The final graphic image display is a result of reflected ultrasound waves that are converted to electrical signals and sent to an external processing system for amplification, filtering, and scan conversion. After leaving the transducer, the beam remains parallel for a short distance (near field; better image quality) and then begins to diverge (far field). The length of the near field is expressed by the equation L = r2/λ, where L is the length of the near field, r is the radius of the transducer, and λ is the wavelength. After encountering a transition between different materials (eg, the interface between blood and the intimal arterial layer), the beam will be partially reflected and partially transmitted, depending on tissue composition and differences in mechanical impedance between materials. Calcium produces nearly complete backscattering of the signal and is displayed as a bright image with a characteristic acoustic shadowing. Ultimately, grayscale IVUS imaging is formed by the envelope (amplitude) of the radiofrequency signal (Fig. 21–2).
The ultrasound signal is generated in a piezoelectric crystal that transmits and receives sound waves (A). Ultrasound reflected by the tissue deforms crystal generating radiofrequency (RF) signal (B). Grayscale intravascular ultrasound (IVUS) is derived from the amplitude of RF signal, discarding information beneath the peaks of the signal (C). Changes in the electric field of the piezoelectric crystal caused by ultrasound reflection is used to generate a gray image (D). IVUS RF analysis uses several additional spectral parameters to identify four plaque components (E). Plaque components that are identified are dense calcium (white), fibrous (green), fibrofatty (greenish-yellow), and necrotic core (red) (F). IVUS palpography takes advantage of RF signals generated by the artery being deformed by blood pressure (BP). Using analysis of RF signals at “low” and “high” BP, the strain (deformation) in the inner layer of atheroma is determined (G). This strain is quantified and superimposed on the IVUS image at the lumen–vessel wall boundary (H). Note that high strain (yellow) is found at the shoulders of the eccentric plaque.
The quality of ultrasound images can be described by spatial resolution and contrast resolution. Spatial resolution is the ability to discriminate small objects and is determined in axial (parallel to the beam; primarily a function of wavelength) or lateral (perpendicular to the beam; a function of wavelength and aperture). Axial resolution is approximately 100 μm, and lateral resolution reaches 200 to 250 μm in conventional IVUS systems (20-45 MHz) (Table 21–1). Contrast resolution is the distribution of the gray scale of the reflected signal and is often referred to as dynamic range. An image of low dynamic range appears as black and white with a few levels of gray; images at high dynamic range are often softer.
TABLE 21–1.Comparison of Time-Domain Optical Coherence Tomography, Fourier-Domain Optical Coherence Tomography and Intravascular Ultrasonography ||Download (.pdf) TABLE 21–1. Comparison of Time-Domain Optical Coherence Tomography, Fourier-Domain Optical Coherence Tomography and Intravascular Ultrasonography
|Specifications ||TD-OCT ||FD-OCT ||IVUS |
|Axial scan/second ||5000–10,00 ||~ 100,000 || |
|Lines/frame ||~ 200 ||~ 500–1000 || |
|Maximum frame rate (frames/second) ||20 ||~ 200 ||30 |
|Maximum pullback speed (mm/s) ||3 ||20–40 ||1 |
|Scan diameter (FOV) (mm) ||6.8 ||~ 6–10 ||5–19 |
|Axial resolution (μm) ||15 ||10–15 ||65–150 |
|Lateral resolution (μm) ||90 ||20–40 ||200–250 |
|Tissue penetration (mm) ||1.5–3.0 ||2.0–3.5 ||4–10 |
|Balloon occlusion ||Highly recommended ||No ||No |
|Coronary infusion required ||Yes; Ringer’s lactate with occlusion balloon ||Yes; iodine contrast ||No |
The IVUS equipment consists of a catheter incorporating a miniaturized transducer and a console to reconstruct and display the image. Lately, IVUS consoles have been incorporated into the catheterization laboratory equipment for easier operation. Current catheters range from 2.6 to 3.2 French (Fr) in size and can be introduced through conventional 6-Fr guide catheters. The Opticross catheter (Boston Scientific) is the only mechanical system that is compatible with a 5-Fr guiding catheter; Eagle Eye Platinum electronic phased-array IVUS catheters (Volcano Corporation) are also 5-Fr guide catheter compatible. Rotational, mechanical IVUS catheters operate at frequencies between 30 and 45 MHz, and electronic phased-array systems operate at a center frequency of approximately 20 MHz. Higher ultrasound frequencies are associated with better image resolution. At 30 MHz, the wavelength is 50 μm, yielding a practical axial resolution of approximately 150 μm. But increasing the frequency beyond 60 MHz has been limited because of decreased tissue penetration.11 The mechanical probes rotate a single piezoelectric transducer at 1800 rpm, yielding 30 images per second. Electronic systems have up to 64 transducer elements in an annular array that are activated sequentially to generate the cross-sectional image.11 In general, whereas electronic catheter designs are slightly easier to set up and use, mechanical probes offer superior image quality and have a better crossing profile. Images may be recorded on videotape or digitally with both systems. Electronic IVUS catheters have the ability to display blood flow in color to facilitate distinction between lumen and wall boundaries.
Autoregressive spectral analysis of IVUS backscattered data has been incorporated into conventional IVUS systems to facilitate image interpretation of different tissue components (see Fig. 21–2). This function displays different plaque components in a color-coded scheme according to tissue signal properties (necrotic core in red, dense calcium in white, fibrous in dark green, and fibrofatty in light green).12 In a postmortem validation study, radiofrequency analysis demonstrated sensitivity and specificity for detection of necrotic core of 92% and 97%, respectively.12 The initial algorithm was validated using the mechanical rotating single-element IVUS system, but the first commercially available IVUS backscattering image analysis, named virtual histology (IVUS-VH), was built on the electronic 20-MHz IVUS platform. Lately, rotational mechanical IVUS systems have also integrated a backscattering regression algorithm.13
IVUS-VH imaging has been extensively used in clinical trials to monitor progression of atherosclerosis and to predict plaque growth, providing for the first time evidence that plaque composition carries significant prognostic information.14,15,16,17,18 Nevertheless, reports have raised concerns about the efficacy of this approach to characterize plaque composition, especially in complex calcified lesions and in stented segments, highlighting the need to develop alternative imaging modalities for more reliable assessment of the composition of the plaque.19,20
IVUS-based imaging has been used to assess the mechanical properties of the plaque and in particular its deformability using the analysis of radiofrequency signals at different diastolic pressure levels, normalized to a pressure difference of 2.5 mm Hg per frame. This allows the construction of a “strain” image, in which harder (low strain) and softer (high strain) regions of the coronary arteries can be identified, with radial strain values ranging between 0% and 2%21 (see Fig. 21–2). Postmortem coronary arteries were investigated with histology and IVUS palpography. The sensitivity and specificity of palpography to detect vulnerable plaques was reported to be 88% and 89%, respectively.21 Although preliminary studies have provided evidence about the ability of palpography in detecting high-risk plaque, a substudy of the Providing Regional Observations to Study Predictors of Events in the Coronary Tree (PROSPECT) study, which investigated the value of IVUS in identifying lesions that will progress and cause cardiovascular events, showed that palpography did not provide additional prognostic information.22 This technology has currently limited applications in the research arena.
The IVUS procedure is performed using standard guide catheters (≥ 5 Fr) under full anticoagulation with an activated clotting time of more than 250 seconds. After intracoronary infusion of nitroglycerin (100-200 μg) to minimize vasospasm, the rapid-exchange IVUS catheters are introduced in the coronary over a standard 0.014-in guidewire. Mechanical IVUS systems require infusion of heparinized saline through a dedicated port to clear air bubbles inside the sheath covering the transducer before inserting the catheter in the guide catheter. The IVUS catheter should be advanced under fluoroscopic guidance approximately 10 mm distal to the segment of interest. The operator should realize that the distal marker in mechanical systems at the tip of the catheter is a radiopaque marker and not the ultrasound probe. If possible, the catheter should be advanced beyond 10 mm and retracted slowly to straighten the catheter shaft, which may have built up some slack during insertion, to minimize nonuniform rotation distortion (NURD) artifacts. Motorized pullback devices should be used to withdraw the catheter at a constant speed (most frequently at 0.5 mm/s) to allow proper examination of the entire coronary artery and calculation of distances. During pullback, the record function should be activated to allow storage of images with concomitant voice narration of interesting findings for off-line detailed analysis. Unless coronary ischemia ensues, the catheter should be withdrawn up to the aortic coronary junction and the guide catheter should be retracted slightly to allow imaging of the coronary ostium. Side branches visualized by angiography or ultrasonography are useful landmarks to facilitate interpretation and comparisons in sequential examinations.
IVUS has been performed safely in a large number of subjects enrolled in research studies with no apparent increase in the incidence of adverse effects. The rate of complications associated with IVUS images was investigated in 2207 patients from 28 centers (including 915 patients studied for diagnostic purposes).23 A total of 87 patients (3.9%) had complications judged to be unrelated to IVUS imaging, and 63 patients (2.9%) had transient vasospasm. In nine patients (0.4%), complications were judged to be related to the IVUS procedure, including five acute vessel occlusions, three dissections, and one embolism. In 14 patients (0.6%), complications were judged to have an “uncertain” relationship to IVUS. Major events (acute myocardial infarction [MI] or emergency bypass surgery) occurred in 3 of 9 and 5 of 14 of these patients, respectively. The complication rate was higher in patients with unstable angina or acute MI and in patients undergoing intervention compared with diagnostic IVUS.
Another report comprising 718 IVUS “examinations” performed at 12 centers reported a 1.1% rate of complications without adverse clinical consequences. There were four cases of transient vasospasm; two dissections; and two guidewire entrapments. All complications occurred in patients with unstable angina undergoing PCI. In 7085 IVUS studies from 51 centers, vasospasm occurred in 3% of patients.24 Major complications (dissection, thrombosis, ventricular fibrillation, and refractory spasm) occurred in 10 (0.14%). There was only one major event.
The safety of multivessel IVUS imaging has been investigated in the PROSPECT and the Prediction of Progression of Coronary Artery Disease and Clinical Outcome Using Vascular Profiling of Shear Stress and Wall Morphology (PREDICTION) studies. These studies utilized three-vessel IVUS imaging to assess plaque characteristics and predict atherosclerotic disease progression and future cardiovascular events in patients admitted with an acute coronary syndrome (ACS). In these high-risk populations, the complications caused by IVUS imaging were 1.6% in the PROSPECT study and only 0.6% in the PREDICTION study.15,25
Limitations and Image Artifacts
The echogenicity and texture of different tissue components may exhibit comparable acoustic properties. The similar appearance of different materials represents an inherent limitation of all gray-scale IVUS systems. For example, an echolucent intraluminal image may represent thrombus, atheroma with a high lipid content, retained contrast, or an air bubble.
Most mechanical limitations of IVUS imaging are specific to the construct of each system. NURD is specific to mechanical catheters and arises from friction of the transducer in the coronary or guiding catheter or from a poor connection of the IVUS catheter in the motor drive unit, which causes a typical “onion skin” image appearance. Common causes of NURD are tortuosity; severely stenotic segments; small guide lumen size; and guide catheters with sharp secondary curves, slack in the catheter shaft, or tightened hemostatic valve. The ring-down artifact, however, is specific to electronic systems and is caused by transducer oscillations that obscure the near-field image.26 The side lobe artifact is an intense reflection that comes from strong reflectors, such as calcium and stent struts. This usually follows the circumferential sweep of the beam. The presence of the side lobes may mask the actual lumen edge or may also be taken as tissue prolapse or dissection flaps. Reverberation is another artifact that comes from strong reflectors and is concentric repetitions at equidistant locations of the same image.
An eccentric or nonperpendicular position of the IVUS catheter produces geometric distortions and an artificially elliptical appearance of the cross-sectional image, leading to overestimation of the lumen area.26 The speed of catheter pullback is also prone to errors, which may lead to incorrect assessment of the length of the segment of interest. In non–sheath-based catheters (electronic system), the pitfalls of automatic pullback are greater and include the presence of catheter slack outside the patient and friction in the coronary artery and guiding catheter during pullback. Sheath-based, mechanical catheter systems allow more uniform pullbacks during image acquisition and more precise length measurements.27 The accuracy of four IVUS pullback systems was evaluated in 180 patients (45 in each group) who had been treated with a single stent of known length ranging from 8 to 33 mm. The correlation between the actual versus IVUS-measured stent lengths was 0.92 for Cardiovascular Imaging Systems (CVIS), 0.83 for Galaxy (both from Boston Scientific Corporation, IVUS Technology Center, Fremont, CA), 0.63 for Endosonics Track-Back, and 0.69 for Volcano Model R-100 (both from Volcano Corporation, Rancho Cordova, CA) pullback devices.27 The pullback speed is least accurate at the beginning of an imaging run (in the distal artery) and most accurate in the proximal artery. In battery-powered pullback devices, the battery charge should be verified at the beginning of the procedure to ensure uniform pullback speed. Ideally, the same pullback device and catheter should be used in repeated longitudinal studies. Finally, one should realize that intravascular imaging only assesses a single coronary segment at a time.
Optical Coherence Tomography
OCT evolved from optical one-dimensional, low-coherence reflectometry, which uses a Michelson interferometer and a broadband light source. The addition of transverse B-scans by James Fujimoto in 1991 enabled a two-dimensional imaging of the retina.28 Intravascular OCT uses a single optical fiber that both emits and records light reflection. The image is formed by the backscattering of light from the vessel wall based on “echo time delay” with measurable signal intensity. The speed of light (3 × 108 m/s) is much faster than that of sound (1500 m/s), thus requiring interferometry techniques because direct signal quantification cannot be achieved on such a time scale. The interferometer uses a fiberoptic coupler similar to a beam splitter, which directs half of the beam to the tissue and the other half to the reference arm. The reference arm of the first-generation, time-domain OCT system (TD-OCT), consists of a mirror that moves to produce variable, yet known echo delays. The second-generation intravascular OCT systems use a fixed mirror with a variable frequency light source or “swept laser.” This method, termed Fourier-domain OCT (FD-OCT), allows the simultaneous detection of reflections from all echo time delays, making the system significantly faster. There are two types of FD-OCT systems that differ in their method of data extraction from the interferometer: optical frequency domain imaging (OFDI), also known as swept-source OCT, and spectral domain OCT. OCT images are formed by the reflected signal returning from the tissue and reference arms “interferes” in the fiber coupler and are detected by a photodetector.
As the fiberoptic rotates, multiple axial scans (A-lines) are continuously acquired to generate the cross-sectional vascular image. Bandwidths in the near-infrared spectrum with central wavelengths ranging from 1250 to 1350 nm are used for intravascular applications. Longer wavelengths would provide deeper tissue penetration but would be limited by absorption and the refractive index caused by protein, water, hemoglobin, and lipids. Thus, the image depth of current OCT systems ranges from 1 to 3 mm into the vessel wall. The axial resolution, which is determined by the light wavelength, ranges from 12 to 20 μm. The lateral resolution in catheter-based OCT is typically 20 to 90 μm (see Table 21–1).
The first clinically available OCT systems used time-domain (TD) technology (M2 and M3 OCT Imaging Systems, LightLab Imaging Inc., Westford, MA). The newer systems are based on frequency domain technology. TD-OCT system consists of a single-mode fiberoptic wire (ImageWire), which is able to rotate inside a fluid-filled polymer sheath. An over-the-wire low-pressure occlusion balloon catheter (Helios Goodman, Advantec Vascular Corp.) with distal flush ports is used to infuse saline or Ringer’s lactate at approximately 0.5 mL/s to displace blood during imaging acquisition.
The St. Jude C7XR (FD-OCT) was the first OFDI-OCT system and is approved for clinical use in most countries, including the United States. It is equipped with a tunable laser light source, with the fiber encapsulated within a rotating torque wire built in a rapid exchange 2.6-Fr catheter compatible with 6-Fr guides. Recently, St. Jude introduced an updated version Optis that is able to acquire 180 frames per second, enabling more accurate axial imaging resolution. Lunawave (Terumo Corporation, Tokyo, Japan) has also developed an FD-OCT system, which has a 2.6-Fr shaft.
General principles of intravascular imaging also apply to OCT. Standard guide catheters (≥ 6 Fr) are used, patients must receive full anticoagulation, the image wire or catheter should be advanced under fluoroscopy guidance approximately 10 mm distal to the segment of interest, and the image is acquired after intracoronary infusion of intracoronary nitroglycerin. Despite its guidewire-like profile, the OCT ImageWire is not designed to be advanced into the coronary artery as a stand-alone device. Thus, a conventional angioplasty guidewire (0.014 in) should be inserted first and exchanged to the ImageWire using the Helios or a microcatheter with an inner diameter larger than 0.019 in. Proper OCT imaging requires a blood-free environment because any amount of residual red blood cells causes signal attenuation. In TD-OCT, this is achieved by the inflation of the occlusion balloon proximal to the segment of interest. After advancing the ImageWire into the distal coronary segment, the occlusion balloon is pulled back and repositioned in a healthy proximal segment. Before inflating the Helios balloon at 0.4 to 0.7 atm using a dedicated inflator, injection of Ringer’s lactate at 0.5 mL/s should be started using an injector pump with a warming cuff. Infusion rates can be increased to up to 1.0 mL/s until blood is completely cleared. In TD-OCT systems, the pullback speed can be adjusted from 0.5 to 3.0 mm/s.
FD-OCT is built in rapid-exchange, 6-Fr–compatible, catheter-based systems that are advanced into the coronary segment over a standard 0.014-in guidewire. These systems can acquire 100 to 180 image frames per second, reaching pullback speeds up to 18 to 40 mm/s with the potential to scan epicardial coronary vessels that are 4 to 6 cm long in less than 3 seconds, with the use of a single, high-rate (~ 4 mL/s) bolus injection of contrast (see Table 21–1).29 For nonocclusion techniques, iodinated contrast media are preferred over saline or Ringer’s lactate because of the advantage of high-viscosity solutions in completely removing blood. Before introducing the catheter, a small amount of iodine contrast must be purged into the covering sheath to clear air bubbles and prevent blood entry. The activation of the pullback is automatic as the blood is cleared from the lumen and triggers recording of images, which are stored electronically. The pullback system also has a function to readvance the fiber within its cover sheath. The fast pullback speed does not allow voice recording during imaging acquisition of FD-OCT systems.
The OCT imaging technique has been adopted in many centers across Europe and Japan, and various centers have presented large series of patients imaged by OCT without major complications. The applied energy in intravascular OCT is relatively low (output power in the range of 5.0-8.0 mW) and is not considered to cause functional or structural damage to the tissue. Safety concerns thus seem mainly dependent on the mechanical properties and need for blood clearance for image acquisition. Studies comparing OCT with IVUS suggest that TD-OCT is safe and can be performed with success rates at least comparable to those of IVUS. The most frequent complication in 76 patients imaged with TD-OCT using the occlusive OCT technique were transient chest discomfort, bradycardia or tachycardia, and ST-T changes on electrocardiography (ECG), all of which resolved immediately after the procedure. Similar transient events were also seen during IVUS imaging procedures. Neither hemodynamic instability nor ventricular tachyarrhythmia was observed.30 In addition, in a multicenter registry, the acute complications associated with OCT were assessed in 468 patients. TD-OCT was performed using a nonocclusive flush technique in 45.3% of the patients with a mean contrast volume of 36.6 ± 9.4 mL (M2 and M3). Transient chest pain and QRS widening or ST-segment depression or elevation were observed in 47.6% and 45.5% of patients, respectively. Major complications were rare; they included five (1.1%) cases of ventricular fibrillation caused by balloon occlusion or deep guide catheter intubation, three (0.6%) cases of air embolism, and one (0.2%) case of vessel dissection. There were no cases of coronary spasm or major adverse cardiovascular events (MACE) during or within the 24-hour period after OCT examination.31 The more recent nonocclusive FD-OCT technique has led to an important reduction in the image acquisition time (3-4 seconds) and in the rate of chest pain and ECG changes during image acquisition. The safety and feasibility of the larger catheter-based FD-OCT systems has been examined in numerous small scale studies, which demonstrated that FD-OCT is safer than TD-OCT, as it is associated with a low risk of complications.32,33 In the study of Lehtinen and colleagues, the procedural success of FD-OCT was 87.4%, whereas the complication rate was 17.4% (chest pain in 10.9% of the cases, myocardial ischemia in 2.6%, and minor bleeding in 4.8%). In contrast, in the study of Imola and coworkers, which included 90 patients who underwent FD-OCT to assess lesion severity and the final results poststent implantation, none of the studied patients experienced major complications (ie, death, MI, emergency revascularization, embolization, life-threatening arrhythmia, coronary dissection, prolonged and severe vessel spasm, or contrast-induced nephropathy).32,33
Recently, Taniwaki and colleagues reported the safety and feasibility of multimodality three-vessel intravascular imaging in the Integrated Biomarkers Imaging Study (IBIS) 4.34 This study included 103 patients admitted with a ST-segment elevation MI who had serial OCT and IVUS-VH imaging at baseline and 13 months follow-up. OCT and IVUS-VH were feasible in the vast majority of the cases at both baseline (OCT: 89.9%; IVUS-VH: 85.7%) and follow-up (OCT: 86.6%; IVUS-VH: 84.8%). Failure of OCT imaging was mainly caused by poor vessel flushing and difficulty of advancing the OCT catheter to the distal vessel; IVUS-VH imaging failure was mainly caused by the inability to gate the ECG or to advance the catheter into the distal vessel. A low complication rate was reported (1.9% at baseline; 1.1% at follow-up); all the events (one ventricular fibrillation, and one dissection at baseline and one ventricular fibrillation at 13 months follow-up) were caused by OCT imaging. There were no differences in the MACE between patients who had PCI with and without multimodality intravascular imaging at 2 years follow-up (16.7% vs 13.3%, P = .39). These findings suggest that multimodality intravascular imaging is feasible and safe, even in high-risk patients.
Limitations and Image Artifacts
Some image artifacts are common to both OCT and IVUS. NURD can result in focal image loss or shape distortion, but this seems to occur less frequently in OCT than in IVUS imaging, perhaps as a result of the smaller profile and simplified rotational mechanics of OCT wires. Sew-up artifact is the result of rapid artery or imaging wire movement in one frame’s imaging formation, leading to misalignment of the lumen border. Eccentricity intraluminal position of the ImageWire may lead to wider distances between A-lines and consequently decrease lateral resolution. This effect has been dubbed the “merry-go-round.” Some catheters have reflecting surface; the reflections may bounce off of multiple facets of the catheter, resulting in one or more circular lines; this artifact is called a multiple reflection artifact. Similarly, one can notice the reflection from metallic stent struts aligning toward an eccentrically positioned ImageWire (“sunflower” effect). The saturation artifact occurs when light reflected from a highly specular surface (usually stent struts) produces signals with amplitudes that exceed the dynamic range of the data acquisition system. Although these artifacts are common to all intravascular imaging technologies, others are specific to OCT imaging. The attenuation artifact occurs when the catheter is against the vessel wall. In this case, the light beam is attenuated as it passes along the surface of the artery wall, and thus the vessel wall may appear as a signal poor region below the luminal surface.35 The foldover artifact is seen when the vessel is larger than the ranging depth; in this occasion the vessel may appear to fold over in the image. Residual blood produces light attenuation and may defocus the beam if the red blood cell density is high. This reduces the brightness of the vessel wall. Residual blood does not appear to affect area measurements in OCT images with a clearly defined lumen-wall interface.35 The operator should be aware that residual blood can mimic the OCT appearance of thrombus. Foldover artifact is specific to the new generation of FD-OCT systems and is the consequence of the “phase wrapping” or “alias” along the Fourier transformation when structure signals are reflected beyond the field of view. This can occur at the site of bifurcations or in large vessels. Air bubbles can form within the silicon lubricant used to reduce friction between the sheath and the optic fiber in TD-OCT systems, causing light attenuation (Fig. 21–3).
Cross-sectional optical coherence tomography images of human coronary arteries depicting the most frequently observed artifacts. A. Vessel “out-of-screen.” B. Sew-up artifact resulting in misalignment of the image. C. Nonuniform rotational distortion, usually produced by a tight bend or by an imperfection in the torque wire or sheath that interferes with rotation speed. D. Incomplete intracoronary blood displacement causing light attenuation. E. The presence of blood inside the image catheter sheath. F. Foldover artifact observed in a large vessel bifurcation. G. Saturation artifact; some scan lines have a streaked appearance. H. Light attenuation caused by air bubbles within the silicon lubricant used to reduce friction between the sheath and the optic fiber in time-domain optical coherence tomography systems. The box is indicating a magnified view of the sheath and the optic fiber. I. Eccentric intraluminal location of the ImageWire can distort stent reflection (circle), causing struts to align toward the imaging wire (“sunflower effect,” left box) or appear elongated (“merry-go-round,” right box).
Other Intervascular Imaging Modalities
Another imaging modality able to detect necrotic core invasively is NIRS. To this aim, the 3.2-Fr NIRS catheter approved by the US Food and Drug Administration is used. This catheter is compatible with a conventional 0.014-in guidewire, contains a rotating (240 Hz) NIRS light source at its tip, and is pulled back by a motor drive unit at 0.5 mm/s. The catheter’s ability to detect lipid core plaques (LCPs) has been validated successfully in a human autopsy study in which the device algorithm prospectively identified LCP with a receiver operator characteristic area of 0.80 (95% confidence interval [CI], 0.76-0.85).36 Nevertheless, a more extensive validation suggested that NIRS may be superior to IVUS in detecting lipid-rich plaques, but it has limited accuracy in characterizing their phenotype and detecting fibroatheroma (FA) and thin-cap fibroatheroma (TCFA).9 Other significant limitations of NIRS that have considerably reduced its application in the clinical arena are its inability to quantify plaque burden; visualize lumen and outer vessel wall; and assess plaque characteristics associated with increased vulnerability, such as plaque erosion, neovascularization, and microcalcification.
To overcome these limitations, combined NIRS-IVUS imaging has been proposed. Recently, a combined NIRS-IVUS catheter has been designed that enables simultaneous acquisition and coregistration of the NIRS and IVUS data. The TVC imaging system (InfraReDx, Burlington, MA) incorporates a monorail catheter that can be inserted through a 6-Fr guide and over a 0.014-in guidewire. The catheter has an entry profile of 2.4 Fr and a shaft profile of 3.6 Fr, and includes a mechanical rotating IVUS probe operating at 40 MHz that is pulled back at a speed of 0.5 mm/s using an automated pullback device. NIRS imaging is performed by a NIRS probe located at the tip of the catheter that acquires 30,000 chemical measurements per 100 mm. Each measurement interrogates 1 to 2 mm2 of a tissue at a depth of 1 mm. The output of NIRS is the chemogram that is a map of the measured probability of the presence of lipid-rich tissue presented as a two-dimensional image with pullback position on the x-axis, rotation on the y-axis (as if the artery were cut longitudinally and laid flat). Yellow color indicates regions with a high probability of lipid-rich tissue, whereas red regions suggest a low probability. A measure of the lipid burden is the lipid core burden index, which ranges from 0 to 1000, and gives a semiquantitative summary of the lipid core within the scanned segment. The block chemogram is the summary of the chemogram that display the presence of the lipid tissue in 2-mm intervals. The TVC system also provides a composite view of the coregistered NIRS and IVUS data, allowing comprehensive assessment of the association between plaque burden and composition (Fig. 21–4).
Output of the combined near-infrared spectroscopy–intravascular ultrasound TVC imaging system (Infraredx, Burlington, MA). RCA, right coronary artery.
Histology-based studies have shown that combined NIRS-IVUS imaging allows more accurate characterization of the composition of the plaque,9,37 while several in vivo studies have examined the reproducibility of NIRS imaging and provided robust data that this enables a consistent assessment of plaque compositional characteristics. Over the past years, numerous studies have implemented this technology to assess the implications of lipid-lowering therapies on the compositional characteristics of the plaque and its role in identifying high-risk plaques and culprit lesions.18,38,39,40,41,42,43
Emerging Imaging Modalities
Raman spectroscopy relies on the spectral analysis of the Raman scattering by a tissue after being illuminated with a laser beam. The laser light is scattered between molecules, and finally the energy of the photons is changed, resulting in a different frequency of the backscattered light. The Raman spectra is unique for a given molecule and, therefore, it can be used to assess the chemical composition of the plaque and differentiate esterified from nonesterified cholesterol (see Fig. 21–1A). A significant limitation of Raman spectroscopy is the increased noise, generated within the fibers of the imaging catheters, which interferes with the Raman spectra. Modification of the initial prototypes and utilization of high-wave number Raman spectroscopy can potentially overcome this limitation and enable its application in the study of atherosclerosis.44,45
Intravascular Photoacoustic Imaging
Intravascular photoacoustic (IVPA) imaging involves irradiation of the vessel wall by short laser pulses with a length of several nanoseconds. The absorbed laser light transfers the optical energy to the tissue, which causes a transient pressure rise46 that propagates through the tissue as an acoustic wave and can be detected with an ultrasound transducer. The time interval between tissue irradiation and acoustic wave detection enables localization of the source and depth-resolved imaging. Studies have demonstrated that IVPA can detect the composition of the plaque and, in particular, lipid component as well as plaque features associated with increased vulnerability, such as intraplaque hemorrhage and inflammation (see Fig. 21–1B).47
Limitations of IVPA include its inability to assess the luminal, outer vessel wall and plaque dimensions; the vessel geometry the distribution of the plaque; and the fact that optimal image quality requires blood clearance and prolonged timing. With the advent of high-speed lasers, the time for imaging is becoming compatible with clinical application.
Near-Infrared Fluorescence Imaging
Near-infrared fluorescence (NIRF) imaging caries a great potential for the study of vascular inflammation. It relies on the use of activatable markers that have the ability to bind molecules associated with plaque vulnerability and fluorescence when they irradiated by near-infrared light (see Fig. 21–1C). Advances in molecular biology have enabled detection of fibrin, of plaque’s inflammation, and identification of neovascularization.48 Recently, a two-dimensional rotational NIRF catheter has been designed that is pulled back by an automated pullback device at a constant speed, enabling generation of a spread out plots of the coronary plaque inflammation and assessment of vessel wall biology.49 This design enables in vivo detection of different molecules, such as thrombin, matrix metalloproteinases 2 and 9, and cathepsin K, D, and S, and permits assessment of plaque biology, inflammation, and neorevascularization.
Significant limitations of this modality are (1) that it requires injection of an activatable agent; (2) that it cannot provide depth-resolved information, and (3) assess the lumen, outer vessel wall, and plaque dimensions and composition.
Time-Resolved Fluorescence Spectroscopy
Time-resolved (lifetime) fluorescence spectroscopic (TRFS) imaging using pulsed ultraviolet light appears able to provide information about the compositional characteristics of the plaque. It relies on the assessment of the time required to resolve the fluorescence emitted after molecules being excited by light. The compositional characteristics of the plaque (ie, the lipid component, macrophages, elastin, and collagen) have different fluorescence properties; thus TRFS has a value in assessing plaque composition and inflammation (see Fig. 21–1D).50
Recent technological advances in fluorescence lifetime imaging have enabled adaptation of these techniques for intravascular scanning.51,52 Experimental studies have shown that TRFS measurements enable detection of diseased segments, visualization of pathological changes in the superficial plaque (~ 250-300 μm depth), and characterization of the phenotype of the plaque and differentiation of TCFA from FA.