Magnetic resonance imaging device with planar gradient coil arrangement

Abstract: Shielded gradient coils are adapted for use in an MR imaging apparatus having spaced apart pole pieces. The x, y, and z coils each include active (40, 38, 36) and shield layers (32, 30, 34). The x-active (40) and shield (32) layers are interconnected by a plurality of vertical electrical connections (44) extending around the periphery of the coil. The y-active (38) and shield (30) layers are similarly connected (42). The z-active (40) and shield (34) layers each have a generally spiral conductor pattern.

Description: The present invention relates to magnetic resonance imaging. It finds particular application in conjunction with gradient coils for use therein. Magnetic resonance imaging (MRI) is a widely used diagnostic imaging method. MRI equipment includes magnet apparatus for generating a very strong, homogeneous static magnetic field. A so-called "open" magnet arrangement includes a pair of pole pieces disposed on opposite sides of an imaging region. Because it is necessary to impose a series of rapidly varying gradient magnetic fields on the static magnetic field, MRI systems typically include a set of gradient coils disposed between the pole pieces and the imaging region to generate the requisite gradients, usually in orthogonal x, y, and z directions. Unfortunately, the time varying gradient fields induce eddy currents in electrically conductive materials, such as the magnet pole pieces, in the vicinity of the gradient coils. These eddy currents have various deleterious effects, most notably opposing rapid variations in the gradient magnetic fields. The interactions between the gradient magnetic flux and the iron pole pieces can cause field transients during changes in the gradients. The eddy currents also create a non-homogeneous magnetic field which decays with a relatively long time constant, for example on the order of hundreds of milliseconds. These effects lead to worsened image quality and can make imaging using relatively rapid sequences difficult, if not impossible. Various attempts have been made to overcome these effects. In a so-called passive shielding technique, materials having desirable magnetic properties have been placed on the surface of the pole piece so as to attenuate the magnetic flux coupled to the bulk iron of the pole pieces. However, it is difficult to find materials having suitable properties, particularly at higher fields strengths. Because the magnetic flux generated by the gradients tends to penetrate deeply into the pole piece, the material needs to be relatively thick. Further, this approach decreases the undesirable effects but does not eliminate them completely. Another technique involves electronic compensation of eddy current effects. Due to the complexity of the various interactions between the gradient fields and the pole pieces, effective compensation is difficult. Moreover, this technique does not account for gradient induced heating of the pole pieces. Still another technique is disclosed in our U.S. Patent No. 5,555,251, entitled Arrangement to Minimize Eddy Currents in MR Imagers. The disclosed technique includes uses a series of electrically insulated layers. A series of slots further reduces the eddy currents in the layers. While this technique has proven effective in reducing eddy current effects, further improvement is nonetheless desirable. A solution to the problems discussed above is achieved with an apparatus as defined in claim 1. The features of the preamble of claim 1 are known in combination from US-A-5581187. Ways of carrying out the invention will now be described in greater detail, by way of example, with reference to the accompanying drawings, in which: Figure 1 depicts a magnetic resonance imaging apparatus according to the present invention; Figure 2 is a side view which depicts the various layers of a gradient coil assembly; Figure 3 is an exploded view of a z-axis shield coil; Figure 4 is an exploded view of a z-axis active coil; Figure 5 is an exploded view of an x-axis shield coil; Figure 6 is an exploded view of an x-axis active coil; Figure 7 is a side view which depicts the spacing between the layers of a gradient coil assembly; and Figure 8 depicts the vertical interconnections between the active and shield layers of an x-axis gradient coil. With reference to Figure 1, an MRI apparatus which produces images of the anatomy of patient 1 includes a generally C-shaped magnet body 3. The patient 1 is placed in an imaging region located between a pair of opposed pole pieces 23. A main magnetic field control 5 supplies energizing current to the driver coils 2 associated with each pole so as to generate the main magnetic field Bo in the z direction. Necks 4 connect the pole pieces 23 to the body 3 of the magnet, thereby providing a return path through the body of the magnet. Gradient coil assemblies 6 located between each pole piece 23 and the imaging region generate time-varying gradient magnetic fields, preferably in three orthogonal directions (e.g., x, y, z). While only the lower gradient coil assembly 6 is shown in Figure 1, it will be appreciated that a second gradient coil assembly is similarly located in relation to the upper pole piece such that the coils are positioned symmetrically with respect to the imaging volume. Stated another way, the upper and lower coil assemblies are symmetrical with respect to reflection in the z=0 plane. The gradient coil assemblies 6 are substantially planar and are positioned so as to lie in the x-y plane (e.g., normal to the main magnetic field Bo). The respective gradient coils are energized by a gradient field controller 9 which contains a gradient field amplifier associated with each coil. The gradient controller 9 is in turn controlled by a computer 11 to generate time varying gradient sequences necessary to carry out a desired magnetic resonance excitation and detection sequence. The apparatus also includes an r.f. coil system 7 which includes a transmitting and a receiving coil arrangement. While only the lower r.f. system 7 is shown, it will be appreciated that a second set of r.f. coils is similarly located in relation to the upper pole piece such that the coils are positioned symmetrically with respect to the imaging volume. The coil 7 is energized by r.f. transmitter 15 under control of the computer 11 to apply an r.f. field. The r.f. coil system 7 also includes an r.f. receiving coil arrangement arranged to detect r.f. signals resulting from magnetic resonance excited in the body of the patient 1. The detected signals are passed via a receiver 19 to an imager 21 which. under control of the computer 11, processes the signal to produce an image representative of the patient's body. These signals are, in turn, passed to display device 24 to provide a human-readable display of the image. In operation of the apparatus, the strong magnetic field provided by the magnet defines an equilibrium axis of magnetic alignment in the body of the patient 1. To obtain an image of a selected region, e.g. a cross-sectional slice of the body, an r.f. field pulse is first applied to the patient 1 using coil system 7 to excite magnetic resonance in the selected region. To this end, the coil system 7 produces a field in a direction orthogonal to the static field direction so as to tip the spins of nuclei in the selected region from the direction of the static field. To restrict excitation to a desired region, the r.f. field pulse is applied in conjunction with magnetic field gradients imposed by the coil arrangement 6, the frequency of the r.f. field being chosen in conjunction with the magnitudes and directions of the imposed gradients so that the Larmor frequency protons in the body, e.g. hydrogen protons, is equal to the r.f. field frequency only in the desired region. The r.f. signals resulting from excitation are then spatially encoded by application of one or more further gradient magnetic fields in known manner, detected by the r.f. coil system 7, and processed to produce an human-readable image, for example on monitor 24. Normally, a number of excitation and signal detection sequences are required to collect sufficient data to produce a satisfactory image. Each gradient coil assembly 6 includes coils for generating gradients in the x, y, and z directions. With reference to Figure 2, a y-axis shield coil 30 is disposed nearest the pole piece, followed by an x-axis shield coil 32, a z-axis shield coil 34, a z-axis active coil 36, a y-axis active coil 38, and an x-axis active coil 40. A series of electrically insulating dielectric layers (not shown) separates the coil layers. Various of the conductors of the of the x-axis shield coil 32 are connected to conductors of the x-axis active coil 40 by vertical electrical connections represented schematically at 44 to form a squirrel cage structure. Similar vertical electrical connections represented schematically at 42 connect various of the conductors of the y-axis shield coil 30 with conductors of the y-axis active coil 38. A cooling layer 45 is disposed between the z-axis shield 34 and the z-axis active 36 coils so as to provide cooling for the gradient coil assembly 6. The z-axis shield 34 and active 36 coils are connected in series external to the assembly 6 such that they are driven in series by the gradient amplifier. With reference to Figures 3 and 4, the z-axis shield 34 and the z-axis active 36 coils each include two layers 46, 48 and 60, 62, respectively. Each layer is fabricated from a copper plate having a thickness of 2 mm. The plates each have a spiral gap 35 formed therein such that each layer forms a flat spiral winding. The plates also have slits 37 which are about 1 mm in width through the copper plate. The slits extend in the direction of current flow so as to reduce eddy currents within the coil generated by flux having a component perpendicular to the plane of the coil. With reference to Figure 3, the z-axis shield coil 34 includes first 46 and second 48 layers, the first layer 46 being disposed nearer to the pole piece. Non-conductive dielectric layers 50 separate the coil layers 46, 48. The interior end 52 of the first spiral coil layer 46 is welded to the interior end 54 of the second spiral layer 48 through an aperture in the insulating layer so that the layers 46, 48 form a continuous spiral coil. The z-axis shield coil is driven through electrical connections at connectors 56, 58. The first 46 and second 48 layers are substantially identical, the second shield layer 48 being "flipped" with respect to the first layer 46 about the x-axis so that the layers are mirror images of each other. A particular advantage of this configuration is that undesirable radial current components arc substantially cancelled. Current is. for example, introduced at the outer edge of the first layer 46 at connection point 56. Viewed from the perspective of Figure 3. this current flows in a counterclockwise direction toward the centre of the first layer 46 where the coil is connected to the second layer 48. There, the current continues to flow in a counterclockwise direction, but toward the outer edge of the second layer 48 and connection point 58. Because the radial components of the currents in the first 46 and second 48 layers are equal and opposite, the undesirable radial current component is effectively cancelled. As shown in Figure 4, z-axis active coil 36 includes first 60 and second 62 layers, the first layer 60 being disposed nearer the pole piece. As with the shield coil 34, the first 60 and second 62 active layers are substantially identical. The second active layer 62 is "flipped" with respect to the first layer 60 about the x-axis so that the layers are mirror images of each other. Non-conductive dielectric layers 64 separate the coil layers 60, 62. Tab 66 at the interior of spiral coil layer 60 is welded to tab 68 at the interior of spiral layer 62 through an aperture in the insulation layer so that the layers 60, 62 form a continuous spiral coil. The z-axis active coil 36 is driven through electrical connections at tabs 70, 72. As with the z-axis shield coil 34, the undesirable radial current components are substantially cancelled. As shown in Figure 5, the x-axis shield coil 32 includes first 74 and second 76 coil layers, the first layer 74 being disposed nearer to the pole piece. Each coil layer is fabricated from copper plate having a thickness of 2 mm to form a plurality of conductors. The conductors also have slits 33 which are about 1 mm in width through the copper plate and extending in the direction of current flow. The slits arc located in regions having a relatively vertical flux component so as to reduce eddy currents within the coil generated by flux having a component perpendicular to the plane of the coil. Non-conductive layers 78 separate the layers 74, 76. The second shield layer 76 is "flipped" with respect to the first layer 74 about the x-axis so that the layers are substantially mirror images of each other. Tab 80 on first layer 74 is welded to tab 82 on the second layer 76 through an aperture in the insulation layer 78. Tab 84 on first layer 74 is likewise welded to tab 86 on the second layer 76. An advantage of this configuration is that undesirable current components, especially those in a direction perpendicular to the desired current flow, may be substantially cancelled. Current is, for example, introduced into a conductor in a spiral region of the first shield coil 74 at connection point 79. Viewed from the perspective of Figure 5, this current flows in a clockwise direction toward the centre of the spiral region and connection point 80 where the conductor in the first shield layer 74 is connected to a conductor in the second shield layer 76 at connection point 82. There, the current continues to flow in a clockwise direction, but toward the outer edge of the layer 76 and connection point 81. Because undesirable current components in the first 74 and second 76 layers are equal and opposite, these components are effectively cancelled. These interconnections also allow the four layers of the x-axis gradient coil to be configured as a squirrel cage having one input and one output terminal. A plurality of electrical connections 88 such as tabs or fingers is disposed around the outer circumference of the first 74 and second 78 layers. As described more fully below, these connections 88 allow conductors in the various layers of the x-axis coil to be connected in series. As shown in Figure 6, the x-axis active coil 40 includes first 90 and second 92 layers, the first 90 layer being disposed nearer to the pole piece. Each layer is fabricated from copper plate having a thickness of 2 mm to form a plurality of conductors. The plates also have slits 33 which are about 1 mm in width through the conductors. The slits extend in the direction of current flow and so as to reduce eddy currents within the coil generated by flux having a component perpendicular to the plane of the coil. The second active layer 92 is "flipped" with respect to the first layer 90 about the x-axis so that the layers are mirror images of each other. Non-conductive dielectric layers 94 separate the coil layers 90, 92. Tab 96 on first layer 90 is welded to tab 98 on the second layer 92 through an aperture in the insulation layer 94. Tab 100 on first layer 90 is likewise welded to tab 102 on the second layer 92. As with the like connections on the shield coil, these connections allow undesirable current components to be effectively cancelled. A plurality of electrical connections 104 such as fingers or tabs is disposed around the outer circumference of the first 90 and second 92 layers. The fingers 104 and 88 on the x-axis shield coil 32 layers and the x-axis active coil 40, respectively, provide the vertical electrical connections shown schematically at 44 of Figure 2. The fingers are located so that, when the gradient coil is assemblies, corresponding fingers are located at the same angular position about the circumference of the coil. The fingers are preferably bent over and welded together to provide the necessary interconnections between the layers. Other suitable methods such as individual wires, conductors, or the like may also be utilized. Preferably, the average current in the x,y direction around the circumference of the coil should be substantially zero. Figure 8 illustrates the various interconnections between the various layers of the x-axis active 40 and x-axis shield 32 coils. For ease of illustration, the coils are shown separated in an exploded fashion. Current is applied to the x-axis shield coil assembly 9 through electrical connections illustrated at 106 and 107. As noted above, the layers are interconnected to form a squirrel cage structure having a single pair of electrical connections 106, 107. As will be appreciated, the interconnections scheme is applicable to coils having any number of conductors. Returning now to Figures 5 and 6, the preferred interconnections between the x-axis shield 32 and active 40 layers depicted therein will be described in greater detail. The various connection points are designated with underlined reference numerals (e.g., 1 , 2 , 3 , 4 , ...). Like numbered connection points are connected together. Thus, for example, connection point 2 on the second shield layer 76 is electrically connected to connection point 2 on the first active layer 90, connection point 118 on the second active layer 92 is electrically connected to connection point 118 on the first shield layer 74, and so on. Electrical connections to the gradient power source may advantageously provided at connection points 0 and 129 on the second active layer 92, it being understood that a break in the coil prevents a direct electrical connection between these two points. Current entering the coil at point 0 may be visualized as proceeding through the coil to point 1, then to point 2 , then to point 3 , and so on in sequence, until exiting at point 129 . As will be appreciated, within each layer 74, 76, 90, 92, current in the central conductors runs in the same direction. Current in the central conductors of the active layers 90, 92 likewise runs in the same direction. Similarly, current in the central conductors of the shield layers 74, 76 runs in the same direction, although this direction is opposite to that in the active layers 90, 92. It will also be appreciated that the squirrel cage connections in each of the active layers are connected by vertical connections to the squirrel cage connections in a corresponding shield layer. Thus, for example, the first active layer 90 is connected to the second shield layer 76 at points 2 through 62 . Similarly, the second active layer 92 is connected to the first shield layer 74 at points 64 through 126 . The current in the vertical connections between the first active layer 90 and the second shield layer 76 and between the second active layer 92 and the first shield layer 74 runs in the same direction. Hence, it is not necessary that the vertical component of the current be cancelled. At the same time, the horizontal components of the current in the vertical connections between the first active layer 90 and the second shield layer 76 and between the second active layer 92 and the first shield layer 74 are in opposite directions. Hence, the horizontal components of the current are cancelled. While the foregoing discussion has focussed on the x coil, it is of course equally applicable to the y coil as well. The vertical interconnections prevent the return flux from extending beyond the edges of the coils. In particular, the vertical wires connect respective wires in the active and shield coils such that the sum of the current in the x, y plane (i.e. the horizontal component of the current) at each position x, y along the outer circumference of the coils is as close to zero as possible: where p is the total number of x-active and x-shield layers. The operational characteristics of the gradient coil system are dependent on a number of interdependent variables. In particular it is desirable that the slope of the magnetic field gradient be constant over the imaging region. For example, the x-gradient field should obey the relationship: dB x dx = k where k is a constant equal to the desired gradient field strength. In order to minimize the power requirements of the gradient amplifier and maximize the gradient slew rate, it is also desirable that the energy stored in the gradient coils be as low as possible. Stated another way, it is desirable that the inductance of the gradient coils be as low as possible. Maximum shielding performance requires that the scattered flux density Bs which couples to the iron of the pole piece should be as low as possible. Thus, B s → 0 It will be appreciated, however, that a practical coil represents a compromise between these requirements. For example, the shield coil is configured to cause the return flux to run between the active and shield coils. Of course, current in the conductors of the shield coil influences the gradient field inside the imaging volume. This effect may be corrected by the configuration of the conductors in the active layer. These corrections in turn cause leakage of flux into the pole pieces. In order to compensate for this leakage, the configuration of the shield coil should be changed, and so on. Similarly. increasing the shielding effectiveness causes an increase in flux density in certain of the regions between the layers as well as on the imaging volume side of the coil. This increases coil inductance. Improving the gradient field linearity inside the imaging volume requires modification of the shape of the conductors (especially in the active layers) in such a way to increase the coil inductance. Hence, the configuration of the coil is preferably optimized using a numerical optimization procedure. In order to reduce the complexity of the optimization and to remove undesirable spherical harmonics from the gradient field decomposition, the shape of all conductors in the active or shield layers is expressed via current distribution function V(r), where r is the radius of a point on a layer. The function V(r) describes the conductors as follows. The boundary between two neighboring conductors is the plurality of points which satisfies the expression U(r,ϕ) = V(r) cos(mϕ) where m=0 for the z coil and 1 for the x-coil, and ϕ is the angular position about the x, y plane. A numerical optimization is utilized to determine the function V(r) using the cost function: C(V a ,V s ) = a L C L (V a ,V s ) + a S C S (V a ,V s ) + a E E(V a ,V s ) where V a is the current distribution function in the active layer; and V s is the current distribution function in the shield layer. C L (V a ,V s ) is a cost function component related to the gradient field non-linearity inside the imaging volume, C S (V a ,V s ) is a cost function component related to the scattered flux touching the pole piece, and E(V a ,V s ) is the total magnetic energy. Coefficients a L , a S , and a E define the relative importance of each of the terms in the cost function. The cost function components C L (V a ,V s ) is defined as follows: where B l = G z R sin(α i ) for the z gradient; B l = G x R cos(α i ) for the x gradient; G z and G x are the required gradient strengths; R is the radius of the imaging volume; and α 1 is a set of m angles between 0 and π / 2 distributed evenly or almost evenly. The cost function component C S (V a ,V s ) is defined as follows: where ρ i is a set of n points located on the surface of the pole piece along the x-axis. These points are distributed evenly or almost evenly. The calculation may be made using conventional finite element software or other similar computational method. The same software may also be used to compute the total magnetic energy. The values of C L (V a ,V s ) and C S (V a ,V s ) are evaluated by direct computation of the flux densities at n+m points created by the current distribution described by functions V a and V s . Suitable values for the coefficients a L , a S , and a E are, for example, as follows: x and y gradient coils z gradient coils a L 1 1 a S 0.03 0.6 a E 0.807 x 10 -8 2.75 x 10 -8 The foregoing discussion has focussed on the gradient coils for the x direction. It will be appreciated that the y-axis shield 30 and active 38 coils are substantially identical in construction to the corresponding x-axis coils, although they are rotated ninety degrees (90°) about the z-axis so as to provide a gradient in the y-direction. Figure 7 is a side view which depicts distance in millimeters between the various layers of the gradient assembly 6 associated with the upper pole piece and the centre of the magnet, it being understood that these distances are also representative of the distances for the lower gradient assembly. A particular advantage of the coil arrangement depicted in Figure 7 is that the sensitivities of the various coil are approximately equal so that, for an equal drive current, each coil produces an approximately equal gradient field. For a given spacing between active and shield layers, the z-axis coils would have a sensitivity greater than those of the x-axis and y-axis coils. Hence, it is advantageous that the z-axis layers are disposed inside the x-axis and y-axis layers. In this way, the distance between the x-axis and y-axis is relatively greater than the distance between the z-axis layers. Similarly, the x-axis and y-axis coils are spaced equally so that each has a substantially equal efficiency. A gradient coil system as described herein has been demonstrated to provide gradient fields in a system having a total gap between the pole of 500 mm and an imaging volume which is a sphere with radius 360 mm. The total diameter of the gradient set is 890 mm, and the thickness of each gradient assembly 6 is 50 mm. Other salient characteristics are listed below: Parameter x and y coils z coil Current @ 10mT/M 142.5 A 119.2 A Slew rate at 300V (theoretical) 25.4 T/m/S 59.2 T/m/S Stored energy @ 10 mT/m 8.414 J 3.018 J Power @ 10 mT/m (theoretical) 1.99 kW 2.03 kW Coil Resistance (theoretical) 0.098 Ohms 0.143 Ohms Leakage flux density @ 10 mt/M centre of pole < 0.2 mT < 0.2 mT edge of coil < 6.3 mT < 1.1 mT Gradient linearity in air 4.98 % 4.85 % with iron 5.3 % 5.3 % Number of turns active layer 44 35 shield layer 36 30 The invention has been described with reference to the preferred embodiment. Obviously, modifications and alterations will occur to others upon reading an understanding the preceding description. It is intended that the invention be construed as including all such modifications and alterations insofar as they come within the scope of the appended claims or the equivalents thereof.

Claims: An apparatus for use in magnetic resonance imaging comprising a pair of opposed pole piece magnets with first and second pole pieces (23) disposed on opposite sides of an examination region in order to generate a main magnetic field within the examination region, and further comprising a first generally planar gradient coil arrangement (40, 32), the first gradient coil arrangement having an outer perimeter and being disposed between the first pole piece and the examination region, the first gradient coil arrangement comprising an active coil (40) and a shield coil (32) situated in respective planes perpendicular to said magnetic field, the shield coil being disposed between the active coil and the first pole piece, and the active coil and shield coil being electrically connected (42, 44) at a plurality of locations along the outer perimeter of the gradient coil arrangement, characterized in that , the active coil (40) and the shield coil (32) further both comprise a respective first layer (90, 74) and a respective second layer (92, 76), the respective first and second layers being mirror images of each other about an axis perpendicular to the magnetic field created by the pole piece magnets. An apparatus for use in magnetic resonance imaging as claimed in claim 1, characterized in that the first gradient coil arrangement generates a first gradient magnetic field in a direction orthogonal to the main magnetic field, and that there is a further generally planar gradient coil arrangement (38, 30), for generating a second gradient magnetic field in a direction orthogonal to the main magnetic field, the further gradient coil arrangement having an outer perimeter and being disposed between the first pole piece and the examination region, the further gradient coil arrangement comprising an active coil (38) and a shield coil (30) situated in respective planes perpendicular to said magnetic field, the shield coil (30) being disposed between the active coil (38) and the first pole piece, and the active coil (38) and shield coil (30) being electrically connected at a plurality of locations along the outer perimeter of the gradient coil arrangement, the active coil and the shield coil of the further gradient coil arrangement further both comprising a respective first layer and a respective second layer, the respective first and second layers being mirror images of each other about an axis perpendicular to the magnetic field created by the pole piece magnets. An apparatus for use in magnetic resonance imaging as claimed in claim 2, characterized in that the active coil (38) of the further gradient coil arrangement is disposed between the active coil (40) of the first gradient coil arrangement and the first pole piece. An apparatus for use in magnetic resonance imaging as claimed in claim 3, characterized in that the shield coil (30) of the further gradient coil arrangement is disposed between the shield coil (32) of the first gradient coil arrangement and the first pole piece. An apparatus for use in magnetic resonance imaging as claimed in claim 4, characterized in that it further includes a third gradient coil arrangement (36, 34) for generating a third gradient magnetic field in a direction parallel to the main magnetic field created by the first and second pole pieces, the third gradient coil arrangement being disposed between the active coil (38) of the further gradient coil arrangement and the shield coil (32) of the first gradient coil arrangement. An apparatus for use in magnetic resonance imaging as claimed in claim 5 characterized in that , the third gradient coil arrangement comprises an active coil (36) and a shield coil (34) the shield coil (34) being disposed between the active coil (36) and the first pole piece, the active coil and the shield coil further both comprising a respective first layer (46, 60) and a respective second layer (48, 62), the respective first and second layers being mirror images of each other about an axis pependicular to the magnetic field created by the pole piece magnets. An apparatus for use in magnetic resonance imaging as claimed in any of claims 1, 2 or 6, characterized in that the respective first and second layers (90, 92, 74, 76, 46, 48, 60, 62) are planar conductors having slits extending therethrough, whereby eddy currents generated by magnetic flux having a component perpendicular to the plane of the gradient coil are reduced.

Inventors: SEMENTCHENKO MIKHAIL G, SEMENTCHENKO, MIKHAIL G.
Applicants: MARCONI MEDICAL SYSTEMS FINLAN, MARCONI MEDICAL SYSTEMS FINLAND, INC.
Application year: 1998
Main classification: G01R33
IPC codes: G01R33, A61B5