Abstract:
A dedicated RF coil for MRI comprising tuned primary and secondary circuits having primary and secondary inductors which are inductively coupled. The primary and secondary inductors have respective spatial regions of sufficiently uniform sensitivity for the MRI study to be carried out which substantially totally overlap. A port in the secondary circuit permits the transfer of RF energy between the tuned secondary circuit and an external device.

Description:
BACKGROUND OF THE INVENTION 
     The present invention relates to dedicated radio frequency coils for use in magnetic resonance imaging (MRI) of biological subjects, and more particularly tuned inductively coupled coils of this type. 
     The development of dedicated radio frequency (RF) coils has long been a topic of interest in the MRI field. A useful introduction to the subject can be found in the article by W. T. Sobel, &#34;Dedicated Coils in Magnetic Resonance Imaging&#34;, Reviews of Magnetic Resonance in Medicine, Vol. 1, No. 2, pp. 181-224 (1986). In his paper Sobel distinguishes between magnetic resonance imaging (MRI) and magnetic resonance spectroscopy (MRS). The same distinction is made in the present specification, in which references to imaging refer to two-dimensional Fourier transform imaging (which may be derived from spectral information) as opposed to spectroscopy which involves the reproduction of a nuclear magnetic resonance spectrum. 
     Several RF coil features and parameters directly affect its suitability for use in magnetic resonance imaging. Ideally, the coil would have a high uniform sensitivity within a particular spatial region of interest, and low sensitivity elsewhere, with a resulting high signal-to-noise ratio (SNR). The coil would be large enough to achieve the spatial coverage desired, but small to achieve a high fill factor, to fit within the MRI system magnet and to conform comfortably to the body of a subject. The RF coil must resonate at approximately the Larmor frequency of the nuclei used to develop the MRI signal, so that neither the coil size nor geometry can create an inductance or self-capacitance which prevents tuning to the desired frequency. The coil must also couple to the MRI system amplifier stages efficiently. 
     One technique for coupling the RF coil to the imaging system amplifier stage is inductive coupling. In this scheme a primary winding is positioned proximate the part of the subject which is to be imaged and a second winding, typically a single loop, is positioned adjacent the primary winding for inductively coupling with it. The secondary loop is coupled to the MRI system amplifier. Magnetic resonance signals are excited within the subject under magnetic field conditions to permit imaging and are received by the primary winding. Current flowing in the primary winding induces a voltage in the secondary winding which is amplified and processed to develop an image. 
     At this juncture there is no comprehensive analysis of inductively coupled RF coils for MRI. The importance of the degree of inductive coupling between the primary and secondary coils, their relative spatial positions and respective geometries, and how these factors fit into the other aspects of RF coil design mentioned above, remain largely unanswered. The various inductively coupled RF coils analyzed in the literature appear to be special cases of the general problem. 
     The article by W. Froncisz et. al., &#34;Inductive (Flux Linkage) Coupling to Local Coils in Magnetic Resonance Imaging and Spectroscopy, Journal of Magnetic Resonance 66, 135-143 (1986) presents an analysis of an inductively coupled coil for use in MRI. The secondary coil is an untuned single loop and its sole purpose is to couple the primary coil to a receiver. The article concludes, among other things, that detuning can be minimized by making the secondary coil with the smallest possible inductance and coupled as tightly as possible to the primary coil. 
     A loop array structure was proposed by C. Leussler et. al. &#34;Optimized RF Coils for Low Field MRI&#34;, Proc. SMRM 1989, p. 938, for applications where solenoids previously had found use. They disclose a head coil comprised of eleven turns and a body coil of eight turns, wherein each turn is a single loop LC resonator. Each resonator is tuned to the same frequency. The article presents data to show that the Q of the loop array is degraded less than the Q of the selonoid when loaded, over a frequency range of about 2.5 to 25 MHz. 
     Signal-to-noise ratio was considered in the article of D. J. Gilderdale et. al. &#34;The Performance of Mutually-Coupled Coils for Magnetic Resonance Signal Recovery&#34;, Proc. SMRM, p. 956 (1989). The authors concluded that in an inductively coupled coil system, the best SNR is obtained when the primary and secondary coils are slightly overcoupled and the lower frequency peak of the coil system frequency response in tuned to the frequency of interest. 
     The SNR of inductively coupled systems in which the secondary coil is significantly larger than the primary coil was investigated in the paper by S. N. Wright, &#34;Estimation of the SNR Loss Due to Inductive Coupling Loops&#34;. Proc. SMRM, p. 955 (1989). The paper concludes that there will be less than a five per cent drop in SNR of the coupled system relative to the primary alone, if the input resistance of the system is greater than ten times the primary coil resistance. 
     Finally, L. Darrasse, et al. &#34;Optimization of Receiver Coil Bandwidth by Inductive Coupling&#34;, Proc. SMRM, p. 1340 (1990) discloses a strongly overcoupled inductively coupled coil having a large bandwidth for fast scanning on a low field MRI system. 
     Significantly, the prior art frequently relies on an analysis of inductively coupled coils which is based on a lumped parameter circuit model. There is little if any consideration of the spatial sensitivity distribution of the primary coil, and usually there is a tacit assumption that the NMR signal is received by only the primary coil, and then inductively coupled to the secondary coil. Additionally, signal-to-noise ratio is evaluated in terms of the NMR signal voltage and the noise voltage which is output from the secondary coil; not the SNR of the image which is ultimately formed. However, the spatial sensitivity of the coil system can play a determinative role in the image SNR that is achieved. 
     SUMMARY OF THE INVENTION 
     It is an object of the invention to provide a dedicated inductively coupled RF coil for magnetic resonance imaging which achieves good signal-to-noise ratio in the resulting image over a desired field of view. 
     Another object of the invention is to provide a dedicated inductively coupled RF coil which can easily be resonated at the hydrogen Larmor frequency for mid-range magnetic field strengths. 
     Another object of the invention is to provide a dedicated inductively coupled RF coil for MRI which can be conveniently configured for a particular anatomical study of humans without substantially sacrificing image signal-to-noise ratio, field of view, and the ability to resonate at the hydrogen Larmor frequency for mid-range magnetic field strengths. 
     According to the invention a dedicated radio frequency coil for MRI is comprised of a tuned primary circuit and a tuned secondary circuit. Each of the tuned circuits is comprised of a capacitor and an inductor connected in series. The primary and secondary inductors are each dimensioned and shaped for defining a spatial region in which the coil sensitivity is sufficiently uniform for the MRI study which it is used to carry out. The primary and secondary inductors are positioned for inductively coupling the primary and secondary tuned circuits, and a port is provided for transferring radio frequency energy between the tuned secondary circuit and an external device. The primary and secondary inductors are dimensioned and shaped for defining respective spatial regions of sufficiently uniform sensitivity which substantially totally overlap for increasing the coil sensitivity relative to a single inductor within a particular field of view. 
     In one embodiment the primary inductor is a pair of spaced conductive loops, each tuned by a respective capacitor, and each inductively coupled to the secondary coil. Another primary structure is a pair of conductive loops disposed cylindrically symmetrically, and spaced from each other a distance approximately equal to the radius of the conductive loops, with each conductive loop comprising a single discontinuous loop having a gap between a pair of confronting ends. A pair of links are each connected between a respective end of one loop and the corresponding end of the other loop for connecting the pair of primary loops in parallel. The secondary circuit inductor is comprised of a pair of loops connected in parallel like the primary inductor. 
     In another preferred embodiment of the invention, the primary and secondary inductors are flexible and disposed on a flexible sheet of insulative material. The insulative sheet and flexible loop combination can be wrapped around a part of the anatomy of a subject conveniently and comfortably while obtaining the desired characteristics of the dedicated coil for imaging. 
    
    
     BRIEF DESCRIPTION OF THE DRAWING 
     The invention will be readily understood from the detailed description and the accompanying drawing in which: 
     FIG. 1 is an isometric view of an inductor used in dedicated RF coils according to the invention; 
     FIG. 2 is an isometric view of a dedicated RF coil according to the invention having primary and secondary inductors with the structure shown in FIG. 1; 
     FIGS. 3 and 4 illustrate another embodiment of the dedicated RF coil according to the invention comprising primary and secondary inductors having the structure shown in FIG. 1; 
     FIG. 5 is an isometric view of another embodiment of the dedicated RF coil according to the invention having a primary inductor comprised of separate loops and a secondary inductor having the structure shown in FIG. 1; 
     FIG. 6 is a plan view of a flexible dedicated RF coil according to the invention laid out flat; 
     FIG. 7 is a plan view of another embodiment of a dedicated RF coil according to the invention laid out flat; 
     FIG. 8 is an isometric view of the dedicated RF coil shown in FIG. 7 folded to form a loop; 
     FIG. 9 illustrates a fastener for use with the embodiments of the invention shown in FIGS. 6-8, and 
     FIG. 10 illustrates structure for adjusting the inductive coupling between the primary and secondary inductors of the invention. 
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
     FIG. 1 illustrates a basic building block of several different embodiments of the invention. The structure is comprised of a first loop 11 and a second loop 12, with the planes of the two loops parallel. The loop 11 is discontinuous with a pair of confronting ends 13, 14 and a gap between them. The loop 12 is likewise discontinuous with a gap between its confronting ends 15, 16. Respective ends of each of the loops 11 and 12 are connected together by conductive links 17 and 18. Thus, loop ends 14 and 16 are connected by the link 17, and loop ends 13 and 15 are connected together by the link 18. As a consequence of this structure, the loops 11 and 12 are connected electrically in parallel. 
     In a preferred embodiment the two loops 11, 12 are circular, and the spacing between them is equal to the loop radius. This is the same spacing as the known Helmholtz pair, however, in the present invention the spacing between the loops is permitted to differ somewhat from a distance equal to the loop radius. If the loop spacing differs from the ideal distance of a loop radius, then it is preferred that the loops be closer rather than further apart. 
     FIG. 2 illustrates a dedicated RF coil which incorporates the parallel connected double loop structure just described. The coil is comprised of a primary inductor having a first loop 21 and a secondary loop 22 spaced from it by a distance equal to about the loop radius. The primary inductor loops 21 and 22 are connected in parallel and a tuning capacitor 23 is connected in series with the parallel loop combination. A secondary inductor is similarly comprised of a first loop 24 and a second loop 25 which are positioned between and adjacent the primary inductor loops. More particularly, the secondary inductor loop 24 is adjacent the primary loop 21, and the secondary inductor loop 25 is adjacent the primary loop 22. As a consequence, the spacing between the secondary inductor loops 24 and 25 is also about equal to the loop radius. A capacitor 26 is connected in series with the secondary inductor loops 24, 25 so as to form a tuned secondary circuit. A port 27 shown in the form of a terminal pair is provided to permit the transfer of RF energy between the secondary circuit and an external device. The inductor comprised of loops 24 and 25 need not be the secondary inductor. Instead, the inductor comprised of loops 21 and 22 could have been the secondary inductor with the port 27 across the capacitor 23. Loops 24 and 25 would then comprise the primary inductor. 
     In FIGS. 3 and 4 the illustrated dedicated RF coil has exactly the same spacing between the primary inductor loops as the secondary inductor loops. Thus, loops 30 and 31 together comprise the primary inductor and loops 32 and 33 together comprise the secondary inductor. The distance between the loops 30 and 31 is equal to the distance to the distance between the loops 32 and 33, and ideally the interloop spacing for both inductors is equal to the loop radius. This structure is shown in longitudinal section in FIG. 8 which illustrates that the interloop spacing for both the primary inductor and the secondary inductor are equal. 
     Because the interloop spacing of both the primary and secondary inductor is equal, either loop pair can be used for the primary or secondary without changing the sensitivity profile of the coil. The fact that the loop pair 32 and 33 is shown as having an output port 34 is arbitrary. The output port could have been equivalently on the loop pair 30 and 31. 
     In the present invention the primary inductor loops are inductively coupled to the secondary inductor. There is no requirement that the primary inductor loops be connected, either directly or inductively, to each other. FIG. 5 illustrates an embodiment of a dedicated RF coil having secondary inductor loops 50 and 51 connected in parallel in the manner previously described. Also, like in the previously described embodiments the secondary inductor is tuned by a capacitor 52. There are two unconnected primary inductor loops 53 and 54. The primary loop 53 is adjacent the secondary loop 50, and the primary loop 54 is adjacent the secondary loop 51. Primary loop 53 is connected in series with a tuning capacitor 55, and primary loop 54 is connected in series with a tuning capacitor 56. The primary tuned circuit is thus comprised of a pair of mechanically separate and unconnected loops 53 and 54, both of which are inductively coupled to the secondary inductor 50, 51. As in the previously described embodiments the interloop spacing of the secondary inductor, or the primary inductors, is no greater than about the loop radius. 
     EXAMPLE 
     An embodiment of the invention like that shown in FIG. 5 was constructed with dimensions particularly well suited for imaging the human knee. The coil windings were made of copper tubing having a 0.25 inch outer diameter. All four coil loops were circular with an inner diameter of 6.0 inches. The primary windings were spaced 3.0 inches center-to-center and were insulated with a plastic sleeve of 0.020 inch wall thickness. The two secondary loops were in contact with the primary loops insulation sleeves and were spaced approximately 2.7 inches center-to-center. The links connecting the two secondary loops were made of the same copper tubing as the coil loops. The capacitors for the two primary loops and the secondary each had the same nominal value which corresponds to the frequency of operation. For a hydrogen Larmor frequency in the 12 megahertz region corresponding to a magnetic field of around 3000 gauss, capacitor values of 200 picofarads were used. Use of this embodiment of the invention resulted in images which exhibited a high signal-to-noise ratio and good contrast across a field of view as wide as the spacing between the primary inductor loops. 
     It is sometimes desirable to have available flexible dedicated RF coils. These are useful, for example, in making abdominal images of humans. 
     FIG. 6 illustrates a flexible dedicated RF coil laid out flat. Flexible insulated sheet 60 has a major surface upon which a closed outer conductive loop 61 is disposed. The loop 61 is comprised of a pair of parallel legs 62, 63 extending lengthwise of the loop, and a pair of shorter legs 64, 65 connecting the longer legs 62 and 63 at the ends of the loop. The coil further comprises an inner loop 71 having a pair of parallel longer legs 72 and 73 which extend adjacent and parallel to the longer legs 62 and 63 of the outer loop 61. The inner loop is completed by shorter legs 74 and 75 which connect the longer legs 72 and 73 at the ends of the inner loop 71. 
     The outer loop 61 has a pair of terminals 66 and 67 at opposite ends of the loop and electrically connected with the loop. The inner loop 71 similarly is provided with terminals 76 and 77 at its opposite ends. Typically, a second flexible insulative sheet (not shown) would overlie the structure shown in the drawing to improve the appearance of the coil and prevent shorting of the inner and outer loops. The terminals 66, 67 and 76, 77 would protrude from such a second sheet. 
     In use, the coil is wrapped around the part of the person to be imaged, such as the abdomen, so that the outer loop end legs 64 and 65 are adjacent. A capacitor is connected between the terminals 66 and 67 for tuning the inductor formed by the outer loop 61 and a second capacitor is connected between the terminals 76 and 77 for tuning the inductor formed by the inner loop 71. In this way an inductively coupled dedicated RF coil having a tuned primary circuit and a tuned secondary circuit, as in the previously described embodiments, is realized. 
     FIG. 7 illustrates another flexible dedicated radio frequency coil which has a structure corresponding to that shown in FIG. 3. A flexible insulative sheet 80 has a first loop 81 disposed on a major surface of the sheet 80. A second loop 82 (illustrated with dashed lines) is identical to the first loop 81 and is disposed on the opposite major surface of the sheet 80. The first loop 81 is provided with terminals 83 and 84, and the second loop 82 is likewise provided with terminals 85 and 86. Moreover, additional insulative sheets (not shown) are typically provided over the first and second major surfaces of the sheet 80 for covering and insulating the first and second loops 81 and 82. 
     The structure shown in FIG. 7 is wrapped so as to form the coil structure shown in FIG. 8, with a pair of tuning capacitors connected across the terminal pairs 83, 84 and 85, 86, respectively. Either terminal pair can be selected as a port for the coil. Structure for holding the coil together and mounting the capacitors is shown in FIG. 9. The structure includes an insulative block 90 having openings for the terminal pairs 83, 84, and 85, 86. Capacitors 91 and 92 are mounted on the block and across the terminal pairs, and a connecting cable 93 is connected for transferring energy between one of the two loops and an external device. 
     The dedicated RF coils according to the invention are used as receiver coils for magnetic resonance imaging in the conventional manner. A subject is positioned with the part of its anatomy to be imaged within the magnetic field of an MRI system. The RF receiver coil is placed surrounding the part of the anatomy to be imaged and connected to the MRI system receiver preamplifier and the MRI system is operated to generate the magnetic field gradients and RF pulses necessary to stimulate the emission of MRI signals. The RF receiver coil receives the emitted MRI signals which are processed by the MRI system to produce an image. 
     Another operating mode for the RF coils according to the invention is to use them for the RF transmitter coil of the MRI system as well as the RF receiver coil. The disclosed RF coils have good spatial uniformity of RF energy distribution and permit the application of RF energy just in the region of the anatomy which is to imaged. 
     One aspect of the imaging process which is necessary to good results is matching of the MRI system receiver to the RF receiver coil. Typically, the RF receiver coil is connected to the MRI system receiver preamplifier through a tuning network that contains a variable capacitance element such as a varactor or a variable capacitor. The value of the capacitance element is varied to compensate for loading of the receiver coil by the subject being imaged after the subject is placed with the receiver coil. 
     FIG. 10 illustrates an embodiment of the RF coil according to the invention which includes means for varying the inductive coupling between the primary and secondary inductors for compensating for coil loading by the subject being imaged. The secondary inductor 100 is comprised of a pair of loops connected electrically in parallel, and each secondary loop is adjacent to a respective one of the primary inductor loops 101 and 102. A lead pair 103 connects the secondary inductor 100 to the MRI system preamplifier. 
     Each of the primary inductor loops 101 and 102 has attached to it a corresponding one of the travelers 104 and 105 which are free to travel lengthwise along the guide 106. Both of the travelers 104, 105 have a threaded bore for receiving a threaded rod 107. The threaded rod 107 has threaded sections 108 and 109 which are threaded with an opposite pitch. A gear 110 is fixed to the threaded rod between the rod sections 108 and 109 and is driven to rotate by the motor 112 through a second gear 111. Because the threads of the rod sections 108 and 109 have an opposite pitch, rotation of the rod 107 will cause the traveler 104 and 105 to move in the direction of the length of the guide 106 either toward each other or away from each other, depending upon which direction the rod 107 is rotated. Consequently, the primary inductor loops 101, 102 will move closer or further from the secondary inductor 100, depending upon the direction of rotation of the rod 107. In this way, the inductive coupling between the primary and secondary inductors can be adjusted. 
     The embodiment of the invention shown in FIG. 10 is used in the same way as the previously described embodiments. Additionally, before collecting data for an image, MRI signal emission is evoked and the spacing between the primary and secondary inductors is adjusted to obtain maximum MRI signal strength. Then, data collection for imaging is carried out. 
     The preferred embodiments of the invention disclosed herein are exemplary and should not be considered exhaustive of the scope of the invention. For example, variations in the structure of the mechanism for adjusting the primary and secondary inductor spacing, or a different mechanism, can be used. Non-circular coil loops, and different loop configuration, can also be used. Accordingly, the scope of the invention should be defined by the following claims.