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. Author manuscript; available in PMC: 2014 May 1.
Published in final edited form as: Magn Reson Med. 2012 Jul 6;69(5):1486–1493. doi: 10.1002/mrm.24380

Increased Vessel Depiction of the Carotid Bifurcation with a Specialized 16-Channel Phased Array Coil at 3T

Quinn Tate 1,2, Seong-Eun Kim 2,3, Gerald Treiman 3,4, Dennis L Parker 2,3, J Rock Hadley 2,3
PMCID: PMC3556366  NIHMSID: NIHMS382507  PMID: 22777692

Abstract

The purpose of this work was to design and construct a multi-channel receive-only RF coil for 3 Tesla magnetic resonance imaging of the human carotid artery and bifurcation with optimized signal to noise ratio in the carotid vessels along the full extent of the neck. A neck phantom designed to match the anatomy of a subject with a neck representing the body habitus often seen in subjects with carotid arterial disease, was constructed. Sixteen circular coil elements were arranged on a semi-rigid fiberglass former that closely fit the shape of the phantom, resulting in a 16-channel bilateral phased array coil. Comparisons were made between this coil and a typical 4-channel carotid coil in a study of 10 carotid vessels in 5 healthy volunteers. The 16-channel carotid coil showed a 73% average improvement in signal to noise ratio (SNR) at the carotid bifurcation. This coil also maintained an SNR greater than the peak SNR of the 4-channel coil over a vessel length of 10 cm. The resulting increase in SNR improved vessel depiction of the carotid arteries over an extended field of view, and demonstrated better image quality for higher parallel imaging reduction factors compared to the 4-channel coil.

Keywords: MRI, carotid imaging, phased array, SNR

Introduction

Disease in the cervical carotid artery is a major cause of stroke and subsequent disability and mortality. With an annual estimate of more than 100,000 deaths per year, stroke is the third leading cause of death in the United States (1). Cardiovascular magnetic resonance imaging (MRI) is a safe and noninvasive method that adds crucial information to that obtained by conventional Doppler ultrasound. Various MRI techniques are used to accurately identify the vessel wall and characterize constituents of plaque (2–6). Unfortunately, the diagnostic utility of MRI in the examination of atherosclerotic disease is reduced because of artifacts, inadequate tissue contrast, reduced signal to noise ratio (SNR), and patient-dependent factors such as body habitus or inability to remain motionless for the length of the scan (7). Due to the very small (0.1 to 0.5 mm) size of some important details in carotid plaques, the images should be acquired with high spatial resolution (8). Unfortunately, there is a difficult trade-off between voxel size and image SNR. Imaging with small voxels (high spatial resolution) can require longer scan times in order to maintain adequate image SNR. Increasing MRI field strength from 1.5T to 3T has increased SNR for carotid plaque imaging, allowing improved spatial resolution, and/or reduced scan time (9).

To more fully utilize the available SNR, it is essential that the receiver coil be designed to cover the region where arterial disease is most likely to occur. In regions of low wall shear stress due to blood recirculation and stasis, such as at the carotid bifurcation, there is an increased chance of atheroma deposition (3,5,7,8,10). However, atherosclerotic plaques can also be found several centimeters in either the superior or inferior direction (5,11) as well as at the origins. Further, because of distinct inter- and intra-patient variations the location of the bifurcation can range from the lower ear to just above the clavicle at an average depth of 2–5 cm.

A significant increase in SNR can allow for shorter scan times, which could potentially reduce artifacts from vessel motion caused by respiration, swallowing and pulsatile blood flow. With sufficient SNR and appropriate RF coil geometries, parallel imaging can be used to reduce acquisition time and reduce sensitivity to motion artifacts (12). High spatial resolution and high SNR is essential for disambiguation of plaque components.

There are limitations in the coils that are currently used for carotid MRI. Neurovascular coils that are designed with a small number of large elements can provide coverage along the entire length of the carotid artery but achieve limited SNR (9). As explained by Wright et al. (13), increasing the number of receiver elements surrounding the volume can yield increased SNR near the coil elements as long as the coil elements remain sample noise dominated as opposed to coil noise dominated (14). The increased number of elements maintains coil sensitivity at greater depths (13). A 4-channel (4ch) bilateral 2-element phased-array (PA) surface coil developed initially by Hayes et al (15), has been considered by many to be the best available coil for carotid bifurcation imaging (9), (see Figure 1); however, the design has a limited Superior/Inferior (S/I) Field of View (FOV) and may require repositioning of the coil to center it over the disease location. Although slight modifications to this design have been made to extend the S/I FOV (9,16), and thereby. increase the vessel coverage, the coil still often requires repositioning for optimal placement over the bifurcation (9). A recently developed eight channel carotid coil (8ch) has been shown to increase SNR at the bifurcation (~70%) as well as increase the S/I FOV of the coil, when compared to the standard four channel coil (17). This coil was not available for comparison at the time of our work on the 16 channel coil, and the relative SNR and FOV improvements are discussed below.

Figure 1.

Figure 1

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Placement of carotid coils used at our institution. a) 16ch carotid coil on a semi-rigid close-fitting fiberglass former and b) 4ch coil with flexible two-paddle design.

The goal of this work was to develop a specialized receive only PA carotid coil that would improve the SNR, over that of our existing 4ch coil, and extend the region of coil sensitivity, particularly in the S/I direction. The result was a custom coil composed of 16 circular loop elements placed on a close fitting fiberglass former (16ch), see Figure 1. This work presents details on coil design and construction and compares imaging performance of our 16ch carotid coil with that of the standard 4ch receive-only RF carotid coil.

Methods

In this work, carotid coil comparisons were made using a homogenous phantom as well as results from in-vivo human carotid MRI studies. All studies were performed using a Siemens 3T TIM Trio MRI scanner (Siemens Medical Solutions, Erlangen, Germany). IRB approval was obtained for all studies described in this work and patients gave their informed consent.

Coil Construction

To achieve the goals of this work, it was decided that a close fitting semi-rigid fiberglass former would provide the best surface for mounting the individual coil elements. The coil former was designed to cover the neck from the lower ear to just above the clavicle, allowing coil element placement that could provide high signal sensitivity over the majority of potential carotid bifurcation positions.

Given the bifurcation depth (2–5 cm with an average depth of 3cm (9)), it was also decided to use circular loop elements of 4.5 cm diameter (14) for the 3T coil. This loop size is the smallest we have found to be sample noise dominated at 3T and still provides good SNR at the 2–4 cm depths. Improvements in SNR at greater depths (>3 cm) were expected due to the synthesis of the smaller loops to form larger loops as a result of the multi-channel reconstruction algorithms. Eight coil elements were placed on the former on each side of the neck (16 total elements) and were arranged in an overlapping geometric pattern, to minimize magnetic coupling of adjacent loops (18,19).

Each loop element was made from 18-gauge wire with 4 equally spaced capacitors (see Figure 2). Adjacent loops were overlapped for minimum magnetic coupling. Each coil element used a parallel capacitor and series inductor match circuit and was connected to its own low input impedance pre-amplifier by a 24 cm RG316 coaxial cable in series with a cable trap and phase shifter network (see Figure 2). The cable lengths were determined in order to keep the cable shields less than 1/10 of a wavelength at 3T frequencies and thereby reduce cable coupling and other antenna affects. The cable traps, constructed from 4 turns of a semi-rigid coax in parallel with a high voltage capacitor across the shield to realize a parallel resonance at our 3T frequency, were used to reduce common mode shield currents, to further reduce unwanted coupling between individual coil cables. When using the short cable lengths in conjunction with the cable traps on the RF receiver board, no unwanted cable coupling was observed. The electrical line length of the cable, cable trap, and phase shifter was 180o for optimal preamp decoupling of the PA loops during the receive portion of the pulse sequence, reducing magnetic coupling of non-adjacent loops in the array. The impedance of the trap at 123 MHz was measured to be about 6.7 kΩ. Active loop decoupling during the transmit portion of the pulse sequence was achieved using DC bias from the scanner to activate a shunt diode at the input of the match circuit. Each loop element was tuned and matched for optimal SNR. Variable capacitors were used to fine-adjust the tune and match of each loop. Return loss measurements were on the order of −30 dB; insertion loss between adjacent loops was approximately −11 to −13 dB. Active decoupling was between −28 and −35 dB, and preamp decoupling was approximately −18 to −20 dB. Each loop also included a passive decoupling circuit that reduced the loop resonant currents by approximately ~ −40 to −45 dB when activated.

Figure 2.

Figure 2

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Schematic of the circuitry used for each individual coil element.

Phantom Construction

As a template for coil former construction, a custom anatomically shaped fiberglass phantom was fabricated to represent the head, neck and shoulders of a volunteer with physical characteristics similar to those of many patients with carotid arterial disease imaged at our institution (see Figure 3a). For imaging studies, the phantom provided a homogenous volume through the sensitive region of the 16ch coil and offered a load similar to that of a human for the coil elements. The phantom was filled with a copper sulfate solution containing 1.955g CuSO4 5H2O per liter of water with table salt added to achieve a coil load similar to that of a human subject. When initial construction and phantom testing were completed the tune and match of each coil element were adjusted for an average human volunteer. This final adjustment was primarily to tune the coil, as the match was similar between the phantom and human volunteers.

Figure 3.

Figure 3

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a) Anatomical Head, neck and shoulders phantom. Custom fiberglass phantom filled with copper sulfate solution. b) Coronal, Axial, and Sagittal SNR maps of the homogeneous phantom. Black lines correspond to representative image slices through the bifurcation region where SNR line profile data was acquired. c) Left/Right SNR line profile plots from phantom SNR images through the expected location of a bifurcation. d) Superior/Inferior SNR line profile plots at 3 cm, 4.2 cm, and 5 cm depth from phantom neck surface.

Coil Comparisons

All carotid coil comparisons were made between the new 16ch carotid coil and a slightly modified version of a widely used 4ch receive PA coil design published by Hayes et al. (15). A limited comparison of these coils to a commercially available 19-channel head/neck coil (19ch) was also included. Although the 19ch coil has a large S/I FOV, fits almost any patient, and is often used for clinical imaging of the carotid arteries, the SNR provided at the bifurcation is relatively low compared to the 16ch and 4ch coils. The 4ch coil has been shown to provide better SNR at the bifurcation than many commercial coils used for carotid imaging (9) and was used as the reference coil for this study. The 19ch coil is included in this work only to show the benefits of coils designed for the specific purposes of imaging the bifurcation over currently available commercial general-purpose neck coils. The 4ch coil used in this study was comprised of two bilateral paddles, each consisting of two overlapped rectangular loops (e.g. 6.5×4.5 cm (9).), and was a slight modification from the Hayes 4ch design (4.5×4.5cm)(15) in order to increase coil coverage along the length of the carotid vessels and reduce the need for coil repositioning to obtain adequate images of the bifurcation. Figure 1b displays the modified 4ch used in this study. Note that the 8ch coil recently published by Balu et al. was not available for comparison at the time of this study. However, relative performance can be compared.

Coil comparisons included phantom studies to assess relative SNR, sensitivity profiles, and g-factors for parallel imaging potential. Here, the term ‘relative’ is used to imply that all other factors, except variations in the coils, were constant and that SNR results are only functions of the coil types compared. The noise used to calculate SNR was obtained by hand-drawn ROI’s in areas of the image with no signal or artifact. Human studies were performed to assess clinical imaging using the 16ch coil and to evaluate arterial SNR profiles and parallel imaging performance.

Phantom Studies

SNR of the 16ch and 4ch coils was compared using the custom phantom. Images from this study produced SNR maps (20) that could then be compared at specific depths. Using volunteers from this study, the range of bifurcation depths was 3 to 5 cm with an average of 4.2cm. Therefore, axial, coronal and sagittal slices were selected from stacks of 2D phantom images that most closely corresponded to these depths. Because the depth of the carotid artery in a human neck varies significantly along the length of the neck (9), the corresponding SNR line plots from a single sagittal plane resulted in a depth ranging from 4.2 cm to 7.5 cm along an S/I length of 10 cm centered at the middle of the neck.

As secondary outcome, the phantom studies were also used to assess the parallel imaging capabilities of these coils (21). Inverse g-factor maps for phase encoding in the AP direction were determined.

Human in-vivo Studies

For the in vivo carotid artery SNR profile comparisons, 5 healthy volunteers with neck diameters varying between 12 – 14.5 cm (weights between 68 to 105 kg) and bifurcation depths from 3.4 to 5.2 cm were imaged. In order to compare SNR profiles of the 16ch and 4ch coils along the axis of the carotid vessels over the full extent of the neck, a 2D TOF sequence was used with the following parameters: 60 slices, TE/TR = 5.5 ms/27 ms, FOV = 30×30 cm2, Matrix = 512×512, Flip Angle = 50°, and slice thickness = 5 mm. Imaging studies for each volunteer were performed on the same day using both the 16ch and 4ch coils. The SNR profiles from 10 carotid arteries were compared over a volume with an S/I length of 30 cm. For an interesting reference comparison, one volunteer was also scanned with the 19ch and body coils, and the average SNR for the two carotid arteries was measured.

The SNR maps were calculated for each individual channel throughout the volume and were then combined using the square root of the sum of squares algorithm (18). For each resulting SNR image slice, ROI’s were drawn around the internal carotid, bifurcation, or common carotid artery and the signal intensities for all voxels within 60% of the peak value inside the lumen were averaged. The average axial vessel SNR values obtained at each 5 mm increment along the artery were plotted for the 19ch, 4ch and 16ch coils. The data provided vessel SNR values from the circle of Willis along the internal carotid artery, through the bifurcation and along the common carotid artery to the aortic arch. The relative SNR profile for each subject was spatially aligned so that the bifurcation of each profile was centered at 0 cm and the mean and standard deviation of the SNR within each group at each location was calculated.

For comparisons of image quality, several clinical patients were scanned with a 2D TSE sequence to produce a volume of 2D T2 weighted images using GRAPPA (22) with a reduction factor of 2 (R = 2) and a 3D TSE sequence for high resolution MRI. The sequence parameters were as follows: 24 axial slices, TE/TR = 64 ms/3500 ms, ETL = 11, FOV = 13×13 cm2, matrix = 256×256, slice thickness = 2 mm, with four averages. To compare the quality of high-resolution images, a 3D T1w TSE sequence was used with 32 slices, TE/TR = 23 ms/700 ms, ETL = 37, Echo Spacing = 5.2ms, FOV = 14×14 cm2, matrix = 640×640, slice thickness = 1.0 mm, and voxel size = 0.22×0.22×1.0 mm3. The same imaging parameters were used for scans performed with the 4ch and the 16ch coils.

Results

Phantom Studies

Figure 3b displays SNR images acquired from the homogenous phantom with the associated line plots of the SNR profiles. These phantom studies demonstrated a significant increase in SNR, from the 16ch over the 4ch, at depths close to the coil surface. The percent improvement decreased with increasing distance from the coil. The plots in Figure 3c and d contain SNR data from both a transverse and longitudinal cross section of the coronal image from Figure 3b. These plots show there is a significant improvement in the S/I SNR profile of the 16ch coil over the 4ch coil and with increasing depth. For example, relative to the peak SNR of the 4ch, the 16ch coil showed an improved SNR over a simulated vessel length of 10 cm.

The inverse g-factor map comparisons for the axial plane of the two carotid coils can be seen in Figure 4. While these plots do not show significant advantages of the 16ch coil over the 4ch coil in the axial plane, the 16ch coil has the same parallel imaging performance along the length of the carotid artery covered by the coil elements. The g-factor maps for the 16h coil are improved for other slice orientations, but those orientations are not typical in non-contrast enhanced carotid imaging due to blood flow directions.

Figure 4.

Figure 4

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Inverse g-factor maps obtained from phantom images. These maps (phase encoding in the Anterior/Posterior direction) demonstrate improvements in parallel imaging of the 16ch over the 4ch coil for non-contrast enhanced imaging of the carotid vessels.

In-vivo Human Studies

Figure 5 compares the vessel SNR along the axis of the carotid vessels, from the Circle of Willis to the aortic arch, between the 4ch and 16ch coils. The mean relative SNR profile for each coil is shown with the associated standard deviation error bars. Note that while the lower limit of the standard deviation of the relative SNR from the 16ch coil extends below the upper limit of the standard deviation of the corresponding SNR of the 4ch coil, the 4ch never outperformed the 16ch coil in any given study. The lower SNR limits for both coils were obtained when imaging a subject with deeper arteries and the upper SNR limits were obtained while imaging shallow arteries. The trends observed in the phantom SNR studies were also observed in the human studies.

Figure 5.

Figure 5

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Average Axial rSNR profile plot through the carotid vessels from the aortic arch to the circle of Willis. Peak 4ch rSNR was normalized to 100 in order to show percent improvement of 16ch SNR. Error bars show the standard deviation of rSNR for the group of 10 carotid vessels (two vessels from each of 5 volunteers). Individual SNR profiles for each vessel were aligned at the location of their bifurcation.

Also plotted in Figure 5 are the relative SNR measurements obtained using the 19ch commercial head/neck coil and the MRI scanner body RF coil. Although the 19ch coil provides a large coverage of the vessels from the arch to the circle of Willis, the available SNR at the bifurcations is significantly less than that provided be either the 4ch or 16ch coil. For these studies the 4ch coil was positioned for best SNR over the bifurcation (9). The 16ch coil provided comparable coverage to the 19ch coil while significantly improving the SNR in the region of the bifurcation over even the 4ch coil.

There were also improvements in parallel imaging with the 16ch compared to the 4ch coil. Figure 6 displays the 2D T2w images of a clinical patient volunteer that had visible disease. These images were acquired with the 4ch and 16ch coils while using a parallel imaging reduction factor of two (R = 2). For R = 2, the 4ch shows a noticeable reduction in image quality compared to the 16ch coil at the location of the carotid vessels. Inspection of the 16ch images in the regions closer to the skin surface, near the bifurcation show a noticeable improvement in vessel wall definition, both in relative SNR and contrast, over the images acquired using the 4ch coil. Although, changing coils resulted in different compression of the neck and required slice repositioning, nine contiguous slices are provided for each coil to improve visual comparison.

Figure 6.

Figure 6

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Clinical image of a patient volunteer obtained with the carotid coils used in this study, demonstrating the appearance of disease. Images were acquired with a 2D GRAPPA (R=2) T2 weighted TSE sequence. a) Image series showing comparison of 4ch and 16ch coils with a patient who has noticeable arterial disease. b) Enlarged images of the same series showing improvement in image clarity.

The increase in SNR of the 16ch coil can also be used to obtain increased image resolution while maintaining and possibly improving image quality. Figure 7 displays 3D high-resolution T1w images acquired using the 16ch and 4ch coils. Inspection of these images shows that in 3D high resolution imaging (acquired at 0.5×0.5 mm2 in plane resolution), the overall image quality, including reduced background noise and increased image clarity, is improved with the 16ch coil.

Figure 7.

Figure 7

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High resolution images (acquired at 0.5 × 0.5 mm2 resolution and interpolated to 0.25×0.25 mm2) comparing the a) 16ch and b) 4ch coils. These images show the improved image quality and reduced image artifact that is obtained using the 16ch coil compared to the 4ch coil.

Discussion

The 16ch coil results show a significant improvement in the vessel SNR profiles over the 4ch coil along the entire length of the carotid vessel. From comparisons at the bifurcation, where there is often significant signal loss due to disordered non-repetitive blood flow, there is an average improvement of 73% in SNR using the 16ch coil. The analysis of the data acquired using a phantom shows that the 16ch provides an appreciable gain in SNR with increasing depth compared to the 4ch coil. On average, nearly twice (1.98) the available SNR was received over a range of eight centimeters centered at the carotid bifurcation. The SNR improvement achieved by the 16ch over the peak SNR of the 4ch extends for 10 cm along the length of the vessel. This demonstrates that the 16ch has a much greater likelihood of being sensitive to regions containing the carotid bifurcation, reducing the need for coil repositioning that is more frequently required when using the 4ch coil. We note that these comparisons were made with a modified version of the original 4ch coil design (15). The differences in SNR between the original 4ch design (15) and the 4ch coil used in these experiments are likely small but have not been quantified.

The 16ch coil also shows a significantly greater SNR than the Siemens 19ch coil over about a 20cm extent centered on the bifurcation. The 19ch coil, which was designed for general purpose neck imaging, is composed of 12 elements around the head and several other elements, not in close proximity to the neck. Although 19ch is not a contender as an optimal coil for high resolution imaging of the carotid bifurcation, as seen in our comparison plots, it was added as an interesting reference comparison because this is a coil that many institutions, including our own, use for clinical imaging of the bifurcation.

Even though the 16ch coil former was not optimized for all volunteer neck sizes or for all potential anatomical locations of the bifurcation, the coil performed well for the wide variety of volunteer and patient anatomy involved in these studies. The 16ch coil was not repositioned relative to the subject for any of these studies. However, in three of the five normal volunteers it was necessary to reposition the 4ch coil for the bifurcation to be within the sensitive region of the coil. To accommodate a larger percentage of possible patients, more than one former size may be needed. Other options include the division of the fiberglass coil former into two pieces (8 coils each), allowing the coil elements to fit a wider range of neck sizes.

A benefit of the 4ch coil is that it has the flexibility to be placed virtually anywhere, regardless of patient anatomy. For cases of unusual bifurcation position, the 16ch coil is limited in that the semi-rigid former may not allow the coil elements to be placed in a specific location. In these instances, the 4ch coil might outperform the 16ch coil in relative SNR. Maximum intensity projection (MIP) images for one volunteer with a relatively high bifurcation, shown in Figure 8, demonstrate the ability of the 4ch coil to be repositioned over the jawbone and the unique ability of the 16ch coil to obtain images over the bifurcation with comparable SNR without repositioning. Note that, in this study, the 4ch coil was repositioned over the carotid bifurcation. When there is disease away from the bifurcation, the 4ch coil without repositioning is much less likely to detect the disease than the 16ch coil.

Figure 8.

Figure 8

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Magnitude MIP images for a single volunteer with a bifurcation at the level of the jaw. Images were obtained using the body, 19ch, 4ch, and 16ch coils. The images all have roughly the same background noise level, and have been scaled by the same factor such that the maximum value in the 4-channel coil is 100.

A limitation of the current 16ch design is the cumbersome pre-amp housing and assembly that is in close proximity to the head of the patient. Future work will include the design of a more ergonomic and patient friendly design allowing for closer proximity of the pre-amps, circuit boards and associated RF circuitry. For example, the preamps might be positioned inside the head support, with short cables connecting the loop elements to the RF circuitry.

The increase in SNR allows for improved parallel imaging performance along the length of the carotid vessels in the neck, as well as high definition imaging with improved vessel definition. Parallel imaging benefits of the 16ch coil were not extensively tested in this work because the coil elements were positioned to increase SNR for conventional (non-parallel) imaging. However, because of the extended positioning of coil elements, the 16ch coil can perform parallel imaging over a much greater S/I extent than the 4ch coil. Future work could investigate the clinical tradeoffs between increased SNR and image acquisition speed as well as other benefits of using parallel imaging as part of carotid artery MRI protocols. The high-resolution images (Figure 7) may allow for a more in depth study of plaque composition. The scanner used for these studies is equipped with a total of 32 channels that could be used to further extend the coverage toward the circle of Willis and the Aortic arch.

Finally, a comparison with the recently published 8ch coil (17) was difficult because the S/I FOV experiments for the 8ch coil were done along the length of a cylindrical phantom at constant depth rather than in actual vessels that vary in depth along the length of the vessel as performed in this work. The 8ch coil SNR results were reported as improvements in vessel wall SNR from black blood images. In this work SNR results are reported based upon blood flow signal from Time-of-Flight (TOF) images.

Conclusion

A specialized 16ch PA coil for imaging the carotid bifurcation was designed to increase SNR and to extend coverage of the carotid artery as compared to an in-house-developed 4ch PA carotid coil. Results from Time of Flight white blood imaging showed that the 16ch coil significantly surpassed the 4ch coil in both vessel coverage and SNR. The 16ch coil showed an average improvement of more than 73% in SNR at the bifurcation while the average improvement ± 4cm distance from the carotid bifurcation was greater than 100% over the 4ch coil. The 16ch coil maintained a vessel SNR improvement over the 4ch along the entire length of the carotid artery from the Circle of Willis to the aortic arch, when the bifurcation was located near the center of the neck, and demonstrated substantial improvements in parallel and high resolution imaging when compared with the 4ch coil. Comparisons against a commercially available 19-channel coil designed for MRI of the neck showed about a 400% improvement in SNR around the bifurcation and the SNR of the 16-channel coil remained greater than the 19-channel coil over a distance of about 20cm centered on the bifurcation.

Acknowledgments

We acknowledge helpful contributions from Sathya Vijayakumar, John Roberts, Melody Johnson, Henry Buswell and Anne Haroldsen in the Utah Center for Advanced Imaging Research.

This work has been supported by the Ben B. & Iris M. Margolis Foundation, a Clinical Merit Review Grant from the Veterans Administration Health Care System, NIH R01-HL057990, and Siemens Medical Solutions.

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