JP3450508B2 - Magnetic resonance imaging equipment - Google Patents

Description

BACKGROUND OF THE INVENTION DETAILED DESCRIPTION OF THE INVENTION [0001] The present invention is magnetic resonance imaging instrumentation
In particular, the present invention relates to signal correction of magnetic resonance imaging . 2. Description of the Related Art A magnetic resonance imaging (MRI) apparatus causes a water molecule in an arbitrary plane crossing a subject to cause nuclear magnetic resonance, and generates a nuclear magnetic resonance signal from the nuclear magnetic resonance signal generated by the nuclear magnetic resonance. 1 is a medical image diagnostic apparatus that obtains a tomographic image at a time. In general, a slice gradient magnetic field for specifying a plane on which a tomographic image of an object is to be obtained is applied, and at the same time, an excitation pulse for exciting spins in the plane is given, so that the excited spins converge. To obtain a magnetic resonance signal (echo signal) generated in the step of performing a phase encoding gradient magnetic field and a readout gradient magnetic field until the echo signal is obtained in order to give positional information to the echo signal. Has become. [0004] Such pulses and each gradient magnetic field are applied based on a preset pulse sequence. Various pulse sequences are known according to the purpose.
A method in which echo signals are continuously measured in accordance with a sequential change of a phase encoding gradient magnetic field in response to the application of a single excitation pulse, whereby a so-called echo planer is obtained in which all echo signals necessary to obtain one tomographic image are obtained. Those that adopt the law are known. In this echo planar method, the sequence is repeatedly operated, for example, as in a spin echo method.
A feature that enables extremely high-speed imaging compared to the method of sequentially obtaining echo signals required to obtain one tomographic image by sequentially changing the phase encoding gradient magnetic field at each repetition. It is. The details of the echo planar method are described in, for example, the document "J. Phys. C: Solid State Phys., Vol. 10,
pp. L55-L58, 1977 ". However, in such an echo planer method, a series of measured echo signals is such that each of these signal intensities has a time constant inherent to the living tissue of the subject and is exponentially sequential. (The relaxation phenomenon)
As a result, it has been pointed out that a so-called blur or false image is generated in the reconstructed image. That is, a series of echo signals are stored in a so-called k-space memory in which the horizontal axis represents the frequency direction and the vertical axis represents the phase direction.
The intensity of each echo signal is not equal, in other words, a signal equivalent to the intensity modulated on the k-space is stored. For this reason, a correct tomographic image cannot be obtained even if image reconstruction is performed by performing an inverse Fourier transform on the signal on the k-space (so-called raw data) as it is, and the above-described blur Alternatively, there is a problem that a false image is generated. The present invention has been made in view of such circumstances, and an object of the present invention is to provide a magnetic resonance imaging apparatus capable of obtaining an image free from blurring or false images. Means for Solving the Problems Of the inventions disclosed in the present application, the outline of a representative one will be briefly described.
It is as follows. That is, in the magnetic resonance imaging apparatus according to the present invention , the application of an excitation pulse for exciting an object, the phase
Includes encoding gradient magnetic field application and echo signal measurement.
Sequencer that controls the pulse sequence,
A computer for processing the echo signal,
The sequencer responds to the excitation pulse by
A plurality of first echoes without applying a phase encoding gradient magnetic field
A first pulse sequence for measuring a signal, and
In response to the application of an excitation pulse,
Measurement of multiple second echo signals by applying different Sakai signals
Control of the second pulse sequence,
The plurality of first echo signals from the plurality of first echo signals.
Calculate the coefficient indicating the intensity attenuation of each echo signal
And a coefficient indicating the respective intensity attenuation.
Performing a process of calculating the corresponding plurality of second echo signals.
It is characterized in that Nau. [0014] [action] Magnetic resonance imaging instrumentation thus constructed
Location, first sequencer for the same test object, a second sea is a sequence for reconstructing a tomographic image
The cans Ru to execute the first sequence is a separate sequence. The first sequence is exactly the same as the second sequence for reconstructing a tomographic image except that no phase encoding gradient magnetic field is applied. Although a series of echo signals obtained by such a first sequence cannot reconstruct a tomographic image, they sequentially attenuate exponentially with a time constant specific to the living tissue of the subject. This is measured as reference data having a phenomenon (relaxation phenomenon). Then, the attenuation of the reference data in each echo signal is determined by the second sequence, which is a sequence for reconstructing a tomographic image.
The same applies to each echo signal obtained by executing the Therefore, the computer first calculates correction coefficients for correcting the intensity attenuation from a series of echo signals of the reference data, and uses the calculated correction coefficients to reconstruct a tomographic image. The image data without attenuation can be obtained by multiplying the series of corresponding echo signals obtained in (1). Next, a computer performs Fourier transform on the image data, so that an image free of so-called blur and false image can be obtained. FIG. 2 is a schematic diagram showing an embodiment of a magnetic resonance imaging apparatus according to the present invention. In FIG. 1, reference numeral 101 denotes a magnet for generating a static magnetic field; 102, a coil for generating a gradient magnetic field;
Reference numeral 3 denotes an inspection target such as a subject, and the inspection target 103 is set in the magnet 101 and the coil 102. The sequencer 104 includes a gradient power supply 1
05 and a high-frequency magnetic field generator 106, and various gradient magnetic fields and high-frequency magnetic fields are generated. The high-frequency magnetic field is applied to the inspection target 103 through the probe 107. That is, the pulse sequence shown in FIGS. 3A and 3B described later is stored in the sequencer 104, and various gradient magnetic fields and high-frequency magnetic fields are generated based on these pulse sequences. It is supposed to. The signal generated from the test object 103 is received by the probe 107 and is detected by the receiver 108. Here, a magnetic resonance frequency (hereinafter, referred to as a detection reference frequency) as a reference for detection is set by the sequencer 104. The detected signal is sent to a computer 109,
Here, signal processing is performed. And
The result is displayed on the display 110. The storage medium 111 can also store signals and measurement conditions as needed. 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 is supplied with current by a shim power supply 113. At the time of 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 to correct the non-uniformity of the static magnetic field. Here, in this embodiment, in particular, the computer 1
09 incorporates at least a circuit as shown in FIG. In the figure, first, echo signals obtained via the receiver 108 are sequentially stored in the image data k-space memory (1) 109A. The echo signal stored in the image data K space memory (1) 109A is obtained by an operation based on the pulse sequence shown in FIG. Similarly, echo signals obtained via the receiver 108 are sequentially stored in the reference data k-space memory 109B. The echo signal stored in the reference data K-space memory 109B is obtained by an operation based on the pulse sequence shown in FIG. Then, the reference data k-space memory 109
The information stored in B is sequentially read and input to the correction coefficient calculation circuit 109C.
Thus, the correction coefficient is calculated from each information. The information corresponding to the correction coefficient is input to the multiplication circuit 109D. On the other hand, the image data k-space memory (1)
The information stored in 109A is sequentially input, and each information is multiplied by a corresponding correction coefficient. The image data k thus corrected
Each information from the spatial memory (1) 109A is rewritten to the image data k-space memory (2) 109E.
After being Fourier-transformed by the Fourier transform circuit 109F, it is output to the display 110. Next, an embodiment of the operation of the magnetic resonance imaging apparatus thus configured will be described. First, the sequencer 104 shown in FIG.
Are stored, and various gradient magnetic fields and high-frequency magnetic fields are generated based on the pulse sequences. This pulse sequence is a so-called echo planar method, which is a method of continuously measuring an echo signal in accordance with a sequential change of a phase encoding gradient magnetic field in response to one application of an excitation pulse. Thus, an echo signal required to obtain one tomographic image can be obtained at high speed. That is, in FIG.
2 is applied via a probe 197. The excitation pulse 402 is a pulse for exciting a spin in the inspection target 103 to generate a signal. A slice gradient magnetic field 401 is applied simultaneously with the excitation pulse 402. This slice gradient magnetic field 401 is a magnetic field for specifying the position of the tomographic plane to be imaged of the inspection object. The phase-encoding gradient magnetic field 405 is a magnetic field that gives positional information to an echo signal to be described later together with the readout gradient magnetic field 407. It is designed to be applied once. The magnitude of the magnetic field corresponds to the storage position of the echo signal in the phase encoding direction of a k-space memory described later. In this case, it is assumed that the magnitude of the magnetic field gradually decreases with time, for example. The echo signal 401 is read out gradient magnetic field 40 every time each phase encoding gradient magnetic field 405 is applied.
7, and as a result, echo 1, echo 2,..., Echo N are continuously received. Here, the intensity of a series of echo signals 403, ie, echo 1, echo 2,..., Echo N, tends to attenuate exponentially with the next constant specific to the tissue of the inspection target 103. This has been a conventional problem. The series of echo signals 403
Is a k-space memory 10 for an image as shown in FIG.
9A. Spatial memory 109
In A, the horizontal axis is the frequency encoding direction kx, and the vertical axis is the phase encoding direction ky, and signals of echo 1, echo 2,..., Echo N are stored sequentially from above the horizontal axis. ing. In this case, the signal strength in the phase encoding direction based on the above-described attenuation is shown in the graph on the right. On the other hand, for the same inspection object 103,
By operating based on the pulse sequence shown in FIG. 3B, each echo signal, that is, echo 1, echo 2,..., Echo N is obtained. The pulse sequence shown in FIG. 3B is exactly the same as the pulse sequence shown in FIG. 3A except that no phase encoding gradient magnetic field 405 is applied. By operating such a pulse sequence, each of the echo signals, that is, echo 1, echo 2,..., Echo N, becomes echo 1, echo 2,. It will be obtained with exactly the same attenuation. The series of echo signals 403
Is a reference k-space memory 10 as shown in FIG.
9B. The signal strength in the phase encoding direction in this case is shown in the graph on the right, which is similar to the characteristic in the case of FIG. Each piece of information stored in the reference data k-space memory 109B is input to the correction coefficient calculation circuit 109C shown in FIG. 1, and correction coefficients are sequentially calculated. That is, as shown in FIG. 5, the peak Pi of each echo signal serving as reference data is detected (step 1).
1), the strength | Pi | is obtained (step 12). The peak of the echo signal is, for example, a position where the signal intensity is the highest, that is, a position where the absolute value of the echo signal has the maximum value. The signal strength correction coefficient Ci of each echo signal is the reciprocal of the echo peak strength | Pi |, that is, Ci = A /
| Pi | (step 13).
Here, A is an arbitrary constant. Finally, the intensity of each echo signal Si of the image data is corrected using the signal intensity correction coefficient Ci (step 14). As for the corrected signal Si ′, a value Sij ′ at the sampling point j (j = 1, 2,..., M; m: the number of sampling points) is obtained by Sij ′ = Sij × Ci. That is, as shown in FIG.
09D, the image data k-space memory (1) 10
Each information read from 9A is multiplied by the correction coefficient. According to the above procedure, the intensity of the image data is corrected, and a good image can be obtained by a normal reconstruction method. Here, this procedure is a simple method of processing because only the echo peak intensity is used for determining the signal intensity correction coefficient. However, depending on the shape of the subject, two or more echo peaks appear, and when there is non-uniformity in the static magnetic field, the echo peaks become duller in the later echo signal, and the signal intensity becomes accurate. In some cases, the correction coefficient cannot be calculated. In this case, the signal strength correction coefficient can be accurately obtained by the following procedure. As shown in FIG. 6, first, reference data is measured in the same manner as in the above-described example (step 21), and then the integrated value Ii of the intensity of each echo signal Ri is calculated according to the following equation (22). [Mathematical formula-see original document] Here, m is the number of sampling points, | Rij
| Is the signal intensity at the j-th sampling point of the echo signal Ri. Then, the signal strength correction coefficient is represented by Ci = A
/ Ii (step 23), and the corrected signal Si ′ of the image data is obtained by calculating the value Sij ′ at the sampling point j by Sij ′ = Sij × Ci (step 24). It is needless to say that a method using power instead of the integrated value of the echo signal strength may be used in a similar manner. FIG. 7 shows the procedure. First, after measuring the reference data (step 31), the power Qi of each echo signal Ri is calculated according to the following equation (32). ## EQU2 ## Here, m is the number of sampling points, | Rij
| Is the signal intensity at the j-th sampling point of the echo signal Ri. Then, the signal strength correction coefficient is represented by Ci = A
/ Qi (step 33), and the corrected signal Si ′ of the image data is obtained by calculating the value Sij ′ at the sampling point j by Sij ′ = Sij × Ci (step 34). In any of the methods, since the entire echo signal is used for calculating the signal intensity correction coefficient, the image data can be accurately corrected regardless of the shape of the subject. According to the embodiment described above, first, a sequence different from a sequence for reconstructing a tomographic image is executed for the same inspection object. This sequence is exactly the same as the sequence for reconstructing a tomographic image except that no phase encoding gradient magnetic field is applied. A series of echo signals obtained by such a sequence cannot reconstruct a tomographic image, but attenuate exponentially with a time constant inherent to the living tissue of the subject (exponential function). This is measured as reference data having a relaxation phenomenon). The attenuation of the reference data in each echo signal is the same in each echo signal obtained by executing a sequence for reconstructing a tomographic image. For this reason, a correction coefficient for correcting the intensity attenuation is calculated from a series of echo signals of the reference data, and the calculated correction coefficients are obtained for reconstructing a tomographic image. By multiplying the series of corresponding echo signals, image data without attenuation can be obtained. Therefore, by performing the Fourier transform on the image data, it is possible to obtain an image free from so-called blur and false image. It is needless to say that the measurement of the reference data in this case is the same regardless of whether it is performed before or after the measurement of the image data. In the above embodiment, the echo planar imaging method is used as the imaging method. However, the signal intensity can be corrected by the same method in the k-space division type echo planar method in which the entire k space is scanned by a plurality of excitations. it can. In this case, the reference data measurement sequence needs to be executed only once, and a correction coefficient is obtained for the measured echo number of each echo.
The obtained correction coefficient can be commonly used for correcting each echo group measured by executing the sequence a plurality of times for acquiring image data. In correcting each echo, a correction coefficient corresponding to the echo number is used. The present invention also relates to a photographing method for measuring a plurality of echo signals by one excitation to obtain image data.
Needless to say, the present invention can be generally applied to an imaging method in which the intensity of the echo signal is not uniform in the k space. For example, the present invention can be applied to a gradient spin echo method in which a 180-degree pulse 504 as shown in FIG. 8 is combined with a readout gradient magnetic field pulse 506 having positive and negative alternating signs. In the gradient spin echo method, as in the case of the spin echo method, an inspection target is placed in a static magnetic field, and a high frequency magnetic field (RF) pulse 502 for magnetization excitation is applied together with the application of the slice gradient magnetic field 501 to generate a magnetic field in the target object. Induce a magnetic resonance phenomenon in the slice. Then, a spin echo signal 510 is generated while applying a 180-degree pulse 504 and applying a readout gradient magnetic field pulse 506 for adding position information in the readout direction. Further, readout gradient magnetic field pulses 508 and 509 with inverted polarity are applied before and after this readout gradient magnetic field pulse 506 to generate gradient echo signals 511 and 507, and a total of three 180 ° pulses are used. It measures the echo signal. In this case, three or more echo signals can be measured by alternately inverting the polarity of the readout gradient magnetic field pulse applied before and after the spin echo signal an arbitrary number of times. Before measuring each echo signal, a phase encoding gradient magnetic field pulse 505 is applied in order to add position information in the phase encoding direction to the phase of the magnetization. In addition,
Reference numeral 503 denotes a dephasing gradient magnetic field pulse for temporarily disturbing the phase of the magnetization in the readout direction. After the above-described application of the 180-degree pulse and the measurement of the echo signal are normally repeated three times or more, after a waiting time of several seconds, the recovery of the magnetization is waited. This sequence is repeated several times while changing the value of the phase encoding gradient magnetic field pulse in order to measure an echo signal that only fills the entire measurement space. Then, the measured signal is subjected to inverse Fourier transform to reconstruct an image. Also in this gradient spin echo method, a series of echo signals are measured without applying the phase encoding gradient magnetic field pulse 505 and used as reference data, and the signal intensity of the image data is determined by the same procedure as described above. Correction can be performed to reconstruct an image free from blur and false images. Also in this case, similarly to the k-space division echo planar method, the echo data obtained by executing the image data sequence a plurality of times is a correction coefficient for each echo number obtained by executing the reference data measurement sequence only once. Groups can be corrected. In addition, a method called a spiral scan method for drawing a locus as shown in FIGS. 9A and 9B in k-space (IEEE Trans. On Med. Imaging MI-5, pp. 2-7, 1
986). In the spiral scan method, one of the gradient magnetic fields applied as Gx or Gy is set to zero so that the echo signal reciprocates in one direction through the origin on the k space as shown in FIG. The measured data is used as reference data. (FIG. 10 shows the case where Gy is set to zero.) Here, the echo signal of the reference data has a wider reciprocating width on the kx axis as the later echo signal. The signal strength correction using the reference data may be performed in exactly the same manner in the case of the above-described method using echo peaks. On the other hand, in the case of the method using the integral value or the power, the signal intensity correction coefficient C
Find i. The number of sampling points of the i-th echo signal of the reference data and the reciprocating width on the kx axis are mi and W, respectively.
i, and the signal at the j-th sampling point is Rij (j =
1, 2, ..., mi). The integral value Ii and the signal strength correction coefficient Ci are obtained by the following equation. [Mathematical formula-see original document] On the other hand, the power Qi and the signal strength correction coefficient Ci
Is calculated by the following equation. (Equation 4) For the correction of the signal intensity of the image data, the same method as described above using the calculated Ci is used. According to the above procedure, signal intensity correction can be performed even by the spiral scan method, and a good image can be obtained. This method also scans the entire k-space with multiple excitations.
It goes without saying that the present invention can be similarly applied to the space division type spiral scan method. As described above, according to the signal correction method for magnetic resonance imaging and the magnetic resonance imaging apparatus for implementing the method according to the present invention, the image reconstruction is performed after the attenuation of the echo signal intensity is corrected. Therefore, it is possible to obtain an image free from so-called blur or false image.

BRIEF DESCRIPTION OF THE DRAWINGS FIG. 1 is a configuration diagram of a main part showing an embodiment of a magnetic resonance imaging apparatus according to the present invention. FIG. 2 is an overall configuration diagram showing an embodiment of a magnetic resonance imaging apparatus according to the present invention. FIG. 3 is an explanatory diagram showing one embodiment of a pulse sequence used in the magnetic resonance imaging apparatus according to the present invention. FIG. 4 is an explanatory diagram of a k-space memory provided in the magnetic resonance imaging apparatus according to the present invention. FIG. 5 is a flowchart showing one embodiment of a signal correction method for magnetic resonance imaging according to the present invention. FIG. 6 is a flowchart showing another embodiment of a signal correction method for magnetic resonance imaging according to the present invention. FIG. 7 is a flowchart showing another embodiment of a signal correction method for magnetic resonance imaging according to the present invention. FIG. 8 is an explanatory view showing another embodiment of the pulse sequence used in the magnetic resonance imaging apparatus according to the present invention. FIG. 9 is an explanatory view showing another embodiment of the signal correction method for magnetic resonance imaging according to the present invention. FIG. 10 is an explanatory view showing another embodiment of the signal correction method for magnetic resonance imaging according to the present invention. [Description of Signs] 101: magnet for generating a static magnetic field, 102: gradient coil, 103: test object, 104: sequencer, 1
05: gradient magnetic field power supply, 106: high frequency magnetic field generator, 10
7: Probe, 108: Receiver, 109: Computer, 11
0: display, 111: storage medium, 112: shim coil, 113: shim power supply, 201: slice gradient magnetic field pulse, 202: high frequency magnetic field pulse for magnetization excitation, 203 ...
Echo signal, 204 ... phase encoding gradient magnetic field pulse,
205 ... 180 degree pulse, 206 ... Readout gradient magnetic field pulse, 401 ... Slice gradient magnetic field pulse, 402 ...
High frequency magnetic field pulse for magnetization excitation, 403 ... echo signal, 4
04: Dephase gradient magnetic field pulse, 405: Phase encode gradient magnetic field pulse, 406: Dephase gradient magnetic field pulse, 407: Readout gradient magnetic field pulse, 5
01: slice gradient magnetic field pulse, 502: high frequency magnetic field pulse for magnetization excitation, 503: gradient magnetic field pulse for dephase, 504: 180 degree pulse, 505: phase encoding gradient magnetic field pulse, 506: readout gradient magnetic field pulse, 507: echo signal.

──────────────────────────────────────────────────続 き Continuation of the front page (72) Inventor Hiroshi Nishimura 1-1-1 Uchikanda, Chiyoda-ku, Tokyo Inside Hitachi Medical Corporation (56) References JP-A-5-168607 (JP, A) JP-A Heisei 2-149251 (JP, A) JP-A-6-105825 (JP, A) JP-A-7-502907 (JP, A) (58) Fields investigated (Int. Cl. ⁷ , DB name) A61B 5/055

Claims (1)

(1) A sequencer that controls a pulse sequence including application of an excitation pulse for exciting an object, application of a phase encoding gradient magnetic field, and measurement of an echo signal. A computer for processing an echo signal, wherein the sequencer does not apply the phase encoding gradient magnetic field in response to the application of the excitation pulse, and outputs an echo number i
= 1, 2,..., N, a first pulse sequence for measuring a plurality of first echo signals, and applying the excitation pulse by changing the phase encoding gradient magnetic field and applying the same . Controlling a second pulse sequence for measuring a plurality of second echo signals of numbers i = 1, 2,..., N , the computer calculates the first echo of echo number i
The first echo signal of echo number i using the
The calculation of the coefficient Ci indicating the intensity attenuation of
1, 2,..., N a plurality of the first echo signals
To reduce the intensity of the first echo signal of echo number i.
The coefficient Ci indicating the decay is represented by the first eco number of the echo number i.
The second echo signal of the echo number i corresponding to the signal
, And N for echo numbers i = 1, 2,.
There is no row, the coefficient is the reciprocal of the integral value of the intensity of the first echo signal, or magnetic resonance, which is a reciprocal of the square root of the square of the integral value of the intensity of the first echo signal Imaging device.