JP5238194B2 - Magnetic resonance imaging system - Google Patents

Description

  The present invention relates to a magnetic resonance imaging apparatus that receives a magnetic resonance signal using an array coil formed by arranging a plurality of element coils.

    A method of efficiently performing parallel imaging using an array coil formed by arranging a plurality of element coils is known.

In this parallel imaging, it is possible to improve performance by considering the position of the array coil. For example, Non-Patent Document 1 discloses a technique for efficiently performing parallel imaging or general imaging using a large number of array coils using “Mode Matrix”.

A technique disclosed in Patent Document 1 is known as a technique for obtaining the position of the receiving coil. Patent Document 1 discloses a technique in which a marker is attached to the center of a receiving coil, an NMR (nuclear magnetic resonance) signal is collected from the marker, and the position of the marker is obtained from the collected NMR signal. A method of providing a special coil position detection mechanism is also proposed in Patent Document 2.
JP 7-124135 A Japanese Patent Publication No. 5-41256 A. Reykowski, M. Blasche, Mode Matrix-A Generalized Signal Combiner For Parallel Imaging Arrays, ISMRM 2004, p. 1587

  According to the technique of Patent Document 1, the position of the receiving coil can be obtained with high accuracy. However, the position of the receiving coil to which no marker is attached cannot be detected. Further, the method of Patent Document 1 is mechanically complicated and is difficult to apply to the position detection of many coils.

  The present invention has been made in consideration of such circumstances, and the object of the present invention is the position of a coil unit that does not include a position detection mechanism such as a marker, or is included in such a coil unit. Another object of the present invention is to provide a magnetic resonance imaging apparatus capable of accurately determining the position of each of a plurality of element coils.

A magnetic resonance imaging apparatus according to an aspect of the present invention includes an array coil in which a plurality of element coils each receiving a magnetic resonance signal from a subject are arranged, and a plurality of magnetic resonance signals received by each of the plurality of element coils. Based on the calculation unit for calculating projection data of each element coil in the arrangement direction of the plurality of element coils based on the above, based on the projection data of each of the plurality of element coils and the known information regarding the arrangement of the plurality of element coils And a determination unit for determining the position of each of the plurality of element coils or the position of the array coil.

  ADVANTAGE OF THE INVENTION According to this invention, it becomes possible to determine correctly the position of the coil unit which is not provided with the mechanism for position detection like a marker, or each position of the some element coil contained in such a coil unit.

  Hereinafter, an embodiment of the present invention will be described with reference to the drawings.

  FIG. 1 is a diagram showing a configuration of a magnetic resonance imaging apparatus (MRI apparatus) 100 according to the present embodiment. The MRI apparatus 100 includes a static magnetic field magnet 1, a gradient magnetic field coil 2, a gradient magnetic field power supply 3, a bed 4, a bed control unit 5, RF coil units 6a and 6b, a transmission unit 7, a selection circuit 8, a reception unit 9, and a computer system 10. It comprises.

  The static magnetic field magnet 1 has a hollow cylindrical shape and generates a uniform static magnetic field in an internal space. As the static magnetic field magnet 1, for example, a permanent magnet, a superconducting magnet or the like is used.

  The gradient magnetic field coil 2 has a hollow cylindrical shape and is disposed inside the static magnetic field magnet 1. The gradient magnetic field coil 2 is a combination of three types of coils corresponding to the X, Y, and Z axes orthogonal to each other. The gradient magnetic field coil 2 generates a gradient magnetic field in which the above three types of coils are individually supplied with current from the gradient magnetic field power supply 3 and the magnetic field strength is inclined along the X, Y, and Z axes. The Z-axis direction is the same as the static magnetic field direction, for example. The gradient magnetic fields of the X, Y, and Z axes correspond to, for example, the slice selection gradient magnetic field Gs, the phase encoding gradient magnetic field Ge, and the readout gradient magnetic field Gr, respectively. The slice selection gradient magnetic field Gs is used to arbitrarily determine an imaging section. The phase encoding gradient magnetic field Ge is used to change the phase of the magnetic resonance signal in accordance with the spatial position. The readout gradient magnetic field Gr is used for changing the frequency of the magnetic resonance signal in accordance with the spatial position.

  The subject 200 is inserted into a space (imaging space) inside the gradient magnetic field coil 2 while being placed on the top plate 41 of the bed 4. Under the control of the bed control unit 5, the bed 4 moves the top board 41 in the longitudinal direction (left and right direction in FIG. 1) and in the vertical direction. Usually, the bed 4 is installed such that the longitudinal direction thereof is parallel to the central axis of the static magnetic field magnet 1.

  The RF coil unit 6a is configured by housing one or more coils in a cylindrical case. The RF coil unit 6 a is disposed inside the gradient magnetic field coil 2. The RF coil unit 6a receives a high frequency pulse (RF pulse) from the transmitter 7 and generates a high frequency magnetic field.

  The RF coil unit 6 b is placed on the top plate 41, built in the top plate 41, or attached to the subject 200. And at the time of imaging | photography, it inserts in the imaging space with the subject 200. As the RF coil unit 6b, various types can be arbitrarily attached. One or three or more RF coil units for reception may be attached. Each of the RF coil units 6b includes at least one element coil. Each of the element coils provided in the RF coil unit 6b receives a magnetic resonance signal radiated from the subject 200. The output signals of the element coils are individually input to the selection circuit 8. The number of element coils that can be connected to the selection circuit 8 at the same time is indicated as m below. This m is, for example, “128”.

  The transmitter 7 supplies an RF pulse corresponding to the Larmor frequency to the RF coil unit 6a.

  The selection circuit 8 selects any 1 to n channel magnetic resonance signals from the maximum m channel magnetic resonance signals output from the RF coil unit 6b. However, n is a positive number smaller than m, for example, “32”. Then, the selection circuit 8 gives the selected magnetic resonance signal to the reception unit 9. Which channel is selected is instructed from the computer system 10.

  The receiving unit 9 includes n channels of a processing system having an amplifier, a phase detector, and an analog / digital converter. A maximum n-channel magnetic resonance signal selected by the selection circuit 8 is input to each of these n-channel processing systems. The amplifier amplifies the magnetic resonance signal. The phase detector detects the phase of the magnetic resonance signal output from the amplifier. The analog-digital converter converts the signal output from the phase detector into a digital signal. The receiving unit 9 outputs digital signals of maximum n channels obtained by each processing system.

  The computer system 10 includes an interface unit 11, a data collection unit 12, a reconstruction unit 13, a storage unit 14, a display unit 15, an input unit 16, and a main control unit 17.

  The interface unit 11 is connected to the gradient magnetic field power supply 3, the bed control unit 5, the transmission unit 7, the selection circuit 8, the reception unit 9, and the like. The interface unit 11 inputs and outputs signals exchanged between these connected units and the computer system 10.

  The data collection unit 12 collects digital signals output from the reception unit 9. The data collection unit 12 stores the collected digital signal, that is, magnetic resonance signal data in the storage unit 14.

  The reconstruction unit 13 performs post-processing, that is, reconstruction such as Fourier transform, on the magnetic resonance signal data stored in the storage unit 14 to obtain spectrum data or image data of the desired nuclear spin in the subject 200. Ask. The reconstruction unit 13 also generates projection data in the arrangement direction of the element coils based on the magnetic resonance signal data related to the magnetic resonance signal received by the specific element coil specified from the main control unit 17.

  The storage unit 14 stores magnetic resonance signal data and spectrum data or image data for each subject.

  The display unit 15 displays various information such as spectrum data or image data under the control of the main control unit 17. As the display unit 15, a display device such as a liquid crystal display can be used.

  The input unit 16 receives various commands and information input from the operator. As the input unit 16, a pointing device such as a mouse or a trackball, a selection device such as a mode change switch, or an input device such as a keyboard can be used as appropriate.

  The main control unit 17 includes a CPU, a memory, and the like (not shown), and controls the MRI apparatus 100 overall. When the array coil formed by arranging a plurality of element coils is used as the RF coil unit 6b, the main control unit 17 performs a scan for determining the position (hereinafter referred to as a position determination pre-scan). And a function for determining the position of the array coil based on the result of the position determination pre-scan. The function for controlling the position determination prescan includes a function for selecting a specific element coil to be used for the position determination prescan from among the element coils included in the array coil, and a condition for the position determination prescan. And a function for setting so as not to cause an overflow in reception. The function of determining the position includes a function of estimating each position of the specific element coil based on projection data obtained for each specific element, a position estimated for each of the specific element coils, and a plurality of element coils. And a function of determining the position of the array coil based on the known information regarding the arrangement state. The known information is typically a physical numerical value. The physical value is typically the distance between the element coils.

  FIG. 2 is a diagram showing an example of mounting the RF coil unit 6b. In addition, the same code | symbol is attached | subjected to FIG. 1 and an identical part, and the detailed description is abbreviate | omitted.

  In the example shown in FIG. 2, five RF coil units 6b are mounted. In addition, when it is necessary to distinguish each of these RF coil units 6b, each shall be described as RF coil unit 6b-1, 6b-2, 6b-3, 6b-4, 6b-5. The RF coil unit 6b-1 is a head coil, and is placed at an arbitrary position on the top board 41 so that the head of the subject 200 is inserted, or attached to the head of the subject 200 and attached to the subject. The sample 200 is placed on the top plate 41. Each of the RF coil units 6b-2 and 6b-3 is an abdominal coil and is attached to the surface side of the body of the subject 200. The RF coil units 6b-4 and 6b-5 are spinal coils, and are placed on an arbitrary position on the top board 41, and the subject 200 is placed on the back or the back of the subject 200. Is mounted on the top board 41 together with the subject 200. Thus, the position where these RF coil units 6b are placed on the top board 41 is indefinite. In addition, the RF coil unit 6b may be optimized for each part such as a knee.

  The RF coil units 6b-2, 6b-3, 6b-4, and 6b-5 each have a plurality of element coils 61 (four in this case) arranged at regular intervals in one direction as shown in FIGS. Is formed. The RF coil units 6b-2, 6b-3, 6b-4, and 6b-5 are used in a posture in which the arrangement direction of the element coils 61 and the Z-axis direction coincide as shown in FIG. FIG. 3 is a perspective view showing an arrangement state of the element coils 61 in the RF coil units 6b-2 and 6b-3.

  Each of the RF coil units 6b-2 and 6b-3 can be individually mounted at an arbitrary position, or connected to each other by a mechanical connection mechanism so as to be mounted at a constant interval. Is possible. The same applies to the RF coil units 6b-4 and 6b-5.

  The element coil 61 may be formed by further combining a plurality of coils. A method called Mode Matrix that uses a combination of about 3 to 4 coils has also been proposed (see Non-Patent Document 1). The coupling unit in this Mode Matrix can also be handled in the same manner as the element coil.

  Now, the code | symbol 21 in FIG. 2 shows the mount frame which accommodated the static magnetic field magnet 1, the gradient magnetic field coil 2, and the RF coil unit 6a. Of the imaging space inside the gantry 21, the imaging usage range that is actually used for imaging is only a part as shown by reference numeral 22. The RF coil units 6 b-2, 6 b-3, 6 b-4, and 6 b-5 are larger than the imaging usage range 22, and only some of the element coils 61 can be located in the imaging usage range 22. In the example of FIG. 2, the second to fourth element coils 61 of the RF coil units 6 b-2 and 6 b-3 are located within the imaging use range 22. That is, in the state shown in FIG. 2, the element coil 61 located within the imaging use range 22 is used for actual imaging. For this reason, as the number of array coils increases, it is necessary to know the position of the RF coil unit or the position of each element coil 61 and to appropriately select the element coil used for imaging.

(1) Determination of position of element coil 61 Therefore, in the MRI apparatus 100, when there is an RF coil unit 6b whose position on the top plate 41 is indefinite, each of the element coils 61 included in the RF coil unit 6b is determined. The position is determined as described below. This position determination may be performed only for the RF coil unit 6b whose position on the top board 41 is indefinite. For the RF coil unit 6b of the type fixed to the top board 41 like a spinal coil, Not applicable.

First, the main control unit 17 executes a position determination pre-scan. The technique disclosed in US Pat. No. 5,936,406 can be used for the position determination prescan. That is, the position determination pre-scan obtains projection data in that direction while applying a gradient magnetic field in the arrangement direction of the element coils 61 in the RF coil unit 6b, that is, in the Z-axis direction by the sequence shown in FIG. In this case, the projection data based on the magnetic resonance signal received by the element coil 61 located within the imaging utilization range 22 represents a rough position of the element coil 61 as shown in FIG. 5, for example. Therefore the main control unit 17 obtains the ends coordinates C1, C 3, for example, using a predetermined threshold value. Further, the main control unit 17 estimates the coordinate C 2 that is the midpoint between the end coordinates C 1 and C 3 as the center coordinate of the element coil 61.

  Note that no magnetic resonance signal is output or only a small magnetic resonance signal is output from the element coil 61 located outside the imaging use range 22. Therefore, the main controller 17 ignores the output signal from the element coil 61 and estimates the position only for the element coil 61 that has output a significant signal. When ignoring the output signal from the element coil 61, the projection data based on the corresponding signal may not be created, or the center coordinates may not be estimated based on the projection data generated from the corresponding signal. Alternatively, the center coordinates estimated for the element coil 61 that has output the corresponding signal may not be used for determining the position of each element coil 61. Further, it is not necessary to estimate the positions of all the element coils 61 that output a significant signal, and the positions may be estimated only for a part of such element coils 61.

  FIG. 6 is a diagram showing an example of projection data obtained for each of the four element coils 61 of the RF coil unit 6b-4. The interval between the four element coils 61 is 120 mm. A healthy person with a standard physique was used as the subject 200, and the pulse sequence shown in FIG. 4 was used and acquired with sagittal imaging (left-right projection) and 50 cm thickness (substantially non-selective excitation). Of course, it is better to use anti-aliasing by oversampling collection in the reading direction.

  The center coordinates estimated as described above for each of the four element coils 61 based on the projection data are 0 mm, 109 mm, 239 mm, and 331 mm as shown in FIG. 7 with the center coordinates of the first channel ch1 as a reference. is there. In this case, however, the center of the peak area after threshold processing is used as the estimated value of the center coordinates of the element coil 61 for the projection data. Here, the above threshold processing is performed with the half width of the peak of the projection data.

  The intervals between adjacent element coils 61 based on the above estimated values are 109 mm, 130 mm, and 92 mm, and do not match the known information of 120 mm. That is, the center position estimated as described above cannot accurately estimate each position of the element coil 61. Therefore, the main control unit 17 uses the coordinates (170 mm) obtained as an average value of the four central coordinates as reference coordinates, and based on the known information that the interval between the element coils 61 is 120 mm, each of the four element coils 61. The center coordinates are determined by the following calculation.

First channel ch1: 170-120 * 1.5 = -10 mm
Second channel ch2: 170−120 * 0.5 = 110 mm
Third channel ch3: 170 + 120 * 0.5 = 230 mm
Fourth channel ch4: 170 + 120 * 1.5 = 350 mm
In this way, each position of the element coil 61 is determined based on the relative relationship of the center coordinates estimated for each of the element coils 61 and the known interval between the element coils 61, so that the accuracy is higher. The position of the element coil 61 can be determined. That is, robust estimation is possible rather than obtaining the position of each element coil 61 individually.

  Although only the position determination of the element coil 61 included in the RF coil unit 6b-4 has been described here, the position determination is performed in the same manner for the other RF coil units 6b. However, regarding the plurality of RF coil units 6b connected to each other by the mechanical connection mechanism, the positions of the element coils 61 included in them can be collectively determined by the above processing.

  For example, it is assumed that the RF coil units 6b-2 and 6b-3 are connected to each other so that the distance between the element coils 61 located at the respective ends is 120 mm. At this time, for example, as shown in FIG. 2, if only the RF coil unit 6b-2 is in the imaging use range 22, information for determining the position of the element coil 61 included in the RF coil unit 6b-3 is obtained. Can't get. However, if the center coordinates of the element coils 61 of the second to fourth channels ch2 to ch4 of the RF coil unit 6b-2 are estimated as 118 mm, -2 mm, and -122 mm, respectively, as described above, the RF coil The center coordinates of the element coil 61 of the first channel ch1 of the unit 6b-2 can be determined to be 238 mm, and the center coordinates of the element coils 61 of each channel of the RF coil unit 6b-3 are -242 mm, -362 mm, -482 mm, Each can be determined to be −602 mm.

  Further, for example, if only the end portions of the RF coil units 6b-2 and 6b-3 are within the imaging use range 22, the estimation is made for at least one element coil 61 included in the RF coil unit 6b-2. Determining the position of each element coil 61 of the RF coil units 6b-2 and 6b-3 based on the position and the position estimated for at least one element coil 61 included in the RF coil unit 6b-3 Can do.

  When the number of element coils 61 included in the RF coil unit 6b whose position is indefinite exceeds the number of reception channels of the reception unit 9, the reception signals of all corresponding element coils 61 are collected in the position determination prescan. I can't. In such a case, a reception channel is assigned to each RF coil unit 6b whose position is indefinite. For example, if the number of receivable channels in the system is “16” and the number of RF coil units 6b whose positions are indefinite is “5”, three channels may be allocated to the five RF coil units 6b.

  The main control unit 17 displays the position of the element coil 61 determined in this way on a user interface for setting imaging conditions such as a scan plan. For this display, for example, a schematic diagram as shown in FIG. 2 can be used. This display may be performed only for the operator to confirm the position of the element coil 61, or may be performed for the operator to select the element coil 61 used for imaging. good.

  By performing such display, the operator can easily and accurately know the position of the element coil 61.

  Here, the position of each element coil 61 is the position information estimated from the received signal of the element coil and known information (physical position information, distance between element coils, distance from the coil unit center to the coil element center, etc.). Therefore, it is determined as a relative position from the center of the static magnetic field. Therefore, the display may be converted into a relative position with respect to the top board 41.

(2) RF power setting Position determination as described above is preferably performed at the earliest possible timing. In the MRI apparatus, generally, after the setting of the subject 200 is completed, the region of interest is moved to the center of the magnetic field, and a normal series of pre-scans for adjusting RF power, receiving gain, and the like are performed. Then, prior to these normal pre-scans, a position determination pre-scan is performed. In this case, the main control unit 17 determines the RF power in the position determination pre-scan according to the weight of the subject 200 according to the weight of the subject 200 input from the input unit 16 and the weight-specific RF table prepared in advance. Set to an appropriate value.

(3) Improvement of robustness by preventing overflow The physique of the subject serving as the subject 200 is various, and the magnitude of the magnetic resonance signal received by the element coil 61 varies greatly. Also, for example, the magnetic resonance signal received by the element coil 61 depending on the object to be imaged, such as the chest including the lung field or the lower extremity that is almost no signal in the midline tends to have a small magnetic resonance signal. The size of changes. In consideration of such circumstances, the slice thickness in the above-described position determination pre-scan is set to 50 cm (substantially non-selective excitation) so that a magnetic resonance signal useful for position determination can be reliably received. However, it is not preferable that the emitted magnetic resonance signal is too large because overflow occurs. Therefore, the level of the magnetic resonance signal used for position determination may be appropriately adjusted by any of the methods described below.

    (3-1) An appropriate reception gain corresponding to the weight of the subject 200 is set according to the weight of the subject 200 input from the input unit 16 and a weight-specific reception gain table prepared in advance.

    (3-2) The magnetic resonance signal is received a plurality of times while changing the reception gain, and a signal that does not overflow is used among the plurality of signals obtained thereby.

    (3-3) The magnetic resonance signal is received a plurality of times while changing the slice thickness, and a signal that does not overflow is used among the plurality of signals obtained thereby.

    (3-4) A magnetic resonance signal is received a plurality of times while changing TR (pulse sequence repetition time), and a signal that does not overflow is used among the plurality of signals obtained thereby.

    (3-5) Collect data for use in position determination after applying pre-saturation pulses and inversion pulses over the entire imaging area.

    (3-6) A multi-echo sequence with a twister is used as the position determination pre-scan sequence.

  In the methods (3-3) and (3-4), it is possible to change imaging conditions other than the slice thickness and TR.

By adopting any of these methods, the level of the magnetic resonance signal used for position determination can be adjusted appropriately. However, the method (3-1) may cause a reception error if the input information about the weight of the subject 200 is not accurate, and the methods (3-2) to (3-4) The time will be longer. In the method (3-4), since it depends on the relaxation time T ₁ related to the subject 200, it is difficult to control the received signal intensity.

  For this reason, the method (3-6) is superior to other methods. Therefore, the method (3-6) will be described in detail below.

  FIG. 8 is a timing diagram of a multi-echo sequence with a twister.

  In this sequence, a minute slice direction twister TW is applied cumulatively. In the SE (spin echo) system, the sign of the twister TW needs to alternate between positive and negative. The size of the twister TW is set to a value equal to or less than one encode of the slice thickness even if the multi-echo is accumulated (referred to as δ encode). Of course, instead of cumulative application, the effect may be reset with an inversion pulse for each echo, and a desired amount of spoiler may be applied with the next echo.

For example, assuming that the slice thickness is W, the gradient magnetic field strength of the twister TW is G, and the application time is T, γGTW = δ holds, so that δ = 0.2, W = 0.5 m, and T = 1 ms.
γG = 0.2 / ((1/1000) × (0.5)) = 400 [Hz / m]
G = 400 / (43.6 × 10 ⁶ ) ≈10 ⁻⁵ [T / m] = 0.01 [mT / m]
Thus, the twister TW may be a weak pulse. Although there is signal attenuation due to TE (echo time), the strength of the echo can be reliably reduced by the twister TW. Then, the main control unit 17 uses the first echo that becomes smaller than a predetermined overflow threshold value for position determination. For example, in the example shown in FIG. 8, the second echo is adopted.

  Incidentally, when the twister TW is used, in the rough approximation when the subject 200 has the same structure, the signal attenuates according to the Sinc function, so the relative signal amount changes as shown in FIG. 9 with respect to δ (attenuation). To do).

  For example, the twister amount as an actual pulse is 0, −0.5, with the difference added to the sign so that the cumulative twister amount is, for example, 0, 0.5, 0.8, 0.9. What is necessary is just to be 0.3 and -0.1.

  Since multi-echo sequences can be used in this way, repeated excitation is not necessary, so that the data acquisition time can be almost ignored. For example, if there are 4 echoes in a TE = 10 ms sequence, data collection can be completed in about 50 ms, which is very efficient.

  Since the signal strength may vary depending on the MRI apparatus, it is desirable that the twister amount and the reception gain are appropriately changed in consideration of this difference. However, even if there is a difference in signal strength, the strength of the received signal is reduced due to the effect of twister TW and the effect of T2 relaxation as the echo progresses, so that it does not overflow, so the difference in signal strength may be ignored. .

  This embodiment can be variously modified as follows.

  (1) The position determination pre-scan can be performed while the top plate is being fed after the setting of the subject 200 is completed at the start of the examination.

  As a typical implementation example, before starting the top board feed, the RF power is set in accordance with the RF table prepared according to weight prepared in advance. Gain is minimized. As shown in FIG. 10, while the top plate 41 is moved at about 18 cm / s, data collection is performed using the above-described multi-echo sequence with a twister every 2 seconds, that is, every time the top plate 41 moves 360 mm. . Although the movement of the top plate 41 may be stopped every time data is collected, it is also possible to collect data without stopping. However, when data collection is performed while the top plate 41 is moved, the top plate 41 moves about 9 mm during data collection, so it is desirable to correct the position for each echo.

  When the top plate 41 is inserted into the gantry 21, data for determining the position up to the first region of interest is prepared. When it is desired to directly confirm the coil mounting position from the top of the head to the tip of the foot in true whole body photography, data collection may be performed while moving the top board 41 over the entire area. The RF coil units 6b-2, 6b-3, 6b-4, and 6b-5 are scanned by the RF coil units 6b-2, 6b-3 by scanning in the states (a), (b), and (c) in FIG. , 6b-4, 6b-5, projection data for determining the positions of all the element coils 61 can be collected. FIG. 10 schematically shows projection data that is arranged on each element coil 61 and obtained based on the magnetic resonance signal received by the element coil 61.

  The main control unit 17 keeps track of the position of the top plate 41, and each element coil based on the position of the element coil 61 determined by the above-described position determination and the position of the top plate 41 when each scan is performed. The position 61 is determined.

  (2) Since the projection data is affected by the coil sensitivity, it has a shape similar to Gaussian. The position where the projection data takes the maximum value can be used as the estimated coordinates of the position of the element coil 61. Alternatively, in order to make it more robust, for example, threshold processing may be performed at about 20% of the maximum value, and the midpoint or center of gravity of the region above the threshold may be used as the estimated coordinates. Alternatively, fitting with Gaussian can be used as the estimated coordinates. In addition, a function having such a unimodal shape may be fitted.

  (3) The gradient magnetic field generally has non-linearity. Particularly in recent years, there is a tendency to allow the gradient magnetic field to some extent as a trade-off such as increasing the gradient magnetic field slew rate, increasing the maximum gradient magnetic field strength, or increasing the diameter. Therefore, in consideration of nonlinearity in the z direction, the position may be determined by adopting only data having a peak in a region of ± 15 cm or less from the center of the magnetic field. In this case, if the position determination pre-scan is performed while the top plate is being fed as described above, data may be collected while moving the top plate 41 by 27 cm every 1.5 seconds. . Of course, the non-linearity may be corrected and used as the estimation data.

  (4) Since an error occurs in the projection position information due to the non-uniformity of the static magnetic field, it may be corrected similarly. Here, since the positional distortion due to the static magnetic field inhomogeneity depends on the strength and polarity of the gradient magnetic field for reading, these are considered.

  (5) In the description so far, the position determination in the Z-axis direction has been described. However, however, if the element coils 61 are arranged in other directions such as the XY plane direction orthogonal to the Z-axis direction. For example, the positions of the element coils 61 can be determined. In this case, the reading direction in the position determination pre-scan may be a gradient magnetic field in the arrangement direction of the element coils 61. Position determination pre-scans may be repeated to determine positions in a plurality of directions among X, Y, and Z.

  (6) The position of each element coil 61 may be determined based only on projection data without referring to known information.

  (7) The physical numerical values relating to the arrangement of the element coils 61 may be various other numerical values such as a numerical value indicating the size of the RF coil unit 6b, a numerical value indicating the size of the element coil 61, A combination of these various numerical values may be used.

  (8) The known information regarding the arrangement of the element coils 61 may include not only physical numerical values but also other information such as information indicating the arrangement state of the element coils 61.

  (9) The position of the RF coil unit 6b may be determined.

  (10) Based on the projection data obtained for each of a part of the plurality of element coils 61 included in the RF coil unit 6b and the known information regarding the arrangement of the element coils 61, the part of the element coils The position of the element coil 61 other than 61 can also be determined. For example, the center coordinates of the element coils 61 of the second and third channels ch2 and ch3 are 118 mm and -2 mm based on the projection data relating to the first to third channels ch1 to ch3 of the RF coil unit 6b-2. If each is determined, the center coordinate of the element coil 61 of the fourth channel ch4 can be determined to be −122 mm.

  Note that the present invention is not limited to the above-described embodiment as it is, and can be embodied by modifying the constituent elements without departing from the scope of the invention in the implementation stage. In addition, various inventions can be formed by appropriately combining a plurality of components disclosed in the embodiment. For example, some components may be deleted from all the components shown in the embodiment. Furthermore, constituent elements over different embodiments may be appropriately combined.

The figure which shows the structure of the magnetic resonance imaging apparatus (MRI apparatus) which concerns on one Embodiment of this invention. The figure which shows the example of mounting | wearing of the RF coil unit in FIG. The perspective view which shows the arrangement | sequence state of the element coil in the RF coil unit in FIG. The figure which shows the basic pulse sequence in a position determination prescan. The figure which shows the principle which estimates the center coordinate of an element coil from projection data. The figure which shows an example of the projection data acquired regarding each of the four element coils contained in the RF coil unit in FIG. The figure showing the various numerical values calculated | required from FIG. The timing diagram of a multi-echo sequence with a twister. The figure which shows the change with respect to (delta) of the relative signal amount at the time of using a twister. The figure which shows a mode that a position determination pre-scan is performed during top-plate feeding.

Explanation of symbols

  DESCRIPTION OF SYMBOLS 1 ... Static magnetic field magnet, 2 ... Gradient magnetic field coil, 3 ... Gradient magnetic field power supply, 4 ... Bed, 5 ... Bed control part, 6a, 6b ... RF coil unit, 7 ... Transmission part, 8 ... Selection circuit, 9 ... Reception part DESCRIPTION OF SYMBOLS 10 ... Computer system, 11 ... Interface part, 12 ... Data collection part, 13 ... Reconstruction part, 14 ... Memory | storage part, 15 ... Display part, 16 ... Input part, 17 ... Main control part, 61 ... Element coil, 100 ... magnetic resonance imaging apparatus (MRI apparatus), 200 ... subject.

Claims (9)

An array coil in which a plurality of element coils each receiving magnetic resonance signals from a subject are arranged;
A calculation unit for calculating projection data of each element coil in the arrangement direction of the plurality of element coils based on a plurality of magnetic resonance signals received by each of the plurality of element coils;
And a determination unit for determining the position of each of the plurality of element coils or the position of the array coil based on the projection data of each of the plurality of element coils and the known information regarding the arrangement of the plurality of element coils. A magnetic resonance imaging apparatus. The determination unit is
An estimation unit that estimates the position of each of the plurality of element coils based on projection data of each of the plurality of element coils;
The position of each of the plurality of element coils or the position of the array coil is determined by correcting the estimated positions of the plurality of element coils based on physical values relating to the arrangement of the plurality of element coils. The magnetic resonance imaging apparatus according to claim 1 , further comprising a correction unit. The estimation unit estimates a center position of the array coil in the array direction based on the estimated positions of the plurality of element coils,
The magnetic resonance imaging apparatus according to claim 2 , wherein the correction unit determines the position of the element coil by correcting the center position based on the physical numerical value. The estimation unit performs a threshold process on the projection data of each of the plurality of element coils, and estimates the center of each area of the plurality of projection data subjected to the threshold process as the position of each of the plurality of element coils. The magnetic resonance imaging apparatus according to claim 2 . The magnetic resonance imaging apparatus can be used in a state where a plurality of the array coils are connected to each other.
The calculation unit calculates the projection data for each of a plurality of element coils included in the plurality of array coils,
The determination unit is configured to determine the position of each of the plurality of element coils included in the plurality of array coils or the plurality of arrays based on the projection data of each of the plurality of element coils and a connection state of the plurality of array coils. The magnetic resonance imaging apparatus according to claim 1 , wherein the position of each of the coils is determined. The magnetic resonance imaging apparatus according to claim 1 , wherein the calculation unit calculates the projection data only for an element coil that has been able to receive a magnetic resonance signal having a prescribed intensity or more. 2. The determination unit according to claim 1 , wherein the determination unit uses only the projection data related to an element coil that has received a magnetic resonance signal having a predetermined intensity or more among the projection data of each of the plurality of element coils for the determination. The magnetic resonance imaging apparatus described. 2. The magnetic resonance according to claim 1 , wherein at least one of the calculation unit and the determination unit does not use data having a large error due to non-linearity of a gradient magnetic field or non-uniformity of a static magnetic field for processing. Imaging device. The determination unit corrects the position determined based on the data with a large error due to the non-linearity of the gradient magnetic field or the non-uniformity of the static magnetic field in consideration of the non-linearity of the gradient magnetic field or the non-uniformity of the static magnetic field. The magnetic resonance imaging apparatus according to claim 1 .

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