Cardiovascular OCT Imaging [2nd ed. 2020] 978-3-030-25710-1, 978-3-030-25711-8

Table of contents :
Front Matter ....Pages i-vii
The Development of Optical Coherence Tomography (James G. Fujimoto, Joseph Schmitt, Eric Swanson, Aaron D. Aguirre, Ik-Kyung Jang)....Pages 1-23
Histology Validation of Optical Coherence Tomography Images (Teruyoshi Kume, Takashi Kubo, Takashi Akasaka)....Pages 25-36
Basic Interpretation Skills (Tom Adriaenssens)....Pages 37-52
Intravascular OCT Imaging Artifacts (Jennifer E. Phipps, Taylor Hoyt, David Halaney, J. Jacob Mancuso, Sahar Elahi, Andrew Cabe et al.)....Pages 53-66
Coronary Plaque Types: Thin Cap Fibroatheroma, Healed Plaque, Calcified Plaque (Francesco Fracassi, Giampaolo Niccoli)....Pages 67-77
Plaque Erosion (Vikas Thondapu, Peter Libby, Ik-Kyung Jang)....Pages 79-89
OCT Imaging of SCAD and Differential Diagnosis (Ashkan Parsa, Jacqueline Saw)....Pages 91-104
How to Use OCT to Optimize PCI? (Teruyoshi Kume, Shiro Uemura)....Pages 105-114
Post-PCI OCT Findings and the Clinical Significance (Taishi Yonetsu)....Pages 115-124
Very Late Stent Thrombosis (Tom Adriaenssens)....Pages 125-137
OCT for Bioabsorbable Vascular Scaffold (Alessio Mattesini, Antonio Martellini, Luigi Tassetti, Carlo Di Mario)....Pages 139-147
Detection of Vulnerable Plaque (Rocco Vergallo, Ik-Kyung Jang)....Pages 149-161
Multimodality Intravascular OCT Imaging (Kensuke Nishimiya, Guillermo Tearney)....Pages 163-174
Future Development (Martin Villiger, Jian Ren, Néstor Uribe-Patarroyo, Brett E. Bouma)....Pages 175-191
Back Matter ....Pages 193-198

Citation preview

Ik-Kyung Jang  Editor

Cardiovascular OCT Imaging Second Edition

123

Cardiovascular OCT Imaging

Ik-Kyung Jang Editor

Cardiovascular OCT Imaging Second Edition

Editor Ik-Kyung Jang Massachusetts General Hospital Harvard Medical School Boston, MA USA

ISBN 978-3-030-25710-1    ISBN 978-3-030-25711-8 (eBook) https://doi.org/10.1007/978-3-030-25711-8 © Springer Nature Switzerland AG 2020 This work is subject to copyright. All rights are reserved by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed. The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use. The publisher, the authors, and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, expressed or implied, with respect to the material contained herein or for any errors or omissions that may have been made. The publisher remains neutral with regard to jurisdictional claims in published maps and institutional affiliations. This Springer imprint is published by the registered company Springer Nature Switzerland AG The registered company address is: Gewerbestrasse 11, 6330 Cham, Switzerland

Foreword

Since the first report on OCT over two decades ago, interest in this imaging technology in medicine has exponentially increased, initially in ophthalmology and recently in cardiology. Introduction of this intravascular imaging modality in cardiology coincided with the emergence of a new concept, namely, the detection of “vulnerable or high-risk plaque.” As vulnerable plaque represents microstructural changes of coronary plaque including thin fibrous cap overlying lipid pool and increased lipid pool, it was thought that this high-resolution imaging modality would be ideal for detecting vulnerable plaques. Indeed, OCT has been proven to be an invaluable tool not only for detecting high-risk plaques but also for studying in vivo vascular biology, particularly in order to understand the pathophysiology of acute coronary syndromes. With a series of OCT studies, our understanding on the pathogenesis of acute coronary syndrome has made remarkable progress and confirmed many findings, which had been possible only in autopsy studies previously. Since OCT can be used in patients repeatedly, in vivo serial studies have been reported, and these studies have provided critical information on the evolution of vascular structure over time and also the vascular response to interventions such as cholesterol-lowering therapy. In addition to the application of OCT to research, OCT has been increasingly used in interventional cardiology, primarily to optimize percutaneous coronary intervention with stenting. Despite gradual adoption of this technology in cardiac catheterization laboratories, the clinical value of OCT remains undefined. How to maximize the information provided by OCT to optimize percutaneous procedures is not widely known. This book will provide background and practical tips for interventional cardiologists. Since his first report on in vivo OCT application in 2001, in my view, Dr. Jang’s group has been the leader in OCT in cardiology. His group brought this technology from bench to bedside through series of preclinical and clinical studies. They for the first time established the OCT criteria for plaque characterization, reported differences in plaque characteristics in patients with various clinical presentations, and recently published the first in vivo demonstration of plaque erosion, one of the important underlying mechanisms for acute coronary syndromes. In addition, his group published the first paper comparing OCT and intravascular ultrasound after stenting, demonstrating the higher sensitivity of OCT in the detection of stent-­related complications. This textbook covers all the practical aspects of OCT in current research as well as clinical practice. With the rapidly growing interest in OCT, this book is extremely timely and very relevant for all cardiologists interested in vascular biology or interventional cardiology. My warmest congratulations to Dr. Jang and his group for such outstanding contributions to the field of cardiovascular medicine. Valentin Fuster Director, Mount Sinai Heart New York, NY, USA Past President American Heart Association Dallas, TX, USA Past President World Heart Federation Geneva, Switzerland

v

Contents

1 The Development of Optical Coherence Tomography���������������������������������������������   1 James G. Fujimoto, Joseph Schmitt, Eric Swanson, Aaron D. Aguirre, and Ik-Kyung Jang 2 Histology Validation of Optical Coherence Tomography Images���������������������������  25 Teruyoshi Kume, Takashi Kubo, and Takashi Akasaka 3 Basic Interpretation Skills�����������������������������������������������������������������������������������������  37 Tom Adriaenssens 4 Intravascular OCT Imaging Artifacts�����������������������������������������������������������������������  53 Jennifer E. Phipps, Taylor Hoyt, David Halaney, J. Jacob Mancuso, Sahar Elahi, Andrew Cabe, Mehmet Cilingiroglu, Thomas E. Milner, and Marc D. Feldman 5 Coronary Plaque Types: Thin Cap Fibroatheroma, Healed Plaque, Calcified Plaque���������������������������������������������������������������������������������������������������������������������������  67 Francesco Fracassi and Giampaolo Niccoli 6 Plaque Erosion �����������������������������������������������������������������������������������������������������������  79 Vikas Thondapu, Peter Libby, and Ik-Kyung Jang 7 OCT Imaging of SCAD and Differential Diagnosis�������������������������������������������������  91 Ashkan Parsa and Jacqueline Saw 8 How to Use OCT to Optimize PCI? ������������������������������������������������������������������������� 105 Teruyoshi Kume and Shiro Uemura 9 Post-PCI OCT Findings and the Clinical Significance������������������������������������������� 115 Taishi Yonetsu 10 Very Late Stent Thrombosis��������������������������������������������������������������������������������������� 125 Tom Adriaenssens 11 OCT for Bioabsorbable Vascular Scaffold��������������������������������������������������������������� 139 Alessio Mattesini, Antonio Martellini, Luigi Tassetti, and Carlo Di Mario 12 Detection of Vulnerable Plaque��������������������������������������������������������������������������������� 149 Rocco Vergallo and Ik-Kyung Jang 13 Multimodality Intravascular OCT Imaging������������������������������������������������������������� 163 Kensuke Nishimiya and Guillermo Tearney 14 Future Development��������������������������������������������������������������������������������������������������� 175 Martin Villiger, Jian Ren, Néstor Uribe-Patarroyo, and Brett E. Bouma Index������������������������������������������������������������������������������������������������������������������������������������� 193

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1

The Development of Optical Coherence Tomography James G. Fujimoto, Joseph Schmitt, Eric Swanson, Aaron D. Aguirre, and Ik-Kyung Jang

1.1

Introduction

verse direction, as shown in Fig.  1.1. This produces a two-dimensional image data set, analogous to a B-scan Optical coherence tomography (OCT) has become a major in ultrasound, which represents the optical backscattering optical imaging modality in biomedical optics and medi- in a cross-sectional plane through the tissue. These crosscine. OCT performs high resolution, cross-sectional and sectional images can be displayed using false color or grey three dimensional volumetric and functional imaging of the scale in order to visualize subsurface tissue structure and internal microstructure in biological tissues by measuring pathology. Three-­ dimensional, volumetric data sets are backscattered light [1]. Tissue pathology can be imaged in generated by acquiring sequential B-scans or cross-secsitu and in real time with resolutions of 1–15  μm, one to tional images while scanning the incident optical beam in a two orders of magnitude finer than conventional ultrasound. raster, spiral pullback, or other pattern. Three-dimensional Imaging can be performed using small fiber-optic probes OCT (3D-OCT) data contains comprehensive volumetric including catheters and endoscopes. The unique features of structural and functional information and can be digitally OCT make it a powerful imaging modality with applications manipulated and visualized similar to 3D ultrasound, MR, spanning many clinical specialties as well as fundamental or CT images. scientific and biological research and many non-medical OCT is a powerful medical imaging technology because applications. it performs an “optical biopsy”; providing real time, in OCT is an optical analog of ultrasound and performs situ visualization of tissue microstructure and pathology, cross-sectional and volumetric imaging by measuring the without the need to excise and process specimens [2, 3]. amplitude and time delay of backscattered light. Cross-­ Histopathology is the gold standard to diagnose pathology, sectional images are generated by performing multiple but it requires tissue excision, fixation, embedding, microtmeasurements of reflected light amplitude versus time oming and staining. OCT has applications in several general delay, analogous to axial scans or A-scans in ultrasound, clinical situations: while scanning the incident optical beam in the trans 1. Where conventional excisional biopsy is hazardous or impossible. Applications include retinal imaging in ophJ. G. Fujimoto (*) thalmology or coronary artery imaging in interventional Department of Electrical Engineering and Computer Science, cardiology. Research Laboratory of Electronics, Massachusetts Institute of 2 . Where excisional biopsy has sampling error. Excisional Technology, Cambridge, MA, USA biopsy and histopathology is the gold standard for e-mail: [email protected] diagnosis of many diseases including cancer, however, J. Schmitt false negatives can result if excisional biopsy misses Health Technologies, Apple, Inc., Sunnyvale, CA, USA the lesion. OCT can guide excisional biopsy to improve E. Swanson sensitivity and reduce the number of biopsies required. Research Laboratory of Electronics, Massachusetts Institute of Technology, Cambridge, MA, USA Since imaging is performed in situ, OCT can assess e-mail: [email protected] much larger regions of tissue than possible with exciA. D. Aguirre · I.-K. Jang sional biopsy. If sufficient sensitivity and specificity Division of Cardiology, Massachusetts General Hospital, Harvard are achieved, OCT can be used for real time Medical School, Boston, MA, USA diagnosis. e-mail: [email protected]; [email protected]

© Springer Nature Switzerland AG 2020 I.-K. Jang (ed.), Cardiovascular OCT Imaging, https://doi.org/10.1007/978-3-030-25711-8_1

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J. G. Fujimoto et al. 1D Axial (Z) Scanning

2D Axial (Z) Scanning Transverse (X) Scanning

3D Axial (Z) Scanning XY Scanning

Backscattered Intensity

Axial Position (Depth)

Fig. 1.1 Optical coherence tomography (OCT) generates cross-­ sectional or three-dimensional images by measuring the magnitude and echo time delay of light. Data is composed of axial scans (A-scans), which are measurements of backreflection or backscattering versus

delay. Cross-sectional images (B-scans) are generated by transverse scanning the OCT beam, acquiring a series of axial scans. Three-­ dimensional volumetric data sets (3D-OCT) are generated by raster scanning, acquiring a series of cross-sectional images (B-scans)

3. For guiding interventional procedures. The ability to see cross sectional and three-dimensional structure enables the guidance of procedures such as stent placement in intravascular imaging. In ophthalmology, OCT can visualize changes in retinal structure and markers of disease, such as neovascularization or edema, to assess progression or pharmaceutical treatment response. The ability to see beneath the tissue surface enables guidance of microsurgical procedures. 4. For performing functional measurements and imaging. Doppler OCT enables quantitative measurement of blood flow. OCT angiography enables visualization of tissue microvasculature using motion contrast from flowing blood.

during coronary stent placement. Applications of OCT are also being developed in other many other clinical specialties such as dermatology and gastroenterology, and its use in fundamental research continues to expand. This chapter reviews the history and early development of OCT including the translation process from ex vivo imaging studies to preclinical imaging in animals and ultimately to the first studies in patients. We present a review of key technological developments that have enabled the use of OCT in intravascular imaging and other applications beyond ophthalmology. Finally, we present an overview of commercial activities in the development of intravascular OCT. Commercialization is a key step in making advances available to the wider clinical community, where new methodologies can ultimately benefit patients and have economic impact.

Although the imaging depth of OCT is limited by attenuation from light scattering, absorption and optical aberration, when coupled with medical devices such as catheters, endoscopes, laparoscopes, or needle delivery instruments, OCT can be used to access luminal organ systems such as the coronary arteries, GI tract and airways, as well as solid organs and masses. OCT has become a standard of care in ophthalmology with an estimated ~30 million imaging procedures performed worldwide every year [4, 5]. Intravascular OCT imaging is also becoming an increasingly powerful tool for understanding the biology of coronary artery disease and for optimizing interventional decision making such as

1.2

 CT Compared to Microscopy O and Ultrasound

OCT has features which are common to both microscopy and ultrasound. Therefore, it is helpful to begin by comparing OCT to these modalities. Figure 1.2 shows the resolution and imaging depth for OCT, ultrasound and microscopy. The resolution of clinical ultrasound imaging is typically 0.1–1 mm and depends on the frequency of the sound wave (3–40 MHz) used for imaging. Sound waves at ultrasound frequencies

1  The Development of Optical Coherence Tomography

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Fig. 1.2  Image resolution and depth. Clinical ultrasound achieves deep imaging depths, but with limited resolution. Higher sound frequencies improve resolution, but increased ultrasonic attenuation reduces imaging depth. Confocal microscopy has submicron resolution, but because it is difficult to reject unwanted scattered light, the imaging depth is a few hundred microns in most tissues. OCT axial image resolution ranges from 1 to 20 μm, according to the coherence length of the light source. In most biological tissues, OCT imaging depth is limited to ~2 mm by attenuation from optical scattering

below 10  MHz are transmitted with minimal attenuation in biological tissue and can image structures deep in the body but have limited resolution. High frequency ultrasound has been used for intravascular imaging as well as research applications. Resolutions of 15–20  μm and finer have been achieved with frequencies above 100 MHz. However, these high frequencies are strongly attenuated in biological tissues and imaging depths are limited to only a few millimeters. Confocal microscopy is an imaging technique which has extremely high transverse image resolution approaching ~1  μm. Transverse resolution Δx is determined by the objective lens numerical aperture (ratio of diameter d to focal length f) and optical diffraction effects (Fig.  1.3). Axial resolution is also determined by the numerical aperture, with finer transverse resolution (higher magnification) associated with shorter depth of field DOF or in-focus range. Microscopes are typically used for high magnification imaging in which en face views rather than a transverse (depth) cross-sections are imaged. Imaging depths in biological tissue are limited to a few hundred microns because image signal and contrast are degraded by unwanted scattered light and optical aberrations. OCT fills an important gap between ultrasound and microscopy. The transverse image resolution in both micros-

High Numerical Aperture focus

Low Numerical Aperture focus

Fig. 1.3  Image resolution in microscopy compared to OCT. The transverse image resolution in microscopy as well as OCT images is determined by the transverse spot size of the optical beam. This spot size is governed by the numerical aperture, the ratio of incident beam diameter d to the objective lens focal length f. Microscopy typically operates with a high numerical aperture focus to achieve a small transverse spot size Δx and a short depth of field (DOF). The axial image resolution Δz in OCT is determined by the bandwidth or sweep range of the light source. OCT typically operates with a low numerical aperture focus with a larger transverse spot size Δx and a long depth of field. OCT can achieve high axial resolution in catheters where the size of the objective lens and the numerical aperture are limited

copy and OCT is determined by the numerical aperture of the objective lens. However the axial resolution in OCT imaging is achieved by measuring the time delay of light using interferometry. OCT imaging is typically performed with low numerical aperture optics and its axial resolution Δz is determined by the bandwidth or sweep range of the light source. Unlike microscopy, OCT typically images with a modest transverse image resolution and long depth of field. OCT can have very fine axial resolution in situations where the numerical aperture of the optics limits conventional axial resolution. These include catheter imaging where limitations on lens sizes and numerical apertures are imposed by the catheter diameter and ophthalmic imaging where the numerical aperture is limited by the pupil of the eye. OCT achieves axial image resolutions ranging from 1 to 15  μm, approximately 10–100 times finer than standard ultrasound imaging, which enables visualization of tissue microstructure at resolutions approaching histology. OCT has become a clinical standard in ophthalmology because the transparency of the eye provides easy optical access to the retina, which makes non-contact high-resolution imaging possible. There is a dearth of other alternative technologies than can acquire such quantitative depth information about the retina. The principal limitation of OCT is that light is highly scattered by most tissues and the attenuation typically limits the imaging depths to ~2 to 3  mm. Although this depth is relatively small compared to ultrasound imaging depth, light

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can be delivered by optical fibers and OCT can be integrated with a wide range of medical instruments and devices such as catheters, guide wires, endoscopes, laparoscopes, or needles to enable imaging in almost any region in the human body. OCT imaging is analogous to ultrasound imaging, except that it uses light instead of sound. There are several different detection methods for performing OCT, but essentially imaging is performed by measuring the magnitude and time delay of back reflected or backscattered light from internal microstructures in materials or tissues. OCT images are two-­dimensional or three-dimensional data sets which represent optical back reflection or backscattering in a cross-sectional plane or 3D volume. When a sound beam or light beam is incident into tissue, it is backreflected or backscattered differently from structures that have varying acoustic or optical properties, as well as from boundaries between structures. These internal structures can be imaged by measuring the “echo” time it takes for sound or light to travel different axial distances. In ultrasound imaging, the axial measurement of distance or depth is called A-mode scanning, while cross-sectional imaging is called B-mode scanning. Volumetric or 3D imaging can be performed by acquiring multiple B-mode images. However, the speed of sound in tissue is ~1500 m/s, while the speed of light is 200,000 times faster, approximately 3  ×  108  m/s. Measuring distances with a 100  μm resolution, a typical resolution for ultrasound imaging, requires a time resolution of ~100 ns. This resolution is well within the limits of electronic detection. Ultrasound technology has

advanced dramatically in recent years with the availability of high performance, low cost analog to digital converters and digital signal processing technology. In contrast, the detection of light echoes requires much higher time resolution. Light travels from the moon to the earth in only ~2 s. Measuring distances with a 10 μm resolution, a typical axial resolution for OCT imaging, requires a time resolution of ~30 fs (3 × 10−15 s). A femtosecond is an extremely short time period; the ratio of 1 fs to 1 s is equal to the ratio of 1  s to the time since the age of dinosaurs. Femtosecond time resolution cannot be achieved by electronic detection using fast clocks/timing and methods such as high-speed optical gating, optical correlation and interferometry are required.

1.3

Measuring Optical Backscatter

1.3.1 Photographing Light in Flight The possibility of using light to visualize subsurface structure in biological tissue was proposed by Michel Duguay in 1971 [6, 7]. These pioneering studies demonstrated an ultrafast optical shutter using the laser induced Kerr effect which could “photograph light in flight”. Figure 1.4 shows a schematic of Duguay’s ultrahigh speed Kerr shutter photographing an ultrashort light pulse propagating though a scattering solution of diluted milk. The Kerr shutter operates by

INFRARED PULSE

POLARIZER 1

POLARIZER 2 CS2

GREEN PULSE

R

TE

T HU

TS

FILTER

AS AF

TR UL

Fig. 1.4  Photographing light in flight. (left) A high speed optical shutter is created using a CS2 cell between crossed polarizers. An intense laser pulse induces transient birefringence (the Kerr effect) to open the shutter. (right) Photograph of an ultrashort laser pulse propagating

through a cell of milk and water. The shutter speed was ~10 ps. These early studies suggested that high speed optical gating could be used to see inside biological tissues by rejecting unwanted scattered light (Duguay [6] and Duguay and Mattick [7])

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using an intense laser pulse to induce birefringence (the Kerr effect) in an optical medium placed between two crossed polarizers. The induced birefringence has an extremely rapid response time and the Kerr shutter can achieve picosecond or femtosecond time resolution. Recognizing that optical scattering limits the ability to image inside biological tissues, Duguay proposed that an ultrahigh speed shutter could remove the unwanted scattered light and detect echoes of light from inside tissue [7]. Ultrahigh speed optical shutters might be used to “see through” tissues and therefore to non-invasively image internal pathology. The principal limitation of the high-speed optical Kerr shutter is that it requires a high intensity, short pulse laser to induce the Kerr effect and operate the shutter.

1.3.2 Femtosecond Time Domain Measurement An alternate method for detecting optical echoes is to use nonlinear optical processes such as harmonic generation, sum frequency generation, or parametric conversion [8–10]. Short light pulses illuminate the tissue and the time delay of backscattered light is detected by nonlinear optical cross correlation, which occurs by mixing the

back reflected light with a time delayed reference pulse in a nonlinear optical material. The nonlinear process can measure the intensity and time delay of the optical signal with a time resolution determined by the pulse duration. Figure  1.5 shows a schematic of how light echoes are detected using nonlinear second harmonic generation cross correlation. The reference pulse is generated by the same laser source and is delayed by a variable time delay ΔT using a mechanical optical delay line. The nonlinear mixing process creates an ultrahigh speed optical gate which is like an optical shutter. The ultrahigh speed optical gate enables measurement of the echo magnitude vs time delay of back reflected or backscattered light, the equivalent of an axial scan in an ultrasound. Figure  1.5 shows a measurement of corneal thickness in an ex vivo bovine eye [10]. Very low scattering from the corneal stroma can be detected. The measurement had a 15 μm axial resolution and was performed using 65 fs duration pulses from a femtosecond dye laser at 625  nm wavelength. Sensitivities of −70 dB or 10−7 of the incident intensity were achieved. However, these sensitivities were still not high enough to image the weak optical backscattering from most biological tissues. Current OCT systems achieve sensitivities 1000× higher, approaching −100 dB or 10−10 of the incident intensity.

DELAY

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AIR

CORNEA

AQUEOUS

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SIGNAL x10

NONLINEAR XTAL (KDP)

FILTER PMT

LOCK-IN AMPLIFIER

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Fig. 1.5  Early demonstration of femtosecond optical ranging in biological systems. (left) Femtosecond echoes of light (signal) are detected using nonlinear second harmonic generation, mixing the signal with a delayed reference pulse. (right) Femtosecond measurement of corneal

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thickness in a bovine eye ex vivo, showing an axial scan of backscattering versus depth. Using a femtosecond dye laser with a 65-fs pulse duration it was possible to achieve a 15-um axial resolution (in air) with a −70 dB or 10−7 detection sensitivity. (From Fujimoto et al. [10])

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1.3.3 D  etecting Backscattered Light Using Interferometry Interferometry is a powerful technique for measuring the magnitude and relative time delay of backscattered (or echoed) light with very high sensitivity which has the additional advantage of not requiring short pulse light sources. OCT is based on interferometric imaging and the first OCT instruments were based on a classic optical measurement technique known as low coherence interferometry, or white light interferometry, first described by Sir Isaac Newton. Low coherence interferometry was used in photonics to measure optical echoes and backscattering in optical fibers and waveguide devices in the 1980s [11, 12]. One of the first biological applications of low coherence interferometry was reported by Fercher et al. in 1988, for the measurement of axial eye length [13]. Different versions of low coherence interferometry were developed for non-invasive measurement in biological tissues [14–18]. Interferometry techniques measure the magnitude and time delay of light by interfering light that is backscattered from the tissue with light that has traveled a reference path length with a calibrated delay. Interferometry measures the electric field of the light rather than its intensity. Figure 1.6 shows a schematic diagram of a classic Michelson type

interferometer. The output from a light source is split into a reference beam and a measurement or signal beam that travels different distances in the two interferometer arms. If the reference path length is scanned, interference fringes will be generated. This process can also be understood by noting that scanning (e.g. linear-translation) of the reference arm produces a Doppler shift of the reference field. A “coherent” light source has a narrow optical wavelength spectrum and will produce interference over a long range of path length differences in the interferometer. However, in order to measure the time delay of optical echoes, a low coherence (broad optical wavelength spectrum or bandwidth) light source is required. When low coherence light is used, interference is only observed when the measurement and reference path lengths have only a small mismatch, defined as the “coherence length” and inversely proportional to the optical bandwidth. The magnitude and echo time delay of light can be measured by scanning the reference arm and detecting the interference signal. The coherence length of the light source determines the axial image resolution. Broadband light sources have shorter coherence lengths and have finer axial image resolution. Figure 1.6 shows an early A-scan measurement of the anterior chamber of an ex vivo bovine eye with a 10 μm axial resolution using an 800 nm wavelength, low coherence diode

Cornea Scanned Reference Path

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Fig. 1.6  Low coherence interferometry for measuring optical echoes. (left) Backreflected or backscattered light signals are interfered with light from a scanning reference path delay. When low coherence length light is used, interference is only observed when the path lengths are matched to within the coherence length. The interferometer output is detected as the reference path delay is scanned (time domain detection).

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Axial scan (A-scan) information is obtained by demodulating the interference signal. (right) Measurement of the anterior chamber in a bovine eye ex vivo. A 10 μm axial resolution was achieved using a low coherence diode light source at ~800 nm. Interferometry enables high sensitivity detection of optical echoes and is the basis for optical coherence tomography. (From Huang et al. [19])

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light source with a 29  nm bandwidth [19]. High detection sensitivities of −100  dB or 10−10 of the incident intensity were achieved. Scanning the optical beam in the transverse direction yielded information on different structures, such as the lens and iris. The axial measurements of backscatter versus depth using low coherence interferometry provided the foundation for optical coherence tomography. Low coherence interferometry has the advantage that it can be performed without the complexity of requiring short pulses of light. Since interferometry measures the electric field rather than the intensity, it is analogous to heterodyne detection in optical communication and can achieve very high sensitivity and dynamic range. Interferometry essentially measures the product of electric fields. Weak signals ESIGNAL (t) are multiplied by a strong reference electrical field EREF (t) producing an interference output proportional to ESIGNAL (t) × EREF (t). This multiplication amplifies the weak signal and very high detection sensitivities can be achieved. In addition, since the intensity is proportional to the square of the electric field I = E2, very high dynamic ranges are possible because a

30 dB dynamic range for detecting the field corresponds to a 60 dB dynamic range for intensity. The combination of high sensitivity and high dynamic range are important for OCT imaging of biological tissue because backscattered light is not only very weak, but can also vary over a wide (>40 dB) dynamic range of signal intensities.

Fig. 1.7  OCT image of the human retina and coronary artery ex vivo and corresponding histology. Imaging was performed with 15 μm axial resolution at 830 nm wavelength. The OCT images are displayed using a log false color scale ranging from −60 to −90 dB of the incident light intensity. (top) OCT shows the optic nerve head contour and vasculature.

The retinal nerve fiber layer is visualized and there is postmortem retinal detachment, with subretinal fluid accumulation. (bottom) OCT shows fibro-calcific plaque (right three-quarters of specimen) and fibroatheromatous plaque (left side of specimen). The fatty-calcified plaque scatters light, limiting the image penetration depth. (From Huang et al. [1])

1.4

The Development of OCT

1.4.1 Early OCT Technology and Systems Optical coherence tomography imaging was first demonstrated in 1991 by Huang et al. [1]. A similar imaging concept was also proposed independently in 1991 by Tanno in Japan [20]. Figure  1.7 shows the first OCT images of the retina and human coronary artery ex vivo with corresponding histology [1]. These examples demonstrated OCT imaging in both ophthalmic and cardiovascular tissues which today are the largest clinical applications of OCT.  Imaging was

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performed with 15 μm axial resolution in tissue at 830 nm wavelength. The image is displayed using a log false color scale with a signal level ranging from ~−60 (red color) to −90 dB (dark blue color) of the incident intensity. The OCT image of the retina in Fig. 1.7 shows the contour of the optic nerve head as well as retinal vasculature near the nerve head. The retinal nerve fiber layer can also be seen emanating from the optic nerve head. The OCT image of the coronary artery in Fig.  1.7 shows fibro-calcific plaque on the right of the specimen and fibroatheromatous plaque on the left. The plaque is easily seen as it scatters light and shadows the structure below it. Optical coherence tomography has the advantage that it can be implemented using fiber optic components and integrated with a wide range of medical devices. OCT systems can be divided into an imaging engine or console (consisting of an interferometer, light source, detector, and support electronics) and patient interfaces such as imaging devices or probes. Early OCT imaging engines employed time domain detection (TD-OCT) with an interferometer using a low coherence light source and scanning reference delay arm. Figure 1.8 shows the first ophthalmic OCT imaging system for imaging patients. The reference path delay is mechanically scanned at high speed to generate interference fringes which are electronically detected to produce A-scan information. The modified slit-lamp directs the OCT beam through Fig. 1.8  OCT uses photonics and fiber optics technology. (top) Schematic of a Michelson interferometer implemented using fiber optics. The probe may be interfaced to a variety of medical imaging devices. (bottom) The first clinical ophthalmic imaging OCT prototype instrument designed at MIT Lincoln Laboratories and used at the New England Eye Center. (bottom left) OCT imaging engine. (bottom right) Slitlamp modified to support in vivo perform OCT (OCT slitlamp photo from August 1992). (From Swanson and Fujimoto [5])

the pupil of the eye, and galvanometer actuated mirrors scan the beam on the retina.

1.4.2 Ophthalmic OCT Imaging The earliest clinical applications of OCT were in ophthalmology because the eye is optically accessible and many optical imaging methods were already widely used for the retina and anterior eye. The first in vivo retinal imaging was performed in 1992 by Swanson et  al. [21]. In vivo retinal imaging was also demonstrated by Fercher et al. around this time [22]. Figure  1.8 shows the MIT prototype instrument with a patient interface designed around a slit lamp biomicroscope. Imaging was performed at 800 nm wavelengths with ~10  μm axial resolution in tissue. This work demonstrated the high speed and signal to noise ratios suitable for clinical imaging, a fiber optical implementation with 2-dimensional galvanometer scanning and imaging optics suitable for the retina, a visible aiming beam, and control/real time imaging software as well as software for correcting axial eye motion. This instrument was used at the New England Eye Center to image over 5000 patients in the first studies of major ophthalmic diseases. Figure 1.9 shows an example of an in vivo OCT retinal image spanning the macula and optic nerve heard [23]. For retinal imaging, safety standards govern the

Fiber Optic Interferometer Low Coherence Light Source

Scanning Reference Delay

Detector

Beam Scanning Mirrors

Electronics Computer

Ocular Lens Operator View

Beamsplitter

1  The Development of Optical Coherence Tomography Fig. 1.9  OCT image of the normal human retina in vivo. An early OCT image of the normal human retina in vivo. OCT has 10 μm axial resolution and was acquired at 800 nm wavelength. The retinal pigment epithelium, choroid, and retinal nerve fiber layers are visible as highly backscattering layers. OCT can noninvasively visualize and quantitatively measure retinal pathology and has become a standard of care in ophthalmology. (From Hee et al. [23])

maximum permissible light exposure and set limits on OCT imaging speeds [21, 23]. Typical retinal images have signal levels −50 to −90  dB below the ­incident intensity, but the high detection sensitivity of OCT enables them to be easily visualized using incident power levels that are well within safe retinal laser exposure limits. To date OCT has had the largest clinical impact in ophthalmology. Early clinical studies investigated OCT for the diagnosis and monitoring of a variety of macular diseases [24], including macular edema [25], macular holes [26] and age-related macular degeneration and choroidal neovascularization [27]. The retinal nerve fiber layer thickness, an indicator of glaucoma, can be quantified in normal and glaucomatous eyes and correlated with conventional measurements of the optic nerve structure and function [28, 29]. Many of the measurement protocols that were developed in these early studies were adopted in the first commercial OCT ophthalmic instruments [30, 31]. OCT is a powerful technique in ophthalmology because it can identify markers of early disease at treatable time points before visual symptoms and irreversible vision loss occurs. Furthermore, repeated imaging can be performed to track disease progression or monitor the response to therapy. The process of translating OCT research into commercial development began in 1992 when an MIT startup company called Advanced Ophthalmic Devices was founded. Advanced Ophthalmic Devices was acquired by Carl Zeiss Meditec in 1994 and Zeiss introduced its first commercial product for ophthalmic diagnostics in 1996. Early instruments had an axial resolution of 10 μm and an imaging speed of 100 A-scans/s. A third generation ophthalmic instrument, the Stratus OCT, was introduced in 2002 which had similar resolution, but a faster speed of 400 A-scans/s. The increased speed enabled an increase in image pixel density. The large number of published clinical studies from early generation instruments, coupled with technological improvement helped drive the clinical adoption of OCT. By the mid 2000s, time domain OCT (TD-OCT) had become a standard of care in ophthalmology and became essential for the diagnosis

9

250 µm

250 µm Log Reflection

and monitoring of many retinal diseases as well as glaucoma with an estimated ~10  M procedures per year [4, 5, 31]. With increases in imaging speed provided by spectral/ Fourier domain detection (see later section of this chapter), the a­ doption of OCT in ophthalmology continued to grow. Today it is estimated that worldwide, ~30 million ophthalmic imaging procedures are currently performed annually [4, 5]. OCT has been especially important for monitoring anti-VEGF treatments for neovascular age related macular degeneration [32, 33]. It is estimated that this has saved in excess of ten billion dollars in health care costs in the USA alone and improved eye care to tens of millions of people every year.

1.4.3 Beyond Ophthalmology to Intravascular and Endoscopic Imaging Detection sensitivity and tissue optical properties are especially important for imaging because backscattered light signals are weak and light is highly attenuated by scattering, absorption and aberration, limiting the imaging depth. In the eye, the path from the cornea to the retina (aqueous, lens, and vitreous) is highly transmitting, although backscattered light from the retina is very weak. Early ophthalmic OCT imaging used 800 nm wavelength light because absorption in the lens and vitreous are minimum at this wavelength. OCT imaging using longer wavelengths was an important advance for other clinical specialties because it reduced scattering and increased image penetration depths [2, 3, 34]. Figure  1.10 shows an early example from Brezinski et  al. 1996 showing OCT imaging in a human epiglottis ex vivo, comparing imaging at 850 nm and 1300 nm wavelengths and demonstrating improvement in imaging depth [3]. At wavelengths shorter than ~1100 nm, the dominant absorbers in most tissues are melanin and hemoglobin, which absorb at visible and near infrared wavelengths. Water absorption has a peak at 1450  nm and is extremely high for wavelengths longer

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850 nm

500 um g

g

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1300 nm g

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Fig. 1.10  OCT in scattering tissues is possible using long wavelengths which are less attenuated by scattering. OCT of the human epiglottis ex vivo at 850 nm and 1300 nm wavelength. Superficial glandular struc-

than 1900  nm. In most tissues, scattering at near infrared wavelengths is 10–20  dB higher than absorption and scattering decreases for longer wavelengths. Therefore, imaging at 1300  nm became a standard wavelength for most OCT applications in clinical specialties other than ophthalmology. Early studies by Schmitt et al. investigated how tissue optical properties produce OCT image contrast [34, 35]. In OCT images, tissue structures are visible because they have different optical scattering, absorption, and aberration properties. OCT images show the true dimensions of tissue (correcting for index of refraction and optical refraction effects) and do not suffer from shrinkage or microtoming artifacts that can occur in histology. However if an OCT image is displayed using a false color scale, the colors represent different optical properties and not necessarily different tissue morphologies. Histological sections are stained in order to produce contrast between different tissue structures such as nuclei vs stroma. OCT relies on intrinsic differences in optical properties of different tissues to produce image contrast. On one hand, this is a limitation because tissue structures, such as cell nuclei, may not have high contrast in OCT images. However, histology requires tissue excision, processing, embedding, sectioning and staining, while OCT imaging can be performed on tissue in situ and in real time, without the need for excision. The concept of an “optical biopsy” was explored in several early OCT imaging studies using ex vivo surgical specimens [36–45]. These studies used the paradigm of comparing OCT imaging to the gold standard of histopathology in order to develop a baseline for OCT image interpretation.

tures (g) are visualized with both 850 nm and 1300 nm wavelength, but underlying cartilage (c) is better visualized with longer 1300 nm wavelength. Scale = 500 μm (From Brezinski et al. [3])

Reflectance

Fig. 1.11  Early OCT image of atherosclerotic plaque ex vivo and corresponding histology. The plaque is highly calcified with relatively low lipid content and a thin intimal cap. This result demonstrated that OCT can resolve morphological features associated with unstable plaques. Scale = 500 μm. (From Brezinski et al. [3])

The concept of using OCT to assess vulnerable plaque was first proposed by Brezinski et al. in 1995 [3, 46]. Figure 1.11 shows an example of one of the first OCT images of plaque ex vivo and corresponding histology from Brezinski et  al. 1996 [3]. The OCT image has an axial resolution of ~15 μm in tissue and imaging was performed at 1300 nm wavelength to optimize imaging depths. The figure shows a potentially unstable plaque characterized by a thin intimal cap layer,

1  The Development of Optical Coherence Tomography

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adjacent to a heavily calcified plaque with low lipid content. The demonstration that OCT could resolve plaque features in ex vivo specimens was an important ­milestone which helped drive the technological, commercial and clinical development of intravascular OCT imaging.

1.4.4 Technology for Catheter and Endoscopic OCT Imaging The development of flexible fiber optic imaging probes, catheter and endoscopic device technology by Tearney, Bouma and colleagues at MIT enabled intravascular as well as endoscopic OCT imaging [47–51]. Figure  1.12 shows one of the first OCT catheter/endoscopes which was a prototype for modern OCT endoscopic probes and intravascular imaging catheters. The catheter/endoscope has a singlemode optical fiber in a hollow rotating torque cable, coupled to a distal lens and microprism that reflects the OCT beam in an orthogonal radial direction. The cable and distal optics are enclosed in a transparent housing. The OCT beam is scanned in a radar like pattern by rotating the cable to generate a cross-sectional image through luminal structures or hollow organs similar to the way IVUS catheters generate images. Imaging can also be performed in a longitudinal plane by push-pull movement of the fiber optic cable assembly [52]. Volumetric 3D imaging can be performed by spiral scanning with a combination of rotation and pull-back. The early catheter/endoscope shown in Figure  1.12 had a diameter of 2.9 French or 1 mm, similar to a standard IVUS catheter. The development of catheter imaging devices is challenging because of the simultaneous mechanical, optical and biocompatibility requirements. The first commercial intravascular devices, the ImageWire™ and Helios™ occlusion balloon catheter, were developed by LightLab Imaging. These devices used micro-optic fabrication methods to create lenses and beam-directing elements with the same diam-

eter as that of a single mode optical fiber (80–250 μm) and which are much smaller than IVUS catheters [53]. In vivo studies in endoscopic OCT imaging are procedurally simpler than intravascular OCT and initially progressed more rapidly [51, 54]. Figure 1.13 shows an example of OCT imaging of the rabbit esophagus in vivo and corresponding histology, from Tearney et  al. in 1997 [51]. This image demonstrates visualization of esophageal layers including the mucosa (m), submucosa (sm), inner muscularis (im), outer muscularis (om), serosa (s) and used the rotary probe design of Fig. 1.13. The first clinical studies of endoscopic OCT imaging in human subjects were reported by Sergeev et al. in 1997 [54] and Feldchtein et  al. in 1998 [55]. These studies were performed using a flexible forward scanning fiber optic probe in the working channel of a standard endoscope, bronchoscope or trocar. The probe was 1.5–2 mm diameter and used a miniature magnetic optical scanner to image in the forward direction. These early studies demonstrated the feasibility of performing clinical OCT imaging of organ systems such as the esophagus, larynx, stomach, urinary bladder and uterine cervix. However, OCT imaging for cancer detection and biopsy guidance proved to be challenging because of limitations in image resolution and difficulties in visualizing nuclei.

1.5

Early Intravascular OCT Imaging

The development of intravascular OCT was challenging because it required imaging technology, medical devices, ex vivo studies, preclinical studies and clinical imaging studies combined with complex regulatory and reimbursement issues. Figure 1.14 shows the first catheter-based image of a human coronary artery ex vivo using an early prototype 2.9 F OCT catheter from Tearney et  al. in 1996 [50]. The figure compares OCT to 30 MHz IVUS. The excellent differentiation of the intima, media, and adventitia and suggested the advantages of intravascular OCT.

500 µm fiber

transparent sheath

guard

microprism microprism aiming beam GRIN lens

speedometer cable

light GRIN lens

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Fig. 1.12  Schematic and photograph of an early OCT catheter/endoscope for intraluminal imaging. A single-mode fiber is housed in a rotating flexible speedometer cable and covered by a transparent plastic sheath. The distal end focuses the beam at ~90° from the catheter axis

and scans a rotary pattern and pullback pattern. The catheter diameter is 2.9 French or ~1 mm and is shown on a United States coin for scale. OCT can be integrated with a wide range of diagnostic and interventional devices

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a

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sm

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im om s a

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2x Fig. 1.13  Endoscopic OCT image of the rabbit esophagus in vivo demonstrating internal body imaging. (a) Esophageal layers including the mucosa (m), submucosa (sm), inner muscular layer (im), outer mus-

OCT

cular layer (om), serosa (s), and adipose and vascular supportive tissues (a) can be visualized. (b) A blood vessel (v) can be seen within the submucosa. (c) Corresponding histology. (From Tearney et al. [51])

IVUS

Fig. 1.14  Early OCT image of a human artery ex vivo and comparison with intravascular ultrasound (IVUS). The OCT image has 15 μm axial resolution and enables the differentiation of the intima, media, and adventitia. Intimal hyperplasia is evident. (From Tearney et al. [50])

1.5.1 Preclinical Feasibility Studies In vivo intravascular OCT imaging required developing improved catheter imaging devices which could be used in animals and human subjects. In addition, since blood scatters and attenuates light, it was necessary to develop saline flushing or balloon occlusion protocols to displace blood, or to significantly dilute the hematocrit in the artery. Preclinical

imaging studies in animals were an important step toward developing technology and understanding the feasibility of clinical applications. Intravascular OCT animal imaging studies were performed in a rabbit model by Fujimoto et al. in 1999 [56]. Imaging was performed using a 2.9 F optical catheter using a time domain OCT system with a broadband femtosecond laser at 1280 nm wavelength to achieve 10 μm axial resolution. Imaging speeds were 4 frames/s with a 512

1  The Development of Optical Coherence Tomography

axial pixels/image. In these studies, saline flushing was used to dilute the hematocrit during imaging. Intravascular OCT in a porcine model was reported by Tearney et  al. in 2000 [57]. The study compared OCT and intravascular ultrasound (IVUS) in five swine using an OCT catheter and saline flushing. The study investigated normal coronary architecture, intimal dissections, and stents demonstrating that OCT could better visualize architectural morphology in dissection and stent apposition than IVUS.

1.5.2 Histological Validation Studies In parallel with technology feasibility studies, it was important to investigate OCT image features and their clinical significance. The first step was to establish correspondence of OCT images with histopathology. The early studies by Brezinski et  al. investigated the correlation between OCT and histology on ex vivo aortic specimens with atherosclerotic plaque [3, 46]. As mentioned previously, the studies demonstrated that the high image resolution of OCT enabled visualization of the thin intimal cap layers associated with vulnerable plaques called thin cap fiberatheromas. They also demonstrated that OCT images had high contrast for lipids with low backscatter compared with fibrous tissues, enabling differentiation of intramural lipid deposits, and that OCT could penetrate calcified plaques. A comprehensive study by Yabushita et al. in 2002 investigated the correlation between ex vivo OCT images and histology on arterial specimens with

a

13

fibrous, fibrocalcific, and lipid rich plaques [58]. A blinded reading of OCT images by two readers using a 50 specimen training set and a 307 specimen study set achieved sensitivities ranging from 71% to 97% and specificities ranging from 90% to 98% for the three plaque morphologies compared with histology. The highest sensitivities and specificities, (>90%) with excellent inter and intraobserver agreement, were obtained for fibrocalcific plaques. This study demonstrated that OCT had high sensitivity and specificity compared to histology for identifying plaque morphologies.

1.5.3 Clinical Studies Intravascular OCT imaging in patients was first reported by Jang et al. in 2001 [59]. This pioneering study used a 3.2 F OCT imaging catheter and compared OCT with IVUS, Fig. 1.15. The OCT image showed tissue prolapse between stent struts, which was not obvious on IVUS imaging. The study represents a landmark and was especially significant because it addressed multiple technological, clinical and administrative challenges. Imaging studies in patients using the commercial LightLab Imaging instrument were also independently reported by Grube et al. at the Segeberg Heart Center in 2002 [60]. This study compared OCT to IVUS imaging at 6  months post intervention, showing the ability of OCT to visualize in-stent restenosis of a drug-eluting stent. A study by Jang et al. in 2002 compared intravascular OCT and IVUS in ten patients [61]. The comparison to

b

Fig. 1.15  First demonstration of intravascular OCT imaging in humans. IVUS (a) and OCT (b) images of the stented right coronary artery in vivo. OCT clearly visualized tissue prolapse between the stent struts (12 to 3 o’clock). (From Jang et al. [59])

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IVUS was important because it was the gold standard for intravascular imaging morphological structure. OCT axial image resolution was 13 ± 3 μm, compared to 98 ± 19 μm for IVUS. The comparison of OCT and IVUS showed that all fibrous plaques, calcification and echolucent regions visualized by IVUS were also visualized by OCT and that OCT could identify features such as intimal hyperplasia and intramural lipid deposits better than IVUS. In order to develop manufacturable instruments and devices, enable multi-center clinical studies, address regulatory and reimbursement issues, and make advances available to the wider clinical community, commercial development is necessary. This process of translating intravascular OCT for commercial development began when the MIT startup company LightLab imaging was co-founded by Mark Brezinski, James Fujimoto and Eric Swanson in 1998. Commercial development of intravascular OCT is described in a later section of this chapter.

1.6

Advances in Imaging Speed

Early intravascular OCT studies demonstrated feasibility, however imaging speed was a serious limitation. Since blood strongly scatters light and attenuates OCT signals, saline flushing or a combination of occlusion with flushing was required to dilute the hematocrit during OCT imaging. However limitations on the injected saline volume as well as concerns about myocardial ischemia, restricted the imaging time and the amount of image data that could be acquired. Therefore, increasing OCT imaging speed was critical for achieving clinically practical intravascular OCT imaging. Fortunately, there were powerful advances in OCT detection technology which enabled dramatic increases in imaging speeds. These techniques are known as Fourier domain OCT and include spectral domain OCT and swept source OCT [62–71]. Swept source OCT has also been termed optical frequency domain imaging (OFDI) [72]. Early OCT instruments used a low-coherence light source and interferometer with a scanning reference delay path. This detection method is known as time domain OCT (TD-OCT). However, it is also possible to perform detection in the Fourier domain using a low coherence interferometer with a broadband light source, measuring the interference spectrum with a spectrometer and a high speed line scan camera [66–69, 73]. This method is known as spectral domain OCT (SD-OCT) or Spectral Radar and was first demonstrated more than two decades ago in 1995–1996 [62, 63]. In 2003, multiple different research groups, working independently, demonstrated that spectral domain detection has a powerful sensitivity advantage over time domain detection, since spectral domain detection essentially measures all of the echoes of light simultaneously [66–69]. This discovery drove a boom in OCT research and

development. The sensitivity is enhanced by the ratio of the axial resolution to the axial imaging depth. For most OCT systems, this corresponds to a sensitivity increase of 50–100 times, enabling a corresponding increase in imaging speeds. Spectral domain detection is well suited for high speed imaging at 800  nm wavelength and has had a powerful impact on clinical ophthalmic OCT.  State of the art commercial SD-OCT instruments operate with 5 um axial resolution and 80 kHz A-scan rates, enabling volumetric imaging and mapping of the retina. SD-OCT is the standard technology for ophthalmic imaging and is used in the majority of commercial instruments. However, the majority of OCT applications in other clinical specialties used SS-OCT. Swept source OCT (SS-OCT), is a second type of Fourier domain detection, uses an interferometer with a frequency-­ swept light source and detectors which measure the interference as a function of time. The basic concept of SS-OCT was described in patents as early as 1991 [47, 48] and experimental studies performed also over two decades ago in 1997 [64, 65]. SS-OCT has the advantage that it does not require a spectrometer and line scan camera as in SD-OCT. It can perform imaging at 1300 nm wavelengths where silicon based camera lack sensitivity. Furthermore, SS-OCT can achieve very fast imaging speeds because it is not limited by camera reading speed and is potentially lower in cost and more compact than SD-OCT systems. Future advances in photonic integration used in optical communications, promise to dramatically enhance the functionality of SS-OCT while dramatically reducing size and cost [74, 75]. The primary technological challenge in SS-OCT is that it requires high speed, frequency swept light source which has a narrow linewidth (high frequency purity).

1.6.1 Swept Source/Fourier Domain OCT Swept source/Fourier domain detection uses an interferometer with a narrowband light source which is frequency swept in time [47, 48, 64–67, 70, 71]. The time dependent frequency essentially labels different time delays in the light beam, which can then be detected by interferometry. These detection techniques were used in the 1990s to measure fiber optical and photonics components [76–78]. The basic concepts of frequency-chirped coherent ranging and imaging, along with the associated fundamental signal-to-noise and performance equations and system tradeoffs, date back to the 1980s when they were widely known and used in coherent laser radar [79]. Figure 1.16 shows a schematic of how swept source detection works. Light from a frequency swept light source is divided into a signal and a reference beam. The signal beam is direct onto the tissue and is backscattered from internal structures at different depths, while the reference

1  The Development of Optical Coherence Tomography

15 Long Delay ∆L

Frequency

Short Delay ∆L

Reference

Sample

time

Beam Splitter

∆L

Detector

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Detector Output

Frequency Sweep

time

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Fig. 1.16 How swept source OCT works. Swept source/Fourier domain OCT uses an interferometer with a narrow band, frequency swept laser and detectors. The Michelson interferometer interferes two frequency sweeps which are time delayed with respect to each other and

generates a beat frequency which is proportional to the path length mismatch ΔL. Fourier transforming the beat signal from the detector yields axial scan information (magnitude and relative time delay). Spatial features are encoded in frequency, somewhat analogous to MR imaging

beam is reflected from a mirror at a fixed delay. Light signals from different tissue depths will be delayed and have a different frequency compared to the reference beam. When the signal is interfered with the reference, the intensity of the interference will vary according to the frequency difference between the signal and reference. Longer delays will produce higher frequency variations. An A-scan measurement of the magnitude and echo delay of light from the tissue time can be measured by digitizing the interference signal during the frequency sweep, correcting any nonlinearities in the sweep versus time, then extracting the interference frequencies by Fourier transforming. The axial image resolution and A-scan rate in SS-OCT are determined by the sweep bandwidth and sweep repetition rate of the laser, respectively. The spectral line width of the laser determines the imaging range of imaging depth of the OCT system. SS-OCT is somewhat analogous to magnetic resonance imaging, where spatial information is encoded in frequency. Fourier domain detection measures all of the optical echoes simultaneously, while time domain detection measures echoes sequentially. Therefore, SS-OCT collects more signal from each scatterer for a given A-scan rate and has a dramatic improvement in detection sensitivity compared with TD-OCT. In addition, SS-OCT can achieve much faster scan speeds than either TD-OCT or SD-OCT, because lasers can be frequency swept much faster than mechanical path length scanning or camera reading rates. Swept source detec-

tion also has the advantage that it can image at wavelengths greater than 1000 nm, where silicon-based cameras lack sensitivity and more expensive InGaAs cameras are required. Because swept source OCT can achieve high imaging speeds at wavelengths of 1300 nm, it has had a powerful impact on intravascular imaging applications. OCT using swept source detection was also demonstrated as early as 1997 by Chinn et  al. [64] and Golubovic et  al. [65], but the sensitivity advantages were not fully recognized and performance was limited by the available laser technology. Pioneering studies by Yun et al. in 2003 demonstrated OCT imaging with 19,000 A-scans/s and 13–14  μm axial resolution in air [70]. Imaging speeds of 115,000 A-scans/s were achieved in 2005 using a novel swept laser technology based on a grating and rotating polygon mirror filter [80]. The performance of swept source OCT systems depends directly on the swept light source, therefore advances in OCT speeds paralleled advances in laser technology. Volumetric SS-OCT imaging was demonstrated in the porcine esophagus and coronary artery at speeds of 10,000 and 54,000 A-scans/s, respectively in 2006 and 2007 using swept lasers tuned with a high speed rotating polygon and diffraction grating filter [81, 82]. However, there is a fundamental limit on the tuning rate of a frequency swept laser because the light at a given frequency must build up from spontaneous emission and be amplified up to a saturation level as the frequency is tuned [83]. Therefore, lasers with short cavities,

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which have a rapid build up time can frequency sweep faster than long cavity lasers. The development of Fourier domain mode locking (FDML) in 2006 by Huber et  al. provided a novel approach for breaking the laser speed limit [84]. FDML lasers use a long optical fiber delay inside the laser, which stores the entire frequency sweep in the cavity and synchronously tunes the filter such that a given frequency arrives at the filter when it is tuned to pass it. This avoids the need to build up lasing at each frequency and enables rapid sweep speeds. Early FDML lasers achieved record imaging speeds of 370,000 A-scans/s [85]. FDML technology was enabling for intravascular/endoscopic as well as ophthalmic applications. Record endoscopic imaging speeds of 100,000 A-scans/s with 5–7  μm axial image resolution were demonstrated in an animal model in 2007 using swept source OCT with FMDL lasers [86]. These high speeds enabled imaging at 50  frames/s and data rates of 61  megavoxels/s. 3D-OCT enables virtual manipulation of tissue geometry, speckle reduction, synthesis of en face views similar to endoscopic images, generation of virtual cross-sectional images with arbitrary orientation, and quantitative measurements of morphology. Swept-source OCT was adopted in commercial instruments where LightLab introduced the first commercial swept-source OCT for intravascular imaging in 2009. Figure 1.17 shows an example of high-speed OCT intravas-

Fig. 1.17 High speed intravascular imaging using swept source OCT. Intravascular OCT pullback image of human coronary artery in vivo using manual contrast injection. Imaging was performed at 100 frames/s, at an axial scan rate of 45,000 axial scans/s. Swept source OCT enabled the acquisition of larger data sets with reduced imaging times and contrast volumes

J. G. Fujimoto et al.

cular imaging in an artery in vivo using a LightLab prototype instrument with an FDML laser that operates at 45,000 A-scans/s, acquiring 100  frames/s with a 15  mm/s pullback speed using contrast injection to remove blood from the imaging field. These high imaging speeds enable rapid pullbacks and mapping of long segments of an artery while minimizing ischemia. In other applications, endoscopic OCT imaging using FDML lasers achieved a 62,000 A-scans/s speed with an axial image resolution of 5  μm in tissue in 2009 [87]. Ophthalmic imaging with speeds greater than 1 million A-scans/s have recently been achieved using swept source OCT with FDML laser technology [88]. Intravascular OCT at records speeds of 2.88 million A-scans/s was demonstrated by the group at Erasmus University Medical Center and the University of Lubeck in 2015 in a technique known as “Heartbeat OCT”, reflecting the fact that the entire pullback was acquired during a single cardiac cycle [89, 90]. Figure  1.18 shows an example of a 3D-OCT image of the coronary artery. The catheter (Fig. 1.18) used a distal micromotor for beam scanning and could achieve frame rates of up to 5600 B-scans/s, more than 10× faster than commercial instruments which use proximal torque cable scanning. Pullback was performed at speeds up to 100 mm/s, acquiring images during the ~0.5 s diastolic phase, dramatically reducing motion artifacts. The frame pitch (space between repeated B-scans) was ~25 to 30 μm, enabling high resolution volumetric imaging. Since the acquisition times are short compared to the 2–5 s times in commercial instruments, the volume of contrast flushing and ischemic times were reduced. These results suggest the future potential of OCT to achieve higher speeds and perform very high resolution 3D volumetric imaging. In addition to research prototype lasers, commercial swept laser technologies also made important advances. The development of a compact and robust short cavity swept laser by Axsun Technologies enabled intravascular as well as ophthalmic applications. The short cavity swept laser uses a tunable filter and a gain medium with a few centimeter cavity length. This laser was used in the LightLab C7XR ™ system introduced in 2009 which operated at 50,000 A-scans/s to achieve imaging speeds of 100 frames/s. This technology is also used for retinal imaging at a 1050  nm wavelength, imaging with ~5  μm axial resolution at 100,000 A-scans/s [91]. Commercial SS-OCT ophthalmic instruments are manufactured by Topcon, Heidelberg, Zeiss, Tomey, and others. Vertical cavity surface emitting laser (VCSEL) technologies developed by Praevium/Thorlabs and others enable even higher performance SS-OCT and are promising for SS-OCT systems in multiple medical specialties [92]. The laser cavity is only a few microns in length and the laser is swept using a miniature micro-electro-mechanical systems (MEMS) mirror enabling MHz sweep repetition rates. The VCSEL can also achieve long imaging ranges and has the advan-

1  The Development of Optical Coherence Tomography

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a

b

c

Fig. 1.18  Heartbeat OCT. (a) Micromotor based catheter performs distal beam scanning at up to 5600 revolutions/s. The total outer diameter is 1.6 mm with 1.2 mm diameter inner probe. (b) OCT images of coronary artery and metal stents acquired at 3000  frames/s and 100 mm/s pullback speed. The high speed enables imaging during dias-

tole, reducing motion artifacts. (c) 3D reconstruction of metal stent. Dense frame spacing generates high quality volumetric images. (Courtesy of Robert Huber, University of Lubeck, Germany and Tianshi Wang, Erasmus Medical Center, Netherlands)

tage that the sweep speed and resolution can be adjusted. VCSELs have been used in ophthalmic SS-OCT, imaging at 1050 nm wavelengths with speeds of 400,000 A-scans/s and for endoscopic imaging at 1300 nm with speeds of 600,000 to 1,000,000 A-scans/s [93–95]. The wide acceptance and large market for ophthalmic OCT as well as developments for other clinical specialties will be synergistic for technological advances in intravascular OCT.

lation and commercial development has many technological and clinical challenges, as well as engineering, manufacturing and regulatory and business hurdles. These factors create a long path to the market, especially for organizations that pioneer a new technology or new method where there is limited prior clinical evidence and no established market. LightLab Imaging, Inc. (initially called Coherent Diagnostic Technology) was founded in 1998 as an MIT startup (co-founded by Mark Brezinski, James Fujimoto and Eric Swanson) and released the first intravascular TD-OCT system in 2004. LightLab also released the first SS-OCT system in 2009. LightLab was acquired by Goodman Corporation in 2002, then acquired by St. Jude in 2010. St. Jude Medical was subsequently acquired by Abbot Labs in 2016. LightLab remains the market leader in intravascular OCT by a substantial margin, followed by Terumo who released their SS-OCT product in 2013. For over a decade, from 2004 to 2013, LightLab/St. Jude Medical was the only commercial manufacturer of intravascular OCT.  This early commercial development made intravascular OCT available to the clinical community and accelerated clinical progress. The majority of clinical studies to date have been performed using LightLab/St. Jude Medical/Abbott Laboratories instruments.

1.7

Technology Translation and Commercial Development

Commercial development is critical for the translation of any new medical technology because it enables widespread clinical access, facilitating multi-center clinical trials which are needed to establish efficacy, support reimbursement and ultimately impact patient care. The translation of intravascular OCT from basic research, to clinical research and patient care has been successful due to a worldwide ecosystem involving government funding, researchers (scientists, engineers, and clinicians), and industry. Startups as well as large corporations have been involved in the development of intravascular OCT. However the process of technology trans-

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Multiple generations of coronary intravascular imaging systems have been developed. The first generation intravascular OCT instrument was the M2 imaging system with which the first clinical studies were performed in Europe in 2002. The M2 system used time domain detection and operated at 15 frames/s (200 A-scans per frame at 3000 A-scans/s) using saline flushing and occlusion. Regulatory approval with CE marking in Europe was obtained in 2004, followed by a release in China in 2006, and Ministry of Health, Labor and Welfare approval in Japan in 2007. The next generation, M3 system, had increased imaging speeds of 20  frames/s (240 A-scans per frame) and was introduced in Japan in 2008. Over 10,000 cases were performed using time-domain OCT with the M2 and M3 systems, providing key insights into engineering design and clinical utility. An important advance came with the commercial release of SS-OCT, which dramatically increased imaging speeds, enabling large pull-back data sets to be rapidly acquired within a much faster acquisition time, eliminating the need for a occlusion balloon and minimizing ischemia. Introduced in Europe in 2009 and in the U.S. in 2010, the C7XR ™ system achieved imaging speeds of 100  frames/s (500 axial scans per frame). The tenfold increase in imaging speed enabled higher frame rates for improved pullback speeds and arterial coverage as well as increased axial scan density for improved image quality. Occlusion-free imaging also became possible by using a radiographic contrast agent rather than a saline injection to dilute hematocrit.

Designed for integration into the cathlab workflow, the Ilumien™ OCT imaging system was introduced in 2011 with wireless fractional flow reserve (FFR). The Ilumien Optis™, a second generation system with FFR and increased imaging speed (180  frames/s) was introduced in Japan in 2012 and is detailed in Figs.  1.19 and 1.20. The Illumien Optis PCI Optimization System™ which received the CE mark and was introduced in Europe in 2013 used advanced software analysis of 3D vessel structure for stent planning. The Optis™ Integrated System received CE Mark and FDA clearance in 2014. This system directly integrated OCT with the cardiac catheterization laboratory, shown in Fig. 1.20, avoiding the need and some negative attributes of a cart-­based instrument, while providing combined angiographic and OCT visualization. Most recently in 2016 the Optis™ Mobil System was introduced in Japan and Europe, providing OCT with co-registration angiography using a portable system which could be transferred between multiple catheterization laboratories. Intravascular OCT has also been developed by other companies through new corporate initiatives and startup companies, both approaches often involving industry academic partnerships [5]. Terumo began development of cardiovascular OCT in 2004, establishing a co-development program with Massachusetts General Hospital. A first-in-man study using the Terumo Lunawave™ was reported by Okamura et  al. in 2011 at the Rotterdam Thoraxcenter, comparing OCT with IVUS and quantitative coronary angiography [96]. In 2013, Terumo announced the first commercial release of

Fig. 1.19  Graphic user interface from the AptiVue™ Software which received FDA clearance in April 2019. The system provides automated measurements and 3D vessel structure for interventional planning along

with advanced software analysis for stent apposition and expansion. (Courtesy of Abbott Laboratories)

1  The Development of Optical Coherence Tomography

19

Fig. 1.20  The OPTIS™ Integrated System received CE Mark and FDA clearance in 2014. This system directly integrates OCT with the cardiac catheterization laboratory, while providing combined angiographic and OCT visualization. (Courtesy of Abbott Laboratories)

its intravascular OCT product the Lunawave™ in Europe and Japan. CardioSpectra was a 2005 startup from the University of Texas that was acquired by IVUS leader Volcano in 2007. However in 2013, after several years of investment and development, Volcano announced the cancellation of its cardiovascular OCT product development effort for a variety of reasons. Avinger is another startup company founded in 2007 that uses OCT for guiding atherectomy in peripheral arteries and received FDA clearance in 2012 and has had over 1000 procedures in peripheral artery applications. Today, there are numerous commercial efforts in the intravascular OCT market space, both established companies and startups with released products or in the process of entering the market. These include: Abbott, Argus, Avinger, Cannon, Conavi Medical, Dyad Medical, Forssman, Gentuity, Horimed, lnnerMedical, OCT Medical Incorporated, SpectraWave, Terumo, VivoLight, and several others. The intravascular OCT market size continues to experience good growth with the number of intravascular OCT procedures ~100,000/year with the potential to grow much larger depending on the outcome of pending clinical trials. With all the new corporate players, continued innovation in areas such as multimodality (e.g. OCT/IVUS, OCT/NIR), higher speeds, higher-resolution, advanced image analysis, software assisted or software automated disease detection powered by artificial intelligence, novel catheters and guidewires, and other areas should also increase market size and further improve the technology and its ability to benefit patients.

development and clinical adoption. The largest publication volume is in ophthalmology. To date OCT has had the most clinical impact in ophthalmology where it has become the standard of care, providing structural, functional, and quantitative information that cannot be obtained by any other modality. OCT improves the diagnosis of retinal disease as well as monitoring disease progression and response to therapy. In ophthalmology, OCT played a major role in the development of new anti-VEGF pharmaceutical therapies for age-related macular degeneration. It is the clinical standard for evaluating anti-VEGF treatment response and has saved well over $10B billion dollars in health care costs in the US [32, 33]. The second largest publication volume at the time of this writing is in cardiology. Intravascular OCT has advanced the understanding of cardiovascular diseases including plaque morphology and its role in myocardial infarction as well as percutaneous coronary interventions such as stent implantation. OCT is a powerful imaging modality for cardiovascular research and is increasingly being used in interventional decision-making. The number of clinical cases has surpassed ~100,000 case per year; however, this is a small fraction of the total yearly PCI procedures and the extent to which intravascular OCT will be used every day for clinical cardiovascular decision making is still to be determined. Multiple clinical studies have been completed or are underway which use OCT to investigate plaque pathogenesis, plan intervention, and assess treatment response. The evidence from these studies will ultimately determine the role that intravascular OCT will have in interventional cardiology. However intravascular OCT has already had a fundamental impact in elu1.8 Conclusion cidating plaque morphology and the mechanisms for major OCT is a powerful imaging technology in biomedical adverse cardiac events and is being used in an increasing research and medicine because it enables high-speed, high-­ number of percutaneous interventions. The underlying techresolution, in situ visualization of tissue structure, function, nology of OCT is continuing to advance, improving funcand pathology that often cannot be obtained by any other tionality and utility. OCT imaging promises to continue to means. Figure 1.21 plots journal publications involving OCT be a key tool for understanding and treating cardiovascular versus year. Journal articles are an indicator of for scientific disease.

20

J. G. Fujimoto et al.

Scientific Publications

5000

Surgery Over 33,000 Cumulative Publications with OCT in Title or Abstract in PubMed

1st Gastroenterology OCT product

Number of Publications

4000

Other Non-Medical Microscopy NDE/NDT

1st Dermatology OCT product 3000

Oral Cavity (not Dentistry) Gynecology Bronchoscopy & Pulmonology Developmental Biology

1st Cardiovascular OCT product 2000

Urology Otolaryngology

1st Ophthalmic OCT product

Other Medical Dentistry

1st “OCT” publication

1000

Neurology Gastroenterology & Endoscopy Dermatology Technology

1991 1992 1993 1994 1995 1996 1997 1998 1999 2000 2001 2002 2003 2004 2005 2006 2007 2008 2009 2010 2011 2012 2013 2014 2015 2016 2017 2018

0

Year

Cardiovascular Ophthalmology

Fig. 1.21  OCT Publication volume versus year (these are for pubmed articles with OCT in the title or abstract). Publications are an indicator of scientific and clinical and economic progress and patient impact. The largest volume of publications is in ophthalmology, where OCT is now a clinical standard. Recently, cardiovascular became the second largest

volume of publications and is growing rapidly. The graph also shows the release dates of key products in the different fields. Commercial introduction of products is a catalyst for advancing clinical research as shown by the growth in publications

References

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1  The Development of Optical Coherence Tomography 94. Liang K, Traverso G, Lee HC, Ahsen OO, Wang Z, Potsaid B, Giacomelli M, Jayaraman V, Barman R, Cable A, Mashimo H, Langer R, Fujimoto JG. Ultrahigh speed en face OCT capsule for endoscopic imaging. Biomed Opt Express. 2015;6: 1146–63. 95. Ahsen OO, Liang K, Lee HC, Giacomelli MG, Wang Z, Potsaid B, Figueiredo M, Huang Q, Jayaraman V, Fujimoto JG, Mashimo H. Assessment of Barrett’s esophagus and dysplasia with ultrahigh-­

23 speed volumetric en face and cross-sectional optical coherence tomography. Endoscopy. 2018;51(4):355–9. 96. Okamura T, Onuma Y, Garcia-Garcia HM, van Geuns RJM, Wykrzykowska JJ, Schultz C, van der Giessen WJ, Ligthart J, Regar E, Serruys PW. First-in-man evaluation of intravascular optical frequency domain imaging (OFDI) of Terumo: a comparison with intravascular ultrasound and quantitative coronary angiography. EuroIntervention. 2011;6:1037–45.

2

Histology Validation of Optical Coherence Tomography Images Teruyoshi Kume, Takashi Kubo, and Takashi Akasaka

2.1

Introduction

High-resolution optical coherence tomography (OCT) can often provide superior delineation of coronary vessel wall structure when compared with intravascular ultrasound (IVUS). OCT can reliably visualize the microstructure (i.e., 10–50 μm versus 150–200 μm for IVUS) of normal and atherosclerotic arteries [1]. However, OCT images are fundamentally different from histology, and it is not clear whether all characteristics identified by histology can also be characterized by OCT. Therefore, careful interpretation of OCT images is required. This chapter reviews the histological validation of OCT images.

2.2

 ormal Vessel Wall Morphology N and Intimal Thickening

In normal coronary artery wall, the media of the vessel appears as a lower signal intensity band relative to that of the intima and adventitia, providing a three-layered appearance (bright–dark–bright) (Fig. 2.1). Intimal thickening occurs during the early phase of atherosclerosis in the coronary artery [2]. Evaluation of intimal thickening of the coronary artery is clinically important, and recent advancements in IVUS have enabled pathological evaluation of the wall structure of the coronary artery. However, intimal thickening is indirectly evaluated as the intima– media thickness by IVUS because the boundary of the

T. Kume Division of Cardiology, Kawasaki Medical School, Kurashiki, Japan e-mail: [email protected] T. Kubo (*) · T. Akasaka Department of Cardiovascular Medicine, Wakayama Medical University, Wakayama, Japan e-mail: [email protected]; [email protected]

intima and media cannot be distinguished completely by this method [3]. On the other hand, OCT is superior to IVUS in terms of qualitative delineation of the vessel wall structure, and OCT can differentiate the intima from media. There is good agreement in the intimal thickness between OCT and histological examination (r  =  0.98, p 300 μm), which is a higher cutoff than suggested by previous investigations [4]. It must be emphasized however that dealing with MA is a relatively difficult subject. First, MA often (but not necessarily) comes along with UE. When following currently available protocols for correcting stent UE, the accompanying MA will be corrected to some extent in the same maneuver in many cases. This will however not always be the case, as there will be circumstances where there will remain some MA struts whatever postdilatation action is taken (e.g. in an asymmetrically expanded stent against a calcified burden in the vessel wall). Furthermore, when a stent is implanted in a segment with some focal aneurysmal dilatation, one will not want to aim for full apposition of stent struts in this specific segment, as this would lead to over dilatation of the device, with inherent disadvantages and risks. Finally, when MA is observed in an acquisition later than the initial stent implantation procedure and no intracoronary imaging data of this first implantation procedure are available, it will be impossible to discriminate late persistent from late acquired MA.  This will then further impact the ability to adequately define the exact mechanism of ST.

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T. Adriaenssens

Fig. 10.5  Severe in-stent restenosis with superimposed thrombus. A severe in-stent restenotic lesion is observed in panel (a). The lumen area is hardly larger than the imaging catheter cross-section. It is hypothesized that in such circumstances, disturbance of bloodflow through the lumen can cause formation of small thrombi (indicated with a red arrow in panel (a)). This can lead to total occlusion of the stented segment, with resulting ST elevation myocardial infarctio (STEMI). Often, the vessel will only be subtotally occluded, resulting in a non ST elevation MI (NSTEMI). It is not always easy do define whether neoatherosclerosis plays a role or is also present in cases of in-stent restenosis with a

thick neointimal layer. In panel (b), the stent struts seem to be surrounded by neointimal tissue with a darker appearance border (indicated with blue arrow) than the tissue more close to the lumen border. As this discrimination is relatively sharp, this is more probably to be considered as a peri-strut low intensity area. Between 12 and 4 o’clock, an accumulation of macrophages, characterized as hyperintense dots casting a sharp shadow is observed (yellow arrow, magnification in panel (c)). A small neointimal neovascularization is observed around 5 o’clock (indicated with a green arrow). The asterisk is indicating the guide wire artifact

10.3.7 In-Stent Restenosis (Fig. 10.5)

10.4 Bioresorbable Vascular Scaffold Thrombosis

Finally, there is an overlap between in-stent restenosis (ISR) and stent thrombosis (ST), when taken together named ‘late stent failure’. In the PRESTIGE registry, in 12.7% of cases, important restenosis was identified as the dominant mechanism of ST [6]. Classical examples are severe cases of ISR, where lumen encroachment gradually becomes so important that thrombus originates at the site of severe lumen narrowing, leading to sudden occlusion of the vessel. On the other hand, some cases of ST can present without symptoms, so that the thrombus later gradually gets incorporated, contributing to ISR later on, or when completely occlusive, the problem is discovered only later at the time of a routine control examination and the occluded vessel is considered a pattern IV ISR lesion. The edges of a stented segment form a particular place with predilection for the development of restenosis, due to plaque overgrowth promoted by the action of the stent edges [25].

10.3.8 Coronary Evaginations, Bifurcation Disease Apart from the earlier described categories of VLST, some less frequent problems can be observed as well. These include, among others, plaque rupture in the edge segments of the stented segment and coronary evaginations (Fig. 10.6). Fig. 10.7 illustrates the origin of ST at the side of a coronary bifurcation.

10.4.1 Scaffold Thrombosis (Fig. 10.8) Fully bioresorbable coronary scaffolds were introduced to overcome the limitations of metallic stents associated with permanent caging of the arterial wall aiming to restore vessel physiology, as well as to mitigate the longterm risk of device-related adverse events, including restenosis and stent thrombosis [26]. It was anticipated that the risks of ST after DES implantation would be solved with these new devices, which offer the possibility of transient scaffolding of the vessel to prevent acute vessel closure and recoil while also transiently eluting an antiproliferative drug to counteract constrictive remodeling and excessive neointimal hyperplasia [27]. However, with a déjà vu resembling the DES storm after early enthusiastic adoption of the technology, reports of (early and late) scaffold thromboses and signals of a safety signal in some trials, confidence in the safety of the technology started to wane. After an increased risk of scaffold thrombosis, compared to the well-known metallic everolimus-­eluting stent, at 3 years had been documented in the ABSORB II trial [28] and a higher risk of target vessel MI and early, late and very late device thrombosis at 2 years in the AIDA trial [29], the device was retracted from the market in 2017.

10  Very Late Stent Thrombosis

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Fig. 10.6  Coronary evaginations (a–d). Coronary evaginations, also termed interstrut cavities, is believed to originate from an inflammatory reaction in the vessel wall, leading to positive vascular remodeling and a retraction of the vessel wall away from the stent struts. This causes evaginations in between stent struts, and, possibly by a tearing force on

the neointima, struts to become uncovered over time. The combination of evaginations between struts, and the uncovered struts themselves, is a nidus for stent thrombosis. Thrombus is indicated with a red arrow. The asterisk is indicating the guide wire artifact

10.4.2 The INVEST Registry

(2.6%). (1) Scaffold discontinuity represents the loss of the circular scaffold strut lining and occurs as a result of device degradation, that is, after the structural integrity is no longer maintained (12 months in Absorb BVS) [31]. Discontinuous struts can penetrate into the lumen and cause (late acquired) malapposition, thus exposing the highly trombogenic remnant scaffold material to the blood flow, with subsequent activation of the coagulation cascade. The mechanisms that lead to discontinuous struts resulting in malapposition (i.e. as opposed to discontinuity embedded in neointima, the expected

The INVEST (Independent OCT Registry on Very Late Bioresorbable Scaffold Thrombosis) registry studied OCT characteristics of BVS thrombosis in 38 lesions in 36 patients [30]. VLScT occurred at a median of 20 months after implantation. The mechanisms underlying VLScT were scaffold discontinuity (42.1%), malappositon (18.4%), neoatherosclerosis (18.4%) underexpansion or scaffold recoil (10.5%), uncovered struts (5.3%), and edge-related disease progression

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consequence of strut resorption) remain poorly understood. Of concern, scaffold discontinuity was also observed in cases preceded by a state of fully apposed and covered struts [30]. Physical forces toward the lumen center such as vessel con-

T. Adriaenssens

traction or external shear stress might be responsible for conditions allowing scaffold strut segments to migrate into the lumen despite the presence of tissue coverage. With the biodegradation process of BVS results in an inflammatory peak reaction approximately 1 year after implantation, a potent inflammatory response in some individuals leading to a necrotizing vasculitis with subsequent penetration of scaffold remnants might represent an alternative biological explanation for the phenomenon of scaffold discontinuity. Because the 2 leading mechanisms underlying VLScT are associated with transient strut material inside the lumen, extension of DAPT until the time of complete scaffold resorption (i.e. 3–4 years) appears to be important to mitigate the risk of scaffold thrombosis among patients with an Absorb BVS while carefully balancing the associated bleeding risk [30].

10.4.3 BVS Neoatherosclerosis

Fig. 10.7  VLST at the level of a coronary bifurcation. The bifurcation level, especially after complex intervention with a double stent technique, is a predilection place for the occurrence of ST. The combination of remnant stent material, often with uncovered struts protruding deeply into the lumen, and altered fluid dynamics, predisposes the bifurcation segment to thrombotic complications. Thrombus is indicated with a red arrow. The side branch ostium is indicated with ‘SB’. Accumulating stent material with a blue arrow. The asterisk is indicating the guide wire artifact

Fig. 10.8  Very Late Scaffold Thrombosis. Very Late thrombosis in a BVS implanted 18 months earlier. In panel (a), clear underexpansion and asymmetrical expansion of the device is observed. From 2 to 7 o’clock, the circular shape of the device is lost. A BVS strut is indicated with a blue arrow. Most plausibly, this is due to asymmetric expansion

Neoatherosclerosis seems to occur earlier after BVS compared to DES.  Otsuka et  al showed that inflammation surrounding struts was greater in BVS at 6 to 36 months than in metallic DES in a porcine model [32]. A more intense inflammation during the absorption process could accelerate the recurrence of atherosclerotic lesions within neointima because of impaired endothelial function, reportedly a key trigger of neoatherosclerosis after DES implantation [13]. The lessons learned from OCT imaging in VLScT have important implications for the design of future bioresorbable scaffold iterations (design, polymer composition, degradation).

at the time of scaffold implantation, due to insufficient lesion preparation at the level of a calcified plaque. In panel (b), more proximal in the scaffolded segment, a region with neoatherosclerosis (indicated with ‘L’) and a small adjacent thrombus (red arrow) is observed. The asterisk is indicating the guide wire artifact

10  Very Late Stent Thrombosis

10.5 Impact on PCI Procedure In the PESTO study, physicians reported that the ST origin could be identified with certainty by angiography in only 12% of cases whereas this percentage rose to 41% with OCT analysis, pointing out the very poor specificity of angiographic tools to assess ST etiology. Balloon angioplasty and medical treatment alone were provided in 37 and 29% of the cases, which is lower than in other published works. In a series of 7135 ST analyzed by angiography, Armstrong et al. reported use of redo stenting in >50% of cases [33]. It is our strong conviction that a detailed assessment and understanding of the origin of stent thrombosis, will importantly impact the therapy delivered to the patient. In cases were an important burden of uncovered struts is seen as the cause for VLST, it would be unwise to implant even more metal into the vessel. Rather, after restoration of patency of the vessel, an adequate lifelong antiplatelet regimen should be installed. In cases of underexpansion and malapposition, adequate balloon dilation measures seems most appropriate. In some cases, OCT will be able to even

Fig. 10.9  Multiple causes of VLST. A patient suffered acute inferior STEMI, after she had been treated with a bare metal stent, and subsequently with a first-generation DES for ISR in the right coronary artery (RCA), many years before. Some days before the acute event, dual antiplatelet therapy had been interrupted abruptly at the time of admission for a serious intraabdominal problem, for which urgent surgery had been performed. After implantation of a temporary PM and restoration of TIMI III flow with a thrombus aspiration catheter, successful OCT image acquisition in the RCA was performed, with adequate clearance of the lumen from blood and excellent image quality. The coronary angiogram is shown in panel (b) with corresponding locations of the OCT frames as marked by white lines. In panel (a), some minor coronary evagination is observed, unrelated to the ST problem. In panel (d),

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define underlying causes for underexpansion, such as important undilatable calcium burdens, which can than possible be tackled by novel therapies, such as lithotripsy. The implantation of additional DES would then be reserved for cases of in-stent restenosis, neoatherosclerosis and stent edge disease.

10.6 Conclusion Establishing an accurate diagnosis for the cause of ST is necessary in order to implement appropriate therapy [5]. Even in the setting of the acute presentation with ST, most often, it is possible to achieve good imaging quality with OCT.  From these acquisitions, a dominant mechanism of ST can be withheld in OCT in the great majority of patients. In many cases, multiple morphological abnormalities, interplaying with each other, are observed (Fig. 10.9). The specific characteristics observed in the OCT images hold promised for a precise and individualized approach towards each ST case.

distal from the thrombus region as well, some residual disease, with moderate lumen encroachment, is observed, with a calcium burden (indicated with ‘c’) between 4 and 7 o’clock. Panel (e) shows some ISR, at the level of the earlier implanted BMS, where no coverage with a subsequent DES had been performed. Panels (c) and (f) show the region with overlapping stent struts, multiple uncovered struts (indicated with blue arrows) and thrombus (red arrow) attached exactly to these struts. In order to avoid further addition of metallic components, the patient was treated with balloon dilation solely, in order to decrease the thrombus burden, and treated with life-long dual antiplatelet therapy. Together with the mechanistic problem of delayed healing, the proinflammatory and prothrombogenic postoperative state certainly contributed to the occurrence of ST in this particular case

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10  Very Late Stent Thrombosis Satogami K, Yamano T, Kameyama T, Orii M, Ota S, Kuroi A, Kitabata H, Tanaka A, Hozumi T, Akasaka T.  Optical coherence tomography predictors for edge restenosis after everolimus-eluting stent implantation. Circ Cardiovasc Interv. 2016;9:e004231. 26. Sabate M, Windecker S, Iniguez A, Okkels-Jensen L, Cequier A, Brugaletta S, Hofma SH, Raber L, Christiansen EH, Suttorp M, Pilgrim T, Anne van Es G, Sotomi Y, Garcia-Garcia HM, Onuma Y, Serruys PW. Everolimus-eluting bioresorbable stent vs Durable polymer everolimus-eluting metallic stent in patients with st-­ segment elevation myocardial infarction: results of the randomized absorb st-segment elevation myocardial infarction-trofi ii trial. Eur Heart J. 2016;37:229–40. 27. Sotomi Y, Suwannasom P, Serruys PW, Onuma Y. Possible mechanical causes of scaffold thrombosis: insights from case reports with intracoronary imaging. EuroIntervention. 2017;12:1747–56. 28. Serruys PW, Chevalier B, Sotomi Y, Cequier A, Carrie D, Piek JJ, Van Boven AJ, Dominici M, Dudek D, McClean D, Helqvist S, Haude M, Reith S, de Sousa AM, Campo G, Iniguez A, Sabate M, Windecker S, Onuma Y. Comparison of an everolimus-eluting bioresorbable scaffold with an everolimus-eluting metallic stent for the treatment of coronary artery stenosis (ABSORB II): a 3 year, randomised, controlled, single-blind, multicentre clinical trial. Lancet. 2016;388:2479–91. 29. Wykrzykowska JJ, Kraak RP, Hofma SH, van der Schaaf RJ, Arkenbout EK, IJsselmuiden AJ, Elias J, van Dongen IM, RYG T,

137 Koch KT, Baan J Jr, Vis MM, de Winter RJ, Piek JJ, JGP T, JPS H. Bioresorbable scaffolds versus metallic stents in routine pci. N Engl J Med. 2017;376:2319–28. 30. Yamaji K, Ueki Y, Souteyrand G, Daemen J, Wiebe J, Nef H, Adriaenssens T, Loh JP, Lattuca B, Wykrzykowska JJ, Gomez-Lara J, Timmers L, Motreff P, Hoppmann P, Abdel-Wahab M, Byrne RA, Meincke F, Boeder N, Honton B, O’Sullivan CJ, Ielasi A, Delarche N, Christ G, Lee JKT, Lee M, Amabile N, Karagiannis A, Windecker S, Raber L.  Mechanisms of very late bioresorbable scaffold thrombosis: the invest registry. J Am Coll Cardiol. 2017;70:2330–44. 31. Sotomi Y, Onuma Y, Collet C, Tenekecioglu E, Virmani R, Kleiman NS, Serruys PW. Bioresorbable scaffold: the emerging reality and future directions. Circ Res. 2017;120:1341–52. 32. Otsuka F, Pacheco E, Perkins LE, Lane JP, Wang Q, Kamberi M, Frie M, Wang J, Sakakura K, Yahagi K, Ladich E, Rapoza RJ, Kolodgie FD, Virmani R.  Long-term safety of an everolimus-­ eluting bioresorbable vascular scaffold and the cobalt-chromium xience v stent in a porcine coronary artery model. Circ Cardiovasc Interv. 2014;7:330–42. 33. Armstrong EJ, Feldman DN, Wang TY, Kaltenbach LA, Yeo KK, Wong SC, Spertus J, Shaw RE, Minutello RM, Moussa I, Ho KK, Rogers JH, Shunk KA.  Clinical presentation, management, and outcomes of angiographically documented early, late, and very late stent thrombosis. JACC Cardiovasc Interv. 2012;5:131–40.

OCT for Bioabsorbable Vascular Scaffold

11

Alessio Mattesini, Antonio Martellini, Luigi Tassetti, and Carlo Di Mario

Abbreviations BRS BVS DES ISA ISR IVUS OCT PCI PLLA RCTs ScT TLF TLR TVR VLScT

Bioresorbable scaffold Bioresorbable vascular scaffold Drug-eluting stent Incomplete strut apposition In-stent restenosis Intravascular ultrasound Optical coherence tomography Percutaneous coronary intervention Poly L-lactic acid Randomized controlled trials Scaffold thrombosis Target lesion failure Target lesion revascularization Target vessel revascularization Very late scaffold thrombosis

11.1 Introduction Bioresorbable scaffolds (BRSs) were designed to overcome the limitations of currently available second-generation drug-eluting stent (DES) [1]. In fact, BRS, after the initial phase when radial force is needed to avoid lesion recoil and scaffolding is crucial to seal plaque rupture and vessel dissections, is progressively resorbed leaving the vessel free from a permanent metallic caging. Thus, a reduction in the risk of late stent thrombosis, which is considered related to delayed arterial healing, incomplete endothelialization, and neoatherosclerosis of DES [2], was expected, together with further advantages such as restoration of the physiologic vasomotion and of positive remodeling, feasibility of coronary non-invasive imaging, and maintenance of the chance

A. Mattesini · A. Martellini · L. Tassetti · C. Di Mario (*) Structural Interventional Cardiology Unit, Careggi University Hospital, Florence, Italy e-mail: [email protected]

for further options of revascularization (percutaneous or surgical). These theoretical benefits were only partially confirmed in the follow-up data. In particular, safety concerns were raised by an unexpected high incidence of scaffold thrombosis with poly-L-lactic acid (PLLA)-made Absorb (Abbott Vascular, Santa Clara, CA, USA) [3]. Intravascular imaging may have a crucial role in improving the outcome of patients treated with BRS. Among the different intravascular imaging techniques, optical coherence tomography (OCT) was considered as the gold standard for BRS evaluation both for percutaneous coronary intervention (PCI) guidance, during the resorption phase follow-up, and for scaffold failure assessment. This pivotal role of OCT is due to multiple reasons. First of all OCT has higher resolution compared with intravascular ultrasound (IVUS) and enables identification of potentially harmful technical pitfalls such as scaffold fracture, malapposition, or underexpansion which are often missed using angiography alone. OCT enables clear identification of the mechanism of BRS failure guiding further interventions. Finally thanks to its unique characteristics OCT has been traditionally the imaging tool used to assess the resorption process of BRS in the clinical practice.

11.2 I maging with OCT and Different Types of BRS Many different resorbable materials have already been tested to produce BRS, but the most intensively used are the poly-­ L-­lactic acid (PLLA) polymers and magnesium.

11.2.1 PLLA Based BRS Most of the scaffolds are PLLA based and, despite the withdrawal of the Absorb stent (Abbott Vascular, Santa Clara, CA, USA)) from the market, Absorb is still undergoing clinical evaluation in controlled studies, and many competitors received CE mark and can still be implanted at least in Europe. PLLA scaffolds generation have some common characteris-

© Springer Nature Switzerland AG 2020 I.-K. Jang (ed.), Cardiovascular OCT Imaging, https://doi.org/10.1007/978-3-030-25711-8_11

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Fig. 11.1  OCT appearance of a partially resorbed ABSORB scaffold and of a DESOLVE scaffold after deployment. In this OCT cross section, the partially resorbed struts of an Absorb implanted since 2 years are visible at outer edge of the lumen vessel as “opened black boxes” (red star) covered by neointimal hyperplasia. The DESOLVE struts of a second scaffold implanted to treat Absorb failure are well apposed to the vessel wall but not embedded (yellow arrow)

tics. Struts are thicker as compared with second-­generation DES with worse rheological abnormalities, especially in smaller vessels [4]. Despite discordant findings in in-vitro studies, the general clinical experience is that polymer-­based stents have a reduced radial strength which translates in more pronounced lesion recoil. Overexpansion cannot be used to overcome this limitation because most polymeric scaffolds have narrow expansion limits and fracture in case of aggressive post deployment overexpansion. Probably the most important limitation of polymeric stents is the difficulty in the identification of an optimal length of polymer chains to maintain full mechanical integrity in the first months after implantation with an acceptable total duration of the resorption process. OCT was selected from the beginning as the essential imaging tool to assess PLLA BRS and monitor the reabsorption process. BRS polymeric struts are optically translucent and appear as a black central core framed by lightscattering borders that do not shadow the vessel wall, with the full thickness of the scaffold strut visualized and even small malappositions detected (Figs. 11.1 and 11.2).

11.2.2 Magnesium Alloy-Based BRS Pure magnesium has low strength and rapid corrosion and it would not be suitable, but the addition of other elements such as zinc and manganese, undergoing a thorough purifica-

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Fig. 11.2  Strut malapposition. Example of malapposed Absorb struts due to scaffold undersizing

tion and anodization process, allowed the development of metallic bioresorbable scaffold with sufficient radial strength to achieve an acute gain comparable with that of other metallic stents. The electronegative properties related to the anodization supply less thrombogenicity to the platform [5]. The degradation of the strut occurs in two phases, initially with an anodic reaction of the magnesium in water to generate magnesium hydroxide subsequently converted into amorphous calcium phosphate [6]. Magmaris (Biotronik AG, Bülach, Switzerland) is the only magnesium BRS currently approved in Europe. Magnesium has completely different structural properties than PLLA, and the OCT appearance of Magmaris immediately after implantation is similar to that of a permanent metallic stent (e.g. a bright strut with well-­ delimited borders and posterior shadowing) [7]. This appearance gradually changes during the absorption process with rapid disappearance of magnesium but frequent persistence of “black” boxes caused by the precipitation of calcium.

11.3 O  CT for Procedural Guidance During BRS Implantation 11.3.1 Basal OCT Assessment Over the past years, intracoronary imaging techniques have been extensively investigated, and growing evidence from observational studies, randomized controlled trials (RCTs), and meta-analyses supports their use in order to achieve better procedural results and to improve clinical outcomes. The

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2018 ESC guidelines on myocardial revascularization consider a Class IIa recommendation for the performance of intracoronary imaging, by IVUS or OCT, to optimize stent implantation [8]. In contrast to permanent metallic stents, where the role of intracoronary imaging to guide and optimize stent implantation have been well established, bioresorbable vascular scaffolds disappeared from the clinical arena too quickly to allow similar observations. As mentioned before, individual studies and meta-analyses revealed an increased risk of scaffold thrombosis at all time points following BRS implantation at least up to 3 years [9]. These findings affected so much the credibility of bioresorbable therapy that the ABSORB BRS has been withdrawn for all clinical use with the exception of controlled study protocols. The embedding force during deployment is low resulting in only partly embedded and largely exposed thick scaffold struts, despite high-pressure post-dilatation. When added to the long persistence of the polymer or of equally thrombogenic biochemical compounds (proteoglycans) and the interruption of double antiplatelet therapy, it is obvious the greater risk of late and very late thrombus formation. Due to these inherent mechanical limitations, accurate lesion preparation especially in the presence of calcium or fibrosis is an even more critical step during BRS implantation. OCT can accurately detect length and circumferential extension of superficial calcium, with many criteria proposed for the use of dedicated devices such as Rotablator or scoring/cutting balloons. After pretreatment, OCT is important to check whether an adequate dilatation and fracture of the calcium plates have

been obtained, allowing safe use of larger and higher pressure balloons if needed. The quantitative measurement of lesion length and identification of position and diameter of proximal and distal references allow correct BRS sizing, avoiding geographical missing and scaffold undersizing which may translate into grossly malapposed struts or scaffold fracture because of their already mentioned limitations.

Fig. 11.3  Quantification of scaffold malapposition. Evaluation of incomplete struts apposition (ISA) is an objective way for estimating the magnitude of scaffold malapposition. This evaluation is feasible

using OCT with PLLA scaffold which are radiolucent. ISA is the result of luminal area (violet contour) minus scaffold area (green contour)

11.3.2 Post Implantation Assessment Several OCT parameters can be collected immediately after scaffold implantation in order to optimize the procedure. Key quantifications include the lumen and scaffold areas, the magnitude of incomplete strut apposition (ISA), in-scaffold prolapse, flow area, eccentricity index, and the presence of strut fracture and edge dissection. When a PLLA scaffold is used, the struts are translucent with OCT, the vessel lumen border can be visualized, and the vessel lumen area delineated along the external (abluminal) side of the struts. The scaffold area is measured by joining the internal middle points of the abluminal side of the black cores of the apposed struts or the abluminal edge of the frame borders of malapposed struts. In the absence of ISA and plaque prolapse, the scaffold area is identical to the lumen area. ISA is defined by a clear separation between the abluminal side of the strut and the vessel wall. ISA area is delineated by the abluminal side of the frame border of the malapposed struts and the endoluminal contour of the lumen (Fig. 11.3). In case of lumen prolapse, the prolapse

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Fig. 11.4  Example of tissue prolapse. This OCT cross section shows tissue protruding between the struts of an Absorb scaffold. The quantitative assessment of tissue prolapse can be assessed as the difference

between the scaffold area (green contour) and the vessel luminal area (violet contour)

area can be estimated by the planimetered difference between the prolapsed contour (i.e., lumen contour) and the scaffold area (Fig. 11.4). An intraluminal defect that is separated from the vessel wall (e.g., thrombus) can also be quantified as an area. Flow area takes into account ISA, plaque prolapse, and intraluminal defects, and it is defined as the difference between the sum of the scaffold and ISA areas and the sum of the areas of intraluminal struts, prolapse, and intraluminal defect. The eccentricity and symmetry of implanted BRSs can be easily assessed by OCT, and these parameters have been shown to be associated with clinical outcomes after metallic stents; the eccentricity ratio is defined as the ratio of the minimum and maximum diameters of the scaffold in each frame. The eccentricity index is obtained by calculating the average of all eccentricity ratios along the length of the scaffold. The diagnosis of short-term strut fracture due to balloon overdilation can be established if two struts overhang each other within the same angular sector of the lumen perimeter. This complication may be observed with or without concomitant strut malapposition. However, if isolated struts are located more or less at the center of the vessel without an obvious connection with other surrounding struts, strut fracture may also be present. Edge dissection is defined by OCT as the disruption of the endoluminal vessel surface at the proximal and distal edges of the BRS (Fig. 11.5). The use of OCT for guidance of BRS implantation has been assessed in different clinical studies. The ABSORB trial has provided the first description of the OCT appearance of the scaffolds. In the ABSORB cohort B study, enrolling 101 patients treated with the ABSORB everolimus-eluting

scaffold (Abbott Vascular, Santa Clara, California), OCT was performed in two groups of patients, respectively at 6 and 24  months (B1, n  =  45) and 12 and 36  months (B2, n  =  56). Baseline imaging was optionally performed in 51 patients. The results showed the maintenance of the scaffold area at distance, with a slight decrease in luminal area, as a consequence on neointimal proliferation inside the BRS. In fact, 97% of the struts were covered at 1 year, with the scaffold displaying initial signs of resorption, but still largely visible. In addition, the comparison of post-procedural and late imaging allowed the discrimination between the acute scaffold disruption, which has been associated with an increased risk of target lesion failure (TLF) and thrombosis, and late discontinuities, representing the natural consequence of the reabsorption process [10]. Gomez-Lara et al. performed an OCT substudy of the ABSORB trial Cohort B, in which only 3-mm-diameter BRS devices were deployed [11]. These investigators found a higher incidence of malapposed struts in vessels with a maximal diameter of >3.3 mm. These data emphasize the importance of correct vessel size measurement to select the appropriate BRS diameter. The role of OCT examination during BRS implantation is also highlighted by a study which showed that an extensive use of OCT for guidance of BRS implantations allowed achieving similar acute performance than second-generation metallic DES even during the treatment of complex coronary lesions [12]. Authors compared 50 complex coronary lesions (all type ACC/AHA B2-C) treated with BRS under OCT guidance matched to an equal number of lesions treated with second-generation DES. Results revealed a similar incidence

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Fig. 11.5  OCT characteristics of BRS strut fractures. (a, b) Cross sections at the level of ostial side branch show a scaffold pattern irregularity with an overhanging strut (white arrows) in the center of the vessel without obvious connection to the expected/adjacent strut pattern

of residual area stenosis and overall percentage of incomplete stent apposition between the DES and BRS groups. Mean and minimal lumen area were similar in the two groups with also a similar mean and minimum eccentricity as well as symmetry index between the two groups. A retrospective analysis of more than 200 consecutive BRS implanted in 101 patients [13] showed that almost half of the OCT examinations led to a change in strategy before and/or after scaffolds implantation. When used before, OCT images suggested additional lesions preparation and allowed fine-tuning of the length and size of BRSs used. When used as a final control, OCT-pullback led to further post-dilatation in almost one-­ third of the cases despite the aggressive systematic angiography-­ guided optimization technique used in the study. To date, although many experts acknowledge the role for intracoronary imaging-guided scaffold implantation to mitigate scaffold failures, there are no evidences from RCTs and one was stopped prematurely (OPTICO BVS) following retraction of ABSORB BVS.

11.4 Resorption Process Evaluation The main chemical process responsible for in vivo molecular weight degradation of BRS is hydrolysis. There are no enzymatic processes and no cellular or inflammatory process involved, given that neither macromolecules nor cells can penetrate the polymer backbone; water alone drives the

hydrolysis reaction. The timing of the whole process is of paramount importance for its performance. In fact the mechanical integrity and the absence of recoil should continue over a period of 6 months, during which time the biological process of restenosis decreases (according to this a permanent device would be possibly unnecessary beyond this time) [14]. As mentioned before, OCT analysis of bioresorbable scaffolds reveals essential differences from metal stents. With metal stents, metal struts are preserved and there is a clear definition between stent and lumen contour with a shadow behind the metal. Poly-lactic acid BRS does not create a shadow behind each strut and at long-term follow-up, are no longer visible, as the struts are filled by fibrous tissue with similar optical properties as the underlying fibrous layer. The vascular structure observed at the scaffolded segment, a product of the solidification of underlying plaque, biodegraded struts, and neointima, is similar to a native atherosclerotic plaque and it was defined as neoplaque. In order to evaluate and separate the different composition of the scaffolded vessel and the underlying neoplaque elements, an indirect assessment has been provided, based on the signal-­ rich layer, which consisted of the neointimal layer, resorbed struts, and pre-existing fibrous tissue [15]. Both pre-clinical and clinical studies described Absorb BVS 1.0 strut features during the different stages of the absorption process [16]. The whole reabsorption process has been demonstrated at OCT, and qualitative assessment of BRS struts over time has been proposed as follows: preserved box, open box, dissolved

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Fig. 11.6  OCT appearance of BRS struts over time. Progressive morphologic phases of BRS resorption process are shown: preserved box (a), up to 12–24 months after implantation; open box (b), at 24–36 months;

bright box, and dissolved black box in order of decreasing reflectivity. Immediately after implantation, the polymeric struts are clearly identified. The bright reflection borders contrast with a black core, giving the typical aspect called “preserved box”. After 6  months, there is a gradual disappearance of this pattern, which is considered to be the initial stages of the bioresorption process. This aspect has been called “open box pattern”, where there is an opening of the extremities of the box in its short axis. The “dissolved bright box” appearance, with partially visible bright spot and poorly defined contours, is considered the further stages of bioresorption. “Dissolved black box”, a black spot with poorly defined contours and no box-shaped, is the fourth and later stage which correspond Absorb BVS degradation and vessel wall integration (Fig. 11.6).

c

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dissolved black box (c) and dissolved bright box (d), at 42–48 months after BRS implantation

lesion failure and device thrombosis between 1 and 3 years after BRS implantation [18]. Several retrospective analyses have been performed with the aim to detect the main pathomechanisms for bioresorbable scaffold thrombosis. In this context OCT, given its high spatial resolution, was demonstrated to play a central role allowing the detection of structural abnormality leading to scaffold thrombosis. These studies revealed that the main mechanisms for both early and late bioresorbable scaffold thrombosis are similar to those of metallic stents and consist of suboptimal implantation with underexpansion, malapposition, and incomplete lesion coverage [19]. A case series by Cuculi et al. revealed that OCT findings are distinctive in early ScT versus late or very late ScT: early ScT was mostly related to underexpansion, while late ScT was associated with lower peristrut light intensity, a surrogate marker of edema which has been associated with malapposition, evaginations, strut fracture, and uncovered 11.5 BRS Failure Assessment: Scaffold struts [20]. Very late scaffold thrombosis (VLScT) is defined if occurring more than 1 year after device implantation and Thrombosis and Restenosis has been described in several isolated reports. A new As BRSs are being increasingly used in real-word popula- bioresorption-­ specific phenomenon, not encountered with tion, several studies have demonstrated high rates of scaffold metallic DES, called strut discontinuity has been described thrombosis (ScT) following implantation with the timing of as a potential cause of VLScT.  Strut discontinuity (or disthe event evenly distributed from acute to very late thrombo- mantling) is present when struts are dislocated into the lumen sis. In the AIDA trial, investigators randomly assigned 1845 despite initial full apposition. Late discontinuity is theoretipatients undergoing percutaneous coronary intervention in cally a benign change during the bioresorption process and routine clinical practice to receive either an everolimus-­ does not cause any problems if the scaffold struts are well eluting bioresorbable scaffold or an everolimus-eluting covered by neointima. However, whenever struts are not covmetallic stent. During follow-up, the 2-year cumulative rate ered by neointima and late discontinuity allows protrusion of of the primary end point (target-vessel failure, which was a part of the struts into the lumen and brings thrombogenic composite of cardiac death, target-vessel myocardial infarc- proteoglycan of the provisional matrix into contact with tion, or target-vessel revascularization) was not significantly blood, it could represent a malignant potential cause of different between the bioresorbable  scaffold and metallic-­ VLScT. The largest cohort study on very late ABSORB BVS stent groups (11.7% vs 10.7%; HR 1.12, 95% CI 0.85–1.48, scaffold thrombosis studied by OCT—the Independent OCT P = 0.43). However, the 2-year cumulative rate of definite or Registry on Very Late Bioresorbable Scaffold Thrombosis probable device thrombosis was significantly higher (INVEST) [21]—included 36 patients, and scaffold discontiwith bioresorbable scaffolds than with metallic stents (3.5% nuity was the most common mechanism underlying VLScT vs 0.9%; HR 3.87, 95% CI 1.78–8.42, P 75% area stenosis. Plaque hemorrhages are more frequently observed at non-­

culprit sites of patients with plaque ruptures than in those dying with severe coronary disease without rupture [20, 21]. Plaque erosion is defined as the presence of acute thrombus overlying a plaque without signs of rupture, in an area of absent or damaged endothelium, and is observed in up to 40% of subjects dying of sudden coronary death [1, 6, 21]. Nevertheless,

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eroded plaques do not a have specific ­morphological substrate such as TCFA for plaque rupture, so that the identification of erosion-prone plaques appears more challenging. The underlying morphology of erosion is frequently a fibrous plaque rich in smooth muscle cells and proteoglycan matrix, but it can be a pathological intimal thickening as well as a thick-cap fibroatheroma [1, 6, 21]. It has been speculated that coronary vasospasm might be involved in the pathophysiology of erosion, as well as local flow disturbances [22, 23] and, more recently, alterations of hyaluronan metabolism [24]. Macrophage accumulation is usually less represented or totally absent in plaque erosions. These lesions tend to be eccentric, and are infrequently associated with calcifications. The most frequent location for plaque erosion is the proximal left anterior descending artery (66%) followed by the right coronary artery (18%) and the left circumflex (14%) [1, 6, 21]. The least frequent thrombotic lesion causing ACS or sudden cardiac death (i.e., about 5–8% of cases) is calcified plaque, which is characterized by superficial substantive calcium without superficial necrotic core. The abluminal area of the plaque shows breaks of the calcified plate, bone formation, and interspersed fibrin with thrombus overlying a disrupted plaque [1, 21].

12.3 I n Vivo Detection of the Vulnerable Plaque by OCT Due to its unprecedented spatial resolution, OCT is often considered the modality of choice for the in vivo identification of what had been defined “vulnerable plaque” by previous histopathology studies [3, 25], enabling the assessment of a number of plaque features, such as lipid content, fibrous cap thickness, and macrophage accumulation, which are not easily detectable by other intravascular imaging modalities [26] (Table 12.1). Historically, the most commonly used definition of TCFA by OCT is a lipid-rich plaque (variably defined as subtending a lipid arc greater than 90° or 1 quadrant) covered by a fibrous cap which is thinner than 65 μm (Fig. 12.3a, b) [25, 27]. However, the debate about the exact definition of a TCFA by OCT remains open, as histological specimens of TCFA are different from OCT cross-sections (largely explained by the anisotropic tissue shrinkage) [28].

It is therefore conceivable that the fibrous cap of a TCFA is thicker in  vivo than on histology [28]. OCT studies have tried to identify the critical fibrous cap thickness characterizing plaques prone to rupture in vivo [29–31]. Pathology studies have demonstrated that ruptured fibrous caps have a mean thickness of 23 ± 19 μm, where 95% of caps are thinner than 65 μm [20]. However, due to the overestimation of fibrous cap thickness by OCT as compared with histology, and to the effect of anisotropic tissue shrinkage during histology sections preparation, these thresholds may not be valid in vivo. Adopting a cut-off value of 70 μm for fibrous cap thickness, only 67% of ruptured plaques in a study including 72 ACS patients resulted to be TCFA [31]. The vast majority of OCT studies focusing on “plaque vulnerability” as well as available consensus documents defined thin cap by using the histology-derived threshold of 65 μm [32– 34]. However, sensitivity and specificity of a fibrous cap thickness of 65 μm for discriminating between ruptured and non-ruptured plaques are relatively low (i.e., 83% and 77%, respectively). As such, not only lipid-rich plaques with a fibrous cap thickness 70%) than in non-critical stenosis (diameter stenosis