- Email: [email protected]
In vivo echo-planar imaging of rat spinal cord
Magnetic Resonance Imaging, Vol. 16, No. 10, pp. 1249 –1255, 1998 © 1998 Elsevier Science Inc. All rights reserved. Printed in the USA. 0730-725X/98 $19.00 1 .00
PII S0730-725X(98)00093-9
● Original Contribution
IN VIVO ECHO-PLANAR IMAGING OF RAT SPINAL CORD DAVID A. FENYES
AND
PONNADA A. NARAYANA*
The University of Texas Medical School at Houston, Department of Radiology, Houston, Texas, USA An integrated approach to echo-planar imaging of rat spinal cord in vivo with a small field of view (FOV) is presented. This protocol is based on a multishot interleaved echo-planar imaging (EPI) sequence and includes: 1) use of an inductively coupled implantable coil for improved signal-to-noise ratio (SNR); 2) three-dimensional (3D) automatic shimming of the magnetic field over the spinal cord; and 3) post-acquisition data processing using a multireference scan for minimizing image artifacts. Some of the practical issues in implementing this protocol are discussed. This imaging protocol will be useful in characterizing the spinal cord pathology using techniques that are otherwise time-consuming, such as diffusion tensor imaging. © 1998 Elsevier Science Inc. Keywords: Echo-planar imaging; Spinal cord; Implanted coil; Autoshim.
as spinal cord pose a number of technical problems. These include poor signal-to-noise ratio (SNR) due to the small size of the spinal cord (typically about 3 mm), the need for a small field of view (FOV) for adequate resolution, and the relatively large susceptibility gradients due to the bony vertebrae. These technical difficulties explain the lack of in vivo EPI studies of rat spinal cord and brain even when EPI would have been appropriate. The purpose of these studies was, therefore, to develop and implement an integrated approach for generating EPI images of spinal cord consistently and routinely. This approach involves the use of an implanted coil for improved SNR, automatic shimming of the magnetic field over the cord, and application of post processing techniques for minimizing image artifacts. A number of practical issues in implementing this integrated approach are discussed.
INTRODUCTION Magnetic resonance imaging (MRI), with its exquisite soft tissue contrast, is an ideal radiologic modality for studying spinal cord pathology. While the sensitivity of MRI is high, its specificity is somewhat limited. However, it is possible to improve the specificity of MRI by measuring multiple tissue properties such as diffusion and magnetization transfer.1– 4 Multiple measurements using conventional spin echo techniques can be quite time consuming. For example, acquisition of a set of apparent diffusion tensor images using a spin-echo sequence can take several hours, depending on the number of diffusion steps. In cases where pathology may evolve over tens of minutes, such as that encountered in spinal cord trauma, higher temporal resolution is desirable. In addition, long scan times are undesirable for longitudinal in vivo animal studies in which prolonged anesthesia times may result in higher mortality rate. Therefore, it is more appropriate to use fast sequences for spinal cord imaging. Echo-planar imaging (EPI) is one of the fastest MR imaging techniques.5 Although EPI studies have been reported for large structures such as human brain, the application of EPI for imaging rat spinal cord has not been reported so far. EPI studies of small structures such
MATERIALS AND METHODS Hardware and Software All imaging was performed on a 2.0 T horizontal bore magnet (Nalorac, Martinez, CA, USA) interfaced to an SMIS console (Surrey, UK), and fitted with a Tesla (Sussex, UK) actively shielded gradient set including 10 shim coils. The coils are capable of producing gradients
RECEIVED 11/22/97; ACCEPTED 2/9/98 Address correspondence to P.A. Narayana, Ph.D., The University of Texas Medical School at Houston, Department of
Radiology, 6431 Fannin, MSB 2.100, Houston, TX, USA. E-mail: [email protected] 1249
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up to 100 mT/m at a peak driving current of 165 A. They were driven by three Techron 8604 gradient amplifiers (Crown Instruments, Elkhart, IN, USA). The shim coils were driven by a Nalorac RTS-50 shim amplifier (Nalorec Marlinez, CA, USA) interfaced to a PCI-3000 analog I/O controller (Burr-Brown, Tucson, AZ, USA). The raw data were transferred to a Sun Ultrasparc I (Sun Microsystems, Mountain View, CA, USA), and the image reconstruction and processing software was implemented in the IDL language (Research Systems Inc., Boulder, CO, USA). Implantable Radiofrequency (RF) Coil Implantable coils have been commonly used to increase the SNR for imaging of the spinal cord and other internal structures.6 – 8 For longitudinal studies lasting several weeks, inductively coupled implantable coils minimize the risk of infection and simplify animal care. An inductively coupled implantable coil for imaging rat spinal cord has been described by Ford et al.8 However, using such coils in studies involving large numbers of animals requires reproducible procedures for tuning the coil, positioning the animals in the scanner, and adjusting the RF transmit power. In this article we address these and other practical issues that arise in routine imaging using the implanted coil. The coils were prepared from 18-gauge soft, bare copper wire (CDA alloy 102, MWS Wire Industries, Westlake Village, CA, USA) which was bent into a 1.1 3 2.7 cm (inner measurements) “rooftop” shape using a custom designed former, and tuned with a 75 pF ceramic chip capacitor (American Technical Ceramics, Huntington Station, NY, USA). The coils were coated with the biologically inert silicone elastomer Silastic MDX4 – 4210 (Factor II, Inc., Lakeside, AZ, USA) and verified to tune to 84.4 MHz when loaded with a 5% NaCl solution. For implantation in animals, coated coils were sterilized by immersion in Cidex disinfectant for 24 h, followed by immersion in sterile water for 2 h prior to implantation. The coil was implanted over the portion of the spine centered about the T7 level. Rats were anesthetized for coil implantation. A midline incision was made from T2 to T10, and the underlying tissues and muscles were dissected away to expose the spine. The spinous processes at T4 and T10 were removed in order to stabilize the coil and allow it to lie flat against the rib cage. The coil was held temporarily in place with a single stitch in the overlying muscle, and tuning was verified before permanently closing the incision. The overlying paraspinal muscles were sutured at the midline, holding the coil in place, and the skin was closed. Animals with implanted coils recovered quickly, with no evidence of impaired mobility or gait abnormalities. For tuning and imaging, the implantable coil was inductively coupled to
a 1.5 3 3 cm external coil, series tuned for minimum impedance at the proton resonance frequency. To ensure reproducible animal positioning, the external coil was mounted on a plexiglas sled so that when the sled is placed inside the magnet, the center of the external coil is approximately 1 cm below the isocenter of the magnet. This ensures that when the implanted coil is properly coupled to the external coil, the spinal cord volume of interest (VOI) is very near the magnet’s isocenter. The animals were placed supine on the sled. Localized Automatic Shimming A phase map based protocol,9 adapted for shimming the magnetic field (B0) over small three-dimensional (3D) volumes, was implemented. An important feature of this implementation, not generally found in many published techniques, is that it allows to specify a volume of any arbitrary size and shape, graphically prescribed over the MR images, over which the B0 field is shimmed. Thus, it is possible to homogenize the field over the spinal cord. 3D phase maps were derived from complex images generated using a 3D dual gradient-echo sequence. The phase for each pixel was determined by the following unwrapping procedure. Low noise regions and sources for phase unwrapping errors were excluded as described by Hedley and Rosenfeld.10 A starting point for phase unwrapping was arbitrarily designated as having phase in the range (2p, p). Next, Dijkstra’s algorithm was applied to determine the minimum distance path for unwrapping, avoiding the previously marked low noise regions and invalidated points.11 The phase was then unwrapped in a one-dimensional fashion along the calculated paths. A complex phase difference image was derived by subtracting the unwrapped phase of the first echo from the unwrapped phase of the second echo. A master shim map was prepared by acquiring a baseline phase difference image for a homogeneous phantom. One by one, shim currents were perturbed enough to generate a significant phase change across the image when the baseline was subtracted, generating magnetic field profiles for each shim coil. A total of 10 shim profiles were acquired, corresponding to each of the available shim coils. On our magnet, these coils provide the X, Y, Z, Z2, XY, XZ, YZ, X2-Y2, Z3, and Z4 harmonics. The field profiles were stored with the master map along with the perturbation values, FOV, and resolution information. A binary map of valid pixels was also stored with the master map, and only valid pixels were used for shim decomposition. For in vivo shimming on the spinal cord, a 3D phase difference image of the cord was generated using the above procedure and the spinal cord VOI was selected. The master map was then interpolated using a trilinear
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interpolation to match the resolution of the phase difference image over the VOI. The interpolated master map, the phase map of the spinal cord, and the shim currents that cancel variations in the spinal cord phase map form an overdetermined set of linear equations. A QR decomposition was performed using the LAPACK DGELS routine12 to find the least-square solution for the unknown shim currents. The solution vector was directly multiplied by the shim perturbations stored in the master map to yield a set of shim current corrections. The current shim settings were read from a shim file, and new shim settings were written to a new shim file, which was then sent to the shim controller. Imaging and Reconstruction To effectively resolve the anatomical details in the spinal cord, a FOV of 25 mm or smaller, with 128 3 128 pixels or greater is needed. For our system, the minimum gradient rise time is between 500 ms and 1 ms, mainly limited by eddy currents. The acquisition window was kept under 60 ms to limit T*2 effects. With a 500 ms ramp time, gradient reversals alone would require 64 ms. It was therefore necessary to split the acquisition into several shots.13,14 We have implemented a multishot, multislice spin-echo-based blipped EPI sequence5 with the ability to partially sample k-space in the phase encode direction. Each shot acquires a single interleave of undersampled k-space, and the entire k-space is reconstructed by interleaving the acquired raw data.13,14 The images were corrected for distortion with the multireference correction technique described by Wan et al.,15 modified for use with multi-shot EPI data. The multireference data are acquired only once at the beginning of the session and are used to correct all subsequent EPI images acquired in the same session. Because of the complexity of the multireference scan, a simpler correction scheme referred to as single reference correction, described by Bruder et al.,16 was also evaluated. The purpose of the single reference scan is to center the echoes in the EPI data for reducing the artifacts due to eddy currents. The single reference scan was acquired as an extra shot per slice after the EPI data has been collected, with only a modest increase in scan time (12.5% increase for an eight-shot image). To further reduce the sampling time and to reduce the echo time, k-space was partially sampled, with 80 phaseencode steps. Once the partial k-space had been corrected by the single reference or multireference scan, a 128 3 128 k-space was synthesized using the basic Fourier correction method described by MacFall et al.,17 which was found to be as effective for these images as the iterative Fourier correction technique also described by MacFall et al.17 On our system, the high duty cycle of the EPI wave-
Fig. 1. 256 3 256 spin-echo images of spinal cord in a supine rat acquired using the implanted coil with RF transmit power optimized for cord imaging (a) and for maximum echo amplitude (b). The improved visibility of the spinal cord can be appreciated in (a). The spinal cord, with its charactersistic H-shaped gray matter, is at the center of the image. The conductors of the coil (marked C) are seen below the cord (dorsal) to either side. The lungs (marked L) are seen as large signal voids ventral to the cord bilaterally. The imaging parameters are: slice thickness of 1 mm, FOV 5 25 mm, repetition time 5 3000 ms with triggering, echo time 5 25 ms, number of excitations 5 2.
form in multislice acquisitions limited the maximum readout gradient to 35 mT/m to avoid gradient amplifier overheating. This limit imposed a maximum bandwidth near 50 kHz for a 25 mm field of view. For imaging rat spinal cord, we acquired the 128 3 80 partial k-space in eight interleaved shots, sampling at 50 kHz, with an echo time of 25 ms and a total sampling time of 32.8 ms per interleave. To minimize motion artifacts, each shot was triggered at a fixed point in the ventilation cycle. Acquisitions were triggered on every second ventilation, yielding a repetition time of 2300 ms. The acquisition time for nine slices with a 2-mm slice thickness and two excitations was 39 s. Six male Sprague–Dawley rats weighing 300 –350 g were used in these studies. For coil implantation and imaging, animals were anesthetized with 4% isoflurane, intubated, and maintained under anesthesia with a mixture of 2% isoflurane, 30% oxygen, and air. While under anesthesia, animals were ventilated with a Harvard Apparatus (South Natick, MA, USA) rodent ventilator (model 683) modified to provide a triggering signal for respiration-gated imaging. All animal studies were approved by the Animal Welfare Committee of the University of Texas at Houston Health Science Center. RESULTS AND DISCUSSION Figure 1a shows a typical 256 3 256, 25-mm FOV density-weighted, 1 mm thick axial spin-echo image
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acquired with the implanted coil. The spinal cord is at the center of the image, with the ventral nerve roots clearly visible. The dorsal aspect of the cord is at the top while the ventral aspect is at the bottom of the image. The clarity of the characteristic central H-band of gray matter can easily be appreciated in this figure. Just below the spine, the two circular voids are the cylindrical conductors of the implanted coil, which are parallel to the cord. The hyperintense area below the coil in the left hand image represents the edematous tissue as a result of incision. Because of the inhomogeneous RF field produced by the implantable coil, special care is needed in setting the transmit attenuation to optimize images from the cord. Improper RF adjustment degrades the image quality as demonstrated in Fig. 1b. This image was generated by adjusting the transmit power for maximum echo amplitude but the RF field was not optimal at the site of the spinal cord. We have empirically determined that, for this coil configuration, the optimum signal from the cord is obtained by increasing the RF power by 3.5 dB over the setting which yields the maximum echo amplitude. The improved image quality resulting from this increased RF power can be appreciated in Fig. 1a. The signal arises primarily from a ring which surrounds the conductors and includes the spinal cord. The signal from the cord is significantly enhanced, and the RF homogeneity over the cord is improved. We have also observed empirically that the slice profiles produced by optimized 90° and 180° pulses such as those designed by the Shinnar–LeRoux (SLR) algorithm are quite sensitive to B1 inhomogeneity. This is because SLR pulse profiles are degraded as the flip angle deviates from the designed value. While it is possible to adjust the RF power so that some portion of the cord experiences the designed flip angle, it is in practice difficult to ensure that the entire cord will experience the same flip angle. The resulting degraded slice profiles adversely affect image quality. However, we have had good results using a 4 ms, 1500 Hz Hamming windowed sinc pulse with two side lobes, suggesting that the slice profile produced by these pulses is less sensitive to the flip angles. The B1 inhomogeneity of the implantable coil is actually advantageous for small FOV imaging, because the rapid signal falloff away from the spinal cord prevents extra-cord tissues from aliasing onto the spinal cord. In these studies we did not find it necessary to use specific techniques for suppressing the lipid resonances. This may be due to the absence of strong lipid signal within the sensitive range of the implanted coil. The B1 inhomogeneity may affect quantitative analyses of the spinal cord data sets. It has been shown that a postprocessing step to correct for RF inhomogeneity18,19 significantly improves the reliability of semiautomatic image segmentation.20 The RF inhomogeneity
correction step will be incorporated in future quantitative studies. We have been routinely using the implantable coils in our laboratory for over 30 months on more than 100 rats. The coils have been constructed by several different students and technicians, and scanning was performed by several different operators, with consistently good SNR and image quality with a standard spin echo sequence. Our experience of roughly fivefold SNR improvement with implantable coils over external RF coils is consistent with that reported by Ford et al.8 Phase unwrapping greatly influences the outcome of the autoshim. The performance of the phase unwrapping algorithm used in the present studies is demonstrated in Fig. 2 which shows the phase of complex images of three contiguous slices generated with the three-dimensional gradient-echo sequence. The phase images generated with arctangent function (top row) show abrupt phase discontinuities. These discontinuities are almost eliminated in the phase unwrapped images shown in the bottom row. The effect of the autoshim on the EPI images is demonstrated in Fig. 3. All the images shown in this figure are from the same section of the spinal cord. Figure 3a and c were acquired prior to autoshimming and were reconstructed with single reference and multireference corrections, respectively. Figure b and d are images acquired after autoshimming, reconstructed with single reference and multireference corrections, respectively. The improved cord shape in Fig. 3d can be appreciated as a reduction in the “notched” appearance of the cord in comparison with Fig. 3c. Although the region of interest for shimming was selected to include only the cord, the more circular appearance of the coil conductor voids in Fig. 3b and d indicate an overall improvement in the shim in the vicinity of the cord. The SNR of the cord in Fig. 3b and d is approximately twice that in Fig. 3a and c. We have observed similar SNR improvements in spinecho images using the autoshim protocol. Without the autoshim step, we find that the SNR of echo-planar images is too low for either the single reference scan or the multireference scan to be effective, This manifests as ghosting and blurring in Fig. 3a and c. Although image distortion is greatly reduced by the autoshim step, it cannot be completely eliminated, since susceptibility-induced gradients in the tissues have a higher spatial frequency that can not be compensated with the available shim coils. Figure 4 demonstrates the effects of the various corrections on the images. Figure 4a is a 128 3 80 spin-echo image reconstructed to 128 3 128 and is included for comparison with the EPI images. Figure 4b shows the EPI image of the cord after shimming and reconstructed without any reference correction. The pinstripe and ghost
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Fig. 2. Illustration of the three-dimensional phase unwrap used by the shimming program. The images in the top row are the wrapped phase of contiguous complex images from the three-dimensional gradient-echo data set of rat spinal cord and surrounding tissues, obtained using the arctangent function. The bottom row contains the corresponding images after phase-unwrapping procedure described in the text. Note the absence of abrupt phase discontinuities in the phase unwrapped images. The images were acquired using the implanted coil. The white arrows indicate the location of the spinal cord in each image. The noisy regions marked with (L) represent air in the lungs.
artifacts are due to shifting and distortion of the echoes by eddy currents. At larger FOV, these effects can be reduced by adding preemphasis to the gradient waveforms to compensate for the eddy current effects.21–23 However, at a small FOV, the gradient system operates at near maximum current, precluding the implementation of pre-emphasis. The echo shifting and distortion are easily corrected by use of a reference scan.16 Figure 4c is the same data, corrected with the single reference scan. The reduction of ghosts and pinstripes can easily be seen in this figure. It is clear that with no correction to align the echoes, these EPI images would be unusable. A comparison of Fig. 4a and c indicates that although single reference scan dramatically improves the image quality, the image still retains some geometric distortion. Unlike the single reference correction, the multireference corrects much of the distortion of anatomical details within the cord, as can be appreciated in Fig. 4d. The favorable comparison between the multireference scan corrected EPI and the spin-echo images can be better appreciated on the zoomed images (33) shown in Fig. 3e and f. Wan et al.15 noted that the multireference correction degrades the SNR This is consistent with our observation that multireference corrected images, have lower SNR compared to the single reference corrected image.
However, the overall effectiveness of the multireference scan is improved by the autoshim step. Motion between shots can generate artifacts in the EPI images. Approaches to the reduction of motion related ghosting include the use of a navigator echo24,25 and synchronizing the acquisitions to patient motion by triggering.26 –28 In imaging the rat spinal cord, the major source of motion is from the lungs. We have tried both navigator echoes and triggering from the ventilator, and it has been our experience that for rat spinal cord imaging, triggering is critical for reducing image artifacts. For best resolution of the anatomical details of the cord, it is desirable to image with the smallest possible slice thickness. The minimum slice thickness is limited by the intrinsically low SNR of the cord. Although we have been able to obtain satisfactory EPI cord images as thin as 1 mm, we have found 2-mm slices yield consistently high quality images in routine EPI imaging of the cord. It is possible that in the future, more sensitive detection methods may allow routine rat cord EPI using a 1-mm slice thickness. When experimental conditions and goals are compatible with long acquisition times, a spin-echo sequence offers higher SNR and minimal distortion. However, fast imaging times are important under certain conditions.
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Fig. 3. Effect of autoshim on the EPI reconstruction using single reference and multireference scans. The imaging parameters are: FOV 5 25 mm, and slice thickness 5 2 mm, repetition time 5 2300 with triggering, echo time 5 25 ms, number of excitations 5 2, 128 3 80 views reconstructed to 128 3 128. Images prior to autoshim with single reference correction (a), and multireference correction (c) exhibit Low SNR and ghosting. After autoshimming, SNR is increased and ghosting is reduced in the image when reconstructed with both single reference correction (b) and multireference correction (d).
When acquiring multiple-image sets from the cord, such as diffusion tensor imaging, the total imaging time may be quite lengthy and inappropriate for in vivo studies. The multi-reference scan adds additional 10 minutes to the EPI scan times. However, the multi-reference scan needs to be acquired only once at the beginning of the scan. Therefore, even with the additional scan time needed for acquiring multi-reference scan, EPI is much faster and better suited than conventional spin-echo imaging. If small image distortions are tolerable, it is possible to use single reference corrected EPI images and further reduce the total EPI acquisition time. CONCLUSION We have implemented an integrated approach that includes the use of an implanted coil for increased SNR, autoshimming of the magnetic field over the cord, and postprocessing using a reference scan in generating EPI images of the spinal cord in rats in vivo. The protocol described here will facilitate examination of cord properties in vivo. This will allow time consuming techniques, such as apparent diffusion tensor
Fig. 4. Effect of single reference and multireference correction. For reference, a spin echo image of rat cord is shown (a). EPI images of the same slice acquired after autoshimming are shown in (b), (c), and (d). The images were reconstructed without any reference scan (b), with the single reference scan (c), and with the multireference scan (d), respectively. Pinstriping is eliminated and ghosting is reduced in both (c) and (d). Reduction in geometric distortion of the cord in (d) compared to (c) can be seen. The zoomed (33) spinal-cord images from (a) and (d) are shown in (e) and (f), respectively, for a better appreciation of the anatomical details. The imaging parameters for all the images are: FOV 5 25 mm and slice thickness 5 2 mm. The k-space is 128 3 80 samples reconstructed to 128 3 128. Repetition time 5 2300 ms with triggering, echo time 5 25 ms, number of excitations 5 2.
mapping, to be used to study both normal and pathological spinal cord in the rat. Although the autoshim and multi-reference scan consume about 15 min at the start of the scan, the subsequent images can be collected at a high temporal resolution, which will permit the study of cord tissue pathology evolving after acute traumatic spinal cord injury. Acknowledgments—This work was supported by the National Institutes of Health Grant NS30821. The authors thank Dr. Shi-Jie Liu for his help with the surgery and assistance with the animal work.
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Hand, C.; Black, P. Characterization diffusion coefficients in a rat spinal cord injury model. Magn. Res. Med. 31: 488 – 494; 1994. Howseman, A.; Stehling, M.; Chapman, B.; Coxon, R.; Turner, R.; Ordidge, R.; Cawley, M.; Glover, P.; Mansfield, P.; Coupland, R. Improvements in snap-shot nuclear magnetic resonance imaging. Br. J. Radiol. 61:822– 828; 1988. Farmer, T.; Cofer, G.; Johnson, G. Maximizing contrast to noise with inductively coupled implanted coils. Invest. Radiol. 25:552–558; 1990. Wirth III, E.; Mareci, T.; Beck, B.; Fitzsimmons, J.; Reier, P. Comparison of an inductively coupled implanted coil with optimized surface coils for in vivo imaging of the spinal cord. Magn. Reson. Med. 30:626 – 633; 1993. Ford, J.; Hackney, D.; Joseph, P.; Phelan, M.; Alsop, D.; Tabor, S.; Hand, C.; Markowitz, R.; Black, P. Method for in vivo high resolution MRI of rat spinal cord injury. Magn. Reson. Med. 31:218 –223; 1994. Schneider, E.; Glover, G. Rapid in vivo proton shimming. Magn. Reson. Med. 18:335–347; 1991. Hedley, M.; Rosenfeld, D. A new two-dimensional phase unwrapping algorithm for MRI images. Magn. Reson. Med. 24:177–181; 1992. Cormen, T.; Leiserson, C.; Rivest, R. Introduction to Alogorhythms. New York: McGraw Hill; 1990: pp. 527–532. Anderson, E.; Bai, Z.; Bischof, C.; Demmel, J.; Dongarra, J.; Du Croz, J.; Greenbaum, A.; Hammarling, S.; McKenney, A.; Ostrouchov, S.; Sorenson, D. Lapack User’s Guide. Philadelphia: SIAM; 1994, p.140. McKinnon, G. Ultrafast interleaved gradient-echo-planar imaging on a standard scanner. Magn. Reson. Med. 30: 609 – 616; 1993. Butts, K.; Riederer, S.; Ehlman, R.; Thompson, R.; Jack, C. Interleaved echo planar imaging on a standard system. Magn. Reson. Med. 31:67–74; 1994. Wan, X.; Gullberg, G.; Parker, D.; Zeng, G. Reduction of geometric and intensity distortions in echo-planar imaging using a multireference scan. Magn. Reson. Med. 37:932– 944; 1997. Bruder, H.; Fischer, H.; Reinfelder, H.; Schmidt, F. Image reconstruction for echo planar imaging with non-
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PII S0730-725X(98)00093-9
● Original Contribution
IN VIVO ECHO-PLANAR IMAGING OF RAT SPINAL CORD DAVID A. FENYES
AND
PONNADA A. NARAYANA*
The University of Texas Medical School at Houston, Department of Radiology, Houston, Texas, USA An integrated approach to echo-planar imaging of rat spinal cord in vivo with a small field of view (FOV) is presented. This protocol is based on a multishot interleaved echo-planar imaging (EPI) sequence and includes: 1) use of an inductively coupled implantable coil for improved signal-to-noise ratio (SNR); 2) three-dimensional (3D) automatic shimming of the magnetic field over the spinal cord; and 3) post-acquisition data processing using a multireference scan for minimizing image artifacts. Some of the practical issues in implementing this protocol are discussed. This imaging protocol will be useful in characterizing the spinal cord pathology using techniques that are otherwise time-consuming, such as diffusion tensor imaging. © 1998 Elsevier Science Inc. Keywords: Echo-planar imaging; Spinal cord; Implanted coil; Autoshim.
as spinal cord pose a number of technical problems. These include poor signal-to-noise ratio (SNR) due to the small size of the spinal cord (typically about 3 mm), the need for a small field of view (FOV) for adequate resolution, and the relatively large susceptibility gradients due to the bony vertebrae. These technical difficulties explain the lack of in vivo EPI studies of rat spinal cord and brain even when EPI would have been appropriate. The purpose of these studies was, therefore, to develop and implement an integrated approach for generating EPI images of spinal cord consistently and routinely. This approach involves the use of an implanted coil for improved SNR, automatic shimming of the magnetic field over the cord, and application of post processing techniques for minimizing image artifacts. A number of practical issues in implementing this integrated approach are discussed.
INTRODUCTION Magnetic resonance imaging (MRI), with its exquisite soft tissue contrast, is an ideal radiologic modality for studying spinal cord pathology. While the sensitivity of MRI is high, its specificity is somewhat limited. However, it is possible to improve the specificity of MRI by measuring multiple tissue properties such as diffusion and magnetization transfer.1– 4 Multiple measurements using conventional spin echo techniques can be quite time consuming. For example, acquisition of a set of apparent diffusion tensor images using a spin-echo sequence can take several hours, depending on the number of diffusion steps. In cases where pathology may evolve over tens of minutes, such as that encountered in spinal cord trauma, higher temporal resolution is desirable. In addition, long scan times are undesirable for longitudinal in vivo animal studies in which prolonged anesthesia times may result in higher mortality rate. Therefore, it is more appropriate to use fast sequences for spinal cord imaging. Echo-planar imaging (EPI) is one of the fastest MR imaging techniques.5 Although EPI studies have been reported for large structures such as human brain, the application of EPI for imaging rat spinal cord has not been reported so far. EPI studies of small structures such
MATERIALS AND METHODS Hardware and Software All imaging was performed on a 2.0 T horizontal bore magnet (Nalorac, Martinez, CA, USA) interfaced to an SMIS console (Surrey, UK), and fitted with a Tesla (Sussex, UK) actively shielded gradient set including 10 shim coils. The coils are capable of producing gradients
RECEIVED 11/22/97; ACCEPTED 2/9/98 Address correspondence to P.A. Narayana, Ph.D., The University of Texas Medical School at Houston, Department of
Radiology, 6431 Fannin, MSB 2.100, Houston, TX, USA. E-mail: [email protected] 1249
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up to 100 mT/m at a peak driving current of 165 A. They were driven by three Techron 8604 gradient amplifiers (Crown Instruments, Elkhart, IN, USA). The shim coils were driven by a Nalorac RTS-50 shim amplifier (Nalorec Marlinez, CA, USA) interfaced to a PCI-3000 analog I/O controller (Burr-Brown, Tucson, AZ, USA). The raw data were transferred to a Sun Ultrasparc I (Sun Microsystems, Mountain View, CA, USA), and the image reconstruction and processing software was implemented in the IDL language (Research Systems Inc., Boulder, CO, USA). Implantable Radiofrequency (RF) Coil Implantable coils have been commonly used to increase the SNR for imaging of the spinal cord and other internal structures.6 – 8 For longitudinal studies lasting several weeks, inductively coupled implantable coils minimize the risk of infection and simplify animal care. An inductively coupled implantable coil for imaging rat spinal cord has been described by Ford et al.8 However, using such coils in studies involving large numbers of animals requires reproducible procedures for tuning the coil, positioning the animals in the scanner, and adjusting the RF transmit power. In this article we address these and other practical issues that arise in routine imaging using the implanted coil. The coils were prepared from 18-gauge soft, bare copper wire (CDA alloy 102, MWS Wire Industries, Westlake Village, CA, USA) which was bent into a 1.1 3 2.7 cm (inner measurements) “rooftop” shape using a custom designed former, and tuned with a 75 pF ceramic chip capacitor (American Technical Ceramics, Huntington Station, NY, USA). The coils were coated with the biologically inert silicone elastomer Silastic MDX4 – 4210 (Factor II, Inc., Lakeside, AZ, USA) and verified to tune to 84.4 MHz when loaded with a 5% NaCl solution. For implantation in animals, coated coils were sterilized by immersion in Cidex disinfectant for 24 h, followed by immersion in sterile water for 2 h prior to implantation. The coil was implanted over the portion of the spine centered about the T7 level. Rats were anesthetized for coil implantation. A midline incision was made from T2 to T10, and the underlying tissues and muscles were dissected away to expose the spine. The spinous processes at T4 and T10 were removed in order to stabilize the coil and allow it to lie flat against the rib cage. The coil was held temporarily in place with a single stitch in the overlying muscle, and tuning was verified before permanently closing the incision. The overlying paraspinal muscles were sutured at the midline, holding the coil in place, and the skin was closed. Animals with implanted coils recovered quickly, with no evidence of impaired mobility or gait abnormalities. For tuning and imaging, the implantable coil was inductively coupled to
a 1.5 3 3 cm external coil, series tuned for minimum impedance at the proton resonance frequency. To ensure reproducible animal positioning, the external coil was mounted on a plexiglas sled so that when the sled is placed inside the magnet, the center of the external coil is approximately 1 cm below the isocenter of the magnet. This ensures that when the implanted coil is properly coupled to the external coil, the spinal cord volume of interest (VOI) is very near the magnet’s isocenter. The animals were placed supine on the sled. Localized Automatic Shimming A phase map based protocol,9 adapted for shimming the magnetic field (B0) over small three-dimensional (3D) volumes, was implemented. An important feature of this implementation, not generally found in many published techniques, is that it allows to specify a volume of any arbitrary size and shape, graphically prescribed over the MR images, over which the B0 field is shimmed. Thus, it is possible to homogenize the field over the spinal cord. 3D phase maps were derived from complex images generated using a 3D dual gradient-echo sequence. The phase for each pixel was determined by the following unwrapping procedure. Low noise regions and sources for phase unwrapping errors were excluded as described by Hedley and Rosenfeld.10 A starting point for phase unwrapping was arbitrarily designated as having phase in the range (2p, p). Next, Dijkstra’s algorithm was applied to determine the minimum distance path for unwrapping, avoiding the previously marked low noise regions and invalidated points.11 The phase was then unwrapped in a one-dimensional fashion along the calculated paths. A complex phase difference image was derived by subtracting the unwrapped phase of the first echo from the unwrapped phase of the second echo. A master shim map was prepared by acquiring a baseline phase difference image for a homogeneous phantom. One by one, shim currents were perturbed enough to generate a significant phase change across the image when the baseline was subtracted, generating magnetic field profiles for each shim coil. A total of 10 shim profiles were acquired, corresponding to each of the available shim coils. On our magnet, these coils provide the X, Y, Z, Z2, XY, XZ, YZ, X2-Y2, Z3, and Z4 harmonics. The field profiles were stored with the master map along with the perturbation values, FOV, and resolution information. A binary map of valid pixels was also stored with the master map, and only valid pixels were used for shim decomposition. For in vivo shimming on the spinal cord, a 3D phase difference image of the cord was generated using the above procedure and the spinal cord VOI was selected. The master map was then interpolated using a trilinear
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interpolation to match the resolution of the phase difference image over the VOI. The interpolated master map, the phase map of the spinal cord, and the shim currents that cancel variations in the spinal cord phase map form an overdetermined set of linear equations. A QR decomposition was performed using the LAPACK DGELS routine12 to find the least-square solution for the unknown shim currents. The solution vector was directly multiplied by the shim perturbations stored in the master map to yield a set of shim current corrections. The current shim settings were read from a shim file, and new shim settings were written to a new shim file, which was then sent to the shim controller. Imaging and Reconstruction To effectively resolve the anatomical details in the spinal cord, a FOV of 25 mm or smaller, with 128 3 128 pixels or greater is needed. For our system, the minimum gradient rise time is between 500 ms and 1 ms, mainly limited by eddy currents. The acquisition window was kept under 60 ms to limit T*2 effects. With a 500 ms ramp time, gradient reversals alone would require 64 ms. It was therefore necessary to split the acquisition into several shots.13,14 We have implemented a multishot, multislice spin-echo-based blipped EPI sequence5 with the ability to partially sample k-space in the phase encode direction. Each shot acquires a single interleave of undersampled k-space, and the entire k-space is reconstructed by interleaving the acquired raw data.13,14 The images were corrected for distortion with the multireference correction technique described by Wan et al.,15 modified for use with multi-shot EPI data. The multireference data are acquired only once at the beginning of the session and are used to correct all subsequent EPI images acquired in the same session. Because of the complexity of the multireference scan, a simpler correction scheme referred to as single reference correction, described by Bruder et al.,16 was also evaluated. The purpose of the single reference scan is to center the echoes in the EPI data for reducing the artifacts due to eddy currents. The single reference scan was acquired as an extra shot per slice after the EPI data has been collected, with only a modest increase in scan time (12.5% increase for an eight-shot image). To further reduce the sampling time and to reduce the echo time, k-space was partially sampled, with 80 phaseencode steps. Once the partial k-space had been corrected by the single reference or multireference scan, a 128 3 128 k-space was synthesized using the basic Fourier correction method described by MacFall et al.,17 which was found to be as effective for these images as the iterative Fourier correction technique also described by MacFall et al.17 On our system, the high duty cycle of the EPI wave-
Fig. 1. 256 3 256 spin-echo images of spinal cord in a supine rat acquired using the implanted coil with RF transmit power optimized for cord imaging (a) and for maximum echo amplitude (b). The improved visibility of the spinal cord can be appreciated in (a). The spinal cord, with its charactersistic H-shaped gray matter, is at the center of the image. The conductors of the coil (marked C) are seen below the cord (dorsal) to either side. The lungs (marked L) are seen as large signal voids ventral to the cord bilaterally. The imaging parameters are: slice thickness of 1 mm, FOV 5 25 mm, repetition time 5 3000 ms with triggering, echo time 5 25 ms, number of excitations 5 2.
form in multislice acquisitions limited the maximum readout gradient to 35 mT/m to avoid gradient amplifier overheating. This limit imposed a maximum bandwidth near 50 kHz for a 25 mm field of view. For imaging rat spinal cord, we acquired the 128 3 80 partial k-space in eight interleaved shots, sampling at 50 kHz, with an echo time of 25 ms and a total sampling time of 32.8 ms per interleave. To minimize motion artifacts, each shot was triggered at a fixed point in the ventilation cycle. Acquisitions were triggered on every second ventilation, yielding a repetition time of 2300 ms. The acquisition time for nine slices with a 2-mm slice thickness and two excitations was 39 s. Six male Sprague–Dawley rats weighing 300 –350 g were used in these studies. For coil implantation and imaging, animals were anesthetized with 4% isoflurane, intubated, and maintained under anesthesia with a mixture of 2% isoflurane, 30% oxygen, and air. While under anesthesia, animals were ventilated with a Harvard Apparatus (South Natick, MA, USA) rodent ventilator (model 683) modified to provide a triggering signal for respiration-gated imaging. All animal studies were approved by the Animal Welfare Committee of the University of Texas at Houston Health Science Center. RESULTS AND DISCUSSION Figure 1a shows a typical 256 3 256, 25-mm FOV density-weighted, 1 mm thick axial spin-echo image
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acquired with the implanted coil. The spinal cord is at the center of the image, with the ventral nerve roots clearly visible. The dorsal aspect of the cord is at the top while the ventral aspect is at the bottom of the image. The clarity of the characteristic central H-band of gray matter can easily be appreciated in this figure. Just below the spine, the two circular voids are the cylindrical conductors of the implanted coil, which are parallel to the cord. The hyperintense area below the coil in the left hand image represents the edematous tissue as a result of incision. Because of the inhomogeneous RF field produced by the implantable coil, special care is needed in setting the transmit attenuation to optimize images from the cord. Improper RF adjustment degrades the image quality as demonstrated in Fig. 1b. This image was generated by adjusting the transmit power for maximum echo amplitude but the RF field was not optimal at the site of the spinal cord. We have empirically determined that, for this coil configuration, the optimum signal from the cord is obtained by increasing the RF power by 3.5 dB over the setting which yields the maximum echo amplitude. The improved image quality resulting from this increased RF power can be appreciated in Fig. 1a. The signal arises primarily from a ring which surrounds the conductors and includes the spinal cord. The signal from the cord is significantly enhanced, and the RF homogeneity over the cord is improved. We have also observed empirically that the slice profiles produced by optimized 90° and 180° pulses such as those designed by the Shinnar–LeRoux (SLR) algorithm are quite sensitive to B1 inhomogeneity. This is because SLR pulse profiles are degraded as the flip angle deviates from the designed value. While it is possible to adjust the RF power so that some portion of the cord experiences the designed flip angle, it is in practice difficult to ensure that the entire cord will experience the same flip angle. The resulting degraded slice profiles adversely affect image quality. However, we have had good results using a 4 ms, 1500 Hz Hamming windowed sinc pulse with two side lobes, suggesting that the slice profile produced by these pulses is less sensitive to the flip angles. The B1 inhomogeneity of the implantable coil is actually advantageous for small FOV imaging, because the rapid signal falloff away from the spinal cord prevents extra-cord tissues from aliasing onto the spinal cord. In these studies we did not find it necessary to use specific techniques for suppressing the lipid resonances. This may be due to the absence of strong lipid signal within the sensitive range of the implanted coil. The B1 inhomogeneity may affect quantitative analyses of the spinal cord data sets. It has been shown that a postprocessing step to correct for RF inhomogeneity18,19 significantly improves the reliability of semiautomatic image segmentation.20 The RF inhomogeneity
correction step will be incorporated in future quantitative studies. We have been routinely using the implantable coils in our laboratory for over 30 months on more than 100 rats. The coils have been constructed by several different students and technicians, and scanning was performed by several different operators, with consistently good SNR and image quality with a standard spin echo sequence. Our experience of roughly fivefold SNR improvement with implantable coils over external RF coils is consistent with that reported by Ford et al.8 Phase unwrapping greatly influences the outcome of the autoshim. The performance of the phase unwrapping algorithm used in the present studies is demonstrated in Fig. 2 which shows the phase of complex images of three contiguous slices generated with the three-dimensional gradient-echo sequence. The phase images generated with arctangent function (top row) show abrupt phase discontinuities. These discontinuities are almost eliminated in the phase unwrapped images shown in the bottom row. The effect of the autoshim on the EPI images is demonstrated in Fig. 3. All the images shown in this figure are from the same section of the spinal cord. Figure 3a and c were acquired prior to autoshimming and were reconstructed with single reference and multireference corrections, respectively. Figure b and d are images acquired after autoshimming, reconstructed with single reference and multireference corrections, respectively. The improved cord shape in Fig. 3d can be appreciated as a reduction in the “notched” appearance of the cord in comparison with Fig. 3c. Although the region of interest for shimming was selected to include only the cord, the more circular appearance of the coil conductor voids in Fig. 3b and d indicate an overall improvement in the shim in the vicinity of the cord. The SNR of the cord in Fig. 3b and d is approximately twice that in Fig. 3a and c. We have observed similar SNR improvements in spinecho images using the autoshim protocol. Without the autoshim step, we find that the SNR of echo-planar images is too low for either the single reference scan or the multireference scan to be effective, This manifests as ghosting and blurring in Fig. 3a and c. Although image distortion is greatly reduced by the autoshim step, it cannot be completely eliminated, since susceptibility-induced gradients in the tissues have a higher spatial frequency that can not be compensated with the available shim coils. Figure 4 demonstrates the effects of the various corrections on the images. Figure 4a is a 128 3 80 spin-echo image reconstructed to 128 3 128 and is included for comparison with the EPI images. Figure 4b shows the EPI image of the cord after shimming and reconstructed without any reference correction. The pinstripe and ghost
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Fig. 2. Illustration of the three-dimensional phase unwrap used by the shimming program. The images in the top row are the wrapped phase of contiguous complex images from the three-dimensional gradient-echo data set of rat spinal cord and surrounding tissues, obtained using the arctangent function. The bottom row contains the corresponding images after phase-unwrapping procedure described in the text. Note the absence of abrupt phase discontinuities in the phase unwrapped images. The images were acquired using the implanted coil. The white arrows indicate the location of the spinal cord in each image. The noisy regions marked with (L) represent air in the lungs.
artifacts are due to shifting and distortion of the echoes by eddy currents. At larger FOV, these effects can be reduced by adding preemphasis to the gradient waveforms to compensate for the eddy current effects.21–23 However, at a small FOV, the gradient system operates at near maximum current, precluding the implementation of pre-emphasis. The echo shifting and distortion are easily corrected by use of a reference scan.16 Figure 4c is the same data, corrected with the single reference scan. The reduction of ghosts and pinstripes can easily be seen in this figure. It is clear that with no correction to align the echoes, these EPI images would be unusable. A comparison of Fig. 4a and c indicates that although single reference scan dramatically improves the image quality, the image still retains some geometric distortion. Unlike the single reference correction, the multireference corrects much of the distortion of anatomical details within the cord, as can be appreciated in Fig. 4d. The favorable comparison between the multireference scan corrected EPI and the spin-echo images can be better appreciated on the zoomed images (33) shown in Fig. 3e and f. Wan et al.15 noted that the multireference correction degrades the SNR This is consistent with our observation that multireference corrected images, have lower SNR compared to the single reference corrected image.
However, the overall effectiveness of the multireference scan is improved by the autoshim step. Motion between shots can generate artifacts in the EPI images. Approaches to the reduction of motion related ghosting include the use of a navigator echo24,25 and synchronizing the acquisitions to patient motion by triggering.26 –28 In imaging the rat spinal cord, the major source of motion is from the lungs. We have tried both navigator echoes and triggering from the ventilator, and it has been our experience that for rat spinal cord imaging, triggering is critical for reducing image artifacts. For best resolution of the anatomical details of the cord, it is desirable to image with the smallest possible slice thickness. The minimum slice thickness is limited by the intrinsically low SNR of the cord. Although we have been able to obtain satisfactory EPI cord images as thin as 1 mm, we have found 2-mm slices yield consistently high quality images in routine EPI imaging of the cord. It is possible that in the future, more sensitive detection methods may allow routine rat cord EPI using a 1-mm slice thickness. When experimental conditions and goals are compatible with long acquisition times, a spin-echo sequence offers higher SNR and minimal distortion. However, fast imaging times are important under certain conditions.
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Fig. 3. Effect of autoshim on the EPI reconstruction using single reference and multireference scans. The imaging parameters are: FOV 5 25 mm, and slice thickness 5 2 mm, repetition time 5 2300 with triggering, echo time 5 25 ms, number of excitations 5 2, 128 3 80 views reconstructed to 128 3 128. Images prior to autoshim with single reference correction (a), and multireference correction (c) exhibit Low SNR and ghosting. After autoshimming, SNR is increased and ghosting is reduced in the image when reconstructed with both single reference correction (b) and multireference correction (d).
When acquiring multiple-image sets from the cord, such as diffusion tensor imaging, the total imaging time may be quite lengthy and inappropriate for in vivo studies. The multi-reference scan adds additional 10 minutes to the EPI scan times. However, the multi-reference scan needs to be acquired only once at the beginning of the scan. Therefore, even with the additional scan time needed for acquiring multi-reference scan, EPI is much faster and better suited than conventional spin-echo imaging. If small image distortions are tolerable, it is possible to use single reference corrected EPI images and further reduce the total EPI acquisition time. CONCLUSION We have implemented an integrated approach that includes the use of an implanted coil for increased SNR, autoshimming of the magnetic field over the cord, and postprocessing using a reference scan in generating EPI images of the spinal cord in rats in vivo. The protocol described here will facilitate examination of cord properties in vivo. This will allow time consuming techniques, such as apparent diffusion tensor
Fig. 4. Effect of single reference and multireference correction. For reference, a spin echo image of rat cord is shown (a). EPI images of the same slice acquired after autoshimming are shown in (b), (c), and (d). The images were reconstructed without any reference scan (b), with the single reference scan (c), and with the multireference scan (d), respectively. Pinstriping is eliminated and ghosting is reduced in both (c) and (d). Reduction in geometric distortion of the cord in (d) compared to (c) can be seen. The zoomed (33) spinal-cord images from (a) and (d) are shown in (e) and (f), respectively, for a better appreciation of the anatomical details. The imaging parameters for all the images are: FOV 5 25 mm and slice thickness 5 2 mm. The k-space is 128 3 80 samples reconstructed to 128 3 128. Repetition time 5 2300 ms with triggering, echo time 5 25 ms, number of excitations 5 2.
mapping, to be used to study both normal and pathological spinal cord in the rat. Although the autoshim and multi-reference scan consume about 15 min at the start of the scan, the subsequent images can be collected at a high temporal resolution, which will permit the study of cord tissue pathology evolving after acute traumatic spinal cord injury. Acknowledgments—This work was supported by the National Institutes of Health Grant NS30821. The authors thank Dr. Shi-Jie Liu for his help with the surgery and assistance with the animal work.
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