JP4502488B2 - Magnetic resonance imaging device - Google Patents

Description

[0001]
BACKGROUND OF THE INVENTION
The present invention relates to a magnetic resonance signal receiving method and apparatus and a magnetic resonance imaging apparatus, and in particular, a magnetic resonance signal receiving method and apparatus using a plurality of RF coils, and a magnetic resonance imaging apparatus using such a magnetic resonance signal receiving apparatus. About.
[0002]
[Prior art]
In a magnetic resonance imaging (MRI) apparatus, an object to be imaged (patient etc.) is placed in an internal space of a magnet system, that is, a space where a static magnetic field is formed, and a gradient magnetic field and a high-frequency magnetic field are applied. Then, a magnetic resonance signal is generated in the object, and a tomographic image is generated (reconstructed) based on the received signal.
[0003]
An RF coil (radio frequency coil) is used to receive the magnetic resonance signal. One of the RF coils is a phased array coil. The phased array coil is constituted by a combination of a plurality of coil loops that do not interfere with each other.
[0004]
The tomographic image is reconstructed using a signal obtained by fully adding the reception signals of a plurality of coil loops. Alternatively, each image is reconstructed based on the reception signals of the plurality of coil loops, and the tomograms are obtained by fully adding the images.
[0005]
[Problems to be solved by the invention]
The distribution of brightness and contrast in the reconstructed image corresponds to the sensitivity distribution of the phased array coil, but the sensitivity distribution of the phased array coil is fixed depending on the shape and positional relationship of multiple coil loops. The brightness and contrast distribution of the reconstructed image are also fixed and cannot be arbitrarily adjusted.
[0006]
Accordingly, an object of the present invention is to realize a magnetic resonance signal receiving method and apparatus capable of adjusting the sensitivity distribution without changing the physical structure of the RF coil, and a magnetic resonance imaging apparatus using such a magnetic resonance signal receiving apparatus. It is to be.
[0007]
[Means for Solving the Problems]
(1) According to another aspect of the invention for solving the above-described problem, a plurality of RF coils that do not interfere with each other receive magnetic resonance signals from a common target, and the plurality of magnetic resonance signals received respectively. The magnetic resonance signal receiving method is characterized in that the phases are adjusted and added to each other.
[0008]
In the invention from this viewpoint, since the plurality of magnetic resonance signals respectively received by the plurality of RF coils are fully added by adjusting the phase between them, the synthesized signal is a vector sum of the individual received signals. Therefore, by adjusting the phase between the plurality of received signals, the magnitude of the vector sum can be adjusted, and the sensitivity distribution can be adjusted.
[0009]
(2) In another aspect of the invention for solving the above-described problem, a plurality of RF coils that do not interfere with each other receive magnetic resonance signals from a common target, and the plurality of magnetic resonance signals received respectively. A method for receiving a magnetic resonance signal, characterized in that the phase and individual amplitude of each other are adjusted and added.
[0010]
In the invention in this aspect, since the plurality of magnetic resonance signals respectively received by the plurality of RF coils are fully added by adjusting the phase and the individual amplitude, the synthesized signal becomes a vector sum of the individual received signals. . Therefore, the magnitude of the vector sum can be adjusted and the sensitivity distribution can be adjusted by adjusting the phase and individual amplitude between the plurality of received signals.
[0011]
(3) According to another aspect of the invention for solving the above-mentioned problem, the plurality of coils form a phased array coil. The magnetic resonance signal reception according to (1) or (2), Is the method.
[0012]
In the invention from this point of view, the plurality of magnetic resonance signals of the phased array coil are fully added by adjusting the phase and the like between each other, so that the combined signal is a vector sum of the individual received signals. Therefore, the magnitude of the vector sum can be adjusted by adjusting the phase or the like between the plurality of received signals, and the sensitivity distribution of the phased array coil can be adjusted.
[0013]
(4) In another aspect of the invention for solving the above-described problem, a plurality of RF coils that do not interfere with each other receive a magnetic resonance signal from a common target, respectively, and receiving means A magnetic resonance signal receiving apparatus comprising: a signal adjusting unit that adjusts a phase between magnetic resonance signals; and an adding unit that adds the plurality of adjusted magnetic resonance signals.
[0014]
In the invention from this viewpoint, since the plurality of magnetic resonance signals respectively received by the plurality of RF coils are fully added by adjusting the phase between them, the synthesized signal is a vector sum of the individual received signals. Therefore, by adjusting the phase between the plurality of received signals, the magnitude of the vector sum can be adjusted, and the sensitivity distribution can be adjusted.
[0015]
(5) According to another aspect of the invention for solving the above-described problem, a plurality of RF coils that do not interfere with each other receive a magnetic resonance signal from a common target, respectively, and receiving means A magnetic resonance signal receiving apparatus comprising: a signal adjusting unit that adjusts a phase and an individual amplitude between magnetic resonance signals; and an adding unit that adds the plurality of adjusted magnetic resonance signals. .
[0016]
In the invention in this aspect, since the plurality of magnetic resonance signals respectively received by the plurality of RF coils are fully added by adjusting the phase and the individual amplitude, the synthesized signal becomes a vector sum of the individual received signals. . Therefore, the magnitude of the vector sum can be adjusted and the sensitivity distribution can be adjusted by adjusting the phase and individual amplitude between the plurality of received signals.
[0017]
(6) In another aspect of the invention for solving the above-described problem, the plurality of coils form a phased array coil. The magnetic resonance signal reception according to (4) or (5), Device.
[0018]
In the invention from this point of view, the plurality of magnetic resonance signals of the phased array coil are fully added by adjusting the phase and the like between each other, so that the combined signal is a vector sum of the individual received signals. Therefore, the magnitude of the vector sum can be adjusted by adjusting the phase or the like between the plurality of received signals, and the sensitivity distribution of the phased array coil can be adjusted.
[0019]
(7) Another aspect of the invention for solving the above-described problems is that magnetic resonance imaging is used to construct an image based on a magnetic resonance signal generated inside a target using a static magnetic field, a gradient magnetic field, and a high-frequency magnetic field. An apparatus for receiving the magnetic resonance signal includes: a receiving means for receiving a magnetic resonance signal from a common target by a plurality of RF coils that do not interfere with each other; and a plurality of the received magnetic resonance signals. A magnetic resonance imaging apparatus comprising: a signal adjusting unit that adjusts a phase between each other; and an adding unit that adds the plurality of adjusted magnetic resonance signals.
[0020]
In the invention from this viewpoint, since the plurality of magnetic resonance signals respectively received by the plurality of RF coils are fully added by adjusting the phase between them, the synthesized signal is a vector sum of the individual received signals. Therefore, by adjusting the phase between the plurality of received signals, the magnitude of the vector sum can be adjusted, and the sensitivity distribution can be adjusted. Thereby, the brightness and contrast distribution of the image can be adjusted.
[0021]
(8) In another aspect of the invention for solving the above-described problem, magnetic resonance imaging is used to construct an image based on a magnetic resonance signal generated inside a target using a static magnetic field, a gradient magnetic field, and a high-frequency magnetic field. An apparatus for receiving the magnetic resonance signal includes: a receiving means for receiving a magnetic resonance signal from a common target by a plurality of RF coils that do not interfere with each other; and a plurality of the received magnetic resonance signals. A magnetic resonance imaging apparatus comprising: a signal adjusting unit that adjusts the phase and amplitude between each other; and an adding unit that adds the plurality of adjusted magnetic resonance signals.
[0022]
In the invention in this aspect, since the plurality of magnetic resonance signals respectively received by the plurality of RF coils are fully added by adjusting the phase and the individual amplitude, the synthesized signal becomes a vector sum of the individual received signals. . Therefore, the magnitude of the vector sum can be adjusted and the sensitivity distribution can be adjusted by adjusting the phase and individual amplitude between the plurality of received signals. Thereby, the brightness and contrast distribution of the image can be adjusted.
[0023]
(9) The magnetic resonance imaging apparatus according to (7) or (8), wherein the plurality of coils form a phased array coil. It is.
[0024]
In the invention from this point of view, the plurality of magnetic resonance signals of the phased array coil are fully added by adjusting the phase and the like between each other, so that the combined signal is a vector sum of the individual received signals. Therefore, the magnitude of the vector sum can be adjusted by adjusting the phase or the like between the plurality of received signals, and the sensitivity distribution of the phased array coil can be adjusted. Thereby, the brightness and contrast distribution of the image can be adjusted.
[0025]
DETAILED DESCRIPTION OF THE INVENTION
Hereinafter, embodiments of the present invention will be described in detail with reference to the drawings. The present invention is not limited to the embodiment. FIG. 1 shows a block diagram of the magnetic resonance imaging apparatus. This apparatus is an example of an embodiment of the present invention. An example of an embodiment relating to the apparatus of the present invention is shown by the configuration of the apparatus. An example of an embodiment related to the method of the present invention is shown by the operation of the apparatus.
[0026]
As shown in FIG. 1, the apparatus has a magnet system 100. The magnet system 100 includes a main magnetic field coil unit 102, a gradient coil unit 106 and an RF coil unit 108. Each of these coil portions has a substantially cylindrical outer shape and is arranged coaxially with each other. The object 300 is mounted on a cradle 500 in the internal space of the magnet system 100 and is carried in and out by a conveying means (not shown).
[0027]
The cradle 500 is provided with a receiving coil unit 110. The reception coil unit 110 has a pair of members facing each other. The object 300 is mounted on the cradle 500 in such a posture that the body portion is sandwiched from the front and rear between a pair of members facing each other. The receiving coil unit 110 will be described later.
[0028]
The main magnetic field coil unit 102 forms a static magnetic field in the internal space of the magnet system 100. The direction of the static magnetic field is generally parallel to the body axis direction of the object 300. That is, a so-called horizontal magnetic field is formed. The main magnetic field coil unit 102 is configured using, for example, a superconducting coil. Needless to say, the present invention may be configured using not only a superconducting coil but also a normal conducting coil.
[0029]
The gradient coil unit 106 generates a gradient magnetic field for giving a gradient to the static magnetic field strength. There are three types of gradient magnetic fields that are generated: a slice gradient magnetic field, a read out gradient magnetic field, and a phase encode gradient magnetic field. The gradient coil unit 106 corresponds to these three types of gradient magnetic fields. Has three gradient coils (not shown).
[0030]
The RF coil unit 108 generates a high-frequency magnetic field for exciting spins in the body of the target 300. Hereinafter, the formation of a high-frequency magnetic field is also referred to as transmission of an RF excitation signal. The receiving coil unit 110 receives an electromagnetic wave generated by excited spin, that is, a magnetic resonance signal.
[0031]
A gradient driving unit 130 is connected to the gradient coil unit 106. The gradient driving unit 130 gives a driving signal to the gradient coil unit 106 to generate a gradient magnetic field. The gradient driving unit 130 includes three systems of driving circuits (not shown) corresponding to the three systems of gradient coils in the gradient coil unit 106.
[0032]
An RF drive unit 140 is connected to the RF coil unit 108. The RF drive unit 140 provides a drive signal to the RF coil unit 108 and transmits an RF excitation signal to excite spins in the body of the target 300.
[0033]
A data collection unit 150 is connected to the reception coil unit 110. The data collection unit 150 takes in the reception signal received by the reception coil unit 110 and collects it as digital data. A portion including the receiving coil unit 110 and the data collecting unit 150 is an example of an embodiment of the magnetic resonance signal receiving apparatus of the present invention.
[0034]
A bias driving unit 120 is also connected to the receiving coil unit 110. The bias driving unit 120 supplies a bias signal to a diode (diode), which will be described later, included in the receiving coil unit 110, thereby switching the receiving coil between enable / disable (enable / disable).
[0035]
A control unit 160 is connected to the bias driving unit 120, the gradient driving unit 130, the RF driving unit 140, and the data collection unit 150. The controller 160 controls the bias driver 120 or the data collector 150, respectively.
[0036]
The output side of the data collection unit 150 is connected to the data processing unit 170. The data processing unit 170 stores the data fetched from the data collection unit 150 in a memory (not shown). A data space is formed in the memory. The data space constitutes a two-dimensional Fourier space. The data processing unit 170 reconstructs an image of the target 300 by performing two-dimensional inverse Fourier transform on the data in the two-dimensional Fourier space. Hereinafter, the two-dimensional Fourier space is also referred to as k space.
[0037]
The data processing unit 170 is connected to the control unit 160. The data processing unit 170 is above the control unit 160 and controls it. A display unit 180 and an operation unit 190 are connected to the data processing unit 170. The display unit 180 displays the reconstructed image and various information output from the data processing unit 170. The operation unit 190 is operated by an operator and inputs various commands and information to the data processing unit 170.
[0038]
Next, the receiving coil unit 110 will be described. In FIG. 2, the schematic diagram of the mutual relationship of the object 300 and the receiving coil part 110 is shown. As shown in the figure, the receiving coil unit 110 has a pair of members facing each other, that is, a pair of coil plates 210 and 220. The coil plates 210 and 220 are provided with coil loops 212 and 222, respectively. The coil loops 212 and 222 are an example of an embodiment of the RF coil in the present invention. The electrical configuration of the coil loops 212 and 222 will be described later.
[0039]
The coil plates 210 and 220 are opposedly supported by the support member 230 at a predetermined interval. The support member 230 can be expanded and contracted in the length direction, whereby the distance between the coil plates 210 and 220 can be adjusted.
[0040]
For example, as shown in FIG. 3, the extendable support member 230 includes a column member 232 fixed to the coil plate 210 side and a sheath member 234 fixed to the coil plate 220 side, and the column member 232 is a sheath. The member 234 is fitted and slidable. There is friction between the two, and the stretched or shortened state to an arbitrary length can be maintained.
[0041]
The support member 230 can be omitted. In this case, the coil plates 210 and 220 are configured to have flexibility so that the target 300 is wrapped from the front and the back. Note that the coil loops 212 and 222 may be arranged side by side on either the front side or the rear side of the target 300 instead of placing the target 300 between the front and rear.
[0042]
FIG. 4 shows an electric circuit of the coil loop 212. As shown in the figure, the coil loop 212 is constituted by a series connection of a capacitor 302 and a conductor 304. Capacitors and conductors are represented in one place. The number of capacitors 302 is not limited to four as shown, and may be appropriate.
[0043]
An input circuit of a preamplifier 306 that amplifies a magnetic resonance signal received by the coil loop 212 is connected to both ends of one capacitor 302 through an inductor 308. As the preamplifier 306, an amplifier whose impedance of the input circuit is sufficiently low, that is, a low input impedance amplifier is used.
[0044]
A diode 310 is connected in parallel to the input circuit of the preamplifier 306. A forward bias voltage or a reverse bias voltage is applied to the diode 310 from the bias driver 120 via an RF choke circuit 312. The diode 310 is sufficiently turned on by the forward bias voltage to configure a parallel circuit of the inductor 308 and the capacitor 302.
[0045]
Since the resonance frequency of the LC parallel circuit is selected so as to coincide with the frequency of the RF signal for spin excitation, the LC coil unit 108 turns on the diode 310 when transmitting the RF signal so that the LC A parallel resonance circuit is configured, and the coil loop is substantially opened by a high impedance thereby, and decoupling with the RF coil unit 108 is performed. A state in which the diode 310 is on is referred to as a disabled state of the receiving coil unit 110.
[0046]
When receiving a magnetic resonance signal, the diode 310 is turned off by a reverse bias voltage. A state in which the diode 310 is off is referred to as an enable state of the receiving coil unit 110. In this case, since the preamplifier 306 is a low input impedance amplifier, an LC parallel circuit composed of the capacitor 302 and the inductor 308 is substantially formed. For this reason, even when a magnetic resonance signal is received, the coil loop is substantially in an open loop state due to the high impedance due to the resonance of the LC parallel circuit.
[0047]
Even if the diode 310 and the input terminal of the preamplifier 306 are connected by a coaxial cable having an integral multiple of 1/2 (λ / 2) of the wavelength of the magnetic resonance signal, the above condition is satisfied. Therefore, this allows the preamplifier 306 to be provided at an appropriate position away from the coil loop.
[0048]
The coil loops 212 and 222 having such a configuration are formed on the coil plates 210 and 220, respectively, and are mounted on the cradle 500 with the loop surfaces facing each other. An electric circuit of a pair of coil loops in the opposed state is shown in FIG. FIG. 5 is drawn from an angle at which the loop surface of each coil is viewed obliquely.
[0049]
A portion including the coil loops 212 and 222, the inductors 308 and 308 ′, and the preamplifiers 306 and 306 ′ is an example of an embodiment of the receiving means in the present invention.
[0050]
Since the pair of coil loops are substantially open loops due to the high impedance of the LC circuit at the time of reception, there is substantially no coupling between them. For this reason, it becomes the same as that each coil loop exists independently, and can receive a magnetic resonance signal individually, without receiving the influence of the coil loop which opposes.
[0051]
Further, since there is no mutual influence, a plurality of pairs of other coil loops having the same configuration may be provided. A plurality of coil loops arranged side by side and not interfering with each other form a so-called phased array coil. As a result, a wider range of the object 300 can be imaged. Hereinafter, the case where there is one pair of coil loops will be described.
[0052]
FIG. 6 shows a block diagram of the magnetic resonance imaging apparatus. This apparatus is an example of an embodiment of the present invention. An example of an embodiment relating to the apparatus of the present invention is shown by the configuration of the apparatus. An example of an embodiment related to the method of the present invention is shown by the operation of the apparatus.
[0053]
The apparatus shown in FIG. 6 includes a magnet system 100 ′ and a receiving coil section 110 ′ that are different from the apparatus shown in FIG. Other than that, the configuration is the same as that of the apparatus shown in FIG. 1, and the same reference numerals are given to the same parts, and the description is omitted.
[0054]
The magnet system 100 ′ includes a main magnetic field magnet unit 102 ′, a gradient coil unit 106 ′, and an RF coil unit 108 ′. Each of the main magnetic field magnet section 102 ′ and each coil section is composed of a pair of facing each other across a space. Moreover, all have a disk-shaped external shape and are arrange | positioned sharing a central axis. The target 300 is mounted on the cradle 500 and carried into and out of the internal space of the magnet system 100 ′ by a conveying means (not shown).
[0055]
The cradle 500 is provided with a receiving coil section 110 ′. The reception coil unit 110 ′ has a pair of members facing each other. The object 300 is mounted on the cradle 500 in a posture in which the body is sandwiched between a pair of members facing each other from the left and right. The reception coil unit 110 ′ is an example of an embodiment of the RF coil of the present invention. The receiving coil unit 110 ′ will be described later.
[0056]
The main magnetic field magnet unit 102 ′ forms a static magnetic field in the internal space of the magnet system 100 ′. The direction of the static magnetic field is generally orthogonal to the body axis direction of the object 300. That is, a so-called vertical magnetic field is formed. The main magnetic field magnet section 102 ′ is configured using, for example, a permanent magnet. Of course, not only permanent magnets but also superconducting electromagnets or normal conducting electromagnets may be used.
[0057]
The gradient coil section 106 ′ generates a gradient magnetic field for giving a gradient to the static magnetic field strength. There are three types of gradient magnetic fields to be generated: a slice gradient magnetic field, a readout gradient magnetic field, and a phase encode gradient magnetic field, and the gradient coil unit 106 'has three gradient coils (not shown) corresponding to these three types of gradient magnetic fields. .
[0058]
The RF coil unit 108 ′ transmits an RF excitation signal for exciting spins in the body of the target 300. The reception coil unit 110 ′ receives a magnetic resonance signal in which excited spin occurs.
[0059]
Next, the receiving coil unit 110 ′ will be described. In FIG. 7, the schematic diagram of the mutual relationship of object 300 and receiving coil part 110 'is shown. As shown in the figure, the receiving coil section 110 ′ has a pair of members facing each other, that is, a pair of coil plates 210 ′ and 220 ′. Coil loops 212 'and 222' are provided on the coil plates 210 'and 220', respectively. The coil loops 212 ′ and 222 ′ are an example of an embodiment of the RF coil in the present invention. The electrical configuration of each coil loop is the same as that shown in FIG.
[0060]
The coil plates 210 ′ and 220 ′ are opposed and supported by a support member 230 ′ at a predetermined interval. The support member 230 ′ can be expanded and contracted in the length direction, and thereby the distance between the coil plates 210 ′ and 220 ′ can be adjusted. The extendable support member 230 'has a configuration similar to that shown in FIG.
[0061]
That is, the receiving coil unit 110 ′ has the same basic configuration as the receiving coil unit 110. However, the mutual relationship with the object 300 is 90 ° different from the receiving coil unit 110 around the body axis. The coil loops 212 ′ and 222 ′ may be arranged side by side on either the right side or the left side of the target 300 instead of placing the target 300 between the left and right sides.
[0062]
Since the plurality of coil loops are not affected by each other, a plurality of pairs of other coil loops having the same configuration may be provided. A plurality of coil loops arranged side by side and not interfering with each other form a so-called phased array coil. As a result, a wider range of the object 300 can be photographed. Hereinafter, the case where there is one pair of coil loops will be described.
[0063]
The data collection unit 150 in the apparatus shown in FIG. 1 or FIG. 6 includes an adder 470 that adds the output signals of the preamplifiers 306 and 306 ′, as shown in FIG. By adding the two reception signals by the adder 470, a reception signal equivalent to that received by the saddle type coil can be obtained.
[0064]
That is, if the sensitivity characteristics of the individual coil loops are given by, for example, the curves a and b shown in FIG. 9, a received signal equivalent to that received by the receiving coil having the sensitivity as the composite curve c is obtained. Can do. The composite curve c is equivalent to the sensitivity characteristic of the saddle type coil.
[0065]
The addition of the received signal may be performed at the stage of an analog signal or may be performed after conversion to a digital signal. The digital signal may be temporarily stored in a memory and performed on the read signal.
[0066]
In addition, as shown in FIG. 10, signal adjustment units (units) 472 and 472 ′ are provided at the subsequent stage of the preamplifiers 306 and 306 ′, and the phases are adjusted after adjusting the phases of the output signals. The signal adjustment units 472 and 472 ′ may be provided in only one of them.
[0067]
The signal adjustment units 472 and 472 ′ are an example of the embodiment of the signal adjustment means in the present invention. Adder 470 is an example of an embodiment of the adding means in the present invention. The signal adjustment units 472, 472 ′ and the adder 470 are constituted by electronic circuits, for example. Note that the present invention is not limited to this, and may be realized by a computer program. Implementation by a computer program is preferable in that a dedicated electronic circuit is not required.
[0068]
Thus, the vector of two signals is combined by adjusting and adding the phases of the individual signals. Therefore, the magnitude of the combined vector can be adjusted by adjusting the phase difference between both signals.
[0069]
By adjusting the magnitude of the combined vector, the apparent signal intensity distribution of the signal received by the receiving coil unit 110 can be adjusted. The apparent signal intensity distribution will be described below.
[0070]
FIG. 11 conceptually shows intensity distributions of signals received by the coil loops 212 and 222, respectively. In the figure, the vertical axis s is the signal intensity, and the horizontal axis d is the distance between the centers of the coil loops 212 and 222. However, both are normalized to 0-1.0. Note that the coil loop 212 is at a position of d = 0, and the coil loop 222 is at a position of d = 1.0.
[0071]
Graphs A and B show intensity distributions on the lines connecting the centers of the coil loops 212 and 222, respectively, of the signals received by the coil loops 212 and 222. The strength of the received signal decreases with distance from the coil loop.
[0072]
For ease of explanation, it is assumed that the signal strength decreases linearly with distance. That is, as shown in graph A, the strength of the received signal of the coil loop 212 linearly decreases from s = 1.0 to 0 over d = 0 to 1.0, and the strength of the received signal of the coil loop 222 is As shown in the graph B, it decreases linearly from s = 1.0 to 0 over d = 1.0 to 0.
[0073]
Thus, the composition ratio of the signals received by the coil loops 212 and 222 is (A, B) = (1.0, 0) when d = 0, and (A, B) = (0 when d = 0.2. .8, 0.2), d = 0.5, (A, B) = (0.5, 0.5), and d = 0.8, (A, B) = (0.2, 0.8). ) And d = 1.0, (A, B) = (0, 1.0).
[0074]
The vector synthesis of received signals will be described with reference to FIGS. In the figure, the received signal is represented by a vector in the polar coordinate plane. Vectors A and B represent the received signals of coil loops 212 and 222, respectively.
[0075]
FIG. 12 shows the received signal at d = 0. Since (A, B) = (1.0, 0), the combined vector is only the vector A. The magnitude of the combined vector is 1.0.
[0076]
FIG. 13 shows the received signal at d = 0.2. When (A, B) = (0.8, 0.2), the vector sum of the vectors A and B is obtained. The magnitude of the combined vector A + B changes according to the phase difference θ between the two vectors. When the phase difference is θ = 0 °, the magnitude of the combined vector A + B is 1.0.
[0077]
As the phase difference θ changes, the position of the tip of the combined vector A + B changes along the circumference of the circle indicated by the broken line. This circle is a circle with the tip of the vector A as the center and the size of the vector B as the radius.
[0078]
FIG. 14 shows the received signal at d = 0.5. When (A, B) = (0.5, 0.5), the vector sum of the vectors A and B is obtained. The magnitude of the combined vector A + B changes according to the phase difference θ between the two vectors. When the phase difference is θ = 0 °, the magnitude of the combined vector A + B is 1.0.
[0079]
As the phase difference θ changes, the position of the tip of the combined vector A + B changes along the circumference of the circle indicated by the broken line. This circle is a circle with the tip of the vector A as the center and the size of the vector B as the radius.
[0080]
FIG. 15 shows the received signal at d = 0.8. When (A, B) = (0.2, 0.8), the vector sum of the vectors A and B is obtained. The magnitude of the combined vector A + B changes according to the phase difference θ between the two vectors. When the phase difference is θ = 0 °, the magnitude of the combined vector A + B is 1.0.
[0081]
As the phase difference θ changes, the position of the tip of the combined vector A + B changes along the circumference of the circle indicated by the broken line. This circle is a circle with the tip of the vector A as the center and the size of the vector B as the radius.
[0082]
FIG. 16 shows the received signal at d = 1.0. Since (A, B) = (0, 1.0), the resultant vector is only vector B. The magnitude of the combined vector is 1.0. With the change of the phase difference θ, the position of the tip of the combined vector changes along the circumference of the circle indicated by the broken line. This circle is a circle with the coordinate origin 0 as the center and the size of the vector B as the radius.
[0083]
As shown in FIGS. 12 to 16, when the phase difference is θ = 0 °, the magnitudes of the combined vector A + B are all 1.0 in the range of d = 0 to 1.0. As a result, a uniform signal intensity distribution as shown by a straight line e in FIG. 11 is obtained.
[0084]
17 to 19 show the magnitudes of the combined vectors A + B at d = 0.2, 0.5, and 0.8 when θ = θ1 (0 <θ1 <90 °). As shown in the figure, the magnitude of the combined vector A + B is smaller than 1.0. The magnitude of the combined vector A + B is also minimum at d = 0.5. Note that the magnitude of the combined vector at d = 0, 1.0 is always 1.0 regardless of θ.
[0085]
Also in the case of θ = −θ1, the magnitude of the combined vector A + B is the same as that in the case of θ = θ1. Hereinafter, the case where the phase difference is positive will be described, but the same applies to the case where the phase difference is negative. As a result, an apparent signal intensity distribution as shown by the curve f in FIG. 11 is obtained.
[0086]
20 to 22 show the magnitudes of the combined vectors A + B at d = 0.2, 0.5, and 0.8 when θ = 90 °, respectively. As shown in the figure, the magnitude of the combined vector A + B is further reduced. The magnitude of the combined vector A + B is also minimum at d = 0.5. Note that the magnitude of the combined vector at d = 0, 1.0 is always 1.0 regardless of θ. As a result, an apparent signal intensity distribution as shown by a curve g in FIG. 11 is obtained.
[0087]
23 to 25 show the magnitudes of the combined vectors A + B at d = 0.2, 0.5, and 0.8 when θ = θ2 (90 ° <θ2 <180 °). As shown in the figure, the magnitude of the combined vector A + B is further reduced. The magnitude of the combined vector A + B is also minimum at d = 0.5. Note that the magnitude of the combined vector at d = 0, 1.0 is always 1.0 regardless of θ. As a result, an apparent signal intensity distribution as shown by the curve h in FIG. 11 is obtained.
[0088]
26 to 28 show the magnitudes of the combined vectors A + B at d = 0.2, 0.5, and 0.8 when θ = 180 °, respectively. As shown in the figure, the magnitude of the combined vector A + B is further reduced. The magnitude of the combined vector A + B becomes 0 when d = 0.5. Note that the magnitude of the combined vector at d = 0, 1.0 is always 1.0 regardless of θ. As a result, an apparent signal intensity distribution as shown by the straight line i in FIG. 11 is obtained.
[0089]
The apparent signal intensity distribution shown in FIG. 11 is a substantial sensitivity distribution of the receiving coil unit 110. Therefore, the sensitivity distribution of the reception coil unit 110 can be adjusted by adjusting the phase difference in which the reception signals of the coil loops 212 and 222 are added.
[0090]
In addition to the adjustment of the phase difference, the amplitudes of the reception signals of the coil loops 212 and 222 may be adjusted by the signal adjustment units 472 and 472 ′, respectively. Since the magnitude of the combined vector changes by adjusting the amplitude of each received signal, the substantial sensitivity distribution can be adjusted.
[0091]
The amplitude adjustment is preferably performed by a computer program in that a dedicated electronic circuit is not required. In particular, when high-speed operation is required, it is preferable to adjust the amplitude with a dedicated electronic circuit to reduce the burden on the computer.
[0092]
The same applies to the coil loops 212 ′ and 222 ′. Accordingly, the sensitivity distribution can be substantially adjusted in the same manner for the receiving coil section 110 ′.
[0093]
Such phase adjustment and amplitude adjustment may be performed on a signal once stored in the memory. For the signal addition, the digital signals subjected to phase adjustment or the like may be stored in a memory, and the read signals may be added.
[0094]
FIG. 29 shows an example of another configuration of the receiving coil unit 110. As shown in the figure, the receiving coil unit 110 has another pair of coil loops 214 and 224 that face each other with the target 300 interposed therebetween. Coil loop 214 is on the same side as coil loop 212 with respect to object 300, and coil loop 224 is on the same side as coil loop 222 with respect to object 300.
[0095]
FIG. 30 shows an electric circuit of these coil loops. As shown in the figure, the electric circuit of each coil loop has the same configuration. The reception signals of the coil loops 212 and 214 are output from the preamplifiers 306 and 326, respectively. The reception signals of the coil loops 222 and 224 are output from the preamplifiers 306 ′ and 326 ′, respectively.
[0096]
FIG. 31 shows an adder circuit for these received signals. As shown in the figure, the output signal of the preamplifier 306, that is, the reception signal of the coil loop 212, is added to the output signal of the preamplifier 326 ', that is, the reception signal of the coil loop 224, by adjusting the phase by the signal adjustment unit 472'. It is added by 470.
[0097]
Further, the output signal of the preamplifier 326, that is, the reception signal of the coil loop 214, is added by the adder 480 after adjusting the phase by the signal adjustment unit 482 ', the output signal of the preamplifier 306', that is, the reception signal of the coil loop 222. .
[0098]
As a result, the four coil loops 212, 222, 214, and 224 shown in FIG. 29 are combined two by two in a so-called brushing relationship. In such a combination, by appropriately setting the signal adjustment amount by the signal adjustment units 472 ′, 482 ′, as shown in FIG. 32, four coil loops 212 ′, 222 ′, 214 ′, A reception signal equivalent to the reception of the magnetic resonance signal by two sets of saddle type coils composed of 224 ′ can be obtained.
[0099]
The operation of this apparatus will be described. The operation is essentially the same between the apparatus shown in FIG. 1 and the apparatus shown in FIG. The operation of this apparatus proceeds under the control of the control unit 160.
[0100]
FIG. 33 shows an example of a pulse sequence used for magnetic resonance imaging. This pulse sequence is a pulse sequence of a spin echo (SE: Spin Echo) method.
[0101]
That is, (1) is a sequence of 90 ° pulses and 180 ° pulses for RF excitation in the SE method, and (2), (3), (4) and (5) are respectively the slice gradient Gs and the lead. This is a sequence of an out gradient Gr, a phase encode gradient Gp, and a spin echo MR. The 90 ° pulse and the 180 ° pulse are represented by center signals. The pulse sequence proceeds from left to right along the time axis t.
[0102]
As shown in the figure, 90 ° excitation of spin is performed by a 90 ° pulse. At this time, the slice gradient Gs is applied, and selective excitation for a predetermined slice is performed. After a predetermined time from the 90 ° excitation, 180 ° excitation by a 180 ° pulse, that is, spin inversion is performed. At this time, the slice gradient Gs is applied, and selective inversion is performed for the same slice.
[0103]
In the period between 90 ° excitation and spin reversal, a readout gradient Gr and a phase encode gradient Gp are applied. Spin dephase is performed by the lead-out gradient Gr. Spin phase encoding is performed by the phase encoding gradient Gp.
[0104]
After the spin inversion, the spin is rephased at the readout gradient Gr to generate the spin echo MR. The spin echo MR is an RF signal having a symmetrical waveform with respect to the echo center. The central echo occurs after TE (echo time) from 90 ° excitation. The spin echo MR is received by the two coil loops 212 and 222, and the received signals are output from the preamplifiers 306 and 306 ′, respectively. The data acquisition unit 150 adds the two output signals by appropriately adjusting the phase difference at the stage of an analog signal or a digital signal.
[0105]
Such a pulse sequence is repeated 64 to 256 times in a cycle TR (repetition time). The phase encoding gradient Gp is changed every time it is repeated, and a different phase encoding is performed each time. Thus, view data of 64 to 256 views (view) is obtained in the k space.
[0106]
Another example of the magnetic resonance imaging pulse sequence is shown in FIG. This pulse sequence is a pulse sequence of a gradient echo (GRE) method.
[0107]
That is, (1) is a sequence of α ° pulses for RF excitation in the GRE method, and (2), (3), (4) and (5) are slice gradient Gs, readout gradient Gr, It is a sequence of a phase encoding gradient Gp and a gradient echo MR. The α ° pulse is represented by a center signal. The pulse sequence proceeds from left to right along the time axis t.
[0108]
As shown in the figure, the α ° excitation of the spin is performed by the α ° pulse. α is 90 or less. At this time, the slice gradient Gs is applied, and selective excitation for a predetermined slice is performed.
[0109]
After the α ° excitation, spin phase encoding is performed by the phase encoding gradient Gp. Next, the spin is first dephased by the readout gradient Gr, and then the spin is rephased to generate the gradient echo MR. The gradient echo MR is an RF signal having a symmetrical waveform with respect to the echo center. The central echo occurs after TE from α ° excitation.
[0110]
The gradient echo MR is collected as view data by the data collection unit 150 as described above. Such a pulse sequence is repeated 64-256 times in a cycle TR. The phase encoding gradient Gp is changed every time it is repeated, and a different phase encoding is performed each time. Thereby, view data of 64 to 256 views is obtained in the k space.
[0111]
View data obtained by the pulse sequence of FIG. 33 or FIG. 34 is collected in the memory of the data processing unit 170. Note that the pulse sequence is not limited to the SE method or the GRE method, and may be one of other appropriate techniques such as a Fast Spin Echo (FSE) method and Echo Planar Imaging (Echo Planar Imaging). Needless to say.
[0112]
The data processing unit 170 reconstructs a tomographic image of the target 300 by performing two-dimensional inverse Fourier transform on the view data. Alternatively, when two systems of view data are collected using the four coil loops shown in FIG. 29, one set of view data is synthesized from the respective view data, and two-dimensional inverse Fourier transform is performed on the synthesized data. An image is reconstructed, or two images are reconstructed from two systems of view data, and a tomographic image is generated by combining these images.
[0113]
The generated tomographic image is displayed on the display unit 180. The displayed tomographic image has luminance and contrast distribution corresponding to the sensitivity distribution of the receiving coil unit 110. When the operator wants to change the brightness or contrast of a desired region such as a region of interest (ROI) by observing the display image, the operator receives the coil loop 222 closest to that region in the object 300, for example, the coil loop 222. The phase and the like of the signal are adjusted through the operation unit 190.
[0114]
Scanning is performed again with the adjustment of the phase and the like, and a reconstructed image based on a new received signal is displayed. In this image, the brightness and contrast of the ROI increase corresponding to the change in sensitivity distribution. The operator continues adjusting the phase and the like while observing the image until the desired brightness and contrast are reached. Each time the phase or the like is changed, scanning is performed, and a new reconstructed image is displayed. In this way, a reconstructed image with optimized brightness and contrast distribution can be obtained.
[0115]
The above is a case where the receiving coil unit 110 has two coil loops. However, in the case where the receiving coil unit 110 is a phased array coil having a number of coil loops, the distribution of desired luminance and contrast is similarly achieved. It goes without saying that it can be obtained.
[0116]
【The invention's effect】
As described above in detail, according to the present invention, a magnetic resonance signal receiving method and apparatus capable of adjusting the sensitivity distribution without changing the physical structure of the RF coil, and such a magnetic resonance signal receiving apparatus are used. A magnetic resonance imaging apparatus can be realized.
[Brief description of the drawings]
FIG. 1 is a block diagram of an exemplary apparatus according to an embodiment of the present invention.
2 is a schematic diagram showing a mutual relationship between a receiving coil unit and a target in the apparatus shown in FIG.
FIG. 3 is a schematic configuration diagram of a part of the receiving coil unit shown in FIG. 2;
FIG. 4 is an electric circuit diagram of a coil loop in a receiving coil section.
FIG. 5 is an electric circuit diagram of a pair of coil loops.
FIG. 6 is a block diagram of an exemplary apparatus according to an embodiment of the present invention.
7 is a schematic diagram showing a mutual relationship between a receiving coil unit and a target in the apparatus shown in FIG. 6;
8 is a block diagram of a part of a data collection unit in the apparatus shown in FIG. 1 or FIG. 6;
FIG. 9 is a graph showing reception sensitivity characteristics;
10 is a block diagram of a part of a data collection unit in the apparatus shown in FIG. 1 or FIG. 6;
FIG. 11 is a graph showing an intensity distribution of a received signal.
FIG. 12 is a diagram in which received signals are displayed as vectors.
FIG. 13 is a diagram in which received signals are displayed as vectors.
FIG. 14 is a diagram in which received signals are displayed as vectors.
FIG. 15 is a diagram in which received signals are displayed as vectors.
FIG. 16 is a diagram in which received signals are displayed as vectors.
FIG. 17 is a diagram in which received signals are displayed as vectors.
FIG. 18 is a diagram showing received signals as vectors.
FIG. 19 is a diagram in which received signals are displayed as vectors.
FIG. 20 is a diagram showing a vector of received signals.
FIG. 21 is a diagram in which received signals are displayed as vectors.
FIG. 22 is a diagram in which received signals are displayed as vectors.
FIG. 23 is a diagram in which received signals are displayed as vectors.
FIG. 24 is a diagram in which received signals are displayed as vectors.
FIG. 25 is a diagram in which received signals are displayed as vectors.
FIG. 26 is a diagram in which received signals are displayed as vectors.
FIG. 27 is a diagram showing a vector of received signals.
FIG. 28 is a diagram in which received signals are displayed as vectors.
FIG. 29 is a schematic diagram showing the arrangement of coil loops in the receiving coil section.
FIG. 30 is an electric circuit diagram of a coil loop in the receiving coil section.
31 is a block diagram of a part of a data collection unit in the apparatus shown in FIG. 1 or FIG. 6;
FIG. 32 is a schematic diagram showing the arrangement of coil loops in a saddle type coil.
33 is a diagram showing an example of a pulse sequence executed by the apparatus shown in FIG. 1 or FIG.
34 is a diagram showing an example of a pulse sequence executed by the apparatus shown in FIG. 1 or FIG. 6;
[Explanation of symbols]
100,100 'magnet system
102 Main magnetic field coil section
102 'main magnetic field magnet
106,106 'gradient coil section
108, 108 'RF coil section
110, 110 'receiving coil section
120 Bias drive
130 Gradient drive
140 RF drive unit
150 Data collection unit
160 Control unit
170 Data processing unit
180 display unit
190 Operation unit
210, 220, 210 ', 220' coil plate
212, 222, 214, 214 'coil loop
300 subjects
302, 302 ', 322, 322' capacitor
304, 304 ', 324, 324' conductor
306, 306 ', 326, 326' preamplifier
308, 308 ', 328, 328' inductor
310, 310 ', 330, 330' diode
312, 312 ′, 332, 332 ′ choke coil circuit
470, 480 adder
472, 472 ', 482' signal conditioning unit
500 cradle

Claims (5)

A magnetic field in which an object to be imaged is placed in a static magnetic field space, a magnetic resonance signal is generated inside the object using a gradient magnetic field and a high-frequency magnetic field by a predetermined pulse sequence, and an image is reconstructed based on the magnetic resonance signal A resonance imaging apparatus,
Receiving means for respectively receiving the magnetic resonance signals from the object by a plurality of RF coils that do not interfere with each other;
Signal adjusting means for adjusting a phase between the plurality of received magnetic resonance signals;
Adding means for adding the adjusted plurality of magnetic resonance signals;
Image reconstruction means for reconstructing an image based on the added magnetic resonance signal;
Display means for displaying the reconstructed image;
Operating means for the operator to perform a predetermined operation,
The operator looks at the image displayed on the display means, adjusts the phase using the operation means, and newly executes a pulse sequence, whereby the image based on the adjusted phase is displayed on the display means. The magnetic resonance imaging apparatus characterized by displaying . The magnetic resonance imaging apparatus according to claim 1.
The signal adjusting means adjusts the phase and individual amplitude between the plurality of received magnetic resonance signals,
The operator adjusts the phase and the amplitude using the operating means and newly executes a pulse sequence, thereby displaying an image based on the adjusted phase and amplitude on the display means. The magnetic resonance imaging apparatus. The magnetic resonance imaging apparatus according to claim 1 or 2,
Wherein the plurality of coils, magnetic resonance imaging apparatus shall be the features and scores Do the phased array coil. The magnetic resonance imaging apparatus according to claim 1 or 2,
The magnetic resonance imaging apparatus , wherein the plurality of coils are arranged at positions facing each other across the object . The magnetic resonance imaging apparatus according to claim 4.
2. The magnetic resonance imaging apparatus according to claim 1, wherein the plurality of coils are four coils, and two coils are arranged at positions facing each other across the object .

Images