(Radiology. 1999;210:241-246.)
© RSNA, 1999
Articular Cartilage in the Knee: Mapping of the Physiologic Parameters at MR Imaging with a Local Gradient Coil—Preliminary Results
Lawrence R. Frank, PhD1,
Eric C. Wong, PhD, MD1,
Wen-Ming Luh, BS1,
Joong Mo Ahn, MD1 and
Donald Resnick, MD1
1 Department of Radiology, Veterans Administration Medical Center, 9114–RI, 3350 La Jolla Village Dr, San Diego, CA 92161.
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Abstract
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The authors designed and constructed a local gradient coil that produces large gradients and short rise times and connects to their clinical system. In a cadaveric patellar cartilage specimen, the coil was used to acquire images of the physiologic parameters T1, T2, proton density, and apparent diffusion coefficient. Variations in these parameters were evident in regions of normal and abnormal cartilage.
Index terms: Cartilage, MR, 45.121411, 45.121412 • Knee, ligaments, menisci, and cartilage, 4521.121411, 4521.121412 • Knee, MR, 45.121411, 45.121412 • Magnetic resonance (MR), coil arrays • Patella, 453.121411, 453.121412
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Introduction
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Cartilage disorders affect a large segment of the population and produce a variety of clinical manifestations. Magnetic resonance (MR) imaging has shown great potential in the assessment of articular cartilage because it is possible to obtain relatively high contrast between cartilage and surrounding structures. However, whereas MR imaging is helpful in the evaluation of osteochondral injuries and osteochondritis dissecans, in which there is usually some involvement of the subchondral bone, depiction of purely cartilage abnormalities, such as those in chondromalacia and inflammatory or degenerative arthritis, poses a more important problem. In part, this is simply a matter of spatial resolution and signal-to-noise ratio. The structure of cartilage is small relative to the spatial resolution of standard imaging, and intrachondral variations are too small to resolve or are obscured by partial volume effects. In addition, there are difficulties associated with the intrinsic signal of cartilage.
As a consequence of the complex structure of cartilage, the MR signal derived from cartilage is not easily quantified, as it depends on a variety of complicated factors related to both the local environment and the imaging parameters. On the other hand, because the MR signal is uniquely sensitive to important physiologic and biochemical parameters of cartilage, it offers the potential for quantitative assessment of cartilage structure and function in both normal and pathologic cases. To date, MR studies of the structural and physiologic characteristics of cartilage have been confined to high-field-strength systems that provide sufficient spatial resolution to allow visualization of the internal structure of cartilage.
We designed and constructed a local gradient and radio-frequency (RF) coil system optimized for high-spatial-resolution imaging of the extremities with a standard clinical system. Herein, we present our initial imaging results with patellar cartilage in a cadaveric specimen. Four types of images were obtained with a standard field of view (FOV) used in clinical imaging: (a) anatomic images with high spatial resolution and high signal-to-noise ratio; (b) projection-reconstruction images acquired with an exceedingly short echo time (TE) that allowed the capture of signal from components with short T2; (c) high-spatial-resolution maps of T1, T2, and proton density; and (d) diffusion-weighted images.
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Materials and Methods
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Pulse Sequences
For high-spatial-resolution imaging of the knee, clinical versions of spin-echo and gradient-recalled-echo (GRE) pulse sequences were modified for use with the local gradient coil to take into account the improved gradient strength and rise time.
Standard protocols for anatomic imaging of cartilage consisted of a two-dimensional GRE sequence (repetition time [TR] msec/TE msec = 600/9 with 20° flip angle); a three-dimensional spoiled GRE sequence (60/5 with 20° flip angle); and a T2-weighted, spin-echo sequence (4,000/80).
T1, T2, and proton density were calculated from the data for a pair of images (long TR, double-echo [4,000/20, 80] and short TR [600/20]). We used a fairly large difference between the TE values because we expected a somewhat wide range of T2 values. The first TE was short enough to provide a sufficient signal-to-noise ratio.
Diffusion-weighted images were acquired by adding diffusion gradients to a standard spin-echo pulse sequence. The gradient waveforms were designed to apply diffusion weighting independently along all three axes, allowing independent determination of the apparent diffusion coefficients in each axis (1). The sequence was modified to allow use of the larger gradient strength available with the local gradient coil. Diffusion-weighted imaging (600/40 with a b factor of approximately 373 sec/mm2) was performed along each of the three axes. The amplitude of the diffusion-weighting gradients was 5 G/cm, and their duration was 10 msec.
Short TE images were acquired with a projection-reconstruction sequence that reduces the TE by using a half-excitation pulse and eliminating the postexcitation compensating lobes from all gradient pulses (2). This allows data acquisition almost immediately after the completion of the excitation pulse, although it requires use of four half-echo excitations to sample a line of k space. We recently found results with this sequence for the detection of cartilage lesions to compare favorably to those with three-dimensional spoiled GRE and magnetization transfer imaging (3).
Local Gradient Coil
We designed the local gradient coil by using a conjugate gradient descent method described previously (4–6) and constructed it in our laboratory. The coil was optimized for high-spatial-resolution imaging of the extremities. The coil is a symmetric, torque-free, three-axes design built on a cylindric 22.86-cm-diameter, 40-cm-long form that produces 6 G/cm at 100 A along all three axes, with rise times of 100 µsec from zero to full scale. The region of gradient linearity is a 15.24-cm-diameter cylinder, with root-mean-square deviation from linearity of less than 3%. The coil creates gradient field strength five times than that created with the whole-body coil, inductance that is six times lower, and, therefore, switching that is six times faster. The coil allows spatial resolution on the order of 100 µm and read periods of approximately 2 msec for conventional imaging and 1 msec for projection-reconstruction imaging.
The local gradient coil was connected to the power amplifiers of the whole-body gradient coils, which were disconnected and insulated to ensure no currents were supported. The feedback loop in the amplifiers was modified to account for the impedance of the local gradient coil. This configuration was stable with respect to both insulation and gradient-waveform distortion.
RF Coil
The RF coil was designed to strongly couple the RF field to the knee so that dominant losses are from the patient. It is a transmit-receive quadrature birdcage design with diameter of 15.24 cm and length of 15.24 cm. To isolate the RF fields from the gradient fields and to maintain the quality factor of the RF coil near its isolated value, the RF coil is shielded from the local gradient coil by using a 0.13-mm-thick copper shield between them. The RF shield is designed for maximal decoupling of the RF and gradient coils at RF frequencies while remaining transparent to gradient fields. With the shield in place, the coil was restored to approximately its isolated quality factor. No noticeable eddy current effects were observed. For identical pulse sequence parameters, the coil gave a signal-to-noise ratio that is 23% higher than that of the commercial (GE Medical Systems, Milwaukee, Wis) clinical knee coil on our system (a linear saddle design).
Human and Cadaveric Studies
All images were acquired on a 1.5-T clinical MR imaging system (Signa; GE Medical Systems), and installation of the local gradient coil required approximately 15 minutes. The knee of a healthy volunteer was imaged (E.C.W.) in the standard supine orientation. This study was performed with institutional review board approval and informed consent from the volunteer.
The excised cadaveric specimen consisting of patellar cartilage was examined by an experienced arthroscopist to determine a normal and an abnormal region. The abnormality corresponded to a region of cartilage softening. After imaging had been completed in the two regions, the specimen was stained with India ink and suspended in a gelatin solution, as previously described (3), to allow comparison of the lesion with the imaging findings (7,8).
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Results
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The local gradient and RF coil system constructed in this study is shown in Figure 1.
Figure 2 shows a high-spatial-resolution image (FOV, 16 cm; section thickness, 3 mm; matrix, 256 x 256; two signals acquired) of a normal knee in a volunteer. The image was obtained with the two-dimensional GRE sequence and was acquired in approximately 5 minutes.
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Figure 2. High-spatial-resolution image (FOV, 16 cm; section thickness, 3 mm; matrix, 256 x 256; two signals acquired) of the normal knee in a volunteer, obtained with the two-dimensional GRE sequence (600/9 with 20° flip angle). The patellar cartilage (arrow) is clearly evident and of normal thickness.
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The cadaveric specimen was imaged with two standard pulse sequences. Figure 3 shows two cartilage images acquired with the local gradient coil. The high-spatial-resolution images were acquired with the three-dimensional spoiled GRE (FOV, 9.4 cm; section thickness, 1.5 mm; matrix, 256 x 256; one signal acquired) and T2-weighted, spin-echo (4,000/80; FOV, 9.4 cm; section thickness, 3 mm; matrix, 256 x 256; one signal acquired) sequences.
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Figure 3. High-spatial-resolution images (FOV, 9.4 cm; matrix, 256 x 256; one signal acquired) of a cadaveric specimen with a region of cartilage abnormality (arrow). Left: Three-dimensional spoiled GRE image (60/5; flip angle, 20°; section thickness, 1.5 mm); the abnormality is apparent as a region of decreased relative signal intensity. Right: T2-weighted, spin-echo image (4,000/80; section thickness, 3 mm); the abnormality is apparent as a region of increased relative signal intensity.
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The cadaveric specimen was also imaged with a projection-reconstruction sequence (FOV, 16 cm; TE, approximately 150 µsec). The local gradient coil offers a distinct advantage in projection-reconstruction imaging in that it facilitates very short ramp times, which improves acquisition of short T2 components. Figure 4 shows fat-suppressed, projection-reconstruction images acquired with the 500-µsec ramp time typical of clinical sequences, and the defect is seen clearly. This sequence was used in our previous study (3) and was found to perform better than the standard three-dimensional spoiled GRE sequence in the detection of lesions. Acquisition of the image with the 200-µsec ramp time was facilitated by the local gradient coil. The defect was again seen clearly, but a region of high signal intensity suggestive of the calcified cartilage region was seen near the subchondral bone.
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Figure 4. Fat-suppressed, projection-reconstruction images of a cadaveric specimen with a region of cartilage abnormality (short arrows). Images were acquired with 500-µsec (left) or 200-µsec (right) ramp times facilitated by the local gradient coil. The abnormality is again clearly evident, but it appears in the image on the right as a bright region (long arrow) near the subchondral bone that is suggestive of the calcified cartilage region.
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Use of the local gradient coil allowed acquisition of images with sufficient spatial resolution and signal-to-noise ratio to depict the spatial distribution of relaxation parameters across the cadaveric specimen. Figure 5 shows the T1 (top) and T2 (bottom) maps and Figure 6 shows the proton-density maps calculated from the data for a set of high-spatial-resolution (FOV, 9.4 cm; section thickness, 5 mm; matrix, 256 x 256; one signal acquired) long TR, double-echo (4,000/20, 80) and short TR (600/20) images. The in-plane spatial resolution was 0.37 mm. The maps were obtained at two locations: a region of normal cartilage (Fig 6a) and the region of the lesion (Fig 6b). These MR maps show striking changes in the proton-density, T1, and T2 values in the lesion, which is consistent with the theory that the structural changes that accompany cartilage degradation are associated with hydration. All maps show a clear laminar structure of the cartilage, but perhaps the most striking lamination is shown on the T2 map, which clearly reveals a shortening of T2 as a function of depth from the articular surface. These results are promising with regard to the use of relaxation times and proton density to indicate cartilage abnormalities and to study the structure of normal cartilage.
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Figure 5. T1 (top) and T2 (bottom) maps calculated from the data for a set of high-spatial-resolution (FOV, 9.4 cm; section thickness, 5 mm; matrix, 256 x 256; one signal acquired) long TR, double-echo (4,000/20, 80) and short TR (600/20) images. The maps were obtained in a region of normal cartilage (left) and in a region of cartilage abnormality (right, arrow). The scale indicates milliseconds. The cartilage abnormality is clearly seen on both maps, which suggests that relaxation rate imaging is potentially useful as a method of detecting cartilage abnormalities before gross structural changes occur and may provide a framework for quantification of these changes. The observed changes in relaxation rate are consistent with the theory that structural changes that accompany cartilage degradation are associated with hydration.
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Figure 6a. Proton-density maps calculated from the data for a set of high-spatial-resolution long TR, double-echo and short TR images. The maps were obtained in a region of normal cartilage (left) and in a region of cartilage abnormality (right, arrow). The scale indicates arbitrary units. The cartilage abnormality is clearly seen, suggesting that these may also be useful as a method for early detection of cartilage abnormalities. These changes are also consistent with the theory that structural changes that accompany cartilage degradation are associated with hydration.
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Figure 6b. Proton-density maps calculated from the data for a set of high-spatial-resolution long TR, double-echo and short TR images. The maps were obtained in a region of normal cartilage (left) and in a region of cartilage abnormality (right, arrow). The scale indicates arbitrary units. The cartilage abnormality is clearly seen, suggesting that these may also be useful as a method for early detection of cartilage abnormalities. These changes are also consistent with the theory that structural changes that accompany cartilage degradation are associated with hydration.
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Figure 7 shows the diffusion maps calculated from the data for a set of high-spatial-resolution, T2-weighted, spin-echo images (600/40; FOV, 9 cm; section thickness, 5 mm; matrix, 256 x 256; one signal acquired). The amplitude of the diffusion-weighting gradients was 5 G/cm and their duration was 10 msec, producing a b factor of approximately 373 sec/mm2 along each of the three axes. Figure 7a depicts the region of normal cartilage, and Figure 7b depicts that of the lesion. In both maps, a spatial variation is evident in the apparent diffusion coefficient. Further, the lesion is clearly seen to have an apparent diffusion coefficient different from that of the surrounding cartilage. The sensitivity and specificity of measurements of apparent diffusion coefficient for detection of early lesions remain to be tested, but the possibility of apparent diffusion coefficient imaging in the clinical setting provides a new metric for the evaluation of cartilage disorders.
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Figure 7a. Maps of apparent diffusion coefficient calculated from the data for spin-echo images (600/40; FOV, 9.4 cm; section thickness, 10 mm; matrix, 256 x 256; one signal acquired; b factor, approximately 373 sec/mm2 along each of the three axes). The maps were obtained at a region of normal cartilage (left) and a region of cartilage abnormality (right, arrow). The scale indicates square microns per millisecond x 1,000. The cartilage abnormality clearly shows an increase in apparent diffusion coefficient, which is again consistent with changes in hydration.
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Figure 7b. Maps of apparent diffusion coefficient calculated from the data for spin-echo images (600/40; FOV, 9.4 cm; section thickness, 10 mm; matrix, 256 x 256; one signal acquired; b factor, approximately 373 sec/mm2 along each of the three axes). The maps were obtained at a region of normal cartilage (left) and a region of cartilage abnormality (right, arrow). The scale indicates square microns per millisecond x 1,000. The cartilage abnormality clearly shows an increase in apparent diffusion coefficient, which is again consistent with changes in hydration.
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Discussion
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The structure of cartilage is quite complex and heterogeneous, consisting of several layers, or laminae, with different compositions and orientations of collagen fibrils and proteoglycans (9,10). This complicated composition has an effect on the intrinsic MR signal derived from normal cartilage owing to the extreme sensitivity of the signal to water content and parameters related to fiber orientation, such as relaxation (11), diffusion effects produced by the microscopic random motions of water (12), magic angle effects (11,13), and magnetization transfer effects related to variations in the structure of the MR-visible water that produce cross-relaxation between water pods (14,15). The content and distribution of hydrophilic proteoglycan molecules and the zonal organization of collagen influence not only the amount of water but also its state (16). Collagen, by virtue of its highly organized structure, tends to restrict the mobility of water protons and thus promotes spin-spin interactions and T2 decay (11). Because oriented fibers are associated with exceedingly short transverse relaxation times (11), visualization of regions of aligned collagen structures can be achieved with specially designed short TE sequences (17,18). Variations in T2 decay are then expected from the various chondral layers, short in the deep radial layers and in the fibrous superficial layers but longer in the transitional zone of disorganized collagen between these two layers. This complex signal behavior may be related to disruption of the normal structure of cartilage. Loss of proteoglycans results in edema of the articular cartilage as the vacated interfibrillar spaces are replaced with free water, resulting in relative signal intensity that is decreased on short TE images and increased on long TE images (19). Collagen loss, such as that seen in osteoarthritis or after injury, tends to increase relative signal intensity on long TE images because of the loss of the T2 shortening effect of collagen and because such loss promotes the hydration of proteoglycans (20,21). The T1 in normal cartilage has been examined (22), but it is poorly understood in degenerative states.
Because cartilage is avascular, transport of nutrients and waste products is believed to occur primarily by means of diffusion (23,24), and it is presumed to be altered in degenerative diseases such as osteoarthritis and in inflammatory diseases such as rheumatoid arthritis. Imaging of diffusion with MR has been studied extensively in the analysis of free motion of water in biologic samples. Burstein et al (12) found that the apparent diffusion coefficient of water in young calf cartilage increases with enzymatic depletion of proteoglycan and possibly other matrix components with trypsin, which shows that proteoglycan content is a major determinant of diffusion in cartilage. These investigators also found changes in the apparent diffusion coefficient associated with compression of the cartilage (12). Xia et al (22) also obtained high-spatial-resolution images of apparent diffusion coefficient, T1, and T2 in excised canine cartilage disks and plugs, and they determined that the apparent diffusion coefficient is not a linear function of the water concentration. Henkelman and co-workers (15) recently reported no anisotropic diffusion in cartilage. The changes in diffusion in degenerative cartilage are not well known, however.
The diagnostic capability of MR imaging for the assessment of cartilage abnormalities holds great promise because of the sensitivity of the MR signal to reflect changes in the complex structure of normal cartilage. The ability to utilize this sensitivity, however, depends on the availability of hardware capable of imaging the complex structure of cartilage. In particular, the demands of high-spatial-resolution imaging, short TE imaging, and diffusion-weighted imaging require use of hardware that is capable of larger gradient strengths and shorter ramp times than are normally needed for standard clinical imaging.
Visualization of intrachondral variations in relaxation requires much higher resolution than the cartilage dimensions. The high-spatial-resolution relaxation images shown in this article suggest their efficacy in both the detection and quantification of cartilage abnormalities. This is not possible, however, without the high spatial resolution and signal-to-noise ratio provided by the local gradient coil. The short ramp times allowed by use of the coil have an important advantage in our short TE, projection-reconstruction sequence because sampling is performed on the ramp, which results in a faster traversal of k space and shorter TE values. Findings on our initial images suggest that short T2 components near the calcified region of the bone-cartilage interface may be visible.
Generation of b factors large enough to produce measurable diffusion effects in cartilage, with spatial resolution high enough to allow visualization of intrachondral spatial variations, is beyond the limits of clinical scanners. Our initial results suggest, however, that our local gradient coil produces gradient strength adequate to produce diffusion weighting sufficient to allow detection of spatial variations in the apparent diffusion coefficient in regions of cartilage abnormalities. Even with use of currently available high-performance clinical systems that produce a gradient amplitude as high as 2.5 G/cm, the TE of the diffusion-weighted sequences would be approximately 15 msec higher than that used in this study, which would be prohibitively long for cartilage imaging.
In conclusion, these preliminary results suggest that use of specialized hardware to improve sensitivity and spatial resolution can greatly improve the sensitivity and specificity of cartilage imaging. In addition, quantitative information regarding cartilage physiology obtained on such MR images will be important clinically in the assessment of cartilage status.
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Footnotes
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D.R. supported in part by Veterans Administration grant SA-360
Address reprint requests to L.R.F.
Abbreviations: FOV = field of view
GRE = gradient-recalled-echo
RF = radio-frequency
TE = echo time
TR = repetition time
Author contributions: Guarantor of integrity of entire study, L.R.F.; study concepts, E.C.W., W.M.L.; study design, L.R.F.; definition of intellectual content, L.R.F.; literature research, L.R.F.; clinical studies, L.R.F.; experimental studies, L.R.F.; data acquisition and analysis, L.R.F.; statistical analysis, J.M.A.; manuscript preparation, L.R.F.; manuscript editing, J.M.A.; manuscript review, D.R.
Received April 21, 1998;
revision requested May 7, 1998; revision received June 11, 1998;
accepted August 24, 1998.
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TOP
Abstract
Introduction
