JP3205061B2 - Magnetic resonance imaging equipment - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to an improvement in image quality of a magnetic resonance imaging apparatus for imaging a desired portion of a subject using magnetic resonance.

[0002]

2. Description of the Related Art Magnetic resonance imaging (hereinafter referred to as MRI)
The apparatus measures the nuclear spin density distribution, relaxation time distribution, and the like at a desired inspection site in the subject using the nuclear magnetic resonance phenomenon, and displays an image of the cross section of the subject from the measurement data. Things. The nuclear spin of the subject placed in a uniform and strong static magnetic field generator performs precession at a frequency (Larmor frequency) determined by the strength of the static magnetic field, with the direction of the static magnetic field as an axis. Then, when a high-frequency pulse having a frequency equal to the Larmor frequency is irradiated from the outside, spins are excited and transit to a high energy state (nuclear magnetic resonance phenomenon). When the irradiation is stopped, the spin returns to the original low energy state with a time constant corresponding to each state, and at this time, an electromagnetic wave (NMR signal) is emitted to the outside. This is detected by a high-frequency receiving coil tuned to that frequency. At this time, a triaxial gradient magnetic field is applied to the static magnetic field space in order to add positional information to the space. As a result, position information in space can be captured as frequency information.

Here, an image reconstruction method in an MRI apparatus will be described. FIG. 6 is an explanatory diagram of a pulse sequence in a commonly used spin echo method. There are two types of irradiation pulses, 90 degrees and 180 degrees.
In addition, it is applied to the subject together with the slice gradient magnetic field, and the cross section to be imaged is determined. A readout gradient magnetic field is applied between these two types of irradiation pulses to promote phase diffusion of the excited spins. Then, the diffusion direction of the spin is reversed by the next applied 180-degree pulse, and when the read-out gradient magnetic field is applied again, the spin converges to 90 ° -1.
A sharp echo signal is generated in twice the time between 80 degree pulses. This time is called an echo time Te. Although the echo signal obtained here has one-dimensional projection image information in the readout direction, a two-dimensional image cannot be formed by this alone. Therefore, during the application of the readout gradient magnetic field for providing phase diffusion, a gradient magnetic field is applied in the encode direction, which is another axis, to give a phase rotation at the position, and information of the encode direction is superimposed on the echo signal as phase information.
Furthermore, the echo signal is repeatedly measured by applying while changing the amount of the encoding gradient magnetic field. This repetition time is called Tr, and the encoding amount is assigned a positive / negative number. Encoding No. zero sets the encoding gradient magnetic field amount to zero and applies a positive / negative encoding amount.

When the echo signal sequence obtained in this way is analyzed by two-dimensional Fourier transform means, two-dimensional image information can be obtained. An image reconstruction method using the two-dimensional Fourier transform is hereinafter referred to as a 2DFT method.

Next, image enhancement that is clinically important in MRI images will be described. There are two types of relaxation phenomena, called longitudinal relaxation and transverse relaxation, that change depending on the environment in which protons spin. By this relaxation phenomenon, the signal strength S is calculated as in the following equation.

S = ρ · (1-exp (−Tr / T1)) · (exp (−Te / T2)) Here, ρ is the density of existing protons. Also, T
1, T2 is a constant peculiar to each tissue, and is a value determined by the environment in which protons exist. In this equation, the second term represents a recovery process of the signal intensity with respect to the repetition period Tr, which is a longitudinal relaxation phenomenon (dependent on the T1 value). The third term is a process of attenuating the signal intensity with respect to the echo measurement time Te, which is due to a lateral relaxation phenomenon (depending on the T2 value).

FIG. 7 illustrates the relationship between the measured echo signal intensity according to Tr and Te. For example, as shown in FIG. Considering this, it is assumed that the signal intensity due to each relaxation phenomenon has the characteristics shown in FIGS. 7A and 7B.

When capturing a T1-weighted image that reflects the T1 value, Te as short as possible is used to suppress the influence of lateral relaxation.
(Short Te, S), Tr is a relatively short value (S, displayed, so that the T1 difference due to the tissue can be drawn)
In general, 500 ms or less is used for the human body. In the case of this example, since the T1 value of the tissue B is shorter, an image in which signal recovery due to longitudinal relaxation is quick and a high signal is obtained is obtained. When capturing a T2-weighted image, a sufficiently long Tr (displayed as long Tr, L, usually 1500 ms) is used to suppress the influence of longitudinal relaxation.
) To recover the signal of each tissue, and set a longer Te (display L, about 100 ms) to perform imaging.

In this example, since the tissue A has a longer T2 value, an image whose contrast is inverted from that of the T1-weighted image is obtained. If Tr is set to be long and Te is set to be short, an image which is not affected by the T1 value and T2 value of each tissue can be obtained. Since this is an image based on the proton density of the tissue, it is called a proton density image. Since the T1 value and the T2 value differ depending on the state (eg, tumor) of the same tissue, they are used for specifying a lesion site. The outline of MRI has been described above. For details, refer to “NMR Medicine” (Basic and Clinical) (published by the Society for Nuclear Magnetic Resonance Medicine, Maruzen Co., Ltd., January 20, 1984).
Day issue).

[0010]

In order to obtain a clinically effective image, its resolution is an important factor. In order to improve the resolution, it is necessary to reduce the pixels constituting the image and increase the image matrix. However, in general, in MRI imaging, an increase in the gradient magnetic field intensity is required to obtain a high-resolution image (called a high-definition image). This principle will be described with reference to FIG.

Higher definition in the readout direction can be realized relatively easily by increasing the sampling time. On the other hand, it is difficult to increase the definition in the encoding direction. FIG.
In order to realize twice the number of pixels in the encoding direction with respect to the standard imaging shown in (a), the relative frequency of the received echo signal can be separated by the 2DFT method only in units of 1 Hz. Must be given twice more to collect high-frequency information, and the applied gradient magnetic field must be increased to twice as shown in FIG.
However, a gradient magnetic field coil that generates a gradient magnetic field is an inductive load to a gradient magnetic field power supply, and it is very difficult to control a large current with high accuracy. Further, such a large changing magnetic field is not preferable in terms of influence on the human body, and it can be said that the smaller the gradient magnetic field intensity, the more desirable.

An object of the present invention is to solve the problems caused by such a gradient magnetic field power supply capacity and strength, and to provide an MRI apparatus capable of performing high-definition imaging with a small gradient magnetic field power supply and obtaining a good image. I do.

[0013]

As shown in FIG. 2, the phase rotation of the spin caused by the application of the gradient magnetic field can be controlled not only by the application level but also by the application time. For example, the phase rotation given to the spin is 90 degrees to 18 degrees.
In the case of doubling the angle to 0 degrees, the same applies even if the application level of the gradient magnetic field is doubled and the application time is doubled. But,
In a normal pulse sequence, level control is used for controlling the phase encoding gradient magnetic field. As described above, this is because the echo acquisition time Te is important for image enhancement, and it is to avoid the elongation or change of Te due to the change of the gradient magnetic field application time.

In the signal measurement by the 2DFT method, the more spin phase rotation in the encoding direction is given, the more detailed information in the encoding direction can be collected. This will be described with reference to FIG. Usually, the number of encodings is set to 128 to 2 in order to secure the resolution of the image.
In this case, the signal data measurement is performed 13 times from encoding No. + 6 to -6 for ease of explanation.
It is assumed that a normal image can be obtained by performing the recirculation process using all the data. Here, the central part (hereinafter,
When the reconstruction process is performed using only seven pieces of encoding information from +3 to -3 of some low encoding parts including the zero encoding (referred to as a center part), only coarse information in the encoding direction (corresponding to a low frequency component) Therefore, the contrast of light and dark of the image can be obtained, but fine sub-information such as contours cannot be obtained, resulting in an unclear image. On the other hand, six encodings (corresponding to high-frequency data) from +4 to +6 and -4 to -6, which are data of the peripheral portion (hereinafter, the encoded portion other than the central portion is referred to as the peripheral portion), are added to the central portion. When a reconstruction process is performed by inserting zero data, an image in which contour information is extracted can be obtained. From this, it can be seen that it is possible to substantially separate the contrast component and the contour component of the image according to the used encoded data range of the reconstruction. For this reason, it can be said that in the peripheral portion that determines the resolution of the image, even if there is an extension of Te, the image enhancement is not significantly affected.

Therefore, in the present invention, as shown in FIG. 1, the control of the amount of the gradient magnetic field is controlled by controlling the amount of applied magnetic field while maintaining the applied time at a predetermined value in the central part which affects image enhancement as in the past (hereinafter referred to as control). , Applied level control), and in the peripheral portion, by switching to control for changing the applied time while maintaining the applied level at a predetermined value (hereinafter referred to as applied time control), the increase of the applied gradient magnetic field level is suppressed. The number of encodes can be increased, and high-definition imaging can be easily realized without strengthening the gradient magnetic field power supply.

[0016]

According to the present invention, encoding No.
By switching the control means of the encoding gradient magnetic field between the level control and the time control in accordance with the above, high-definition imaging can be efficiently performed with a small gradient magnetic field current, and a good MRI image can be captured.

[0017]

DESCRIPTION OF THE PREFERRED EMBODIMENTS Embodiments of the present invention will be described below in detail with reference to the accompanying drawings. FIG. 5 is a block diagram showing an example of the overall configuration of the MRI apparatus according to the present invention. The MRI apparatus obtains a tomographic image of a subject 6 by utilizing a nuclear magnetic resonance (NMR) phenomenon. The MRI apparatus includes a static magnetic field generating magnet 10, a central processing unit (hereinafter, referred to as a CPU) 11, a sequencer 12, and a transmission unit. The system 13 includes a system 13, a gradient magnetic field generation system 14, a reception system 15, and a signal processing system 16.

The static magnetic field generating magnet 10 generates a strong and uniform static magnetic field in the subject 6, and a permanent magnet type, a normal conducting type, or a superconducting type is provided in a space around the subject 6 with a certain width. Are provided. The sequencer 12 operates under the control of the CPU 11 and sends various commands necessary for data collection of tomographic images of the subject 6 to the transmission system 13, the gradient magnetic field generation system 14, and the reception system 1.
5 to send.

The transmission system 13 includes a high-frequency oscillator 17, a modulator 18, a power amplifier 19, and an irradiation coil 20 on the transmission side.
A high-frequency pulse output from the high-frequency oscillator 17 is modulated by a modulator 18 in accordance with a command of the sequencer 12, and the modulated irradiation pulse is
The electromagnetic wave is applied to the irradiation coil 20 disposed in close proximity to the subject 6 after the amplification by the electromagnetic wave, so that the subject 6 is irradiated with the electromagnetic wave.

The gradient magnetic field generating system 14 comprises a gradient magnetic field coil 21 wound in three directions of X, Y and Z, and a gradient magnetic field power supply 22 for driving each coil. By driving the gradient magnetic field power supplies 22 of the respective coils in accordance with, the gradient magnetic fields Gx, Gy, Gz in the X, Y, and Z directions are applied to the subject 6. The slice plane for the subject 6 can be set by the method of applying the gradient magnetic field.

The receiving system 15 includes a receiving coil 2, a receiving circuit 23, a quadrature phase detector 24, and an A / D converter 25. The response electromagnetic wave (NMR signal) is detected by the receiving coil 2 arranged close to the subject 6, input to the quadrature phase detector 24 via the receiving circuit 23, and output from the high-frequency oscillator 17 by the quadrature phase detector 24. The waveform is shaped in synchronization with the output signal, and is separated into two signals of a sin component and a cos component and output. The two signals are then converted into digital signals by an A / D converter 25 and converted into digital signals. The data is sent to the processing system 16.

The signal processing system 16 includes a CPU 11, recording devices such as a magnetic disk 26 and an optical disk 27, and a CRT.
The CPU 11 performs processing such as Fourier transform, correction coefficient calculation, and image reconstruction on the data input from the receiving system 15 to obtain a signal intensity distribution of an arbitrary cross section or an appropriate operation for a plurality of signals. Is formed into an image and displayed on the display 28. In this figure, the irradiation coil 20, the receiving coil 2, and the gradient magnetic field coil 21 are
Static magnetic field generating magnet 10 arranged in a space around subject 6
Are arranged in the magnetic field space.

Here, the relationship between the applied level (absolute value is referred to as a level here) of the encoding gradient magnetic field according to the present invention and the application time will be described with reference to FIG. The application of the encoding gradient magnetic field is controlled by the CPU 11 so that the sequencer 1
2 is performed by giving application timings and levels to the X, Y, and Z gradient magnetic field power supplies 22 and the high frequency pulse modulator 18. In a normal pulse sequence, the encoding number is 1
Although the number is 28 to 256, 512 measurements are performed in high-definition imaging in which the resolution of an image is increased. Conventionally, as shown by the broken line, the control of the encoding gradient magnetic field is performed by the applied level. Therefore, when the encoding number is increased (improvement of the resolution), the required gradient magnetic field level increases, and the load on the gradient magnetic field power supply is increased. The limit of refinement was determined by the ease of the gradient power supply.

Therefore, in the pulse sequence according to the present invention, the amount of encoding is divided into a central part and a peripheral part as shown by a solid line, and the control of the gradient magnetic field is switched between control based on the applied level and control based on the applied time. The burden on the magnetic field power supply was reduced, eliminating the limitations of high-definition imaging, and enabling high-definition imaging with ease. In FIG. 1, for the sake of simplicity, the number of encodes is described as a total of 13 times within a range of ± 6, but the level control is changed to time control in the peripheral portion where the gradient magnetic field level increases (± 4 encodes or more). I try to switch. As a result, the maximum output current of the gradient magnetic field power supply can be reduced to about 1/2 of the conventional level, and the efficiency can be improved. Alternatively, it is possible to execute twice the number of encoding times as in the related art without changing the current capacity, thereby enabling twice as high definition imaging.

FIG. 4 is a timing chart when the present invention is applied to a pulse sequence of the spin echo method. FIG.
Is the encoding N that switches from level control to time control
o. +3 and +4 are drawn. Encoding No. +4
In this case, the application time t4 is set to be longer by ３ of the application time t of the encode No. 3, that is, by +1 minute so that the encode amount can be increased while the encode gradient magnetic field level remains the encode No. 3. ing. +5
The encoding of +6 and -4 to -6 is performed according to the same principle. In the encoding No. adopting the control of the application time of the encoding gradient magnetic field, the echo time Te is extended with the extension of the application time. For example, the central portion Te within ± 3 encoding, which is the level control, is constant for the same time as Te3. In the encoding No. + 4, the echo time Te is extended to Te4.

The amount of extension of Te is 100 when the current value for applying the maximum applied gradient magnetic field to the gradient coil in the encoding direction is 100.
[A], assuming that the application time is 10 [ms] and the level of the center ± 128 encoding is controlled, the amount of gradient magnetic field per encoding is 0.78, which is 1/128 of 100 [A] × 10 [ms]. X 10 [A x ms], and the increase time per encoding in the application time control is 0.078 [ms], which is 1/128 of 10 [ms]. In general, T
The amount of extension of e is twice as large (because the interval between the 90 ° pulse and 180 ° pulse which is Te / 2 is extended), and is 0.156 [ms] per encode. Accordingly, Te is 0.156 at the longest in double-resolution imaging (± 256 encoding).
This is extended by about 20 [ms] of [ms] × 128. Of course, when Te is long from the beginning and there is a margin for controlling the application time, Te can be kept constant. It is desirable to take measures such as changing the application level of the read-out gradient magnetic field, changing the application time, changing the application timing, etc., in response to the extension of the echo time Te.

In order to obtain a correct emphasized image, such a T
Although the change of e is not preferable, the influence is small in the peripheral portion and can be ignored. For this reason, if the image enhancement is not important, the gradient magnetic field control switching point should be set closer to the center to further improve the efficiency.If image enhancement is important, the switching point should be closer to the periphery by the maximum power supply capacity. The switching point should be selected according to the situation, such as setting and increasing the number of encodings to improve the limit of high definition imaging.

Although the present invention has been described above from the viewpoint of improving the resolution of a captured image, this content also applies to the reduction of the field of view (FOV) related to the capacity of the gradient magnetic field power supply, and the setting range of the field of view is changed. It is also possible to enlarge and improve the usability of the MRI apparatus.

While the above description has been made with reference to the spin echo method sequence, the present invention is, of course, also applicable to other pulse sequences. FIG. 9 shows a sequence applied to the gradient echo method, and has the same effect as that described in the spin echo sequence.

[0030]

As described above, according to the present invention, high-definition imaging can be performed with a small gradient magnetic field power supply capacity by switching the control parameter of the encoding gradient magnetic field between the application level and the application time during the entire encoding. Can be realized,
Further, there is an effect that the setting range of the imaging field of view can be expanded, and the utilization efficiency of the gradient magnetic field power supply can be further improved.

[Brief description of the drawings]

FIG. 1 is an explanatory diagram of a phase encoding gradient magnetic field application method.

FIG. 2 is an explanatory diagram of a phase rotation by applying a gradient magnetic field.

FIG. 3 is an explanatory diagram of high-definition imaging by increasing an applied gradient magnetic field.

FIG. 4 is an explanatory diagram of a spin echo method sequence according to the present invention.

FIG. 5 is an overall configuration diagram of an MRI apparatus.

FIG. 6 is an explanatory view of a sequence of a spin echo method.

FIG. 7 is an explanatory diagram of image contrast depending on a relaxation time.

FIG. 8 shows a change in a reconstructed image depending on a measurement data range.

FIG. 9 is an explanatory diagram of a gradient echo method sequence according to the present invention.

[Explanation of symbols]

 11 CPU 12 Sequencer 18 Modulator 22 Gradient magnetic field power supply

Claims (2)

(57) [Claims] A magnetic circuit for applying a static magnetic field to a subject; a gradient coil for applying a slice gradient magnetic field, a readout gradient magnetic field, and an encoding gradient magnetic field to the subject; An irradiation coil for applying an irradiation pulse for causing magnetic resonance to the nucleus, a reception coil for detecting a magnetic resonance signal, and image reconstruction means for obtaining an image representing a physical property of the target object using the detection signal, In a magnetic resonance imaging apparatus including a control unit that controls each gradient magnetic field of a gradient coil and an application timing of the irradiation coil, the control unit further changes an application time of an encoding gradient magnetic field during a series of pulse sequences. control performs, encoding
To maintain the applied time of the
Control to be changed and the applied level of the encoding gradient magnetic field.
Control to change the application time while maintaining the constant value
Magnetic resonance imaging apparatus according to claim and this <br/> controlled by switching the code gradient magnetic field application timing. 2. A magnetic circuit for applying a static magnetic field to a subject,
Slice gradient magnetic field, readout gradient magnetic field and
A gradient magnetic field coil for applying an encoding gradient magnetic field,
Magnetic resonance occurs in the nuclei of the atoms that make up the tissue of the subject
An irradiation coil for applying an irradiation pulse
And a target coil using the detection signal
Image reconstruction means for obtaining images representing physical properties of objects
And each of the gradient magnetic fields of the gradient coil and the irradiation coil.
Magnetic sharing device with control means for controlling the
In the sound imaging apparatus, the control unit further includes:
Control to change the application time of encoding gradient magnetic field
At the same time, the application level of the encoding gradient magnetic field was changed.
Measurement data with image contrast information and encoding
Holds the contour information of the image in which the application time of the gradient magnetic field is changed
After obtaining the measurement data, the image reconstruction means
Characterized by reconstructing an image using measurement data
Air resonance imaging device.