Plaque Characterization With Optical Coherence Tomography
Debra Stamper, PhD*,
,
Neil J. Weissman, MD, FACC
and
Mark Brezinski, MD, PhD*,,a,*
* Department of Orthopedic Surgery, Brigham & Women’s Hospital, Boston, Massachusetts
Harvard Medical School, Boston, Massachusetts
Cardiovascular Research Institute, Washington Hospital Center, Washington, DC
Manuscript received August 18, 2005;
revised manuscript received September 12, 2005,
accepted October 17, 2005.
* Reprint requests and correspondence: Dr. Mark Brezinski, Orthopedics Research, MRB105, Brigham & Women’s Hospital. 75 Francis Street, Boston, Massachusetts 02115 (Email: mebrezin{at}mit.edu).
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Abstract
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The identification of unstable plaque is central in risk-stratifying patients for acute coronary events. Optical coherence tomography (OCT) is a recently introduced imaging modality that has shown considerable promise for the identification of high-risk plaques. Advantages of OCT include its high resolution (4 to 20 µm), high data acquisition rate, small and inexpensive guidewires/catheters, and ability to be combined with adjuvant optical techniques. This article summarizes the current state of intravascular OCT imaging, focusing on potential markers of instability and current limitations.
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Abbreviations and Acronyms
| | ACS = acute coronary syndrome | | AMI = acute myocardial infarction | | IV = intravenous | | IVUS = intravascular ultrasound | | OCT = ocular coherence tomography | | PS-OCT = polarization-sensitive OCT | | TCFA = thin-capped fibroatheromas |
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Optical coherence tomography (OCT) is a recently developed method of catheter-based, high-resolution intravascular imaging that has demonstrated considerable potential as a method for assessing unstable plaque (1). It is analogous to ultrasound, measuring the backreflection of infrared light rather than sound. The advantages of OCT include its resolution, which is higher than any currently available imaging technology. Resolutions of catheter-based systems are in the range of 10 to 20 µm. Furthermore, resolutions as high as 4 µm have been achieved with more sophisticated systems that might eventually be applied to catheter-based approaches (2). Second, acquisition rates are near video speed, an advantage relative to many other technologies for assessing plaque (3). Third, unlike ultrasound, the typical OCT catheters contain no transducers within their frame, which makes them both small and inexpensive. An example of the current smallest catheter is the 0.017-inch OCT imaging guidewire from Lightlab (Westford, Massachusetts) used in our laboratory and shown in Figure 1B (4). Fourth, OCT systems are compact and portable, with the Lightlab system shown in Figure 1A. Finally, because OCT uses light, a variety of spectroscopic techniques are available, including polarization spectroscopy, absorption spectroscopy, elastography, OCT Doppler, and dispersion analysis.
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Figure 1 (A) Photograph of an ocular coherence tomography (OCT) device. This is a photograph of the Lightlab (Westford, Massachusetts) imaging engine. (B) Photograph of an OCT catheter. This figure is a photograph of a 0.017-inch OCT imaging catheter/guidewire consisting of relatively simple fiber optics. This includes a gradient index (GRIN) lens, single mode optical fiber, prism, and speedometer cable. (C) Schematic of a traditional OCT system. Ocular coherence tomography works via the technique of low coherence interferometry as described in the text.
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Theory
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With OCT, ultrashort light pulses or low coherent light is directed at the sample (1,5). The time for the light to be reflected back, or the echo delay time, is used to measure distance. The intensity of backreflection is plotted as a function of depth. The beam is then scanned across the sample to produce two- and three-dimensional data sets; however, unlike ultrasound, the echo delay time cannot be measured electronically owing to the high speed associated with the propagation of light. Therefore, a technique known as low coherence interferometry is used.
In one embodiment of low coherence interferometry, light from the source is split by a beam splitter as demonstrated in Figure 1C. One-half of the light is directed at the sample and one-half at a moving mirror. Light reflects from within the sample and off the mirror. If light in both arms travels the same optical distance, interference will occur when the light from each arm is recombined. Optical coherence tomography measures the intensity of interference and uses it to represent backreflection intensity.
The axial resolution of OCT is dependent on the bandwidth of the source or the range of wavelengths within the beam (1,5). For illustrative purposes, the beam can be viewed as consisting of a series of pulses. Owing to a mathematical relationship known as a Fourier transform, the broader the bandwidth, the shorter the "pulse duration." The duration of the "pulse" is critical in determining resolution. With short durations, for interference to occur, the distance traveled by light in each arm must be matched more precisely. Therefore, when the mirror is fixed in one position, light that reflects off the reference arm mirror will only interfere with light that has traveled the same distance in the sample arm to within the width of the pulse; the smaller the "pulse" duration, the smaller the region of backscattering intensity that is measured. This is high-resolution ranging. By moving the mirror, backreflection intensity can be obtained within different depths of the sample.
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Early in vitro studies
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In the mid-1990s, we postulated that the high resolution of OCT made it a potentially useful technology for imaging atherosclerotic plaque; however, significant technological limitations existed at this time that eventually needed to be overcome, including its poor penetration in nontransparent tissue, the lack of an OCT imaging catheter, slow data acquisition rate, high noise reduction in the system electronics, and unsynchronized system components (3,6). Although initial studies had only limited success, we ultimately successfully performed micron scale imaging of both in vitro human aorta and coronary arteries (1,7). The morphology of atherosclerotic plaque was accurately determined by OCT when compared with the corresponding histopathology. The imaging was performed at 16 ± 1 µm, 10x higher than high-frequency ultrasound. In Figure 2, an OCT image of a heavily calcified atherosclerotic plaque is seen with the corresponding histopathology (1); a thin layer of intima <50 µm in diameter, which is identified by the arrow, is overlying the plaque. In top image of Figure 3, a collection of lipid is seen in an atherosclerotic plaque with a thin intimal cap <50 µm in diameter over the region (arrow) (1). The ability to identify small plaque microstructure, such as thin intimal caps, might be a powerful tool for patient risk stratification. In the bottom portion of the same figure, fissuring is occurring to the intimal-medial border, another indication of instability.
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Figure 2 Thin intimal cap. (A) The OCT image. (B) The histopathology normal, native, in vitro human artery. The plaque is on the left, the relatively normal tissue is on the right. Bar is 500 µm. The most important feature is the shown by the arrow, where the intima is <50 µm in diameter. Reprinted, with permission, from Brezinski et al. (1).
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Figure 3 Plaque instability. The image in panel A shows a large collection of lipid, shown by the dark area, with a thin layer of intima above it from a native in vitro coronary artery. The arrow shows a small area of intima <50 µm in diameter that is susceptible to rupture. The image in panel B shows another unstable plaque with fissuring to the intimal medial border (arrows). Bar is 500 µm. Reprinted, with permission, from Brezinski et al. (1).
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Both OCT and intravascular ultrasound (IVUS) have undergone direct comparisons qualitatively and quantitatively with histology as the gold standard (8,9). The axial resolution of both OCT and IVUS (30 MHz) has been measured directly from the point-spread function, with the axial resolution of OCT being 16 ± 1 µm compared with 110 ± 7 µm for IVUS. In qualitative comparisons, OCT consistently demonstrated superior delineation of structural detail. In Figure 4, in vitro imaging with OCT and IVUS (30 MHz) is performed of an aortic atherosclerotic plaque (8). In the IVUS image (top), the presence of the plaque is suggested, but no other structural detail is apparent. The OCT image (middle) not only demonstrates the large plaque, but also a distinct layer is evident within the intima that appeared normal by ultrasound. By histopathology (bottom) the layer is confirmed to represent smooth muscle proliferation. Similar comparisons between IVUS (30 MHz) and OCT have been performed with a 2.9-F OCT catheter. Figure 5 is a comparison between OCT and IVUS of an in vitro coronary artery (6). The intimal hyperplasia is clearly seen with OCT but not with IVUS.
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Figure 4 Intravascular ultrasound (IVUS) versus ocular coherence tomography (OCT) comparison (smooth muscle proliferation). (A) An IVUS image (30 MHz) of an in vitro aortic plaque. In the OCT image (B), in addition to the plaque on the left, a layer of smooth muscle proliferation is seen (red arrow). The smooth muscle proliferation is confirmed by the histopathology (C). The bar in the OCT image is 500 µm. Reprinted, with permission, from Brezinski et al. (8).
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Figure 5 Intravascular ultrasound (IVUS) versus ocular coherence tomography (OCT) comparison (intimal hyperplasia). This image demonstrates an OCT image (A) and IVUS image (30 MHz) (B). Both images were generated with 2.9-F catheters. The IVUS catheters were not reused. In the OCT image the intimal hyperplasia is seen, whereas in the IVUS image, it is difficult to detect. The bar in the OCT image is 500 µm. Reprinted, with permission, from Tearney et al. (6).
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With optical technologies, the ability to image completely through the plaque is a concern. In Figure 6, near occlusive coronary plaques are imaged with both OCT and IVUS, demonstrating the penetration of OCT. The OCT images are the two images on the left of the figure. Both the ability of OCT to image through the width of the plaque and the superior resolution compared with IVUS (40 MHz) are noted (9). Quantitative measurements from OCT were also superior to IVUS.
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Figure 6 Intravascular ultrasound (IVUS) (40 MHz) versus ocular coherence tomography (OCT) comparison. The images in A and C are OCT images, whereas the images in B and D are IVUS (40 MHz) of a native in vitro coronary artery. The IVUS catheters were not reused. In the OCT images, not only is structure better defined but OCT can penetrate completely through the near occlusive vessel to the media. Limitations of OCT penetration depth described in detail in the text. Reprinted (modified), with permission, from Patwari et al. (9).
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Experimental in vivo studies
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In vivo OCT imaging was initially performed of aortas in New Zealand White rabbits before human studies (10). Because light-scattering occurs from red blood cells (discussed in more detail in following paragraphs), saline flushes were required during imaging at a rate of 2 to 3 cc/s. In Figure 7, an OCT image of a normal aorta is seen. Because the vessel does not contain atherosclerosis, the amount of structural detail is minimal but is outlined in the figure legend. Data was obtained at 4 frames/s and saved on super VHS format (state-of-the-art systems now preserve the data digitally). Data was also performed in vivo comparing the ability of IVUS (40 MHz) and OCT to assess stent approximation in the rabbit, where OCT was determined to be superior (11).
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Figure 7 In vivo imaging of rabbit aorta—at 4 frames/s through a 2.9-F ocular coherence tomography (OCT) imaging catheter. The OCT image is in panel A, the histology is in panel B. The OCT image was performed in the presence of saline flush (2 to 4 cc/s). Arrow indicates clot. A = adventitia; M = media; V = vena cava. Reprinted, with permission, from Fujimoto et al. (10).
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In vivo human studies are currently underway by Lightlab Imaging in addition to several other groups (12–14). These studies have confirmed the ability of OCT to characterize plaque. In Figure 8, contributed by LightLab Imaging, an intimal tear is noted in vivo in humans with a 0.017-inch imaging guidewire in the presence of a saline flush (14). In Figure 9, provided by the same group, neointimal hyperplasia is noted in the OCT image on the right but is poorly defined on the IVUS image on the left. Additional studies have included the assessment of stent placement and estimation of macrophage concentration (12,13).
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Figure 8 Presence of an in vivo intimal flap. In this patient, an intimal flap is identified with the ocular coherence tomography imaging guidewire. Reprinted, with permission, from LightLab and Prof. Grube et al. (15). A demonstrates the entire artery while B is an enlargement of the square in A.
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Figure 9 In-stent restenosis. In this patient, in-stent restenosis is noted as imaged with the 0.014-inch imaging guidewire. IVUS = intravascular ultrasound; OCT = ocular coherence tomography. Reprinted, with permission, from LightLab and Prof. Grube et al. (15).
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Limitations.
Major limitations of OCT are its attenuation by blood and its limited penetration in tissue. Blood significantly attenuates the ability of OCT to image through blood. Approaches to overcome this limitation are saline flushes, balloon occlusion, and index matching. Saline flushes have been used successfully both in animals and in patients in vivo (10,12), although several concerns exist over the use of these flushes. These include fluid overloading the patient, leading to pulmonary edema, and whether the field can be cleared completely in all clinical scenarios. The use of occlusion balloons has been performed by Lightlab Imaging (15). Future studies are required to assess the ultimate viability of this approach.
We have proposed an alternative technique for imaging through blood: index matching (16). We hypothesized that the attenuation from blood was produced by scattering due to a refractive index mismatch between the cytoplasm of the red blood cell and serum (16). The hypothesis and technique for reducing scattering were examined in vitro. Imaging is performed of the reflector and intensity is used to represent penetration. In the top image of Figure 10, saline is circulated over the reflector. In the second image, the reflector can no longer be seen because blood is circulated over it. In the bottom image, the blood has been removed, lysed, and reintroduced for circulation. After lysis, penetration returned to a value 95% of that of the saline control, suggesting that the reduced penetration was not due to either absorption or the presence of cell membranes. This is consistent with the hypothesis that the decreased penetration through blood was due to the refractive index mismatch.
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Figure 10 Qualitative proof of mismatch as the source of scattering. In the top image, saline is being circulated over a reflector. The top line is the inner layer of the tubing. The second line is the reflector. Saline is between them. In the middle image, blood is being circulated rather than saline through the system. The reflector is no longer seen. In the bottom image, blood has been removed, lysed, and reintroduced into the system. The signal off the reflector returned to values not significantly different than the control subject. Reprinted, with permission, from Circulation 2001;103:2000–4.
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We further postulated that by increasing the refractive index of the serum near that of the cytoplasm, penetration through blood could be substantially improved. In Figure 11A, the effects of increasing the serum refractive index with dextran and intravenous (IV) contrast are noted. Dextran resulted in a 69 ± 12% increase in penetration relative to the saline control, whereas IV contrast resulted in a 45 ± 4% increase. Both effects were highly significant (p < 0.005).
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Figure 11 (A) Increased penetration with index matching. In this figure, the difference in penetration after addition of compounds or control is seen. The saline control had no significant effect (7 ± 3%). Dextran resulted in a 69 ± 12% increase in penetration. Intravenous (IV) contrast resulted in a 45 ± 4% increase. Both were significantly different than the control. (B) Influence of compounds on hematocrit and red cell counts. There was no significant difference between the effects of dextran and saline on either hematocrit and red cell count. Therefore, the improved penetration was not due to either a change in cell size or number. There was a slight but significant difference in red cell size after IV contrast compared with saline control. This might have contributed slightly to the effect. Reprinted, with permission, from Circulation 2001;103:2000–4.
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To demonstrate that the effect was not due to changes in red cell volume or number, hematocrits and red cell counts were measured. Neither dextran nor saline resulted in a significant difference with respect to either of these parameters; however, as seen in Figure 11B, addition of IV contrast significantly decreased red blood cell volume (as compared with saline). This change in volume might have contributed to some of the increased penetration observed with the IV contrast. New compounds continue to be examined to find superior performance.
A second potential limitation of OCT is that its penetration through the arterial wall is in the range of 2 to 3 mm. Although this is generally sufficient for imaging of most arteries, with some large necrotic cores, the entire length of the core can not be imaged (7,9) Several approaches to overcome this are proposed, which might be of interest to those with a technical background. These approaches include increase of the dynamic range, improved choices in median wavelength, use of a parallel ultrasound beam, and image processing techniques.
Among the approaches to increase dynamic range, and therefore penetration, include spectral radar and source swept OCT, which are discussed elsewhere (11,17–19).
A water absorption peak exists approximately at 1,380 nm that might attenuate penetration. Most groups are using sources with median wavelengths above 1,300 nm. There might be some benefit to imaging below 1,300 nm (1,250 to 1,280 nm), owing to reduced water absorption.
Multiple scattering reduces resolution and therefore image penetration through a reduction in contrast. By the use of a parallel ultrasound beam, it has been shown that multiple scattering can be substantially reduced and ultimately might allow deeper imaging in tissue (20). This approach is still under investigation.
We have used several image processing techniques to improve the identification of deeper structures within tissue. One example that has been used by our group is the use of the sticks technique that has been shown to enhance imaging in coronary arteries (21).
Current state of the intravascular OCT imaging.
Current efforts with OCT focus on the identification of markers of plaque vulnerability to improve patient risk stratification and technology development. Work focusing on overcoming the attenuation from blood and increasing penetration depth has been addressed in the previous section.
Current OCT markers for in vivo studies.
Because OCT is now being introduced for in vivo human imaging at a resolution higher than any current imaging technology, allowing the identification of thin-capped fibroatheromas (TCFA), studies are underway to further define OCT plaque markers that aide in patient risk stratification. The reason is that most acute coronary syndromes (ACSs) result from TCFA, but most TCFA do not lead to ACS, resulting in a need to further risk stratify TCFA (22–29). In one study, in up to 8% of patients dying of noncardiovascular causes, ruptured plaques were noted in the coronary arteries (25). In another IVUS study, plaque rupture of target lesions occurred in 80 acute myocardial infarction (AMI) patients (66%) and in 31 stable angina (SA) patients (27%) (p < 0.001). Non–infarct-related or non-target artery plaque ruptures occurred in 21 AMI patients (17%) and 6 SAP patients (5%) (p = 0.008). Multiple plaque ruptures were observed in 24 AMI patients (20%) and 7 SAP patients (31%) (p = 0.004). Therefore, at least one plaque rupture in any coronary artery was noted in 84 AMI patients (69%) and 35 SAP patients (31%) (p = 0.001) (27). Whereas OCT can identify plaques with intimal caps <70 µm (i.e., TFCA) and small lipid collections, representing a substantial advance over other imaging modalities, a need exists for OCT advances to further risk stratify these plaques beyond just the identification of TCFA. Therefore, expanding OCT markers of coronary pathology remains critical.
The presumed general determinates of a plaque’s vulnerability to rupture are the size/composition of the atheromatous hemorrhagic core (when the core is more than 40% of the plaque volume), the thickness of the fibrous cap covering the core, the collagen content/type of the cap, angiogenesis into the intima, and the degree of inflammation within the cap. In current in vivo OCT human studies, there are currently three OCT markers being used to risk-stratify plaques in humans: cap thickness, cap-intimal borders, and granularity suggestive of macrophage/foam cell content. The first is an accurate measurement of intimal cap thickness, which goes largely undisputed and has been demonstrated by our group and others.
The second marker is the ability of OCT to identify plaque macrophages (13,30). This is an intriguing concept and is an important marker to be evaluated in future studies; however, there is only one study that attempts to quantitate this, and because of several of the following concerns, confirmation of these results seems appropriate. First, the authors neither used a "training set" nor predetermined criteria for what constituted a positive or negative predictive value in the OCT image but determined the criteria after examining the measured data set (30). They established the value for raw OCT data between 6.15% and 6.35%, which gave them a 100% sensitivity and specificity. A different statistical approach might be useful to examine. It is of further note that they did not appear to use these parameters for interpreting a subsequent in vivo study, which did not have histological validation (13). Second, they performed a median filter over a 3 x 3 square kernel that corresponds to an effective axial resolution for macrophages worse than 30 µm. Therefore, the macrophages would either need to be densely packed or >30 µm in diameter. Third, they used the mean background noise for their calculation. The problem with this is best illustrated with a simulation, where we take the case of two different background variations. In the first case, the background has low pixel variations, between 1 and 3. In the second case, the pixels vary from 1 to 9. If we make the concentration of macrophages on both images the same and macrophages have a constant value of 50, the normalized standard deviation will be 50.31% on the first image and 46.77% on the second, which means that the concentration of macrophages on image 1 is higher than on image 2, and this is not true. Finally, when the same group applied the approach in vivo, there was minimal difference between the culprit plaque and remote regions (5.54 ± 1.48 vs. 5.38 ± 1.56) and in culprit lesions in unstable versus stable plaque (5.91 ± 2.06 vs. 4.21 ± 1.74). Particularly when no training set has been established with histologic correlations and the concern over variation due to background noise, it is unclear how these small variations will be diagnostic when spread over a large population. Furthermore, as discussed previously, it is unclear whether macrophages concentrations increase in the high-risk plaque before or after rupture (31). Although the work on OCT assessments of macrophages remains intriguing, future work in this area is needed to confirm the results.
The ability of OCT to distinguish: 1) fibrous, 2) fibrocalcific, and 3) lipid laden plaque by the backreflected signal—the latter being those plaques of interest—is the third marker being currently used in clinical studies (32). In a recent study, the three plaques were respectively differentiated as: 1) homogeneous, signal-rich region, well-delineated; 2) signal-poor region with sharp borders; and 3) signal-poor region with diffuse borders. Therefore, fibrocalcific and lipid laden plaques are distinguished by the sharp versus diffuse borders; however, this study was associated with a high false positive rate, with fibrous caps being diagnosed as lipid laden plaques (sensitivity 71% to 79%). The difficulty in identifying plaque by the diffuse nature of the plaque border is exemplified by Figure 12, from our initial paper published in Circulation on plaque rupture (1). In this figure, the yellow arrows demonstrate a cap-lipid interface that is diffuse and is covered by a cap with high light scattering (yellow reflections). The white arrow, which is also identifies a region over lipid, identifies a cap-lipid interface that is sharply defined, but the cap is associated with lower scattering. Finally, the black arrow shows an intimal cap that is highly scattering over the intimal-elastic interface (no lipid present). The interface is poorly defined in spite of lack of lipid. The diffuse nature of the intimal-elastin border appears to be that multiple scattering by the cap that is obscuring the border (which would also explain a more rapid local exponential decay). It is unclear, therefore, whether the lipid interface is diffuse in some OCT images because of the core composition, as previously suggested, or multiple scattering from the cap, the latter of which could reduce its predictive power. Future studies are required to establish the mechanism as to why some plaques have diffuse cap/core interfaces.
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Figure 12 Examining intima boundary sharpness as a marker for lipid. In this figure, the yellow arrows demonstrate the cap-lipid interface that is diffuse and is covered by a highly scattering cap (yellow reflections). The white arrow, which is also over lipid, identifies cap with lower scattering, and the cap-lipid interface is sharply defined. Finally, the black arrow shows the intimal-elastic layer interface (no lipid present) that is diffuse with an intima that is highly scattering. The reason for the diffuse nature of the intimal-elastin border is likely that multiple scattering is obscuring, owing to high scattering within the intima. It is unclear, therefore, whether the lipid interface is diffuse because of the core composition, as previously suggested, or multiple scattering from the cap. Reprinted (modified), with permission, from Circulation 1996;93:1206–121.
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Potential additional OCT markers.
Additional markers show potential for risk-stratifying plaque. These include polarization-sensitive OCT (PS-OCT) for assessing collagen, OCT elastography, and OCT Doppler to assess angiogenesis, as well as dispersion analysis, which is not discussed here (33).
One of the most critical factors in determining plaque stability is the collagen content (22,23,34). We have recently used single detector PS-OCT to identify organized collagen in arteries as well as other tissue (35–37). An example single detector PS-OCT image of a coronary artery is shown in Figure 13. Strong correlation was noted between collagen concentrations in the cap measured by PS-OCT and collagen as assessed by picrosirus staining. The largest OCT changes had a positive predictive power for collagen of 0.889, whereas the smallest change was 0.929 for a negative predictive power. Furthermore, the study was performed with a partially polarized source, and superior results are anticipated when the source is completely polarized. With single detector PS-OCT, changes in each a-scan of the image are measured with changes in source polarization as opposed to dual detector approaches, which measure the absolute value of the backreflected light polarization state. The advantage of the single detector approach is that it is not altered by catheter bending or compression (during rotation or pullback), which is inevitable during interventional procedures. With the dual detector approach, as the catheter rotates or is pulled back, the polarization state changes independent of collagen concentration, making uniform tissue appear heterogeneous by OCT imaging.
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Figure 13 Single detector polarization-sensitive ocular coherence tomography (PS-OCT) imaging. This image illustrates the principles behind single detector PS-OCT. In the trichrome stained section (A) of this in vitro coronary artery, the artery appears relatively homogeneous. In the picrosirius stained section (B), where brightness identifies organized collagen, it can be seen that the artery is heterogeneous by PS-OCT. The PS-OCT images (C and D), at two different polarization states, show different changes in backreflection intensity in different areas of the arteries. This is consistent with the heterogeneity of the picrosirius section. The largest OCT changes had a positive predictive power for collagen of 0.889, whereas the smallest change had a negative predictive power of 0.929.
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Ultrasound elastography has been applied to arteries to assess their tensile strength (38); however, it is associated with only a limited resolution for assessing plaque. Ocular coherence tomography elastography has recently been introduced for evaluating the elastic properties of tissue (39). Owing to its high resolution, OCT elastography has the potential of quantifying the cap mechanical strength of TCFA (40,41), although technical obstacles exist that need to be overcome before the approach can be applied effectively in vivo.
Angiogenesis into plaque, with the subsequent rupture of its thin capillaries, might be an important mechanism of plaque expansion and rupture (42). Ocular coherence tomography Doppler has been demonstrated to identify blood flow in capillaries within tissue (43,44), and OCT Doppler has the potential to identify angiogenesis in plaque.
Technological investigations.
Although a wide range of technological advances are being pursued by multiple groups, three areas of study are of particular importance for intravascular imaging: spectral based OCT for high-speed imaging; microelectrical mechanical devices, to remove the need for moving parts within the catheter; and the development of an OCT guidewire. More detailed discussions can be found in the cited references (11,17–19,45,46).
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Conclusions
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Optical coherence tomography represents a promising new technology for the assessment of vulnerable and unstable plaque. The advantages of OCT include its high resolution, fast data acquisition rate, small inexpensive designs, and ability to be combined with adjuvant techniques. Future work will focus on improving plaque risk stratification, particularly the identification of reliable markers within the images.
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Acknowledgments
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The authors would like to acknowledge the efforts of the large, talented group of investigators whose contributions were invaluable but because of length constraints were not included here. We would also like to recognize all our previous collaborators, students, postdoctoral candidates, and technicians.
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Footnotes
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Dr. William A. Zoghbi acted as guest editor.
a Dr. Brezinski’s research is sponsored by the National Institutes of Health, contracts R01HL63953, R01-HL55686, R01-AR44812, R01-EB000419, R01 AR46996, R29 HL55686, and R01-EB002638. 
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