Abstract:
A quadrature coil suitable for use with an open frame MRI system provides crossing pairs of arrays of parallel conductor elements, respectively. Compact configuration is provided through use of an isolating circuit for incorporating parasitic capacitances at the resonance frequency of the coil into a blocking parallel resonance. Termination of the parallel conductor elements may be accomplished by equal impedance node connectors formed from branching pairs of conductors or a triangular least resistance connection form. RF shields are provided by pairs of conductive sheets containing eddy current reducing slots aligned with the parallel conductors elements of the coil.

Description:
BACKGROUND OF THE INVENTION 
     The present invention relates generally to magnetic resonance imaging (MRI) systems and, more specifically, to a design for a radio frequency coil for an open magnet MRI system. 
     Magnetic resonance imaging systems provide images of internal structures of the human body and the like by detecting radio signals produced by the precessing spins of the atomic nuclei of the structure when the structure is placed in a strong polarizing magnetic field. The nuclear spins are first excited into precession by a radio frequency (RF) stimulation pulse. Next the spins are isolated spatially by application of one or more gradient magnetic fields that cause their precession frequency to deviate from that provided by the polarizing magnetic field alone. The isolated resonance signals produced by these precessing nuclear spins are detected and processed according to techniques well known in the art to produce tomographic or volumetric images. 
     A single antenna may be used to transmit the stimulating RF pulse and to receive the weaker resonance signals from the precessing nuclei although often separate antennas are used for these two purposes. 
     In a prior art “closed” MRI system, a polarized gradient magnetic field is produced by a cylindrical, annular magnet having a bore for admitting a patient along the axis of the cylinder aligned with the magnetic field B 0 . Nuclei precession within the patient is induced by an RF field providing a magnetic vector in a plane perpendicular to the B 0  axis. 
     For certain procedures, particularly surgical procedures, an “open” MRI system may be desired in which the annular magnet of the closed MRI system is replaced by opposed magnetic pole faces providing therebetween a relatively unobstructed opening into which a patient may be placed while preserving greater access to the patient than in a closed MRI system. In the open MRI system, the B 0  field extends between the pole faces and the RF field is kept perpendicular to the B 0  field. 
     In open MRI systems, to avoid unduly restricting access to the patient through the opening between opposed magnetic pole faces, one or more arrays of parallel conductors positioned near the pole faces are used to provide the RF field. These conductors are energized in a manner that produces a net RF vector in the desired plane perpendicular to the B 0  axis. 
     While a single opposed pair of RF coils may be used for producing an oscillating RF field along a single line, preferably each such RF coil is matched to a second array having perpendicularly running conductor elements. For the RF stimulating pulse, the two matched RF coils are energized with signals having a 90 degree phase difference so as to create a rotating RF field. For reception of the resonance signal, signals detected at the crossing RF coils are combined with the appropriate 90 degree phase difference to produce a signal with superior signal-to-noise ratio. Coils providing for perpendicular reception or transmission patterns are known generally as “quadrature” coils. 
     A radio frequency shield may be placed between the RF coils and coils that produce the gradient magnetic field described above, so as to prevent signal from the gradient coils from interfering with reception of signals by the RF coils. Such radio frequency shields may be used as a return conductor path for an RF coil. 
     While open frame MRI systems provide greater access to a patient for surgical and other procedures than closed MRI systems, providing a high degree of homogeneity for the radio frequency and magnetic fields necessary for high quality imaging is still a challenge. In this regard, it is important that the pole faces be as close as possible to each other, and therefore that the RF coils and radio frequency shield be as close as possible to each other as well. Providing this homogenous RF reception and transmission field with a compact coil structure remains an important area of development. 
     BRIEF SUMMARY OF THE INVENTION 
     A number of improvements to the design of quadrature coils suitable for open frame MRI systems are set forth herein. 
     While it is not possible to produce the ideally desired perfectly uniform RF field between the pole faces, conductor patterns designed to approximate the geometry of uniform current sheets parallel to the magnet pole faces are herein used to achieve a high degree of approximation to the desired RF field over the central imaging region. 
     Although the conductor elements of each coil array of a quadrature coil will be perpendicular and therefore theoretically isolated, in fact there exists significant capacitive coupling between such elements, particularly when the elements are placed in close proximity as is desired in an open frame MRI system. A first feature of the invention is an isolation circuit canceling out this capacitive effect. 
     Conventional termination of the conductor elements of the arrays is unduly resistive and/or promotes unequal current flow through these elements, limiting homogeneity of the resulting field. Accordingly, a second feature of the invention is an improved termination for these conductor elements that provides greater and more equal current flow. Additionally, a series connection between the coil arrays ensures identically matching current flows through the upper and lower corresponding conductor elements. An effective RF shield is provided for such quadrature coils which accommodates both transmission of magnetic field gradients and reduction of interaction between the gradient coils and the RF coil. 
     Specifically, a quadrature RF coil for an open MRI system is provided. The MRI system includes a polarizing magnet with opposed pole faces for establishing a polarizing field axis. The coil includes a first conductor array having separate and substantially aligned conductor elements positioned along a first conductor axis and extending across the polarizing field axis between opposed common connection points. A second conductor array includes separated and substantially aligned conductor elements positioned along a second conductor axis extending across the polarizing field axis between opposed common connection points, and extending perpendicularly to the first conductor elements. A combiner/splitter electrically coupled to a connection point of each of the first and second conductor arrays joins them with a common signal line so that a signal path between the common signal line and the connection point of the first conductor array is substantially 90 degrees out of phase with a signal path between the common signal line and the connection point of the second conductor array. 
     An isolation circuit joins the connection points of each of the first and second conductor arrays to create between the first and second conductor arrays a blocking parallel resonance at the operating radio frequency. The isolation circuit may comprise an adjustable inductor for providing parallel resonance in combination with a parasitic capacitive coupling between the overlying conductors of the first and second conductor array. For flexibility in tuning this circuit, a fixed or variable capacitor may be added between the first and second conductor arrays so as to be coupled in parallel with the parasitic capacitance. 
     Thus the invention, in one embodiment, constitutes an extremely compact planar coil suitable for use in open MRI systems providing high signal-to-noise ratio and quadrature detection. Because an extremely low profile RF coil may be constructed if parasitic capacitance between the elements is overcome, insertion of the inductor to convert this parasitic capacitance into a blocking parallel resonant circuit at the RF frequency, effectively eliminates its effect at the frequencies of interest. 
     Ideally, the radio-frequency body coil would produce a perfectly uniform magnetic field with a direction perpendicular to the static magnetic field produced by the magnetic pole faces. The direction perpendicular to the pole faces is parallel to the static magnetic field and is taken as the direction of the z-axis in a Cartesian coordinate system. A uniform, infinite, y-directed sheet of current with surface current density λ y  does not produce any magnetic field in the y or z directions. The field in the x-direction is given by the expression 
     
       
           B   x =μ o λ y  for  z&gt;z   o   
       
     
     and 
     
       
           B   x =−μ o λ y  for  z&lt;z   o . 
       
     
     Therefore, two such current sheets with equal but oppositely directed current densities, one located at z=z o , slightly below the upper pole face, and the other at z=−z o , slightly above the lower pole face, will produce a magnetic field 
     
       
           B   x =2μ o λ y  for − z   o&lt;z&lt;z   o   
       
     
     and 
     
       
           B   x =0 for  z&lt;−z   o  or  z&gt;z   o . 
       
     
     In theoretical terms this idealized pair of current sheets is an optimized source for the radio-frequency field of an open MR scanner from two points of view: 
     (1) The field between the current sheets is completely uniform and independent of position. 
     (2) The current sheets provide no obstruction to the region of the gap between the pole faces (−z o &lt;z&lt;z o ). However, because of its infinite extent, a coil consisting of such a pair of current sheets is not a practical design for an MR scanner. Furthermore, a large area conducting sheet of metal such as copper would shield the imaging region from the fields of the switched gradient coils which are typically required in MR imaging and which are located in the space between the RF coils and the magnetic pole faces. In a preferred embodiment of the invention, practical coil designs are provided which approximate the desirable properties of the pair of infinite uniform current sheets as just described. 
     A pair of coils, each with its primary conducting elements located within a rectangular region near to and parallel with the magnet pole faces, can form a practical approximation to the ideal pair of current sheets. This region is taken to be of width W in the x-direction and length L in the y-direction. A number N of equally spaced conductor strips, each parallel to the y-axis and extending from y=−L/2 to y=L/2 and each carrying the same y-directed current, are placed within this rectangle and arranged symmetrically around, and parallel to, the y-axis. The same pattern, but with oppositely directed currents, is placed on the lower pole face. By increasing the number of strips so that the space between them becomes negligible and allowing W and L to become arbitrarily large, the magnetic field pattern of this coil pair approaches that of the ideal pair of conducting sheets discussed above. If N is odd, there will be a conducting strip on each coil at the x-location given by x o =0 and an additional (N−1)/2 pairs of conducting strips located at x o (n)=±n W/(N−1) for 1≦n≦(N−1)/2. If N is even there will be on each coil N/2 pairs of strips at x o (n)=±(n−1/2) W/(N−1) for 1≦n≦N/2. 
     Because the parallel sets of conductors just described do not form closed electric circuits, it is necessary to provide additional conducting elements whose purpose is not primarily to produce the magnetic field in the imaging region but, rather, to close the conducting circuits of each of these two conductor arrays. A number of alternative possibilities are available for completing the circuit paths and the most desirable means of doing this will depend on the particular imaging application and system design being utilized. If the current elements closing the path are located remotely from the region of imaging, the field in the imaging region will be substantially that of the linear conductor arrays. This field is described below. 
     Applying the Biot-Savart law to a single linear current element which extends in the y-direction from y=−L/2 to y=L/2 and is located at x=x o  and z=z o  leads to the following expressions for the magnetic field components at the field point (x,y,z).                B   x     =                    μ   o       4      π                         z   -     z   o             (     x   -     x   o       )     2     +       (     z   -     z   o       )     2                                    [           L   /   2     -   y             (       L   /   2     -   y     )     2     +       (     x   -     x   o       )     2     +       (     z   -     z   o       )     2           +         L   /   2     +   y             (       L   /   2     +   y     )     2     +       (     x   -     x   o       )     2     +       (     z   -     z   o       )     2             ]                   B   y     =              0                 B   z     =                    μ   o       4      π                         x   -     x   o             (     x   -     x   o       )     2     +       (     z   -     z   o       )     2                                      [           L   /   2     -   y             (       L   /   2     -   y     )     2     +       (     x   -     x   o       )     2     +       (     z   -     z   o       )     2           +         L   /   2     +   y             (       L   /   2     +   y     )     2     +       (     x   -     x   o       )     2     +       (     z   -     z   o       )     2             ]     .                                  
     A complete coil pair will contain N linear conductors at x o =x o (n) and z=z o  and N additional conductors at x o =x o (n) and z=−z o  where n runs from n=1 to n=N. The total field produced by the two linear arrays is then given by          B   x     =         ∑     n   =   1     N            B   x          (         x   o          (   n   )       ,     z   o       )         +       ∑     n   =   1     N            B   x          (         x   o          (   n   )       ,     -     z   o         )                     B   y     =   0             B   z     =         ∑     n   =   1     N            B   z          (         x   o          (   n   )       ,     z   o       )         +       ∑     n   =   1     N              B   z          (         x   o          (   n   )       ,     -     z   o         )       .                                
     For a single conducting element, if the length L of the conductor becomes very long compared to the quantities (x−x o ) and (z−z o ), then          B   x     →         μ   o       2      π                         z   -     z   o             (     x   -     x   o       )     2     +       (     z   -     z   o       )     2                     B   y     =   0             B   z     →         μ   o       2      π                           x   -     x   o             (     x   -     x   o       )     2     +       (     z   -     z   o       )     2         .                              
     At the center of the imaging volume (x,y,z)=(0,0,0) and the central field of an individual conducting strip is given by          B   x     =       -                  μ   o       4      π                           L                   z   o           (       x   o   2     +     z   o   2       )            (         L   2     4     +     x   o   2     +     z   o   2       )       1   /   2                       B   y     =   0             B   z     =         μ   o       4      π                           L                   x   o           (       x   o   2     +     z   o   2       )            (         L   2     4     +     x   o   2     +     z   o   2       )       1   /   2           .                              
     If N is even in the symmetry of the inventive coil, all current elements can be grouped in groups of four wires with positive currents at (x o ,z o ) and (−x o ,z o ) and negative currents at (x o ,−z o ) and (−x o ,−z o ). This group of four wires produces a central field given by          B   x     =       -                  μ   o     π                         L                   z   o           (       x   o   2     +     z   o   2       )            (         L   2     4     +     x   o   2     +     z   o   2       )       1   /   2                       B   y     =   0             B   z     =   0.                          
     The total central field is determined by summing over all of the groups of four wires that are present in the coil. If N is odd, there is an additional contribution from the pair of wires at (x o =0, z o ) and (x o =0,−z o ) which must also be added to the field of the other conductors. The central field of this wire pair is          B   x     =       -                  μ   o       2      π                           L                   z   o           (       z   o   2          (         L   2     4     +     z   o   2       )       )       1   /   2                     B   y     =   0             B   z     =   0.                          
     Therefore, this coil geometry, as desired, produces a magnetic field that is predominately in the x-direction near the center of the magnet gap. 
     The first and second conductor arrays may be comprised of copper foil laminated to opposite sides of a planar insulating substrate, as typified in conventional printed circuit technology, and the isolation circuit may be coupled to adjacent common connection points on opposite sides of the planar insulating substrate. This simplifies fabrication of extremely compact quadrature coils for open frame MRI systems. The printed circuit technology registers the first and second conductor arrays precisely with respect to each other and allows the isolation circuit to operate by connecting to adjacent coil ends through a small aperture in the insulating substrate. 
     The conductor elements of each conductor array may be connected together via a first node connection connecting the first ends of the conductor elements to a first node and a second node connection connecting the second ends of the conductor elements to a second node. The nodes may in turn be connected to an RF signal line for driving the conductor array or receiving signals from the conductor array. The first and second node connections may provide equal impedance paths between each of the ends and the respective nodes. This may be done by providing equal path links between each end and the respective node and, more particularly, by providing a set of separate equal length branches from a signal node, each branch branching again into a second set of separate equal length branches which ultimately connect to the ends of the conductor elements. The connection of the conductor elements of the arrays thus promotes equal current through each conductor element, simplifying construction of the resulting field and improving its homogeneity. 
     In a second preferred embodiment, the first and second node connections provide substantially non-overlapping straight line paths between the respective ends and the node. This may be realized by a substantially continuous isosceles triangular conductor having its node at the apex and the ends of the conductor elements distributed along the base of the isosceles triangular conductor. In this manner a lowest possible resistance connection between each of the conductor elements and the node is provided. 
     The coil of the invention may include an RF shield for a quadrature coil, the latter having a first conductor array and a second perpendicular conductor array. The RF shield provides a conductive surface interrupted by channels substantially aligned with the conductor elements of both the first and second conductor arrays. The channels of this RF shield prevent eddy current formation caused by excitation of gradient coil fields such as might interfere with the RF coil and/or reduce the power or affect the shape of the gradient coils. 
     The channels aligned with the conductor elements of the first conductor array may be on a first conductive sheet and the channels aligned with the conductor elements of the second conductor array may be on a second conductive sheet adjacent to the first conductive sheet. The channels may be bridged by capacitors sized to provide low admittance at the operational radio frequency. The RF shield is thus easily manufactured. 
     A quadrature coil set comprised of four crossing conductor arrays may be placed at the pole of the open frame MRI magnet, with a first and third conductor array being at opposite poles and having parallel conductor elements, and a second and fourth conductor array being at opposite poles and having parallel conductor elements perpendicular to the conductor elements of the first and third array. Interconnection leads may connect the first and third conductor arrays in series through their connection points and may connect the second and fourth conductor arrays and series through their connection points, thus promoting opposite current flows through the first and third conductor arrays and through the second and fourth conductor arrays. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 is a simplified view of the architecture of a conventional closed MRI system showing orientation of the polarizing field and the perpendicular RF field; 
     FIG. 2 is a simplified view, similar to that of FIG. 1, showing the architecture of an open MRI system and the relative positions of the polarizing and RF fields; 
     FIG. 3 is a cross sectional view along line  3 — 3  of FIG. 2 showing superposition of the fields of the conductor elements of the conductor arrays of FIG. 2 which provide a perpendicular RF field; 
     FIG. 4 is a view of four conductor arrays arranged in quadrature configuration and connected in series in accordance with a first embodiment of the invention; 
     FIG. 5 is a detailed fragmentary view of one of the quadrature coil sets of FIG. 4 showing use of a combiner/splitter to produce the quadrature phased excitation signals and to combine received signals in quadrature for improved signal to noise ratio; 
     FIG. 6 is a detailed fragmentary view of a node connector providing for equal length connections between the conductor elements of one conductor array and a node for receiving or transmitting signals from the conductor elements; 
     FIG. 7 is a view similar to that of FIG. 6 showing an alternative embodiment of the node connector with reduced resistance between the node and the conductor elements; and 
     FIG. 8 is an exploded perspective view of one quadrature coil set showing an RF shield suitable for use with the coil of the invention. 
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     FIG. 1 illustrates a prior art “closed” MRI system  10  wherein a polarized magnetic field B 0  is produced by a cylindrical, annular magnet  12  having a bore  14  for admitting a patient along the axis of the cylinder aligned with the magnetic field B 0 . Precession of nuclei within the patient in bore  14  is induced by an RF field providing a magnetic vector in a plane perpendicular to the B 0  axis. This RF field may be produced by a so called “bird cage” coil  16  having an electrically resonant antenna structure that produces a rotating magnetic field within the desired plane. Exemplary bird cage coils are taught in Hayes U.S. Pat. No. 4,692,705, issued Sep. 8, 1987, and U.S. Pat. No. 4,694,255, issued Sep. 15, 1987, and Edelstein et al. U.S. Pat. No. 4,680,548, issued Jul. 14, 1987, each of which patent is assigned to the instant assignee and hereby incorporated by reference. The term “coil” is used herein synonymously with the term “antenna” and does not require a coil shape. 
     FIG. 2 illustrates an “open” MRI system  18  which is useful for certain procedures, and particularly for surgical procedures. In this open MRI system, annular magnet  12  of the closed MRI system shown in FIG. 1 is replaced by opposed magnetic pole faces  20  providing therebetween a relatively unobstructed opening  22  into which a patient may be placed while preserving greater access to the patient. In the open MRI system, the B 0  field extends between pole faces  20  (vertically as depicted) and the RF field is likewise rotated to remain perpendicular to the B 0  field. While a bird cage coil, such as coil  16  of FIG. 1, could be used in this application, such coil is undesirable to the extent that it may unduly restrict the newly opened access to the patient. Therefore, in open MRI systems, one or more arrays of parallel conductors positioned near pole faces  20  may constitute an RF coil  24 . FIG. 3 illustrates individual conductor elements  26  of the opposed RF coils  24  energized with current flowing in opposite directions so as to produce a net RF vector in the desired plane perpendicular to the B 0  axis. 
     While a single opposed pair of RF coils may be used for producing an oscillating RF field along a single line, preferably each such RF coil  24  is matched to a second array having perpendicularly running conductor elements  26 . For the RF stimulating pulse, the two matched RF coils  24  are energized with signals having a 90 degree phase difference so as to create a rotating RF field as depicted generally in FIG.  2 . For reception of the resonance signal, signals detected at the crossing RF coils  24  are combined with the appropriate 90 degree phase difference to produce a signal with superior signal-to-noise ratio. Coils providing for perpendicular reception or transmission patterns are known generally as “quadrature” coils. 
     A radio frequency shield  28  may be placed between RF coils  24  and gradient coils  30 , as shown in FIG. 3, so as to prevent signal from the gradient coils from interfering with reception of signals by RF coils  24 . Each such radio frequency shield  28  may be used as a return conductor path for a respective RF coil. In order to maintain a high degree of homogeneity for the radio frequency and magnetic fields, pole faces  20  must be situated as close as possible to each other, and therefore RF coils  24  and radio frequency shield  28  should also be as close as possible to each other. 
     As shown in FIG. 4, an open MRI system in accordance with a preferred embodiment of the invention includes opposed RF coils  24   a  and  24   b  situated, respectively, at each pole face  20 , and each coil is composed of two conductor arrays. Thus RF coil  24   a  includes conductor arrays  32   a  and  32   b  while RF coil  24   b  includes conductor arrays  32   c  and  32   d . Each of conductor arrays  32   a - 32   d  is composed, respectively, of a planar set of substantially parallel conductor elements  34   a - 34   d  arranged to extend substantially perpendicularly to the polarizing axis B 0 . Conductor elements  34   a  and  34   c  of conductor arrays  32   a  and  32   c , respectively, are parallel to each other and perpendicular to conductor elements  34   b  and  34   d  of conductor arrays  32   b  and  32   d , respectively, so that the conductor arrays of each of RF coils  24   a  and  24   b  may produce or detect RF signals in quadrature along mutually perpendicular axes. 
     FIG. 5 shows each of the ends of conductor elements  34   a  being connected together by two independent node connectors  36   a  respectively, with one of the node connectors electrically joining first ends of conductor elements  34   a  together and the second of the node connectors electrically joining opposing second ends of conductor elements  34   a  together so that the conductor elements may be attached to nodes  38   a  and from there to signal lines  40   a . Similar connections are made for array  32   b  of coil  24   a  with similar reference numbers and a “b” suffix, and similar connections are also made for arrays  32   c  and  32   d  of coil  24   b  (not shown in FIG.  5 ). 
     Each node connector, such as connector  36   a , may provide for an equal impedance connection between its node  38   a  and each of the conductor elements, such as elements  34   a , by way of a branching structure as shown in FIG. 6, in which equal length constant width branches  42  extend from node  38   a  to secondary nodes  44  and branch again at secondary nodes  44  into secondary branches  46  also of equal length and width and hence equal impedance. Each of secondary branches  46  may proceed to tertiary nodes  48  to create tertiary branches  50  that ultimately communicate with ends of the conductor elements. Each of the primary, secondary and tertiary branches need only be of equal length within their rank (as primary, secondary and tertiary) so that any path between node  38   a  and an end of conductor element  34   a  has the same path length (and cross-sectional configuration) and hence the same impedance. In this way, current is evenly divided among the conductor elements and does not disproportionately flow through the center conductor element of the RF coil. This greatly simplifies construction of the coil, eliminating any need for variable width conductor elements or variation in the spacing of the conductor elements. These latter variations may be reserved for correcting for higher order errors rather than for fundamental differences in current flow. It will also be understood that this branching approach may be used for any number of conductor elements equal to a power of 2. 
     In an alternative embodiment, shown in FIG. 7, node connector  36   a  may be a solid sheet of conductor in an isosceles triangle pattern where the apex of the triangle is node  38   a  and the base connects to the ends of conductor elements  34   a . This creates a direct path in a straight line, and hence the shortest possible path, between node  38   a  and the ends of conductor elements  34   a . By allowing these paths to be non-overlapping, resistance between node  38   a  and each of conductor elements  34   a  is minimized. For this non-overlapping situation to occur, the height of the isosceles triangle measured from apex to base is substantially greater than three times the width of an average conductor element  34   a.    
     As shown in FIG. 4, opposed conductor arrays  32   a  and  32   c  having conductor elements  34  oriented in parallel, may be connected in series by interconnection leads  52   a  and  52   b  joining signal lines  40   a  and  40   c  so as to place conductor arrays  32   a  and  32   c  in series in a continuous loop, thus providing for countervailing currents as described with respect to FIG.  3 . An excitation signal  53   a  may be applied across a capacitor  54   a  positioned along this loop formed by interconnections  52   a  and  52   b , or alternatively, a resonance signal may be extracted across capacitor  54   a  during acquisition of the resonance signal. Similarly, conductor arrays  32   b  and  32   d  having parallel oriented conductor elements  34   b  and  34   d  may be interconnected by leads  52   c  and  52   d  and a signal  53   b  may be inserted across a series connected capacitor  54   b  or received across that capacitor in a manner similar to that described above. Conductor arrays  32   a  and  32   c , and  32   b  and  32   d , may alternatively be driven in parallel using properly phased signals, as known in the art, with the advantage of not requiring any direct cabling across the magnet gap. 
     Conductors  52   a - 52   d  may comprise coaxial cable and be routed so as to provide greater accessibility to the area between pole faces  20 . 
     In an alternative embodiment, capacitor  54   a  or  54   b  may be centered within one of conductor elements  34   a  and  34   b , or  34   c  and  34   d , respectively, near the edge of the conductor arrays. In this instance, the current conducted by the other conductor elements runs counter to the conductor element having the capacitor for introducing the signal. In this configuration, a separate voltage with 180° phase shift may be provided to the lower conductor arrays  32   c  and  32   d . This configuration has the disadvantage that the conductor element used for the connection across the capacitor carries current in the opposite direction to the rest of the conductor elements in the array, reducing the strength and uniformity of the generated RF field. 
     Signals  53   a  and  53   b  will generally have a 90° phase separation and so may be combined by a combiner/splitter  58 , shown in FIG. 5, to provide a single quadrature signal on a signal line  60  having an improved signal-to-noise ratio. Combiner/splitter  58  may be a hybrid circuit of a type well known in the art. Likewise, an excitation signal received from line  60  may be split by combiner/splitter  58  to excite the coils in quadrature for the reverse effect. 
     As shown in FIG. 8, RF coil  24   a  may be fabricated by producing conductor arrays  32   a  and  32   b  (or  32   c  and  32   d ) as etched copper traces on opposite sides of an insulating substrate  62  according to well-known printed circuit techniques. The conductor arrays thus may comprise copper foil laminated on substrate  62 . 
     In FIG. 5, the close proximity of two orthogonal conductor arrays  32   a  and  32   b  results in a parasitic capacitance  64  at each point where conductor elements  34   a  of conductor array  32   a  cross conductor elements  34   b  of conductor array  32   b . These parasitic capacitances impair the electrical isolation of conductor arrays  32   a  and  32   b , the condition of isolation being required for maximum signal-to-noise ratio in a quadrature coil. Accordingly, conductor arrays  32   a  and  32   b  may be re-isolated by connecting an isolation circuit  66  between conductor arrays  32   a  and  32   b , the isolation circuit having a impedance exactly sufficient to match parasitic capacitance  64 , in this instance through an inductance that creates with capacitance  64  a parallel resonant circuit having a peak impedance at the RF frequency matching the Larmor frequency for which the RF coil is intended. In the event that parasitic inductance (not shown) dominates, isolation circuit  66  may be an adjustable capacitor. Alternatively, isolation circuit  66  may provide a fixed inductance, which is easier to fabricate, and may be adjusted by a parallel or series connected adjustable capacitor to provide the same effect. 
     In the embodiment shown in FIG. 8, the fixed inductance of the isolation circuit may be connected between node connectors  36   a  and  36   b  of conductor arrays  32   a  and  32   b , respectively, through a small hole (not shown) cut in substrate  62 . Conductor arrays  32   a  and  32   b  are mounted to substrate  62 , which may comprise a polyester resin impregnated fiberglass board, and may then be attached to an outer surface of a planar support structure  67  providing sufficient rigidity to resist the force of a patient placed thereupon (for the lower coil) but having cut-out sections  68  to provide that strength with a reduced amount of material so as to maximize its dielectric constant. Positioned at one side opposite conductor arrays  32   a  and  32   b  are RF shields  70   a  and  70   b , fabricated as copper foil traces laminated to an insulating substrate  72  in much the same manner as conductor arrays  32   a  and  32   b  are attached to substrate  62 . RF shields  70   a  and  70   b  provide a conductive plane that prevents interference between gradient coils (shown as coils  30  in FIG. 3) and conductor arrays  32   a  and  32   b . While a continuous copper shield would provide such reduction in interference, it would promote conduction of eddy currents, distorting and sapping energy from the gradient coils. Accordingly, shields  70   a  and  70   b  include channels  73  of removed copper material extending parallel to conductor elements  34   a  of conductor array  32   a  for shield  70   a  and parallel to conductor elements  34   b  of conductor array  32   b  for shield  70   b  with the channels positioned in each instance to lie approximately midway between the respective conductor elements. Channels  73  are essentially parallel on each side of substrate  72 , but may converge at their ends if the shields are used for return paths per the parallel connection of the conductor arrays. 
     The RF shielding capability of shields  70   a  and  70   b  is increased by providing small capacitances  74  bridging channels  73  and adjusted so as to be an open circuit (or high impedance) at the frequencies associated with the gradient coils but a closed circuit (or low impedance) at the much higher frequencies associated with the RF signals. An equatorial channel  76  may be added, cutting perpendicularly midway across the channels  73  for further reduction of eddy currents. 
     The effect of eddy currents from gradient coils  30  on conductor arrays  32   a  and  32   b  may be further reduced by a second equatorial channel  69  severing each of conductor elements  34   a  and  34   b  midway along their length. The gaps are again bridged by capacitive elements  71  selected to be substantially open circuits at the frequencies associated with the gradient coils but closed circuits at the higher RF frequencies associated with RF coil Capacitive elements  71  provide a resonance with the coil at the Larmor frequency of the nuclear spins, turning each conductor element  34   a  and  34   b  into a half-wave resonance conductor. In this way, capacitive element  71  also limits build-up of free charge along the conductor surfaces, thus limiting undesirable effects of capacitive coupling between the conductor elements and any of the subject being imaged and other conductive surfaces within the MRI machine. 
     Increasing the thickness of support structure  67 , which is of a low dielectric material to eliminate dielectric losses, reduces interaction between RF coil  24   a  and the gradient coils, as does increasing the dielectric constant of support structure  67 . However, if the thickness of support structure  67  is too great, the overall coil structure will intrude upon the imaging volume. 
     Accordingly, it will be understood that an extremely compact coil may be provided for an open frame MRI system with the principal dimension being determined by the support structure  67  separating conductor arrays  32   a  and  32   b  from shields  70   a  and  70   b  on one side of the opening for patient access and by similar support structure separating a second pair of conductor arrays from a second pair of shields on the other side of the opening for patient access. 
     While only certain preferred features of the invention have been illustrated and described, many modifications and changes will occur to those skilled in the art. It is, therefore, to be understood that the appended claims are intended to cover all such modifications and changes as fall within the true spirit of the invention.