6 is a schematic diagram illustrating an example of a reconstructed image.

FIG. 7 is a diagram showing another embodiment of a photographing sequence according to the features of the present invention.

FIG. 8 is a diagram showing an example of a conventional imaging sequence using an RF burst.

[Explanation of symbols]

 Reference Signs List 1 excitation RF pulse 2 gradient magnetic field in x-axis direction 3 gradient magnetic field in y-axis direction 4 gradient magnetic field in z-axis direction 6 echo 31 180-degree pulse 101 static magnetic field generating magnet 102 gradient magnetic field generating coil 103 subject 104 sequencer 105 gradient magnetic field power supply 106 RF pulse generator 107 Probe 115 RF power amplifier 108 Receiver 109 Computer 110 Display 111 Storage medium 112 Shim coil 113 Shim power supply

Continued on the front page (72) Inventor Hiromichi Shimizu 1-1-1 Uchikanda, Chiyoda-ku, Tokyo Inside Hitachi Medical Corporation

Claims (5)

[Claims] 1. An object arranged in a predetermined magnetic field space includes:
After applying an excitation RF pulse composed of an RF burst amplitude-modulated by a sinc function and a gradient magnetic field in a predetermined direction almost simultaneously, an echo accompanying nuclear magnetic resonance generated in the subject is obtained.
A magnetic resonance imaging apparatus for applying a gradient magnetic field in substantially the same direction as the gradient magnetic field in the predetermined direction, detecting the same as a group of echoes of spin echo or field echo, and creating a tomographic image of the subject based on the detected group of echoes In the method, a gradient magnetic field in a predetermined direction, which is applied simultaneously with the application of the excitation RF pulse to the subject, is generated in two of three orthogonal axes.
Two gradient magnetic field generators for generating an axial gradient magnetic field are driven to generate a combined magnetic field, and the echo groups are generated by inverting the polarities of the gradient magnetic fields of the two gradient magnetic field generators. Magnetic resonance imaging device. 2. The magnetic resonance imaging apparatus according to claim 1, wherein the amplitudes of the gradient magnetic fields generated by the two gradient magnetic field generators when the echo group is generated are substantially the same. 3. An amplitude of the first axial gradient magnetic field generated by one of the two gradient magnetic field generators when the echo group is generated is applied simultaneously with the application of the excitation RF pulse. The second gradient magnetic field generated by the other gradient magnetic field generator, which is larger than the amplitude of the gradient magnetic field in the first axial direction.
The amplitude of the gradient magnetic field in the axial direction when an echo group is generated is set to be smaller than the amplitude of the gradient magnetic field in the second axial direction applied simultaneously with the application of the excitation RF pulse. 2. The magnetic resonance imaging apparatus according to 1. 4. The apparatus according to claim 1, wherein the time when the amplitude of each of the gradient magnetic fields generated from the two gradient magnetic field generators applied when the echo group is first generated becomes zero is different between the two. 4. The magnetic resonance imaging apparatus according to any one of 3. 5. The time during which the output of the first axial gradient magnetic field applied when the echo group is first generated becomes zero,
5. The magnetic resonance imaging apparatus according to claim 4, wherein the output of the gradient magnetic field in the second axial direction applied when the echo group is first generated is set to be later than the time when the output becomes zero.