WO2007081805A2

WO2007081805A2 - System and method for high-resolution magnetic resonance imaging using inductively-over-coupled coils

Info

Publication number: WO2007081805A2
Application number: PCT/US2007/000281
Authority: WO
Inventors: Mehmet Bilgen
Original assignee: University Of Kansas Medical Center
Priority date: 2006-01-06
Filing date: 2007-01-08
Publication date: 2007-07-19
Also published as: WO2007081805A3

Abstract

The present invention is directed to a tunable electromagnetic device for use in magnetic resonance imaging comprising a primary element (20) electrically coupled to an electromagnetic field generator (32), and a secondary element (10) positioned proximate the primary element, such that the two are inductively coupled. The secondary element is non- tunable, and preferably provided with a coating (16) to allow the element to be implanted within tissue. The primary element comprises tuning and matching circuitry (34, 36, 38) to vary the resonance frequency between the two elements to provide an inductive over- coupling between the two to provide a high resolution magnetic resonance image. Also disclosed is an associated method of using the system.

Description

SYSTEM AND METHOD FOR HIGH-RESOLUTION

MAGNETIC RESONANCE IMAGING USING

INDUCTIVELY-OVER-COUPLED COILS

Field of the Invention The present invention relates generally to magnetic resonance imaging. More particularly, the invention relates to a system and method of using inductively over-coupled coils to obtain high-resolution magnetic resonance images. Description of Related Art

Inductively coupled radio frequency (RF) probes have long been used in magnetic resonance imaging (MRI) studies. These probes typically consist of two coils (a primary coil (PC) and a secondary coil (SC)) with no physical connection therebetween. Configurations with a primary matching coil coupled to a secondary volume coil are widely used in traditional applications. Typically, a circuit loop with tuning and matching elements is positioned centrally above the rung of self-resonating low-pass birdcage coil or directly over the window formed by the two rungs and the two end-ring segments of high-pass birdcage coil for mutual coupling. Thus, a matching surface coil with tuning and matching circuitry was used as the PC and coupled to a volume coil with fixed elements serving as the SC In other arrangements, the SC is configured as a surface coil, stent coil, wireless catheter coil or an implantable coil. In those known configurations, the resulting combined coil system provides a system for locally imaging the underlying tissue at increased signal-to- noise ratio (SNR) and spatial resolution. In those known configurations, the stent and wireless coils are used to amplify the excitation field generated by large body coil during transmission and couple the signal detected from the resulting magnetization to a surface coil during the receiving phase. Brief Summary of the Invention

The present invention is directed to a tunable electromagnetic device for use in magnetic resonance imaging comprising a primary element electrically coupled to an electromagnetic field generator, and a secondary element positioned proximate the primary element, such that the two are inductively coupled. The secondary element is non-tunable, and configured to be implanted within tissue. The primary element comprises tuning and matching circuitry to vary the resonance frequency between the two elements to provide an inductive over-coupling between the two to provide a high resolution magnetic resonance image. The system of the present invention uses the primary element for both transmitting and receiving, and thus does not require complicated electronics for switching or detuning purposes.

Using the adjustable tuning and matching capability of the primary coil, the two coils can be inductively overcoupled, allowing reliable and repeatable acquisitions of magnetic resonance data. This overcoupling is useful in experimental studies, such as those aimed at longitudinally imaging the spinal cord. Brief Description of the Drawings

The present invention will be described in greater detail in the following detailed description of the invention with reference to the accompanying drawings that form a part hereof, in which:

FlG. Ia is a plan view of an implantable secondary coil in accordance with an exemplary embodiment of the present invention.

FIG. Ib is a perspective view of a primary coil in accordance with an exemplary embodiment of the present invention. FIG. 2a is an equivalent circuit diagram of the primary and secondary coils of

FIGS. Ia and Ib.

FIG. 2b is a schematic diagram of an exemplary embodiment of the system of the present invention.

FIGS. 3a-3d are graphs of the frequency responses of the primary and secondary coil currents of a second exemplary embodiment of the present invention.

FIGS. 4a-4d are graphs of the frequency responses of the primary and secondary coil currents of a third exemplary embodiment of the present invention.

FIGS. 5a-5d are graphs of the frequency responses of the primary and secondary coil currents of a fourth exemplary embodiment of the present invention. FIGS. 6a-6d are graphs of the frequency responses of the primary and secondary coil currents of a fifth exemplary embodiment of the present invention.

FIGS. 7a-7d are graphs of the frequency responses of the primary and secondary coil currents of a sixth exemplary embodiment of the present invention.

FIGS. 8a-8d are graphs of the frequency responses of the primary and secondary coil currents of a seventh exemplary embodiment of the present invention.

FIGS. 9a-9d are graphs of the frequency responses of the primary and secondary coil currents of an eighth exemplary embodiment of the present invention. FIG. 10a is an axial in vivo spin-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention.

FIG. 10b is a coronal in vivo spin-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention. FIG. 10c is a sagittal in vivo spin-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention.

FIG. 11 a is an axial in vivo gradient-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention.

FIG. 1 Ib is a coronal in vivo gradient-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention.

FIG. 1 1 c is a sagittal in vivo gradient-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention.

FIG. 12a is an axial spin-echo image of a 15cc tube filled with 0.9% NaCl water solution acquired using a system according to an exemplary embodiment of the present invention.

FIG. 12b is a gradient-echo image of a 15cc tube filled with 0.9% NaCl water solution acquired using a system according to an exemplary embodiment of the present invention.

FIG. 13a is an axial in vivo spin-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention.

FIG. 13b is a coronal in vivo spin-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention.

FIG. 13c is a sagittal in vivo spin-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention. FIG. 14a is an axial in vivo spin-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention.

FIG. 14b is an axial in vivo spin-echo image of a spinal cord acquired using a system according to an exemplary embodiment of the present invention. Detailed Description of Exemplary Embodiment Looking first to FIGS. Ia and Ib, a system in accordance with an exemplary embodiment of the present invention includes an implantable secondary coil 10, and an external primary coil 20. As shown in FIG. Ia₅ implantable secondary coil 10 comprises two wires 12a, 12b formed generally into a loop, with capacitors 14a, 14b connected between ends of wires 12a, 12b to complete the loop. The entire secondary coil 10 is encapsulated within a coating 16, to protect the components of the coil and electrically isolate the coil from tissue when implanted. ^"Wires 12a, 12b may be any type of conductive wire known in the art. Preferably wires 12a, 12b are 18 gauge, copper wire, such as 18 gauge, CDA 102 wire available from MWS Wire Industries. Similarly, capacitors 14a, 14b may be any type of capacitor known in the art. Preferably they are ceramic chip capacitors, such as those available from American Technical Ceramics. Secondary coil 10 is preferably formed by soldering wires 12a, 12b to capacitors 14a, 14b, with the wires extending downwardly on each side of the capacitors so that secondary coil 10 is "A" or rooftop shaped when viewed from the end of the coil. Coating 16 may be any type of resilient, insulating coating. Preferably coating 16 is a biologically inert silicone elastomer, such as MDX4-4210 available from Factor II, Inc. Preferably, secondary coil 10 is small in size, allowing it to be implanted within animal tissue. Most preferably, secondary coil 10 measures approximately 25 millimeters by 13 millimeters. Of course, other types of wires, capacitors, and coatings may be used, and variations in the layout or configuration may be employed without deviating from the scope of the present invention.

Looking now to FIG. Ib, primary coil 20 comprises an outer case 22, with external terminals 28a, 28b located at one end of the case. Interior to the case, two wires 24a, 24b are connected at one end via a capacitor 26, with the opposite ends of wires 24a, 24b terminating at terminals 28a, 28b. Thus, the coil formed by wires 24a, 24b and capacitor 26 is contained completely with case 22. Preferably, primary coil 20 is formed as a 6 centimeter inner diameter high-pass quadrature birdcage coil with 16 elements. Of course, other configurations will be apparent to those skilled in the art and are within the scope of the present invention.

In use, secondary coil 10 is implanted within an area (i.e., tissue) in which it is desired to obtain a magnetic resonance image. Primary coil 20 is located external to the area, proximate secondary coil 10. With the coils thus proximately located, the primary and secondary coils are inductively coupled, as shown in the equivalent circuit diagram of the coils in FIG. 2a. Looking to FIG. 2a, the equivalent primary coil 30 includes resistive 34, inductive 36, and capacitive 38 components, and is connected to an electromagnetic field generator 32. Current flow through the coil is indicated by arrow 39. Similarly, the equivalent secondary coil 40 includes resistive 44, inductive 46, and capacitive components, with current flow through the coil indicated by arrow 49. The primary and secondary coils are in electrical communication through mutual inductance M between the coils, with M related to the coefficient of coupling k between the coils via the formula M -

. Looking to FIG. 2b, a schematic diagram of a system in accordance with an exemplary embodiment of the present invention comprises secondary coil 50, including wires 52a, 52b and capacitors 54a, 54b, located proximate primary coil 60, the primary coil including wires 64a, 64b and capacitor 60. As shown in FIG. 2b, wires 64a, 64b of primary coil 60 connect to an electromagnetic field generator. In an exemplary embodiment of the present invention designed to provide an imaging system having a 9.4 Tesla (T) magnetic field strength, secondary coil 50 is formed with two 8.5 pF ceramic chip capacitors 54a, 54b soldered to two segments of 18 gauge wires 52a, 52b to form a pitched rooftop-shaped rectangular loop with dimensions of approximately 20 mm X 15 mm. The primary coil is rectangular in shape with dimensions of approximately 25 mm X 40 mm, with a 2 pF ceramic chip capacitor 66 attached between wires 64a, 64b, and the other ends of the wires attached to an electromagnetic field generator.

Theory of Operation and Use

The simplified electrical circuit in FIG. 2a represents the primary coil (PC) 30 and secondary coil (SC) 40 coupled inductively using equivalent lumped-elements for the resistive 34, 33, inductive 36, 46, and capacitive 38, 48 components of each coil. Each coil in the figure consists of the resistive R, inductive L and capacitive C elements connected in series. The notations R, L, and C will be used in the formulas herein to denote the resistive, inductive, and capacitive components, with the subscripts p and s used to denote the primary and secondary coils, respectively. The mutual inductance M between the coils is defined by the coefficient of coupling k and the inductive elements according to the formula M = kJL L_s . Each coil is a self-resonant series circuit tuned at an angular frequency ω_Ql =

for i=p or s. The PC is connected to a voltage source v, represented by electromagnetic field generator 32 in FIG. 2a. After applying circuit analysis, the PC and SC currents i_p and /_Λ. in this arrangement can be written as a function of frequency /via the relation ω — 2π/"as

(2) where j is an imaginary unit. To analyze the frequency response characteristics of these currents and the resulting magnetic fields during transmission, we consider, without loss of generality, that the coil system, i.e., the combined coil set, is to be used for operations at 400 MHz, corresponding to the magnetic field strength of 9.4 T, the SC is loaded and|v| = 1.

Also, we scale the currents and conveniently study

(both have units of Ampere X Henry) as a function of the parameters k, ω, Q₁- and ωo/ for i=p or s.

We first analyze the coil system when the PC and SC are tuned individually to a different self-resonant frequency, but

where both coils are formed by fixed capacitive and inductive elements. The SC is tuned to ./_OJ,=4OO MHz₅ but the PC is tuned to a much higher off-resonance frequency. By assuming that both coils have the same quality factor, the frequency response of the currents J_p and I_s are plotted for frequencies between 350 and 450 MHz in FIG. 3 for five different coupling values. (FIO. 3 presents frequency response curves (amplitude and phase) of the primary (a,c) and secondary (b,d) coil currents for five different values of coupling coefficient AH).1,0.3,0.5,0.7 and 0.9, when Q_p=Q_s=50, fos^~ 400 MHz and fn_p = 1000 MHz. Arrows indicate the direction of increased k values.) For weak coupling A=O.1, current is induced in the SC and the response curve of this current peaks at its self-resonance frequency of 400 MHz. When the coupling is increased the position of this peak shifts towards lower frequencies with increasing amplitude. The PC curves exhibit peaks that are aligned with that of the SC curves. The phase difference between the PC and SC currents is initially about 90° at the peak frequency of the SC current and increases with coupling since the unwrapped phase of the SC current increases while the phase of the PC current decreases. The coil system becomes purely resistive at frequencies where the phase of the PC current attains zero. According to the phase data presented in FIG. 3 for the PC₅ not every coupling produces purely resistive input impedance. If the degree of coupling is high enough, then the system becomes resistive at two different frequencies where the phase of the PC current attains zero.

In practice, we desire adequate coupling to optimally match the impedance. Under the conditions of strong coupling, the curves in FIG. 3 indicate that the current induced in the SC attains higher values than that of the PC. Since currents in both coils collectively determine the pattern of the excitation field (Bi)₅ the coil with larger current dominates this pattern. When Q_s is improved, the PC and SC current response curves can be shown to become sharper, while the behaviors described above remain unchanged.

Combining these observations, the sensitive region of the probe, as defined by the tissue excitation profile and signal reception, under the arrangement in FIG. 2b can be seen as mostly determined by the features of the SC. Typically, the SC is tuned to a higher frequency to start with, so that the combined coil system would yield the SC to resonate at 400 MHz after the impedance matching. When the SC is implanted, it cannot be accessed physically unless a new surgery is performed. This makes it impractical to correct any deviation from the desired resonance frequency if the same PC in FIG. 2b is used as the external coil. However, coupling the implanted coil to an external coil with tuning and matching circuitry can offer retuning ability for the coil system, as explained below.

Next, we consider when both the PC and the SC are tuned to same resonance, i-e., fo_p= f Os =400 MHz and possess the same quality factor Q_p~Q_s ^~50. The frequency response characteristics of the probe for this case have been described in detail. Under this configuration, the resulting response curves for the currents in the coupled coils depend very largely on the coupling, as shown in FIG. 4. (FIG, 4 presents frequency response curves of the primary (a,c) and secondary (b,d) currents for 7 different values of coupling coefficient /c=0.005, 0.0125, 0.02, 0.0275, 0.035, 0.0425 and 0.05 when Q_P=5Q, Q_s=50, andf_Os=fo_p =400 MHz. Arrows indicate the direction of increased k values.)

When coupling is weak k~0, using Eq. (1), we can show that I_p asymptotically

current approximates the series resonance curve of the primary circuit considered alone. But, from Eq. (2), I_s approaches to -

(3) indicating that the response curve of the SC current is governed by a shape approximating the product of the resonance curves of the PC and SC circuits taken alone. Other important noticeable features of the graphs in FlG. 4 are that as the coupling is increased, the curve of the PC current becomes broader and its peak value is reduced. At the same time, the curve of the SC current becomes larger and its sharpness is reduced. These trends continue until the coupling reaches a critical value where the peak of the SC current reaches a maximum. By this time, the peak of the PC current has already separated into two peaks. Beyond this critical coupling, the peak of the SC splits. With greater coupling, the double humps on the curves of both PC and SC currents separate farther apart, but their amplitudes and spread remain nearly the same. Under the condition of overcoupling (for our purposes, we define overcoupling as the condition beyond the break point where the SC current peak splits), this behavior in the peaks still remain unchanged, even when the quality factors of the coils are different, i.e., Qp≠Q_s.

Under the condition of weak coupling, it is possible to shift the resonance peak of the SC by manipulating the tuning properties of the PC, as shown in FIG. 5. (FIG. 5 presents frequency response curves of the primary (a,c) and secondary (b,d) currents under the weak coupling &=0.005 while φ=50, Q_Λ=50, and /_fo =400 MHz and /_0/J =390 (dashed line), 400 (solid line) and 410 ( dotted line) MHz.) However, this reduces the magnitude of the SC current (and hence the Bi field that it produces) while broadening its peak. For these reasons, weak coupling may not be the preferred choice for tuning this coil system. Nevertheless, Hoult and Tomanek considered two circular coils (one for matching and the other acting as a surface coil) arranged coplanar, and their asymptotical analysis on the coil currents predicted 90° phase difference between the currents. According to curves in FIGS. 5c and 5d, the phases of the primary and secondary currents at 400 MHz respectively read 0° and 90° (with difference of 90°) under the condition of weak coupling and the coils tuned to the same resonance. This phase behavior agrees with the prediction by Hoult and Tomanek. The magnetic flux lines produced by these currents also exhibit quadrature phase relationship at resonance. When the flux lines are not parallel geometrically, elliptically polarized, instead of circularly polarized, field with spatially varying properties is created. This ultimately leads to asymmetry in the excitation field.

Strong coupling between the coil components is, in general, an undesired feature in coil design, and is should be avoided whenever possible, as in the case of array coils for parallel imaging. However, in our application, peak-splitting resulting from strong coupling offer an opportunity to externally tune the probe when the SC contains fixed circuit elements. As shown below, either of the two new peaks after the split can be employed for this purpose. One possibility is to initially tune the PC and SC to a frequency higher than the operation frequency of 400 MHz independently, i.e. while the coils are uncoupled Λ=0. Coupling the coils split the peak of the SC. By manipulating the degree of coupling, the location of the first hump in the frequency axis can be tuned to 400 MHz. If the second hump is desired for the operation, the coils are initially tuned to lower frequency to start with, and then this hump in the SC current is tuned to 400 MHz by changing the coupling strength. Although, these two are equally valid approaches that grant tuning ability to the combined coil system, they still lack taking full advantage of the matching and tuning abilities that may be available in the PC.

We now consider cases where the PC is tuned at frequencies different than the SC, and demonstrate how this provides additional tuning ability for the combined coils. First, we select the first peak after the split to be the resonance peak, and would like to shift the position of this peak by manipulating fi_p, which can easily be accomplished using the tuning and matching scheme available in the PC. This goal can be achieved under two conditions on the free-ringing properties of the coils:_./^^ orf_0s<fo_p. As shown in FIG. 6, for the case fo<>fop, coupling the SC tuned independently tofo_s = 430 MHz to the PC tuned independently tofop = 415 results in the first peak centered at 400 MHz. (FIG. 6 presents frequency response curves of the primary (a,c) and secondary (b,d) currents when k=QΛ, Q_p=50, Q_s=50, and f_Os =430 MHz and/φ =405 (dashed line), 415 (solid line) and 425 (dotted line) MHz.) Changing the tuning properties of the PC from 405 to 425 MHz provide a tuning range for the first peak of the SC between the frequencies 390 and 410 MHz. Or as in FIG. 7, representing the case fo^ fop_t coupling the SC tuned to f_Os = 4\O MHz to the PC tuned to f_Op =44O MHz again centers the first peak at 400 MHz. (FIG. 7 presents frequency response curves of the primary (a,c) and secondary (b,d) currents when

,/φ=430 (dashed line), 440 (solid line) and 450 (dotted line) MHz.) Retuning the PC from 430 to 450 MHz shifts the first peak of the SC between the frequencies 395 and 405 MHz. Comparing the curves in FIGS. 6 and 7, the arrangement with fo_s> fo_p can be seen as providing increased tuning range of 20 MHz as compared to 10 MHz achieved with the arrangement fo_s<fop- On the other hand, the current in the PC with the former arrangement reads about four times greater (~ 0.08 versus 0.02) while the SC current is about 50. This difference in the PC current may have significant implications in the imaging performance. It is also important to note that the phase differences between the PC and SC currents around 400 MHz in Figs. 6 and 7 are about 180°, i.e. out of phase, implying that the flux lines produced by these coils are opposing. Therefore, choosing the first peak as resonance yield the coils to produce Bi fields that are destructive.

If the second peak is selected as the resonance peak of the combined coil system, manipulatingybp, using approaches similar to those described above, can also shift the position of this peak. FIG. 8 demonstrates the second peak tuning between 390 and 410 MHz is varied from 375 to 395 MHz while jo.y-370 MHZ </Q_P. (FIG. 8 presents frequency response curves of the primary (a,c) and secondary (b,c) currents when fc=0.1, Q_p=50, Q_s ^~50, and /o_/=37O MHz and /_0/?=375 (dashed line), 385 (solid line) and 395 (dotted line) MHz.) FIG. 9 shows that the second peak tunes between 395 and 405 MHz when Jop is varied from 350 to 370 MHz while^=390 MHz >fc_p. (FIG. 9 presents frequency response curves of the primary (a,c) and secondary (b,d) currents when k=0Λ, Q_p=50, Q_s=50, and ^=390 MHz and fo_p=35O (dashed line), 360 (solid line) and 370 (dotted line) MHz.) The former approach in this arrangement provides a greater tuning range, but it results in about 4-fold increase in the PC current. From FIGS. 8 and 9, the PC and SC currents around 400 MHz can be seen to be in phase (phase difference ~ 0°) implying that the Bi fields produced by these currents are in the same direction and therefore constructive. Test Results

From the information provided within the theoretical computations and analysis, the tuning and matching enterprise of the inductively coupled coils of the present invention can be seen as diverse in nature. In the following, we select one of the over- coupled configurations analyzed in FIGS 6-9 and apply it into practice to image a rat spinal cord in vivo. The tuning and matching range available in our volume coil was the main criterion on the selection of the specific coil configuration. This resulted in the coil configuration with the frequency response characteristics described in FIG. 8. In the tests, the implantable secondary coil (SC) was built from two 8.5 pF ceramic chip capacitors (American Technical Ceramics, Huntington Station, NY) soldered to two pieces of circular wire (18 gauge, CDA 102) (MWS Wire Industries, Westlate Village, CA) to form a rooftop-shaped rectangular loop with dimensions 25 mm X 13 mm. The SC was coated with biologically inert silicone elastomer MDX4-4210 (Factor II, Inc., Lakeside, AZ) for electrical isolation from its surroundings. The primary coil (PC) was a 6 cm inner diameter high pass quadrature birdcage coil with 16 elements, provided by the manufacturer of the scanner. Only one channel of this coil was used, the other unused channel was terminated with a 50 ohm resistor. It is important to distinguish this coil setup from those employed in previous studies where a matching surface coil with tuning and matching circuitry was used as the PC and coupled to a volume coil with fixed elements serving as the SC. In our application, the roles and configurations of the coils are reversed. Also, in previous studies, stent and wireless coils were used to amplify the excitation field generated by large body coil during transmission and couple the signal detected from the resulting magnetization to a surface coil during the receiving phase. In contrast to these coil setups, our coil system is simpler in design since the PC is used for both transmit and receive, and does not require complicated electronics required for switching or detuning purposes.

The SC was implanted subcutaneously in a Sprague Dawley rat adjacent to the spinal cord at the thoracic level by following the procedures described earlier. The rat was maintained under isoflurane anesthesia delivered through a nose mask and monitored using an MR-compatible small animal monitoring and gating system (Model 1025, SA Instruments, Inc., Stony Brook, NY). This system was also used for respiratory- gated acquisition to minimize the breathing-related image artifacts. The temperature of the rat was kept at 37 ⁰C by circulating warm air with 40 % humidity using a 5 cm diameter plastic tubing fitted at the back door of the magnet bore.

The resonance frequency of the implanted SC was measured using an external rectangular loop attached to a frequency sweeper (Morris Instruments, Inc., Ottawa, Ontario, Canada). For imaging, the rat was placed supine on a Plexiglas tube that was cut half along the long axis, and the tube was inserted into the volume coil such that the implanted coil stayed at its center. Next, to improve the coupling, the volume coil was rotated slightly with respect to the tube until two peaks appeared near the proton resonant frequency of 400 MHz on the sweeper's display. The tuning and matching rods of the volume coil were then engaged to further improve the impedance matching and frequency tuning properties of the second peak observed on the display.

MRI was performed on a 9.4 T horizontal bore scanner (Varian Inc., Palo Alto, CA) using 12 cm ID gradient coil. The Plexiglas sled, supporting the animal and the volume coil, was inserted into the scanner bore. Scout images were first acquired to confirm the placement of the rat at the magnet isocenter. Next, the transmit power of the volume coil was optimized using standard spin echo (SE) sequence to position the 90° excitation band on the spinal cord. Gradient echo (GE) and SE sequences were then employed to demonstrate the excitation field of the combined volume and implanted coils in large field-of-view (FOV) selected in axial, coronal and sagittal planes. The acquisition parameters for the SE data were T_R/T_E = 2500 ms/10 ms, image matrix = 128 X 256, slice thickness = 1 mm and NEX=2, FOV=45 mm X 45 mm for the axial, 35mm X 85 mm for the coronal and 45 mm X 85 mm for the sagittal views. The parameters for the GE data were T_R/TE — 40/3 ms, flip angle=45°, image matrix = 128 X 128, slice thickness = 2 mm and NEX=2, FOV=45 mm X 45 mm for the axial, 35mm X 85 mm for the coronal and 45 mm X 85 mm for the sagittal views. Also, high-resolution SE images were acquired in the same planes but in smaller FOV using TR/TB = 2500 ms/10 ms, image matrix = 128 X 256, slice thickness = 1 mm and NEX=2, FOV=I 5 mm X 20 mm for the axial, 24 mm X 33 mm for the coronal and 15 mm X 33 mm for the sagittal views. These scans were repeated on days 7 and 14 of the coil implantation to show the repeatability of the scans at the same level of imaging quality. These studies were performed under a protocol approved by the institutional animal care and use committee at the University of Kansas Medical Center. Additional scans were performed on a uniform phantom, consisting of a 15 cc plastic tube filled with 0.9% NaCl water solution, by placing the SC on the surface of the tube. The implantable SC exhibited self-resonance at 407 MHz in bench tests in air

(i.e., unloaded) measured with the coil configuration shown in FIG. 2b. The resonance peak was sharp with quality factor of about 150. When implanted into the rat, the resonance frequency of the loaded coil shifted down to 388 MHz and its quality factor dropped to about 30. The volume coil was initially tuned to 400 MHz when it is unloaded. After inserting the rat into the volume coil, a single peak was observed on the frequency response curve, which indicated weak coupling between the implantable SC and the volume coil. Rotating the volume coil with respect to the rat increased the coupling and produced double peaks. The second peak was then tuned and matched at 400 MHz by varying the matching and tuning capacitors of the volume coil. At this time, the first peak was positioned at 379 MHz, was broader and had lower amplitude as compared to the second peak. These resonance properties closely resembled the features of the theoretical case presented in FIG. 8. After placing the rat into the scanner's bore, the peaks on the response curve shifted about 1 MHz to lower frequencies, and were readjusted again Finely.

For imaging the spinal cord properly with the setup described above, the transmit power was optimized at 23 dB when the 90° rf pulse was 2 ms long. This power level was comparable to the power used when the rectangular loop in FIG. 2b was employed as the external coil. If the volume coil were used alone in quadrature mode to image the rat body, the optimal power for the same pulse would typically be achieved with much higher 42 dB power in our scanner.

As discussed previously, the currents in both coils were expected to produce excitation fields in the rat's body during transmission. The images in FIGS. 10 and 11, obtained specifically with large FOV, delineate tissue not only in the footprint of the implanted coil but also in the background regions. (FIG. 10 presents in vivo spin-echo images of spinal cord in (a) axial (b) coronal and (c) sagittal planes acquired on the day of coil implantation using the parameters 7_R/T_E = 2500 ms/10 ms, image matrix = 128 X 256, slice thickness = 1 mm and NEX=2, FOV=45 mm X 45 mm for the axial, 35mm X 85 mm for the coronal and 45 mm X 85 mm for the sagittal views. The intensities in the images were windowed and scaled to enhance the background signal. Notice that, during transmission, the volume coil generates an excitation field that leads to the hypointensity (long arrow) seen in the background. The implanted coil, indicated by IC on the images, generates 90° excitation field that produces a band of hyperintensity (short arrows) surrounding it. The square boxes are used to localize the regions where the images in FIG. 14 below were acquired.) (FIG. 11 presents in vivo gradient echo images of spinal cord in (a) axial, (b) coronal and (c) sagittal planes acquired on the day of coil implantation using the parameters T_R/ T_E ^~ 40 ms/3 ms, flip angle=45°, image matrix = 128 X 128, slice thickness = 2 mm and NEX=2, FOV=45 mm X 45 mm for the axial, 35mm X 85 mm for the coronal and 45 mm X 85 mm for the sagittal views. The intensities in the images were windowed and scaled to enhance the background signal.) Notice that, during transmission, the volume coil generates an excitation field that leads to the hypointensity (long arrows) seen in the background. The implanted coil generates excitation field that produces the hyperintense footprint (short arrows).) This behavior can better be understood from the images shown in FIG. 12, which were produced during studies with the uniform phantom. (FlG. 12 presents axial spin- echo (a) and gradient-echo (b) images of a 15 cc tube filled with 0.9% NaCl water solution. IC denotes implantable coil. The circles and squares respectively denote the regions of interest selected to measure the mean signal intensities in the excitation field of the implanted coil and the background regions excited by volume coil. The intensities in the images were windowed and scaled to enhance the background signal.) The homogeneous weak signal in the background is induced as a result of the excitation by the volume coil. But, the inhomogeneous field seen near the site of the implanted coil is formed by contributions from both coils, but mainly by the implanted coil. We selected two regions of interest; one within the 90° field of the implanted coil and the other in the background field produced by the volume coil, as indicated in FIG. 12, and measured the mean signal levels in these zones. The ratios of the means were 18 on the SE image and 6 on the GE image.

The images in FlG. 13 were acquired with smaller FOV to visualize the anatomy and structure of the rat spinal cord in greater detail. (FIG. 13 presents in vivo high resolution spin-echo images of spinal cord in (a) axial (b) coronal and (c) sagittal planes acquired on the day of coil implantation using the parameters TR/TE = 2500 ms/10 ms, image matrix = 128 X 256, slice thickness = 1 mm and NEX=2, FOV=IS mm X 20 mm for the axial, 24 mm X 33 mm for the coronal and 15 mm X 33 mm for the sagittal views. Arrows on the axial image point to wraparound artifacts.) These images clearly demonstrate the capability of the coil system to acquire adequate quality of data to accomplish the high- resolution imaging task while maintaining good contrast between gray matter and white matter. The data inherently contain wraparound signals from the tissue excited by the volume coil outside the selected FOV. As discussed above, these signals are significantly low in intensity when the SE acquisition is employed, and subsequently produce negligible artifacts. But, wraparound artifacts produced by the signals from the tissue near the implanted coil remain strong and are visible on the images as fold over. If the FOV is selected large enough, these artifacts move towards the edges, allowing the spinal cord be visualized at the center of the image in a nearly artifact free fashion. On days 7 and 14 of the coil implantation, the self-resonance peak of the implanted coil respectively read 385 MHz and 381 MHz, which were lower than 388 MHz measured on the day of implantation. When the rat was placed in the volume coil outside the magnet, double peaks were observed. The second peak was tuned and matched at 400 MHz for imaging. FIG. 14 shows axial images of the spinal cord on day 14, acquired with the same parameter values used to produce the images in FIGS. 10a and 13-a. (FIG. 14 presents in vivo axial spin-echo images of spinal cord acquired on day 14 of the coil implantation in (a) larger view of 45 mm X 45 mm and (b) smaller view of 15 mm X 20 mm using the parameters 7_R/7Ε = .2500 ms/10 ms, image matrix = 128 X 256, slice thickness = 1 mm and NEX=2. The image intensity in (a) was windowed and scaled to enhance the background signal.) Qualitatively, the axial images acquired on both days can be seen as comparable, demonstrating the capability of the coil system to provide consistent data from the same subject scanned two weeks apart even when the tuning frequency of the implanted coil was shifted. If such frequency shift were to occur while using the coil setup in FIG. 2b, the imaging performance would be compromised.

Besides local field uniformity and high SNR, performance, practicality, reliability and repeatability are the key properties of the rf probes, that are highly-desired in longitudinal studies. Using the above coil system, the implantable coil produces a narrow strip of 90° excitation field that is sufficiently wide enough to uniformly image the spinal cord. The small footprint of the excitation allows high resolution imaging of the local tissue. But the resulting images inherently contain wraparound artifacts from the field generated by the volume coil. The sequence and the parameters used for imaging influence the severity of these artifacts. Our data from FIG. 13 indicate that the SE sequence produces relatively little artifacts while maintaining the sensitivity and specificity of the coil intact. A less obvious drawback of the coil system is that the large size volume coil produces more noise in the received signal as compared to a small pickup coil. Also, the effective bandwidth of the coil system increases with overcoupling. Noise (produced externally or thermally) that would be filtered otherwise can therefore leak though the frequency band at the second peak not used for the operation. However, the benefits of the over-coupled coil system overweight its shortcomings.

The results from the systemic investigations presented above demonstrate that the system satisfies the properties expected from rf coils. The system performs well in acquiring high quality in vivo MR data longitudinally. Readily available standard volume coils enhance the system by providing tuning and matching capabilities. This provides practicality by minimizing the time spent in preparing the animal for the scans, reliability and repeatability by ensuring that the data will be acquired continuously in longitudinal studies. In other aspects, additional benefits of using this coil system include simplicity, low-cost and flexibility. It does not require active/passive detuning of the SC during transmission or coil combinations that involve coupling the SC to a surface coil or a like during reception. Also, manufacturing the implantable coil in different shapes and sizes by properly cutting and forming the wires and soldering the capacitors allows local imaging of internal structures when the coil is implanted or placed on the surface. Given the benefits of this type of coil system, it is likely to prove to be highly useful in gathering longitudinal MR imaging data in experimental studies.

As can be seen, the invention described herein provides a system and method for obtaining high-resolution magnetic resonance images using inductively over-coupled coils. Embodiments or configurations of the device other than those specifically described will be apparent to those skilled in the art, and are contemplated by and within the scope of the present invention.

The terms "generally", "substantially" or "approximately" as used herein may be applied to modify any quantitative representation which could permissibly vary without resulting in a change in the basic function to which it is related. For example, secondary coil 10 is described as having dimensions of approximately 25mm x 13 mm, but may permissibly vary from that dimension if the variance does not materially alter the capability of the invention.

While the present invention has been described and illustrated hereinabove with reference to various exemplary embodiments, it should be understood that various modifications could be made to these embodiments without departing from the scope of the invention. Therefore, the invention is not to be limited to the exemplary embodiments described and illustrated hereinabove, except insofar as such limitations are included in the following claims.

Claims

CLAIMSWhat is claimed and desired to be secured by Letters Patent is as follows:

1. A tunable electromagnetic device for use in magnetic resonance imaging, wherein the device is in electrical communication with an electromagnetic field generator for generating an electromagnetic field, the device comprising: a primary element having a resonance frequency electrically coupled to said electromagnetic field generator, said primary element comprising tuning and matching circuitry operable to vary said resonance frequency; and a secondary element having a resonance frequency disposed proximate said primary element, said secondary element inductively coupled to said primary element and reactive to said electromagnetic field.

2. The tunable electromagnetic device of claim 1, wherein said tuning and matching circuitry comprises variable reactance elements.

3. The tunable electromagnetic device of claim 2, wherein said variable reactance elements are selected from the group consisting of variable capacitors, variable inductors, and combinations thereof.

4. The tunable electromagnetic device of claim 1, wherein said primary element and said secondary element are movable relative to each other to vary the inductive coupling between said primary and secondary elements.

5. The tunable electromagnetic device of claim 1, wherein said primary element is positioned proximate said secondary element such that said elements are inductively over- coupled.

6. The tunable electromagnetic device of claim 5, wherein said secondary element is positioned within said primary element.

7. The tunable electromagnetic device of claim 1, wherein said secondary element comprises a bio-compatible material adapted to be implanted within organic tissue.

8. The tunable electromagnetic device of claim 1 , wherein said resonance frequency of said secondary element is suitable for use with a magnetic resonance imaging device.

9. A method of tuning a frequency at which a secondary element inductively coupled to a primary element will be responsive, comprising: providing a tunable primary element; creating an electromagnetic field at a predetermined frequency using an electromagnetic field generator; coupling said electromagnetic field to said first element; positioning a secondary element proximate said primary element such that said elements are inductively coupled; and adjusting a position of said primary element relative to said secondary element such that said elements are inductively over-coupled.

10. The method of claim 9, wherein said secondary element is positioned within said primary element.

11. The method of claim 9, further comprising: tuning said primary element to achieve a desired resonance frequency.

12. The method of claim 11, wherein said tuning comprises matching a resonant frequency of said primary element to a predetermined frequency.

13. The method of claim 11, wherein said tuning comprises varying a reactance of said primary element.

14. The method of claim 11₅ wherein said adjusting comprises: rotating said primary element with respect to said secondary element; monitoring a frequency response of said electromagnetic field generator to determine overcoupling between said primary and secondary elements.

15. The method of claim 14, further comprising: tuning said primary element to match a frequency determined by said monitoring step.

16. The method of claim 15, wherein said electromagnetic field generator generates a radio frequency electromagnetic field suitable for use with an MRI device.

17. A method of using inductively over-coupled elements responsive to the presence of an electromagnetic field for in vivo resolution of a spatially localized biological tissue, comprising: implanting an implantable secondary element in a targeted biological tissue, said element being responsive to an electromagnetic field over a range of frequencies; providing a an electromagnetic field generator; positioning a tunable primary element proximate said secondary element, said primary element coupled to said electromagnetic field generator and positioned such that said primary and secondary elements are inductively coupled; creating an electromagnetic field at a predetermined frequency using said electromagnetic field generator; and adjusting a position of said primary element relative to said secondary element such that said elements are inductively over-coupled.

18. The method of claim 17, further comprising: tuning said primary element to achieve a desired resonance frequency of said- electromagnetic field generator.

19. The method of claim 18, wherein said tuning comprises matching a resonant frequency of said primary element to a predetermined frequency.

20. The method of claim 18, wherein said tuning comprises varying a reactance of said primary element.

21. The method of claim 18, wherein said adjusting step comprises: rotating said primary element with respect to said secondary element; monitoring a frequency response of said electromagnetic field generator to determine overcoupling between said primary and secondary elements.

22. The method of claim 21, further comprising: tuning said primary element to match a frequency determined by said monitoring step.

23. The method of claim 21, wherein said tuning comprises matching a resonant frequency of said primary element to a predetermined frequency.

24. The method of claim 17, wherein said electromagnetic field generator generates a radio frequency electromagnetic field suitable for use with an MRI device.

25. The method of claim 17, further comprising: obtaining images of said targeted tissue.

26. The method of claim 25, wherein said obtaining images comprises: obtaining repetitive images using said implanted secondary element.

27. The method of claim 25, wherein said images comprise high resolution images.

28. The method of claim 27, wherein said high resolution images comprise data having a high signal-to-noise ratio.