High-Tc superconducting rf receiver coils for magnetic resonance imaging of small animals

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Physica C 341-348 (2000) 2561-2564 www.elsevier.nl/Iocate/physc

High-Tc s u p e r c o n d u c t i n g r f r e c e i v e r coils for magnetic resonance imaging o f small animals * J. Wosik, I K. Nesteruk, 2 L.-M. Xie, l M. Strikovski, t F. Wang, 1 J. H. Miller, Jr., 1 M. Bilgen, a and P. A. Narayana 3 ~Texas Center for Superconductivity at University of Houston, and Electrical and Computer Engineering Department, University of Houston, Houston, TX 77204, USA 2Institute of Physics, Polish Academy of Sciences, AI. Lotnikow 32/36, PL-02 668 Warszawa, Poland 3Radiology Department, University of Texas-Houston Medical School, 6431 Fannin, Houston, TX 77030, USA

We report on an HTS rfreceiver surface probe designed for 2-Tesla MRI imaging of spinal cord injuries in small animals. The 2-T probe is used in lieu of an implanted copper coil being currently used in research on spinal cord injuries. The HTS probe was designed with a virtual ground plane, thus reducing the coil-to-ground losses and making its unloaded quality factor and resonant frequency less sensitive to body proximity. Each coil was fabricated using patterned double-sided YBa2Cu3Ox (YBCO) films deposited either on sapphire or LaAIO3 substrates. The signal-to-noise ratio (SNR) was analyzed numerically using complete solutions to MaxweU's equations and the reciprocity principle for a r e c t a n ~ o i l next to a finite Iossy dielectric cylinder. A comparison of images obtained with superconducting and cooled copper probes is shown.

I. INTRODUCTION Magnetic resonance imaging (MRI) is an ideal modality for investigating spinal cord injuries due to its excellent soft tissue contrast. A decade ago, the prevailing thought among most scientists was that treatment aimed at promoting recovery from a spinal cord injury was a hopeless task. However, new advances in pharmaceutical development and tissue grafts are eliciting new optimism among the spinal cord injured patients. For a number of reasons, rats are generally chosen for experimental spinal cord injury studies. Unfortunately, the size of rat's spinal cord is small (about 3 mm in diameter) and places a premium on signal-to-noise ratio (SNR). This problem has been overcome, to some extent, through the use of implanted coils [1]. The implanted coils have produced the best image quality and are considered as the "gold standard". However, despite these advantages, there are a number of disadvantages with the use of such implanted coils.

For example, the probe requires a large blunt dissection of muscle from the spine, ot~en resulting in inflammation of the adjacent musculature. In addition, coil failures due to broken solder joints are common in some animals, which result from additional stresses on the coil. Finally, this approach clearly can not be used in the case of human subjects. The ability to non-invasively characterize and follow the progression of the injury is critical for successful outcome of any therapeutic intervention. A preferred method of performing such studies is to improve the SNR with use of high-To superconducting (HTS) receiver coils, which would eliminate the need for implanted coils. In the case of small-volume, high-resolution imaging, the noise of the probe coil and/or preamplifier sets the system noise floor and hence the MRI performance; body noise no longer dominates the SNR [2]. Thus, it is desirable to reduce thermal coil noise to improve the image resolution and reduce the acquisition

This work was supported by the State of Texas, via the Texas Center for Superconductivity at the University of Houston and the Texas Higher Education Coordinating Board (ARP and ATP grants). *

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time. Since the Johnson noise is a function of the product of resistance and temperature, reduction of either or both of these parameters will enhance the SNR values. High-T¢ superconductors are extremely attractive for such applications due to their extremely low losses. Epitaxial, high quality superconducting YBCO films have a surface resistance at least four orders of magnitude lower than that of copper at I00 MHz [3]. Several demonstrations have shown that, for low-field and high-resolution applications, HTS MR/ receiver coils perform significantly better than comparable copper coils [4-9]. Even in high-field MRI, when the coil and the corresponding region of interest (ROI) are reduced in size, there is a crossover point beyond which the receiver noise becomes the dominant source of noise, which must be reduced to enhance SNR.

Both a fine frequency-tuning copper paddle and a matching system were designed to be placed in the cover of the cryostat. The coupling system preserves the SNR gain afforded by the HTS coil and provides an efficient way to match the input impedance of the preamplifier to the HTS coil.

2. P R O B E D E S I G N

Surface receiver coils should be operated as resonant circuits. Properly designed rfreceiver coils should be sensitive to r f magnetic fields and insensitive to r f electric fields (dielectric losses). Electric and rfmagnetic fields can induce currents in the body tissue, which has a relatively high conductivity, causing significant losses in the resonant circuit of the MR/ probe. One method of reducing electric field influence is to design a symmetrically balanced coil to introduce a virtual ground plane [10]. In this case, the maximum voltage produced in the coil with respect to ground is half of what it would have been if only one end of the coil had been grounded. The unloaded Q and resonant frequency of such a coil are thus less affected by proximity to the body. The sandwichtype coil, discussed in this work, is of this type. We have designed and fabricated a probe consisting of patterned, double sided YBCO films deposited on 2" LaAIO3 or sapphire substrates. Measurements of the unloaded Q of a copper probe at room temperature and 77 K yielded Q = 170 and 550, respectively. The superconducting coil has an unloaded Q, at 77 K, about an order of magnitude higher. A custom-made Tristan Technologies G-10 liquid nitrogen cryostat was modified in our lab for rfuse. The HTS probe is inductively coupled to a copper wire loop, which is connected to the 2-Tesla scanner. This loop also works as a transmit coil.

Figure 1. Photograph of 2" in diameter HTS receiver probe on a LaAIO3 substrate. Double sided 0.5 ~m YBCO film was patterned on both sides of the substrate.

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Figure 2. Experimental arrangement for MRI of the rat spine. HTS pick-up probe is placed above the rat's body. Transmittal coil works also as the coupling coil. A superconducting 2 Tesla magnet is used.

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In Fig. 2 an experimental configuration is shown schematically. The HTS coil is placed on the bottom of the liquid nitrogen cryostat, about 10-13 mm above the rat's body. 3. SIGNAL-TO-NOISE RATIO The body was modeled as a 60 mm long dielectric cylinder of conductivity o (Fig. 3) to calculate SNR for the experimental configuration. We assumed that a rectangular (15 x 30 mm) coil is placed a distance Zp from the cylinder. A voxel of volume AV was taken to be 5 mm deep in the body. This configuration provides an adequate representation of the actual situation for spinal cord imaging in rats, using the 84.4 MHz MRI scanner. The rms thermal noise voltage of the coil is assumed to be caused by both the resistance of the coil Rcml and the effective body resistance Rbotly, and can be expressed as: V~ = ~/4kAffRcoilTcoil + RbodyTbody )

copper coil at room temperature, (2) a copper coil cooled to 77 K, (3) an HTS coil with Rs = 69 ~tf2, and (4) an idealized material with R~ = 0. The HTS surface resistance was chosen very conservatively for the simulation. 69 ~tf~ is a rather a high Rs value for high-T¢ superconductors at 84.4 MHz.

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where k is Boltzmarm's constant and T and Afare the temperature and bandwidth, respectively, of the preamplifier. Signal and noise voltage calculations are carried out using the reciprocity principle and line integrals along the coil contour. The body noise is simulated by losses in the coil due to the existence of the body, and defined as an equivalent body resistance Rbody for the coil. The induced signal Vs in the coil can be expressed as: S oc a~AV(Bl)xy, where Bt is the calculated sensitivity of the coil, and o~ and A V are the frequency and voxel volume, respectively. In this paper, SNR is defined as the signal-to-noise voltage ratio and not the signal-to-noise power ratio. The SNR is calculated by dividing the magnitude of the signal voltage by the thermal noise voltage of the coil for a coil located a distance zp from the dipole source, assumed to be 5 mm deep in thebody. Thus, the SNR as a function of the probe position zp can be expressed as:

SNR(zp) = V+!zp ) .

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V,(z~) Fig. 4 shows the predicted variations of SNR with coil position at 84.4 MHz for a 15 x 30 mm rectangular coil and tissue of conductivity 1 (~m) "~. The calculations were done for four cases: (1) a

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Body-to-CoilDistance(mm) Figure 4. Predicted SNR as a function of distance of the probe from the region being imaged. The curves are calculated for room temperature copper, copper at 77 K, HTS at 77 K, and for the case when the material has no Johnson noise. It can be seen from Fig. 4 that there is an optimum position away from the body which maximizes the SNR while imaging a voxel deep in the body. This is an extremely important result for HTS applications, since the cryostat introduces a gap between the body and the coil. It shows that the HTS probe can still be placed about 10-15 mm away from the body while maximizing the SNR, Also note that the SNR of the HTS probe is substantially larger as that of the warm copper coil at 84.4 MHz.

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4. IMAGING All magnetic resonance studies were performed in a 2-Tesla, custom-built scanner [11] interfaced to an SMIS Console (Surrey, UK). The cryostat with the HTS probe was inserted into a horizontal 20-cm bore superconducting magnet. The MR/protocol included the acquisition of multi-slice density-weighted or T2weighted images, with a 5 mm slice thickness, a small field-of-view (FOV) of 25 mm 2, and a 256 x 256 matrix size. The imaging bandwidth was 20 kHz. Prior to spinal cord imaging, the magnetic field was always homogenized over the region-of-interest using a custom-developed auto shimming routine. In Figure 5 examples of axial spin echo densityweighted images of a rat's spine acquired using the HTS probe and an equivalent copper probe of the same design are shown. These images represent raw data without any filtering or processing for rf inhomogeneity correction.

Note that gray matter can clearly be seen in the images acquired using the HTS probe. Moreover, the already very good resolution can be improved even further by using better quality films with lower Rs. The observed discrepancy between calculated (see Figure 3) and measured gains of the HTS coil over copper equivalent design coil, is due to better design of the measured coils. In simulations a simple rectangular non-balanced coil was used. 5. SUMMARY An HTS rf receiver probe has been designed, constructed and tested for M R / o f a rat's spine. The probe is found to non-invasively provide comparable or better quality images than an implanted copper coil currently being used in research on spinal cord injuries. The images acquired with superconducting probe have shown a great improvement, over copper external coil, in SNR and reduction of imaging time. REFERENCES

Figure 5. Images of rat's spine acquired using normal metal probe at room temperature (upper image) and HTS probe at 77 K (bottom image). Images were acquired using room temperature and 77 K (not shown in this Figure) copper rfprobes for comparison of SNR purposes. The 77-K HTS receiver coil was found to increase the SNR by 4 dB relative to that of an equivalent copper coil at room temperature.

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