JP4551946B2 - MRI gradient magnetic field generating coil - Google Patents

Description

  The present invention relates to an MRI gradient magnetic field generating coil including a winding portion in which a conductor is spirally wound around a bobbin. In particular, it relates to the improvement of the conductor winding position determined along the axial direction on the outer surface of a cylindrical bobbin, and is suitable for a Z-channel coil assembly of a gradient magnetic field coil of a magnetic resonance imaging (MRI) apparatus. The present invention relates to a gradient magnetic field generating coil.

In general, a spiral coil ("zonal" coil) in which a coil conductor is spirally wound around a cylindrical bobbin is used in various devices. A medical magnetic resonance imaging (MRI) apparatus is also equipped with this helical coil. A magnetic resonance imaging apparatus usually includes a static magnetic field coil that generates a static magnetic field, a shim coil that corrects the uniformity of the static magnetic field, a gradient magnetic field coil that generates a gradient magnetic field superimposed on the static magnetic field, and an RF coil for transmitting and receiving high-frequency signals. I use it. Among them, the helical coil is used for the Z channel coil of the gradient magnetic field coil, the Z ² , Z ³ , Z ⁴ , Z ⁵ , and Z ⁶ channels of the shim coil.

  When forming this spiral coil, two methods are known as the winding method of the coil conductor. One winding method is, as shown in FIG. 8, a method in which the winding is performed while changing the lane at each discrete winding position in the bobbin axial direction obtained analytically without actually winding in a spiral shape. is there. That is, when the coil conductor is wound approximately one turn (approximately 360 degrees) at a certain winding position, the lane change is made to the adjacent winding position at a position near the beginning of the winding when entering the next turn.

  As shown in FIG. 9, the other winding method is actually a spiral winding. That is, when the winding start end is aligned with a certain winding position for one turn with respect to the discrete winding position in the bobbin axis direction obtained analytically, the winding end end is the next turn at the next winding position. It is wound in a spiral that becomes the winding start end.

  However, the conventional spiral coil wound as described above has various problems. In the case of a spiral coil using the lane change winding method of FIG. 8, the coil winding itself is technically very difficult and complicated, and the manufacturing process of the spiral coil is remarkably increased. On the other hand, there is a problem that the magnetic field distribution as designed is often not obtained due to the difficulty of the winding work.

  This situation will be described in detail. Winding lane changes usually need to be done in a relatively narrow range (for example, about 10 cm) in the bobbin circumferential direction, so lane changes near the imaging area near the center of the bobbin axis are particularly congested. Work became difficult. As the winding density increases, it becomes difficult to provide sufficient space for lane changes. Further, since the actual winding work is performed while tension is applied to the conductor, it is difficult to bend the coil conductor, and there is a tendency to draw a larger arc at the bending position of the lane change. For this reason, the range in the circumferential direction required for the lane change tends to be enlarged, and there is a disadvantage that it affects the desired magnetic field distribution.

  On the other hand, as shown in FIG. 9, in the case of the coil winding method in which the winding start position and the winding end position are matched from the analytically determined winding position to the next winding position, the winding operation itself is performed. Is simpler than the method of FIG. 8, but has a problem in terms of the performance of magnetic field generation. That is, when the coil conductor is discretely wound, it is ideally wound one turn at each winding position analytically determined. However, as shown in FIG. 9, when the coil is wound spirally from the obtained winding position to the adjacent winding position, the equivalent magnetic field generation position based on the coil of one turn is the axis of the helical coil ( In the direction of the central axis), the winding position deviates from the analytical winding position. For this reason, the spatial magnetic field distribution generated by the entire spiral coil is greatly deviated from the design value, and the leakage magnetic field becomes large, so that the performance shortage is conspicuous in generating the magnetic field.

  The present invention has been made to overcome various difficulties associated with the prior art described above, and can facilitate the manufacturing work when the coil conductor is spirally wound around the bobbin, and the generated magnetic field. An object of the present invention is to provide an MRI gradient magnetic field generating coil capable of reducing a deviation from a design value analytically determined for distribution.

In order to solve the above-described problem, the MRI gradient magnetic field generating coil of the present invention includes a cylindrical bobbin around which a coil conductor is wound, and a winding portion in which the coil conductor is wound around the bobbin in a spiral manner. In the coil for generating an MRI gradient magnetic field having the above-described structure, the winding portion has a plurality of discrete analysis values obtained by reflecting a desired magnetic field distribution along the axial direction of the bobbin. The winding start position and winding end of one turn are aligned with the two winding positions sandwiching the analysis value, and the coil conductor position in the axial direction of the bobbin wound by 180 degrees of the one turn is between the two winding positions. The coil conductors are spirally wound so as to coincide with the analysis values respectively, and the winding position of the bobbin between the two analysis values located on both sides of the winding position is provided. axis It is characterized in that the position of the middle point obtained by equally dividing the distance direction.

  On the other hand, in the MRI gradient magnetic field generating coil, preferably, the outermost winding position of the analysis value is adjacent to the outermost analysis value located on the outermost side in the axial direction of the bobbin and the outermost analysis value. The bobbin is positioned axially outside of the outermost analysis value by a distance equal to the axial separation distance of the bobbin between a winding position located in the middle of the next analysis value to be matched. It is.

  The MRI gradient magnetic field generating coil is formed as a Z-channel coil assembly of the gradient magnetic field coil of the magnetic resonance imaging apparatus, and the gradient magnetic field coil is an active shield type gradient magnetic field coil. The coil assembly is both a Z-channel main coil and a shield coil.

  Further, for example, the MRI gradient magnetic field generating coil has a spiral coil pattern formed on the bobbin, and the turn portion is formed by winding the coil conductor according to the coil pattern. It is what.

  In the MRI gradient magnetic field generating coil according to the present invention, the coil conductor has the same coil characteristics as the lane change winding without winding the coil conductor, and the coil conductor is spirally wound in the axial direction of the bobbin. Since the coil conductor is wound one turn at a time so as to pass through the discrete analysis positions obtained analytically, the coil conductor prevents the deviation from the analytical analysis position and the winding position. The manufacturing process is simplified by simplifying the operation, and the deviation of the generated spatial magnetic field distribution from the analytically determined design value can be reduced to improve the gradient magnetic field generation performance.

  Embodiments of an MRI gradient magnetic field generating coil according to the present invention will be described below with reference to the accompanying drawings.

  In this embodiment, a Z-channel coil assembly of an active (self) shielded gradient coil (ASGC) of a magnetic resonance imaging (MRI) apparatus is illustrated as an MRI gradient magnetic field generating coil.

  FIG. 1 shows a schematic cross section of a gantry 1 of a magnetic resonance imaging apparatus. The entire gantry 1 is formed in a cylindrical shape, and a bore 2 at the center functions as a diagnostic space, and a subject P can be inserted into the bore 2 at the time of diagnosis.

  The gantry 1 includes a substantially cylindrical static magnetic field magnet 11, a substantially cylindrical gradient magnetic field coil 12 disposed in the bore 2 of the magnet 11, a shim coil 13 attached to, for example, the outer peripheral surface of the gradient magnetic field coil 12, And an RF coil 14 disposed in the bore 2 of the gradient coil 12. The subject P is placed on a couch top (not shown) and loosely inserted into a bore (diagnostic space) 2 formed by the RF coil 14.

  The static magnetic field magnet 11 is formed of a superconducting magnet. That is, a plurality of heat radiation shield containers and a single liquid helium container are housed in an outer vacuum container, and a superconducting coil is wound and installed inside the liquid helium container.

  Here, the gradient coil 12 is formed in an active shield type. Since this coil 12 generates a pulsed gradient magnetic field in each of the X-axis direction, the Y-axis direction, and the Z-axis direction, it has a coil assembly for each of the X, Y, and Z channels, and the coil assembly is provided for each channel. The shield structure hardly leaks the gradient magnetic field to the outside.

  Specifically, as shown in FIG. 2, the active shield type gradient coil (ASGC) 12 is insulated by the X, Y and Z channel X coil assemblies 12X, Y coil assemblies 12Y and Z coil assemblies 12Z for each coil layer. However, it is laminated | stacked and the substantially cylindrical shape is comprised as a whole. Each of the X coil assembly 12X, the Y coil assembly 12Y, and the Z coil assembly 12Z includes a main coil having a plurality of winding portions that generate gradient magnetic fields in the respective axial directions, and a gradient magnetic field generated by the winding portions of the main coils. And a shield coil having a plurality of winding portions for shielding (pulse) so as not to leak magnetically to the outside. Each coil assembly 12X, 12Y, 12Z is connected to a gradient magnetic field power source for each channel.

  Among them, the Y coil assembly 12Y includes four saddle type winding portions wound around the bobbin B on both the main coil side and the shield coil side. In other words, two sets of coil elements each having two saddle-type winding portions that are juxtaposed in the Z-axis direction and connected in series are arranged opposite to each other. As a result, a magnetic field linearly inclined in the Y-axis direction can be generated.

  The X coil assembly 12X is similarly arranged with the Y coil assembly 12Y rotated 90 degrees with respect to the Z axis.

  Furthermore, the Z-channel Z coil assembly 12Z includes a main coil 12Zm and a shield coil 12Zs whose shapes are schematically shown in FIG. As shown in FIG. 2, the main coil 12Zm and the shield coil 12Zs are stacked in a layered manner with a predetermined distance therebetween in the Z-axis direction.

  Each of the main coil 12Zm and the shield coil 12Zs includes a hollow cylindrical bobbin B and a flat coil conductor C as shown in FIG. The coil conductor C is wound on the bobbin B along a spiral coil pattern, and winding portions CRa and CRb are formed symmetrically on the left and right sides in the Z-axis direction. The winding portions CRa and CRb are electrically connected in series to each other by passing from one to the other winding portion at the center in the Z-axis direction of the bobbin B. A Z-channel gradient magnetic field is formed by the current flowing through the winding portions CRa and CRb.

  A plurality of spiral turns of the coil conductor C forming the winding portions CRa and CRb are formed based on the winding position setting method according to the present invention. This setting method is shown in FIG. The figure illustrates one winding portion CRb of the winding portions CRa and CRb, but the other winding portion CRa is wound symmetrically with respect to the XY plane.

In order to form the winding portion CRb (CRa), as shown in FIG. 5, first, the analysis value of the winding position in the Z-axis direction as the first winding position for discretely winding the coil conductor C is used. Ask. This analysis value is obtained, for example, by determining a desired magnetic field distribution for the Z channel of the gradient magnetic field, obtaining a streamline function of the magnetic field distribution reflecting the magnetic field distribution, and further, based on the streamline function and the coil current value. It is obtained by determining discrete winding positions. The analysis value of the first winding position obtained in this way is now the analysis value = z ₁ , z ₂ , z ₃ , z ₄ , z for one winding part CRb, for example, as shown in FIG. _5, and _{z 6.}

  Next, as shown in FIG. 5, the winding position in the Z-axis direction is obtained from the analysis value as the second and third winding positions.

  First, an intermediate value in the Z-axis direction as the second winding position is obtained. The “middle” of the intermediate values here is intended to be “located between” the analysis values.

Specifically, in this case, five intermediate values = z ₁₂ , z ₂₃ , z ₃₄ , z ₄₅ , z ₅₆ (corresponding to the second winding position of the present invention) are calculated between the analysis values. Set each. For example, the intermediate Z-axis direction of the analysis values _z 1, _{z 2,} calculates the Z-axis direction of the distance 2.alpha ₁ of their analysis values _z 1, _{z 2} a bisecting middle point, the middle point set position as the intermediate value _{= z 12.} That is, the intermediate value = z ₁₂ is the equidistant α ₁ from both the analysis values z ₁ and z ₂ . Similarly, the intermediate Z-axis direction of the analysis value z _2, z _3, Z-axis direction of the distance 2.alpha ₂ of their analysis value z _2, z ₃ 2 calculates the equally divided midpoint, in this to set a point located intermediate value _{= z 23.} The same applies to the middle of the analysis values z ₃ and z _{4 in} the Z-axis direction, the middle of the analysis values z ₄ and z _{5 in} the Z-axis direction, and the middle of the analysis values z ₅ and z _{6 in} the Z-axis direction. Those intermediate values = z ₃₄ , z ₄₅ and z ₅₆ are set. It should be noted that the accuracy of the midpoint position only needs to be within a predetermined error range in light of the actual operation of coil winding at the time of manufacture.

Further, outer values in the Z-axis direction as third winding positions are respectively determined on both outer sides in the Z-axis direction of the series of analysis value sequences. On the outside of _one analysis value = z _{1 in} the Z-axis direction, an outer value = z ₁₀ is set. According to this embodiment, by using the distance alpha ₁ from the analysis value z ₁ of the next to an intermediate value = z _12, and calculates the position of moving outside the analysis value z ₁ by the distance alpha ₁ minute This position is set as one outer value = z ₁₀ (one winding position). Further, outside value = z ₆₇ is set outside the other analysis value = z _{6 in} the Z-axis direction. Also in this case, the distance α ₅ from the next analysis value z ₆ to the intermediate value = z ₅₆ is used, and the position moved closer to the origin than the analysis value z ₆ by this distance α ₅ is calculated. This position is set as the other outer value = z ₆₇ (one winding position).

  Next, the coil conductor C is wound around the bobbin B in accordance with the winding position (analyzed value, intermediate value, outer value) set in this way. Since the coil conductor C is a flat plate here, the coil conductor C is wound so that the center position in the direction orthogonal to the longitudinal direction of the flat plate coincides with the winding position.

More specifically, first, in the reference position on the circumferential surface of the bobbin B, 1 present one of the outer values of winding starting end of the coil conductor C of (third winding position) = z Fit to _10. Thereafter, the coil conductor C is spirally wound, and the coil conductor position in the Z-axis direction when the coil conductor C reaches 180 degrees in the circumferential direction is made to coincide with the analysis value of the first winding position = z ₁ . The coil conductor position in the Z-axis direction when winding further 180 degrees from here, that is, the winding end end of one turn (and the winding start end of the next turn) is the intermediate value of the first second winding position = z Wound to match ₁₂ Thus, the first turn centering on the analytic value = z ₁ is completed.

The coil conductor C is continuously spirally wound as it is, and the next turn centered on the analytic value of the first winding position = z ₂ is wound. In the case of this turn, the winding position at the beginning of the winding is the first intermediate value = z ₁₂ , that is, the second winding position. Therefore, the coil conductor position when the winding is turned 180 degrees is set to the second first winding position. the analysis value = z ₂ winding position, align each coil conductor position when one round wound further 180 degrees in the second second winding position intermediate position = z _23. In the same manner, the turns with the analysis values of the first winding position = z ₃ , z ₄ , and z ₅ as the centers are sequentially and continuously wound. Finally, a turn centered on the analysis value of the first winding position = z ₆ is wound. This winding is carried out in the same manner with the last intermediate value = z ₅₆ and the remaining outer value z ₆₇ aligned with the winding start end and winding end end, respectively.

The coil conductor C wound up to the position of the outer value z ₆₇ is passed as it is in the Z-axis direction and sent to the other winding part CRa side. This winding portion CRa is similarly wound in a spiral shape.

  The active shield type gradient magnetic field coil formed in this way can generate a predetermined gradient magnetic field pulse by energization. In particular, since the Z-channel coil assembly 12Z is manufactured by employing the method of winding a gradient magnetic field generating coil according to the present invention on both the main coil side and the shield coil side, there are various advantages.

  First, as compared with the conventional coil winding method for performing lane change (see FIG. 8), the coil winding work is simplified, and the manufacturing cost can be reduced. In the case of this embodiment, lane change is not required as in the prior art, so the inconvenience associated with lane change described above is eliminated, and coil winding work is facilitated.

Second, there is an advantage that the magnetic field generation performance is improved as compared with the conventional coil winding work method (see FIG. 9) based on the analysis value of the winding position. Qualitatively, the equivalent winding position of each turn (spiral) around each of the analysis values = z ₁ ,..., Z ₆ in FIG. It can be explained by matching. In the case of the conventional FIG. 9, such an equivalent coil position for one turn is between the analysis values and deviates from the position of the analysis value itself. However, in the case of the method of the present invention, the equivalent coil position of one turn can be matched or brought closer to the analysis value determined as the most ideal under the coil position discretization method, thereby improving the magnetic field generation performance. Can be made.

  FIG. 6 and FIG. 7 compare and show graphs for explaining the improvement of the magnetic field generation performance. Both figures show a predetermined distance from the gradient magnetic field coil 12 when the winding method of the coil for generating MRI gradient magnetic field is applied to the Z-channel coil assembly (main coil and shield coil) of the active shield type gradient magnetic field coil 12. : Simulates the leakage magnetic field distribution of the Z channel at a position 2 cm away from the outside. The graph of FIG. 6 represents an example of the conventional coil winding method of FIG. 9 described above, and the graph of FIG. 7 represents an example of the MRI gradient magnetic field generating coil winding method. The horizontal axis represents the distance [cm] from the center in the Z-axis direction, and the vertical axis represents the leakage magnetic field strength [millitesla: mT]. The number of coil turns is 30 to 40, and the coil current value is 200 [A].

  The curve in FIG. 6 varies greatly up and down, and it can be seen that the leakage magnetic field is large in the conventional method. On the other hand, the vertical fluctuation of the curve in FIG. 7 is significantly smaller than that in FIG. 6, and the leakage magnetic field can be greatly reduced by employing the winding method of the MRI gradient magnetic field generating coil. I understand. In the position where the effect of preventing the leakage magnetic field is the greatest, it is reduced to about 1/3. In particular, the central portion whose position in the Z-axis direction is about 40 cm ± 20 cm corresponds to a magnetic resonance imaging (MRI) imaging region. According to the present invention, the leakage magnetic field at this central portion is also the conventional coil winding shown in FIG. It can be seen that there is a significant decrease compared to the packaging method.

Note that the MRI gradient magnetic field generating coil according to the present invention may be implemented only on either the Z-channel main coil or the shield coil of the active shield type gradient magnetic field coil 12. The MRI gradient magnetic field generating coil of the present invention can be similarly applied to a Z-channel coil assembly of a gradient magnetic field coil that is not an active shield type. Furthermore, the MRI gradient magnetic field generating coil and winding method according to the present invention can be widely applied to spirally wound coils. For example, Z ² , Z ³ , Z ⁴ of shim coils for static magnetic field correction in a magnetic resonance imaging apparatus. , Z ⁵ , Z ⁶ ,... And other channels other than Z ¹ can be preferably implemented, and similar effects can be obtained.

  Further, in the above-described embodiment as the second and third winding positions of the present invention, the midpoint of the distance between the positions of the analysis values (first winding positions) is used. The second and third winding positions are not necessarily limited to this, and can be positions that are positively shifted by a predetermined distance from the midpoint.

  The coil for MRI gradient magnetic field generation according to the present invention analytically determines a plurality of discrete first winding positions along the axial direction of the bobbin on the outer surface of the bobbin. A second winding position is determined for each of the windings, and the winding start and the winding end of one turn of the coil conductor are aligned with two adjacent second winding positions, and the winding is wound by 180 degrees. Since the coil conductor is wound so that the coil conductor position coincides with the first winding position between the two second winding positions, the coil conductor is spirally wound around the bobbin. Manufacturing process can be made much easier than before, and at the same time, the deviation of the gradient magnetic field distribution generated from the analytically determined design value can be reduced to improve the shielding performance or to increase the shielding. Such as maintaining performance Thereby improving the magnetic field generation performance.

The schematic sectional drawing of the gantry of the magnetic resonance imaging apparatus which implemented the coil for MRI gradient magnetic field generation of this invention. FIG. 2 is a schematic cross-sectional view of a surface orthogonal to the Z-axis direction of the active shield type gradient coil (a fractured surface along the line II-II in FIG. 1). The schematic diagram explaining the winding state of the coil of the Z channel of an active shielding type gradient magnetic field coil. The schematic diagram explaining the winding method of the conductor of the coil assembly of a Z channel. The rough flowchart which shows the process of the winding method. The figure which shows the simulation result of the leakage magnetic field distribution which concerns on the winding method of the conventional spiral winding. The figure which shows the simulation result of the leakage magnetic field distribution which concerns on this invention. The figure which shows an example of the conventional winding method. The figure which shows another example of the conventional winding method.

Explanation of symbols

12 Active shield type gradient magnetic field coil 12Z Z channel gradient magnetic field coil 12Zm Main coil (magnetic field generating coil)
12Zs shield coil (coil for magnetic field generation)
CRa, CRb Winding part C Coil conductor B Bobbin

Claims (6)

In an MRI gradient magnetic field generating coil having a cylindrical bobbin around which a coil conductor is wound, and a winding portion in which the coil conductor is spirally wound around the bobbin.
The winding portion analyzes the start and end of one turn of the coil conductor for each of a plurality of discrete analysis values obtained by reflecting a desired magnetic field distribution along the axial direction of the bobbin. The coil coil position in the axial direction of the bobbin wound 180 degrees of one turn is matched with the analysis value between the two winding positions in accordance with the two winding positions sandwiching the value. Provided with a turn part spirally wound with a coil conductor ,
The winding position is a midpoint position obtained by equally dividing the axial separation distance of the bobbin between two analysis values located on both sides of the winding position . coil. The coil for MRI gradient magnetic field generation according to claim 1, wherein the analysis value is determined based on a streamline function of a desired magnetic field distribution and a coil current value. The outermost winding position of the analysis value is an outermost analysis value located on the outermost side in the axial direction of the bobbin and a winding position located in the middle of the next analysis value adjacent to the outermost analysis value. 2. The MRI gradient magnetic field generating coil according to claim 1, wherein the MRI gradient magnetic field generating coil is located outside the outermost analysis value by a distance equal to an axial separation distance between the bobbins therebetween. The MRI gradient magnetic field generating coil according to any one of claims 1 to 3 , wherein the MRI gradient magnetic field generating coil is formed as a Z-channel coil assembly of the gradient magnetic field coil of the magnetic resonance imaging apparatus. The MRI gradient magnetic field generating coil according to claim 4 , wherein the gradient coil is an active shield type gradient coil, and the coil assembly is both a Z-channel main coil and a shield coil. A spiral coil pattern formed on the bobbin;
The MRI gradient magnetic field generating coil according to claim 1, wherein the turn portion is formed by winding the coil conductor according to the coil pattern.

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