JP2011056087A - Conductive member, and magnetic resonance imaging apparatus using the same - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To reduce unevenness in distribution of a rotary magnetic field B1 generated by an RF coil of a magnetic resonance imaging apparatus so as to prevent image irregularity in obtained images by using the simplest structure as much as possible. <P>SOLUTION: The conductive member 116 is disposed between a subject 103 and the RF transmission coil 114 to reduce unevenness in the rotary magnetic field B1 generated by the RF transmission coil. Specifically, it is disposed in such a way that a center part of the conductive member is close to a range of a section of the subject where B1 intensity is larger, and an end part of the conductor member is close to a range with smaller B1 intensity, so that unevenness in B1 intensity in the section of the subject is reduced, thereby B1 unevenness is reduced. <P>COPYRIGHT: (C)2011,JPO&INPIT

Description

  The present invention relates to an MRI (Magnetic Resonance Imaging) apparatus (hereinafter referred to as an MRI apparatus), and more particularly, to a magnetic resonance including a conductor member that reduces non-uniform spatial distribution of a rotating magnetic field that induces a magnetic resonance phenomenon. The present invention relates to an imaging device and a conductor member thereof.

The MRI apparatus is a medical image diagnostic apparatus that causes magnetic resonance to occur in nuclei in an arbitrary cross section that crosses an examination target, and obtains a tomographic image in the cross section from a generated magnetic resonance signal. A radio wave (Radio Frequency wave, hereinafter referred to as RF), which is a type of electromagnetic wave, is transmitted to the inspection object, and the spin of the atomic nucleus in the inspection object is excited, and then a nuclear magnetic resonance signal generated by the nuclear spin is received. The inspection object is imaged. RF is transmitted to the subject by the RF transmitting coil, and a nuclear magnetic resonance signal is received by the RF receiving coil.
In recent years, with the aim of improving the SNR (Signal to Noise ratio) of images, research and development has been progressing such as increasing the static magnetic field strength, and a high magnetic field MRI apparatus (3T MRI apparatus) with a static magnetic field intensity of about 3 T (Tesla). Has begun to spread. However, there is a problem that unevenness is likely to occur in the captured image as the static magnetic field strength increases. This is because the frequency of the RF used for inducing the magnetic resonance phenomenon increases as the magnetic field increases. The 3T MRI apparatus uses an RF having a frequency of 128 MHz, but the wavelength of the RF in the living body is about 30 cm, which is approximately the same scale as the abdominal cross section, and the phase of the RF changes in the living body. For this reason, the irradiation RF distribution and the spatial distribution of a rotating magnetic field (hereinafter referred to as B1) generated by the RF and inducing a magnetic resonance phenomenon are non-uniform, resulting in image unevenness. Under such circumstances, there is a need for a technique for reducing non-uniform distribution of the rotating magnetic field B1 in RF irradiation performed by an ultrahigh magnetic field MRI apparatus.

  As a method for reducing the nonuniformity of the B1 distribution, a method for devising the RF irradiation method has been proposed. Specifically, there is a technique called “RF parallel transmission technology”, which uses a gradient magnetic field and a plurality of RF transmission coils to independently control the RF pulse waveform itself applied to each coil. This is a technique of controlling the B1 distribution in the imaging region (for example, Patent Document 1 and Non-Patent Document 1).

Further, as one of RF parallel transmission techniques, a technique called “RF shimming” has recently appeared and attracted attention. (For example, Patent Document 2 and Non-Patent Document 2). RF shimming is a method of reducing B1 nonuniformity by controlling the phase and amplitude of an RF pulse applied to each coil.
Other than the method of devising the RF irradiation method, use of a “dielectric pad” is mentioned (for example, Non-Patent Document 3). This has the effect of changing the B1 distribution in the abdomen and shifting the position of the low B1 intensity by placing a pad having a certain dielectric constant on the imaging region, for example, the abdomen.

  In addition, studies have been made to place a coupling coil on the abdomen (Patent Document 3, Non-Patent Document 4). This has the effect of increasing the B1 strength by placing a coupling coil in the vicinity of a portion having a small B1 strength in the abdominal section.

US Pat. No. 6,900,656 US Pat. No. 7,078,901 WO200008100536

Katscher U et al., Transmit SENSE, Magnetic Resonance in Medicine, Vol. 49, pp. 144-150, 2003 Nistler J et al., Homogenity Improving Usage A 2 Port Birdcage Coil, Proceedings of International Society of Medicine, Med. 1063, 2007 Schmitt M et al., Improved uniformity of RF-distribution in clinical wise body-in-the-metic-in-the-bound-in-the-mapping-in-the-bound-in-the-mapping. 197, 2004 Schmitt M et al., B1-Homogenization in abdominal imaging 3T by means of coupling coils, Proceedings of International Society of Magnetic Res. 331, 2005

  According to Patent Document 1 and Non-Patent Document 1, in RF parallel transmission, a plurality of RF transmission coils that can be independently driven are prepared, and the waveform of an RF pulse applied to each coil is controlled. Also, the method of applying a gradient magnetic field during RF irradiation is controlled. Although this method can reduce the non-uniformity of B1, it has the disadvantage that the irradiation time of the RF pulse is as long as several tens of ms and has a problem that the pulse sequence is limited. The situation still remains. In addition, since it is necessary to arrange a plurality of RF transmission coils, the plurality of RF transmission coils, and a waveform generator and an RF amplifier that independently drive the RF transmission coils are more suitable. It costs a lot.

  The RF shimming disclosed in Patent Document 2 and Non-Patent Document 2 has an advantage that the pulse sequence is not limited because only the amplitude and phase of the RF pulse applied to a plurality of RF transmission coils are controlled. However, as compared with the RF parallel transmission technique described above, the current situation is that the effect of reducing the non-uniformity of B1 is inferior. Further, the cost required to equip a plurality of waveform generation devices and RF amplifiers is the same as that of the RF parallel transmission technology.

  According to Non-Patent Document 3, the dielectric pad exerts the effect of reducing the nonuniformity of B1 by placing it on the abdomen. However, since the dielectric pad has a weight of several kilograms, there is a problem that a burden on the patient arises. Have. In addition, depending on the sequence, the dielectric pad itself appears in white in the image, which has a demerit that obstructs diagnosis.

  According to Patent Document 3 and Non-Patent Document 4, the coupling coil is lighter (˜0.5 kg) than the dielectric pad, and there is a problem that it appears white in the image like the dielectric pad. It has the advantage of not occurring. However, when using the coupling coil, the resonance frequency of the coupling coil may change due to the difference in the positional relationship between the coupling coil and the abdomen, which may cause a difference in the effect of reducing B1 nonuniformity. Conceivable. The resonance frequency of the coupling coil can be changed by changing the values of resistors, capacitors, inductors, etc. arranged in the coupling coil. In this case, the B1 distribution is acquired in advance before imaging. Since the values of resistors, capacitors, inductors and the like are determined based on the distribution, it is considered that more time is required. In addition, the use of resistors, capacitors, and inductors increases the cost.

  An object of the present invention is to reduce the non-uniformity of the rotating magnetic field B1 distribution generated by the RF transmitting coil in consideration of the above-described problems of the prior art, and is as simple as possible without burdening the patient. It is to provide a conductor material having a simple structure.

  In order to solve the above-described problem, in the present invention, a conductor member is arranged between the subject and the RF transmission coil so as to reduce nonuniformity of the rotating magnetic field B1 generated by the RF transmission coil. When a conductor member is placed in a space where a magnetic field is generated by RF transmission, a phenomenon in which the magnetic flux in the vicinity of the conductor member becomes sparse or dense is greatly changed. In the present invention, the B1 distribution is controlled by actively utilizing this phenomenon. Specifically, the conductor member is arranged so that the central portion of the conductor member comes to the vicinity of the region where the B1 strength is high in the cross section of the subject, and the end portion of the conductor member is located near the region where the B1 strength is low By arranging so that the B1 comes to be small, the variation of the B1 intensity in the cross section of the subject is reduced, and the B1 nonuniformity can be reduced.

  According to the present invention, it is possible to reduce B1 nonuniformity by arranging a conductor member between a subject and an RF transmission coil. Therefore, it is not necessary to control the RF pulse, and the nonuniformity of the rotating magnetic field B1 can be reduced in a situation where the pulse sequence is not restricted at all. Further, it is not necessary to use a plurality of waveform generators and RF amplifiers for independently driving the RF transmitting coil, and the cost can be reduced. In addition, since the conductor member in the present invention exhibits an effect without contact with the subject, it is not necessary to place a member having a weight of several kg such as a dielectric pad on the patient. Further, in the present invention, the effect of changing the magnetic field by using only the conductor member is exhibited, so that it is not necessary to use circuit components such as resistors, capacitors, and inductors in principle. Therefore, no effort is required for adjusting resistors, capacitors, and inductors, and costs can be reduced.

It is a block diagram which shows the outline | summary of the MRI apparatus with which this invention is applied. It is a figure which shows the outline | summary of the RF coil of the MRI apparatus with which this invention is applied. It is a schematic diagram of an RF transmission coil and a phantom. It is a figure which shows the simulation result which shows distribution of the rotating magnetic field B1 in a phantom. It is a simulation result which shows the effect of the conductor member in the coil for RF transmission. It is a schematic diagram which shows the influence on the magnetic flux line of a conductor member. It is a figure which shows the mode of the simulation which confirms the effect of this invention by use of a phantom. It is a simulation result which shows distribution and B1 uniformity of rotating magnetic field B1 in a phantom at the time of changing the distance of a conductor member and a phantom. It is a simulation result which shows distribution of rotating magnetic field B1 in a phantom, and B1 uniformity when changing the coverage with a conductor member. FIG. 8 is a characteristic diagram that graphs changes in the B1 uniformity index with respect to changes in the distance between the conductor member and the phantom in FIG. 7. FIG. 9 is a characteristic diagram in which changes in the B1 uniformity index with respect to changes in the coverage ratio in FIG. 8 are graphed. It is a characteristic view which shows the change of the B1 uniformity in a phantom, and B1 average value when the coverage by a conductor member is changed. It is a figure which shows the principal part of 1st embodiment (Example) of this invention. It is a simulation result which shows the difference in distribution of the rotating magnetic field B1 in a phantom when not arranging the conductor member of the said Example. It is B1 distribution of a human lower abdomen obtained by the imaging experiment of a human lower abdomen using 3T MRI apparatus. It is a figure which shows the principal part of 2nd embodiment (Example) of this invention. It is a figure which shows the principal part of 3rd Embodiment (Example) of this invention. It is a figure which shows the principal part of 4th Embodiment (Example) of this invention. It is a figure which shows the variation of the shape of the conductor member of this invention. It is a simulation result which shows the effect of a plate-shaped and loop-shaped conductor member. It is a figure which shows the principal part of 5th Embodiment (Example) of this invention. It is a figure which shows the principal part of the 6th Embodiment (Example) of this invention.

  Embodiments of an RF coil and a magnetic resonance apparatus using the same according to the present invention will be described in detail below. Note that the present invention is not limited thereby.

<< First Embodiment >>
Hereinafter, a first embodiment of the present invention will be described. First, the overall configuration of the MRI apparatus of the first embodiment will be described. FIG. 1 is a block diagram of the MRI apparatus 100 of the present embodiment. As shown in this figure, the MRI apparatus 100 of the present embodiment includes a magnet 101 that generates a static magnetic field, a coil 102 that generates a gradient magnetic field, a shim coil 112 that adjusts the static magnetic field uniformity, a sequencer 104, and a high frequency An RF transmission coil 114 that generates a magnetic field, an RF reception coil 115 that receives a magnetic resonance signal generated from the subject 10, and a table 107 on which the subject 103 is placed. The gradient coil 102 and shim coil 112 are connected to a gradient magnetic field power source 105 and a shim power source 113, respectively. The RF transmission coil 114 and the RF reception coil 115 are connected to the high-frequency magnetic field generator 106 and the receiver 108, respectively. The sequencer 104 sends commands to the gradient magnetic field power supply 105, the shim power supply 113, and the high frequency magnetic field generator 106 to generate a gradient magnetic field and a high frequency magnetic field, respectively. The high frequency magnetic field is applied to the subject 103 through the RF transmission coil 114. A magnetic resonance signal generated from the subject 103 by applying a high-frequency magnetic field is detected by the RF receiving coil 115 and detected by the receiver 108. A magnetic resonance frequency used as a reference for detection by the receiver 108 is set by the computer 109 via the sequencer 104. The detected signal is sent to the computer 109 through an A / D conversion circuit, where signal processing such as image reconstruction is performed. The result is displayed on the display 110. The detected signals and measurement conditions are stored in the storage medium 111 as necessary. The sequencer 104 normally performs control so that each device operates at a timing and intensity programmed in advance.

The magnetic resonance imaging apparatus according to the present embodiment includes a conductor member 116 disposed in the vicinity of the subject 103. In the present invention, it has been newly found that the B1 non-uniformity in the subject 103 can be reduced by the conductor member 116. Hereinafter, the principle of reducing the B1 nonuniformity by the conductor member 116 will be described using the results of electromagnetic field analysis simulation.
First, FIG. 2 shows an example of a general form when the subject 103 is irradiated with RF by the RF transmitting coil 114. As shown in FIG. 2, when an image is picked up by the MRI apparatus, the subject 103 is inserted into the RF transmission coil, and the image is acquired by performing RF transmission. The RF transmission coil shown in FIG. 2 is called a birdcage coil and is one of the general shapes as an RF transmission coil in an MRI apparatus. The simulation result in the present embodiment uses a birdcage coil.

Next, FIGS. 3A and 3B show a rotating magnetic field B1 generated in the phantom 117 when the phantom 117 simulating the subject 103 is irradiated with RF by the RF transmission coil. 3A is a schematic diagram of a birdcage coil and a phantom, and FIG. 3B is a B1 distribution in the phantom 117. Note that the B1 intensity in FIG. 3B is dimensionless so that the maximum B1 intensity in the phantom is 1. The phantom 117 has a rectangular parallelepiped shape, and the dimensions in the x, y, and z axis directions are 300 mm, 200 mm, and 300 mm, respectively. This is a simplified shape assuming a cross section of the abdomen of the living body. In addition, the physical property value of the phantom is an electric conductivity of 1.0 S / m and a relative dielectric constant of 80, which is determined on the assumption of a water phantom close to the physical property value of a living body. A 24-lang birdcage coil was used to provide magnetic flux to the phantom. The dimensions of the birdcage coil are a diameter of 615 mm and a rung length in the z-axis direction of 400 mm. A simulation was performed by installing a cylindrical shield having a diameter of 655 mm and a z-axis dimension of 900 mm outside the birdcage coil. The frequency of RF irradiated from the birdcage coil was set to 128 MHz assuming a 3T MRI apparatus. Further, the birdcage coil is provided with feeding points 124 at two places, and by feeding a sine waveform to each feeding point 124, two orthogonal magnetic fluxes Bx = A1sin (ωt + φ1).
By = A2sin (ωt + φ2)
Is generated. At this time, the rotating magnetic field B1 is
B1 = (Bx + iBy) / 2
In order to generate B1 most efficiently, the amplitude ratio (A2 / A1) between Bx and By was set to 1, and the phase difference (φ2−φ1) was set to π / 2. This is an RF irradiation method called QD (Quadrature Drive), which is a standard RF irradiation method.

  As can be seen from FIG. 3B, it can be seen that in the phantom 117, the B1 intensity is greatly varied and non-uniform. This is the B1 non-uniformity that is currently a problem in the high magnetic field MRI apparatus. This phenomenon occurs because the RF frequency is as high as 128 MHz and the wavelength in the phantom is as short as about 30 cm, so the RF wavelength is on the same scale as the dimensions of the phantom. Specifically, when RF is incident from around the phantom and propagates inside the phantom, the phase of the RF changes greatly. Such RF non-uniformity is caused by the interference of these RFs. In this calculation result, regions where the B1 intensity is low are seen at two locations above and below the phantom, and regions where the B1 intensity is high are seen at the four corners of the phantom. This non-uniform B1 distribution represents an example of a general tendency of non-uniform B1 that occurs in the abdominal section of the subject.

  FIG. 4 shows changes in the magnetic field applied by the conductor member 116 when the conductor member 116 is disposed with respect to the phantom 117. The conditions for this simulation are shown below. The shape of the phantom is a rectangular parallelepiped similar to FIG. 3B, and the dimensions in the x, y, and z axis directions are 300 mm, 200 mm, and 300 mm, respectively. The physical property values of the phantom were the same as in FIG. 3B. The shape of the conductor member 116 is a rectangle existing on a plane parallel to the xz plane, and the lengths in the x and z axis directions are 120 mm and 300 mm, respectively. The distance between the phantom 117 and the conductor member 116 is 20 mm. In addition, as a physical property value of the conductor member, the conductivity is set to 5.8 × 10 ^ 7 S / m, assuming copper.

  In order to give a magnetic flux to the phantom, in FIG. 3A, power is supplied to two places of the birdcage coil. However, here, in order to confirm the effect of the conductor member 116 as clearly as possible, the power is supplied to only one place. A linear magnetic field is generated by supplying power. Specifically, power is supplied so as to generate a magnetic field By only in the y-axis direction, and the magnetic field shown in the figure is also a component By in the y-axis direction. The By magnetic field distribution shown in FIG. 4 is displayed by subtracting the By magnetic field distribution value when the conductor member 116 is not arranged from the By magnetic field distribution value when the conductor member 116 is arranged, The change in the By magnetic field distribution due to the arrangement of the conductor member 116 is shown.

  From FIG. 4, it can be confirmed that the By in the phantom 117 is largely changed in the vicinity of the conductor member 116. Specifically, the value of By is negative for the phantom region covered with the conductor member 116, indicating that the By value is reduced by the arrangement of the conductor member 116. On the other hand, in the region near both ends of the conductor member 116 in FIG. 4, the By value is positive, indicating that the conductor member 116 has an effect of increasing the By value.

  5A and 5B are schematic diagrams showing magnetic flux lines generated by the RF transmitting coil. FIG. 5A shows magnetic flux lines when the conductor member 116 is not present, and FIG. 5B shows magnetic flux lines when the conductor member 116 is present. 5A shows the linear magnetic field generated by the RF transmitting coil, while FIG. 5B shows that the magnetic flux lines are greatly bent by the conductor member 116. FIG. This is because if the conductor member 116 is disposed in the magnetic field, a current flows in the conductor member 116 in such a direction as to cancel the magnetic field. Due to this phenomenon, the magnetic flux density is sparse at the center of the conductor member 116, and the magnetic flux density is dense at the end of the conductor member 116. The magnetic field is changed also in FIG. 4 due to the generation of the density of the magnetic flux density by this phenomenon. As a result, the value of the rotating magnetic field B1 also changes.

  From these results, it was confirmed that the spatial distribution of the rotating magnetic field B1 generated by the RF transmitting coil can be controlled by arranging the conductor member 116 in the vicinity of the phantom 117. Therefore, it has been devised to reduce the B1 non-uniformity by effectively arranging the conductor members 116. An example is shown in FIG. In the B1 distribution in FIG. 3B, regions where the B1 intensity is low are seen at two locations above and below the phantom, and regions where the B1 intensity is high are seen at the four corners of the phantom, but the conductor member 116 is placed near the four corners of the phantom where the B1 strength is large. It is considered that the conductor member 116 is positioned so that the end of the conductor member 116 is in the vicinity of two places above and below the phantom having a low B1 strength. As a result, FIG. 6 has a structure in which four conductor members 116 are arranged. Considering the effects shown in FIGS. 4 and 5, it is expected that the B1 non-uniformity in the phantom 117 is reduced by the arrangement as shown in FIG.

  Therefore, next, the difference in the B1 non-uniformity reduction effect depending on the arrangement of the conductor members 116 will be considered. Specifically, the effect of the B1 uniformity due to the difference in size and position of the conductor member 116 will be considered based on the result of simulation.

  FIG. 7 shows the difference in B1 distribution and the values of the uniformity indices UNEMA and USD when the distance between the conductor member 116 and the phantom 117 is changed. The phantom is the same as in FIG. 3B. Looking at the B1 distribution, it can be seen that the larger the distance between the conductor member 116 and the phantom 117, the closer the B1 distribution is when there is no conductor member. This is because the influence of the magnetic flux density change given by the conductor member becomes smaller as the distance becomes larger. Under these calculation conditions, it can be seen that the B1 distribution is most uniform when the distance between the conductor member 116 and the phantom 117 is 20 mm.

  Next, a uniformity index is defined for the B1 intensity uniformity in the phantom 117 as follows.

Note that max (B1), min (B1), m (B1), and σ (B1) are the maximum value, minimum value, average value, and standard deviation of B1, respectively. UNEMA is an index using the maximum and minimum values of B1, and when max (B1) = min (B1), UNEMA = 100%, and the difference between max (B1) and min (B1) is As the value increases, the value of UNEMA decreases. In other words, the larger the value of the first uniformity index UNEMA shown in (Equation 1), the higher the uniformity of B1. The second uniformity index USD shown in (Expression 2) is a value obtained by dividing the standard deviation by the average value. This means that the smaller the variation in B1, the smaller the value of USD, and the smaller the value of USD, the higher the uniformity of B1.

  Looking at the values of UNEMA and USD shown in FIG. 7, UNEMA is maximum when the distance between the conductor member 116 and the phantom 117 is 20 mm, and USD is minimum when the distance is 20 mm. That is, it can be seen that the B1 uniformity is the highest when the distance between the conductor member 116 and the phantom 117 is 20 mm. Therefore, under this calculation condition, the closer the conductor member 116 is to the phantom 117, the greater the effect.

  FIG. 8 shows the difference in B1 distribution and the values of the uniformity indices UNEMA and USD when the area of the conductor member 116 is changed. Here, the coverage (%) shown in FIG. 8 is obtained by calculating the ratio of the area of the conductor member, assuming that the area of each of the yz plane and xz plane of the phantom is 100%. FIG. 8 shows the cases where the coverage is 25, 50, and 75%. Looking at the B1 distribution and the uniformity index, it can be confirmed that the uniformity is improved in any value of the coverage of the conductor member 116 as compared with the case where the conductor member 116 is not provided. When the uniformity is compared in the case where the conductor member 116 is used, the uniformity is highest when the coverage is 50%. This is because the end portion and the central portion of the conductor member are arranged at a more appropriate position with respect to the non-uniform distribution of B1 generated in the phantom 117. As a result, under the present calculation conditions, it was shown that the uniformity is highest when the coverage of the conductor member is about 50%.

  FIG. 9A is a characteristic diagram in which the change in the uniformization index with respect to the change in the distance between the conductor member and the phantom without the conductor member shown in FIG. 7 is graphed. FIG. 9B is a characteristic diagram in which changes in the uniformity index with respect to the coverage by the conductor member shown in FIG. 8 are graphed. As described above, it can be seen that the uniformity varies depending on the distance and the coverage of the conductor member. The point to be noted here is that the uniformity is improved by the conductor member 116 in any case within the range of the distance and coverage ratio (distance: 20 to 60 mm, coverage ratio: 25% to 75%) simulated this time. Is. Assuming that the conductor member 116 is actually set on the subject 103, it can be said that even if there is such a difference in distance and coverage, the effect of nonuniformity of B1 is exhibited. In order to further reduce the B1 nonuniformity, it can be said that it is desirable to shorten the distance between the conductor member 116 and the subject 103 and to set the coverage of the conductor member 116 to about 50% under this calculation condition. . However, when the B1 distribution varies due to the difference in the size, shape, and physical property value of the imaging region of the subject 103, the optimum position of the conductor member 116 also varies. It is desirable to determine the position of the conductor member 116.

  FIG. 10 is a plot of changes in the second uniformity index USD and the B1 average value m (B1) with respect to the change in coverage by the conductor member 116. The coverage was examined every 10% with finer steps than in the case of FIGS. Regarding the B1 uniformity, it can be seen that the variation in the B1 distribution becomes small when the coverage is in the range of 20% to 80%, that is, it is effective to arrange a conductor member having a uniform B1 distribution. Similarly to FIG. 9, when the coverage is about 50%, the uniformity index USD is the smallest and the B1 uniformity is high. On the other hand, the B1 average value m (B1) hardly changes until the coverage is about 0% to 60%. This is because, when the coverage is about 0 to 60%, the overall B1 intensity is kept almost constant by the conductor member 116 due to both the effect of decreasing the magnetic flux density and the effect of increasing the density. it is conceivable that. A decrease in B1 is observed when the coverage is around 70%. When the coverage exceeds 90%, the decrease in B1 is significant. From this, when the coverage by the conductor member is in the range of 20 to 80%, the decrease in the B1 average value is not a problem level, but the practical effect of the conductor member arrangement is exhibited, and in particular, the coverage is arranged to be about 50%. Can be said to be optimal. From the above, it has been confirmed that the B1 uniformity can be improved by reducing the B1 average value by the arrangement of the conductor member 116.

  FIG. 11 shows the main part of the MRI apparatus according to the first embodiment of the present invention. The subject 103 to be imaged lying on the table 107 is inserted into the bore of the RF transmission coil 114. Note that the subject 103 is shown in a schematic view in which only the imaging region is cut out. In the case of this example, the RF transmission coil 114 also serves as an RF reception coil. Four conductor members 116 are arranged between the RF transmission coil 114 and the subject 103. The size and arrangement of the four conductor members 116 are determined based on the result of electromagnetic field analysis simulation of B1 distribution with a phantom shown in FIG. The phantom 117 set here has a shape close to a rectangular parallelepiped, and the dimensions in the x, y, and z axis directions are 350 mm, 200 mm, and 300 mm, respectively. The xy section has a shape with curvature at four corners, and the radius of curvature is 30 mm. This is for simulating a shape closer to the abdominal section of the living body. In addition, as physical properties of the phantom, the electrical conductivity is 0.55 S / m and the relative dielectric constant is 50, which is a physical property closer to that of a living body. FIG. 12A shows a simulation result of the B1 distribution when there is no conductor member. Of the phantom surface area corresponding to the body surface of the subject, the central part of the conductor member is in the vicinity of the area where the magnetic field is large, and the end of the conductor member is located near the area where the magnetic field is small Determine the size and placement of the conductor members. FIG. 12B shows a simulation result of the B1 distribution in the process of determining the size and arrangement of the conductor members. If the arrangement of the conductor members in FIG. 12B is also used in imaging of the subject, it can be expected that the B1 distribution in the subject will be uniform. In the arrangement of FIG. 12B, the distance between the phantom 117 and the conductor member 116 is set to 20 mm for the conductor member 116 based on the results of FIG. 6 to FIG. The area of the conductor member 116 was set so as to cover 50% of each of the total four faces of the phantom yz plane and xz plane. The conductor member 116 is mounted around the imaging region of the subject 103 so that the arrangement of the conductor member 116 determined as described above is reproduced when the subject is imaged. In the embodiment of FIG. 11, the conductor member 116 is set on a shell 131 having a shape matched to the subject 103, and the shell 131 is mounted around the imaging region of the subject 103. In order to mount the conductor member 116 on the subject, a flexible mounting member may be used instead of the solid shell. A part of the conductor member 116 may be mounted on the table 107 on which the subject lies. In addition, as a method for adjusting the position of the conductor member, first, a method in which a conductor member is attached to the shell in advance before being attached to the patient, and various types of shells are prepared according to the size of the patient, the imaging region, and the like. Is mentioned. Secondly, a method of adjusting the position of the conductor member by providing a mechanism that allows the conductor member to be attached to and detached from the shell, or by providing a mechanism that slides the attached conductor member can be considered.

  So far, the effect of the conductor member 116 by electromagnetic field analysis simulation has been examined. Next, the results of an imaging experiment using a 3T MRI apparatus in the human lower abdomen will be described.

  FIG. 13 shows the B1 distribution in the lower abdomen of a male weighing 59 kg. FIG. 13A shows the result when the conductor member 116 is not provided, and FIG. 13B shows the result when the conductor member 116 is provided. As for the arrangement of the conductor members, as shown in FIG. 6, the conductor members 116 were arranged at four locations. The acquisition of the B1 distribution was performed using the α-2α method. When the conductor member 116 of FIG. 13 (a) is not present, there are two places where the B1 strength is low in the upper and lower portions in the abdomen, but FIG. 13 (b) shows that B1 is not present when the conductor member 116 is disposed. It can be seen that the uniformity is greatly reduced. From this result, it was found that the present invention has the effect of reducing B1 non-uniformity even in human abdominal imaging. In this experiment, the effect of the conductor member 116 in the lower abdomen was observed, but the prostate, uterus, spinal cord, etc. exist in the lower abdomen, and the conductor member 116 enables highly accurate diagnosis of these organs. .

  Although all the drawings show examples using four conductor members so far, the number of conductor members that achieve uniform B1 magnetic field strength with a subject or a setting phantom that simulates it is not limited to four. Of course.

<< Second Embodiment >>
Next, a second embodiment to which the present invention is applied will be described. FIG. 14 shows the main part of the MRI apparatus of the second embodiment. The shape of the RF transmission coil 114 is the same as that of the first embodiment. However, in this embodiment, the RF transmission coil 114 is dedicated to transmission. Separately, an RF receiving coil 115 is provided. The RF receiving coil 115 is mounted on a shell that fits the subject 103. Inside the shell, a conductor member 116 for making the B1 distribution of the subject 103 uniform is mounted. That is, imaging is performed in a state where the conductor member 116 for uniformizing the B1 distribution, the RF reception coil 115, and the RF transmission coil 114 are arranged in this order from the side closer to the body surface of the subject.

  In the MRI apparatus of this embodiment, even when the conductor member 116 for uniformizing the B1 distribution is not employed, it is necessary to attach a shell in which the RF receiving coil 115 is set to the subject as preparation for imaging. . By adopting a configuration in which the conductor member 116 for uniforming the B1 distribution is integrated in this shell, both the burden on the subject and the labor for imaging preparation by the engineer operating the MRI apparatus are both reduced. It becomes possible to do. The distribution of B1 intensity in each part of the subject also depends on the type and structure of the RF receiving coil. Therefore, it is reasonable to integrate these into the shell in order to realize the optimal arrangement of the conductor members for uniforming the B1 distribution for each RF receiving coil. In the case of this structure, the conductor member 116 can be arranged after the RF receiving coil 115 is set. It is also possible to adopt a structure in which the RF receiving coil 115 and the conductor member 116 are integrated. When integrated, it can be placed on the subject 103 with the same effort as the RF receiving coil 116 currently used, and there is no burden on the subject 103 and the MRI engineer. Has merit. In FIG. 14, the conductor members 116 are arranged at four locations, but this is an example, and the arrangement method is optimally arranged according to the cross-sectional shape of the subject 103 and the non-uniform distribution of B1. Is desirable.

<< Third Embodiment >>
Next, a third embodiment to which the present invention is applied will be described. FIG. 15 shows the main part of the MRI apparatus of the third embodiment. The point that the RF transmitting coil 114 and the RF receiving coil 115 are independent and the point that the RF receiving coil 115 is mounted on the shell mounted on the subject is the same as in the second embodiment of FIG. is there. The difference from FIG. 14 is that the conductor member 116 for equalizing the B1 distribution is set on the outer surface of the shell equipped with the RF receiving coil 115. Therefore, in this embodiment, the conductor member 116 for uniforming the B1 distribution, the RF receiving coil 115, the conductor member 116 for homogenizing the B1 distribution, and the RF transmitting coil from the side closer to the body surface of the subject. They are arranged in the order of 114. In the first and second embodiments, the conductor member 116 has been arranged to make the distribution of the magnetic field generated by the RF transmission coil 114 uniform. However, the conductor member 116 is inside the RF reception coil 115. In some cases, the magnetic resonance signal from the subject 103 is also received by the RF receiving coil 115.

  When considering the reception sensitivity distribution when receiving a magnetic resonance signal, a rotating magnetic field opposite to the rotating direction of the rotating magnetic field B1 generated by the RF transmitting coil is considered on the principle of signal acquisition. Become. For example, when the clockwise rotating magnetic field B 1 is generated by the RF transmitting coil 114, the reception sensitivity distribution when the RF receiving coil 115 receives the magnetic resonance signal is opposite to that of the RF receiving coil 115. It becomes equal to the intensity distribution of the clockwise rotating magnetic field. Therefore, in the subject 103, the distribution of the rotating magnetic field B1 generated by the RF transmitting coil and the distribution of the rotating magnetic field received by the RF receiving coil 115 are different. In that case, when the conductor member 116 is arranged, there is a case where an adverse effect is caused such that the distribution of the rotating magnetic field B1 is made uniform at the time of RF transmission, while the reception sensitivity is made non-uniform at the time of RF reception. On the other hand, the arrangement as shown in FIG. 15 has an advantage that the conductor member 116 does not adversely affect the reception sensitivity.

<< Fourth Embodiment >>
Next, a fourth embodiment to which the present invention is applied will be described. FIG. 16 shows the main part of the MRI apparatus of the fourth embodiment. The configuration of the MRI apparatus is basically the same as that of the second embodiment. The difference from the second embodiment is that the conductor member used for uniformizing the B1 distribution is changed to a loop-shaped conductor member 123.

  FIG. 17 shows various shapes of the conductor member. 17A shows the plate-like conductor member 116 employed in the embodiments so far, FIG. 17B shows the mesh-like conductor member 122, and FIG. 17C shows the loop-like conductor member 123. In order for the conductor member 116 to bring about an effect of reducing B1 nonuniformity, it is only necessary that a current flows through the conductor member so as to cancel the magnetic field generated by the RF transmitting coil 114. Therefore, the conductor member is not limited to a plate shape, and the effect can be obtained by a mesh structure or a loop structure. When adopting a mesh-like structure, it is possible to bring about an effect of reducing B1 non-uniformity by selecting a mesh size, a mesh material, and the like so that a sufficient current flows to generate a density of magnetic flux density. Become. Specifically, there is a practical effect in controlling the density of the magnetic flux density when the lattice size is in the range of about 1 mm to 100 mm. Also, assuming a 3T MRI apparatus, the RF frequency used is 128 MHz, and the wavelength in the air is 2.3 m. Therefore, if the loop has a sufficiently small scale with respect to this wavelength, the effect of generating the density of magnetic flux density necessary as a function of the conductor member can be secured, and the same effect as that of the plate-like conductor member can be obtained. I can expect that.

  FIG. 18 shows a simulation result for the change in the magnetic field applied by the conductor member when the plate-like and loop-like conductor members 123 are arranged. The conductor member is present on the xz plane, and the dimensions are 120 mm in the x-axis direction and 300 mm in the z-axis direction in the case of a plate, and 120 mm in the x-axis direction and z-axis direction in the case of the loop-shaped conductor member 123. The loop has a width of 300 mm and a width of 20 mm. In addition, as a physical property value of the conductor member, the conductivity is set to 5.8 × 10 ^ 7 S / m, assuming copper. Regarding the magnetic field to be applied, as in the case of FIG. 4, power was supplied to one place of the birdcage coil, and RF irradiation was performed so as to generate a linear magnetic field By in the y-axis direction. The By magnetic field distribution shown in FIG. 18 is displayed by subtracting the By magnetic field distribution value when the conductor member 116 is not disposed from the By magnetic field distribution value when the conductor member 116 is disposed. The change in the By magnetic field distribution due to the arrangement of 116 will be shown.

  From FIG. 18, it can be confirmed that the By in the phantom 117 is largely changed in the vicinity of the conductor member regardless of the conductor member. Further, the magnetic field distribution is substantially the same as compared with the plate shape and the loop shape. Specifically, for the phantom region covered with the conductor member, the By value is negative, which indicates that the By value is reduced depending on the arrangement of the conductor member. On the other hand, in the region close to both ends of the conductor member, the change in the By magnetic field is positive, indicating that the conductor member 116 has the effect of increasing the By value. Comparing (a) and (b) of FIG. 18, the loop-shaped conductor member 123 is slightly less effective in giving the density of magnetic flux density than the plate-shaped conductor member. It has been shown that the conductor member 123 can also be provided with a density of magnetic flux density.

  As shown in FIG. 16, it is possible to adjust the degree of changing the density of the magnetic field density by making the conductor member into a loop shape. A mesh-like conductor member has the same effect. In addition, there is an advantage that the weight can be reduced as compared with the plate shape. In FIG. 16, the conductor member 123 is arranged inside the RF receiving coil 115, but may be outside the RF receiving coil 115. It is also possible to adopt a structure in which the RF receiving coil 115 and the conductor member 116 are integrated.

<< Fifth Embodiment >>
Next, a fifth embodiment to which the present invention is applied will be described. FIG. 19 shows the main part of the MRI apparatus of the fifth embodiment. The configuration of the RF receiving coil 115, the use of a loop-shaped conductor member, and the like are the same as in the fourth embodiment shown in FIG. A difference from FIG. 16 is that a switch unit 118 is provided in the middle of a loop formed by the conductor member 123. By turning on / off the switch unit 118, there is an advantage that the effect that the conductor member 116 brings about the density of the magnetic field distribution can be turned on / off.

  The switch unit 118 includes a switch that can be manually switched on and off, for example, so that the MRI engineer can turn on and off the conductor member 116 as necessary when setting the conductor member 116 on the subject 103. It becomes. Other variations of the switch unit 118 include a diode switch, a switch composed of a diode and an inductor, and a MEMS (microelectromechanical system) that can electrically turn the switch unit 118 on and off. In the third embodiment, it has been described that the conductor member 116 may adversely affect reception sensitivity at the time of RF reception. However, when the switch unit 118 is electrically turned on / off, an adverse effect is caused. Can be eliminated. That is, when irradiating the subject 103 with RF by the RF transmission coil 114, the conductive member 116 has an effect of reducing B1 nonuniformity by turning on the switch unit 118. When receiving a magnetic resonance signal, it is possible to prevent the conductive member 116 from having an adverse effect by turning off the switch unit 118.

  Further, a modification using a variable resistor instead of the switch unit 118 is possible. That is, by changing the resistance value of the variable resistor inserted in the middle of the conductor loop, the effect of changing the density of the magnetic flux density by the loop-shaped conductor member shown in FIG. 18 (the degree of change in the B1 intensity distribution of the subject) ) Can be controlled.

<< Sixth Embodiment >>
Next, a sixth embodiment to which the present invention is applied will be described. FIG. 20 shows a main part of the MRI apparatus of the sixth embodiment. The basic configuration of the main body portion of the MRI apparatus of this embodiment is the same as that of the above-described embodiments. 20 is characterized in that a knitted conductor member 123 is formed so as to form a plurality of loops over the entire inner circumference of the shell equipped with the RF receiving coil 115. Each loop is not electrically completed as a conductor loop, but a switch unit 118 is inserted into a conductor branch constituting each mesh. In the example shown in the figure, switch portions are inserted in all the branches of the conductor extending in the circumferential direction. As these switch units, various switches such as an electrical switch that can be manually turned on / off can be applied as in the switch unit of the embodiment of FIG. In the body surface of the imaging region of the subject 103, the switch is turned on so that the loop is turned on in the vicinity of the region where the B1 intensity is large, so that the magnetic flux generated by the RF transmitting coil is canceled and the magnetic flux density becomes sparse. Make it work. In the vicinity of the region where the B1 intensity is small, the switch is turned off so that the loop is turned off, and the magnetic flux density is made to function. Thereby, the nonuniformity of the B1 distribution, which varies depending on the shape and physical property value of the subject 103, can be reduced.

  Compared to each of the above-described embodiments, this embodiment does not require the trouble of determining the width and arrangement of the conductor member corresponding to the B1 distribution, selecting from various conductor members, and setting it in the shell. is there. Instead, it is possible to bring out the effect of reducing the nonuniformity of B1 by selectively turning on and off the switches of each loop. That is, the time and effort for prior adjustment are greatly reduced. Moreover, versatility increases as a member having the function of a shield for adjusting the B1 distribution.

  Furthermore, the switch part 118 of each loop of the conductor member can be used by switching the switch between RF transmission and RF reception. Although described in the third embodiment, the distribution of the rotating magnetic field B1 generated by the RF transmitting coil and the distribution of the rotating magnetic field received by the RF receiving coil 115 are different. Therefore, with the structure of the conductor member shown in FIG. 20, by turning on and off the switch of the conductor portion 116 at the time of transmission and at the time of reception, the conductor necessary for homogenizing the magnetic field at the time of transmission and at the time of reception respectively. A highly uniform rotating magnetic field distribution can be realized by determining the position distribution of the loop and selectively switching on and off the switch section 118 of the conductor member 116 for transmission and reception. It becomes possible.

  In addition, the RF receiving coil 115 and the conductor member 123 can be integrated. The RF receiving coil is usually composed of a loop-shaped coil having an RF frequency as a resonance frequency, but the circuit of the RF receiving coil is devised so that the resonance frequency completely deviates from the RF frequency. By imparting a mode having, it is possible to have an effect of canceling the magnetic field generated by the RF transmitting coil, that is, it is possible to bring about an effect of changing the density of the magnetic flux density. As an example of circuit design, during RF reception, an inductor or conductor is placed in the loop of the RF reception coil to realize a mode having the RF frequency as the resonance frequency. During RF transmission, current flows through the inductor or conductor. It is conceivable to realize a mode in which a mode completely deviated from the resonance frequency is realized by switching the circuit using a switch such as a diode switch.

  According to the present invention, the non-uniformity of the rotating magnetic field strength of the magnetic resonance imaging apparatus can be reduced at a smaller apparatus cost than before, and stable imaging without image unevenness can be achieved. Is likely to be implemented.

101 Static magnetic field magnet 102 Gradient magnetic field coil 103 Subject 104 Sequencer 105 Gradient magnetic field power source 106 High frequency magnetic field generator 107 Table 108 Receiver 109 Computer 110 Display 111 Storage medium 112 Shim coil 113 Shim power source 114 RF transmission coil 115 RF reception coil 116 Conductor member 117 Phantom 118 Switch section 119 Human lower abdominal section 120 Magnetic flux line 121 Plate-like conductor member 122 Mesh-like conductor member 123 Loop-like conductor member 124 Feed point 131 Shell

Claims (20)

  Static magnetic field forming means for forming a static magnetic field, gradient magnetic field forming means for forming a gradient magnetic field, high frequency magnetic field forming means for forming a high frequency magnetic field, a magnetic field transmitting coil for irradiating the subject with the high frequency magnetic field, and the subject Magnetic field receiving coil for detecting magnetic resonance signals generated from a specimen, receiving means for receiving the magnetic resonance signals, magnetic field forming means, high-frequency magnetic field forming means, and control means for controlling the receiving means A resonance imaging apparatus, comprising: a conductor member disposed between the subject and a magnetic field transmission coil at a position that reduces non-uniformity of a magnetic field distribution in the subject. apparatus.   2. The magnetic resonance imaging apparatus according to claim 1, wherein the magnetic field transmitting coil and the magnetic field receiving coil are shared by the same coil, and the conductor member is installed in a bore of the shared coil.   The magnetic resonance imaging apparatus according to claim 1, wherein the conductor member is disposed on a shell attached to the subject.   The magnetic resonance imaging apparatus according to claim 1, wherein the conductor member is disposed on a shell attached to the subject integrally with the magnetic field receiving coil.   The magnetic resonance imaging apparatus according to claim 4, wherein the conductor member is disposed inside the magnetic field receiving coil.   The magnetic resonance imaging apparatus according to claim 4, wherein the conductor member is formed outside the magnetic field receiving coil.   In the magnetic field distribution in the subject generated by the magnetic field transmitting coil, the conductive member has a central portion of the conductive member near the region where the magnetic field is large and an end portion of the conductive member near the region where the magnetic field is small. The magnetic resonance imaging apparatus according to claim 1, wherein the magnetic resonance imaging apparatus is disposed in a magnetic field.   The conductor member is arranged so as to cover an area of 20 to 80% with respect to a subject surface area of a region in which a magnetic field is irradiated from the magnetic field transmitting coil to the subject. Item 2. The magnetic resonance imaging apparatus according to Item 1.   The magnetic resonance imaging apparatus according to claim 1, wherein the conductor member has a plate-like structure. The magnetic resonance imaging apparatus according to claim 1, wherein the conductor member has a mesh structure. The magnetic resonance imaging apparatus according to claim 1, wherein the conductor member has a loop structure. The magnetic resonance imaging apparatus according to claim 11, wherein the conductor member has a loop-like structure, and includes a conduction control unit that controls conduction of the conductor member in the loop.   The magnetic resonance imaging apparatus according to claim 12, wherein the conduction control unit is a switch.   The magnetic resonance imaging apparatus according to claim 13, wherein the switch includes a diode and an inductor.   The conductor member according to claim 13, wherein the switch is based on a microelectromechanical system.   The magnetic resonance imaging apparatus according to claim 13, wherein the switch is switched on and off during magnetic field transmission and magnetic field reception.   The magnetic resonance imaging apparatus according to claim 12, wherein the conduction control unit is a variable resistor. Used in a bore of a magnetic field transmission coil of a magnetic resonance imaging apparatus for transmitting a radio wave to a subject in a high magnetic field space to generate magnetic resonance, and between the transmission coil and the subject, the subject A magnetic resonance imaging apparatus characterized in that the magnetic resonance imaging apparatus is disposed at a position that covers a part of the periphery of the magnetic field and reduces non-uniformity of the magnetic field distribution in the subject generated by the transmission coil. The magnetic resonance imaging apparatus according to claim 18, wherein the magnetic resonance imaging apparatus is formed on a shell attached to the subject. The magnetic resonance imaging apparatus according to claim 18, wherein the magnetic resonance imaging apparatus has a loop structure.