Optical coherence tomography
Optical coherence tomography (OCT) is a non-invasive, non-contact imaging technique that employs low-coherence interferometry with near-infrared light to generate high-resolution cross-sectional images of biological tissues, achieving axial resolutions typically between 1 and 15 micrometers.[1][2] Analogous to ultrasound imaging but using light instead of sound waves, OCT measures the echo time delay and intensity of backscattered or reflected light from tissue microstructures to reconstruct detailed subsurface images in real time.[1][3] This method enables visualization of tissue layers at a micron-scale resolution, far surpassing the 150-micrometer resolution of conventional ultrasound, though imaging depth is limited to about 2–3 millimeters due to optical scattering in tissues.[1] Developed in the early 1990s, OCT was first demonstrated in 1991 by researchers using a Michelson interferometer setup to image biological samples, marking a significant advancement in optical biopsy techniques.[1] The technology gained clinical traction in ophthalmology with the introduction of the first commercial device in 1996, initially based on time-domain OCT (TD-OCT), which scanned light sources mechanically.[2] Subsequent innovations, including spectral-domain OCT (SD-OCT) in the early 2000s and swept-source OCT (SS-OCT) later on, improved imaging speed, sensitivity, and quality by leveraging Fourier-domain detection to eliminate mechanical scanning components.[2] In technical realization, OCT systems typically incorporate a low-coherence light source, such as a superluminescent diode, with bandwidths determining resolution—for instance, a 40-nanometer bandwidth at 880 nanometers yields about 7-micrometer axial resolution in devices like the SPECTRALIS system.[2] Modern implementations often integrate OCT with confocal scanning laser ophthalmoscopy for enhanced motion tracking and multimodal imaging, enabling three-dimensional volumetric scans and functional extensions like OCT angiography (OCTA), which visualizes retinal blood vessels without dye injection.[2][3] OCT's primary applications are in ophthalmology, where it serves as a standard diagnostic tool for assessing retinal conditions such as macular degeneration, diabetic retinopathy, glaucoma, and macular edema by quantifying layer thicknesses and detecting subtle structural changes.[3][2] Beyond the eye, it supports cardiology for imaging atherosclerotic plaques in coronary arteries, gastroenterology for evaluating tumor margins during endoscopy, and other fields like dermatology and oncology for non-invasive optical biopsies and guiding minimally invasive procedures.[1] Its non-invasive nature, high resolution, and real-time capabilities have made OCT indispensable for early disease detection, treatment monitoring, and research in biomedical imaging.[1][3]Introduction and Fundamentals
Definition and Overview
Optical coherence tomography (OCT) is a non-contact, low-coherence interferometry technique that enables micron-level cross-sectional imaging of biological tissues using near-infrared light.[2] It provides high-resolution images with axial resolutions typically ranging from 1 to 15 μm, allowing for detailed visualization of tissue microstructures without the need for physical contact or excision.[1] In medical diagnostics, OCT serves as a primary tool for real-time imaging, particularly in ophthalmology, where it assesses retinal and anterior segment structures with exceptional clarity and speed.[2] Unlike techniques involving ionizing radiation such as X-rays, OCT is entirely non-invasive and safe for repeated use, making it ideal for longitudinal monitoring of conditions like glaucoma and macular degeneration.[4] Conceptually analogous to ultrasound imaging, OCT employs light waves instead of sound for precise optical ranging and depth-resolved mapping of scattering media.[1] The general workflow of OCT involves a broadband light source emitting near-infrared light, which is split by a beam splitter into a reference arm and a sample arm directed toward the tissue of interest.[2] The backscattered light from the sample interferes with the reference beam, and the resulting interference pattern is detected and processed to reconstruct the tissue's internal architecture in cross-sectional or volumetric formats.[5] Developed in the early 1990s, OCT has evolved into a standard of care in ophthalmology, with over 150 million scans performed annually worldwide as of 2024.[6]Historical Development
Optical coherence tomography (OCT) was invented in 1991 by David Huang and colleagues at the Massachusetts Institute of Technology (MIT), who demonstrated its potential for noninvasive cross-sectional imaging of biological tissues, with an initial focus on ophthalmic applications such as retinal imaging.[7] This breakthrough built on earlier work in low-coherence interferometry, enabling micron-level resolution without physical contact, and marked the shift from laboratory prototypes to a viable medical imaging tool.[8] Early clinical adoption followed rapidly, with the first in vivo human retinal OCT images acquired in 1993 by Swanson et al., showcasing the technology's ability to visualize retinal layers in real time. The first commercial OCT system, a time-domain device from Carl Zeiss Meditec, received FDA approval in 1996, facilitating broader use in ophthalmology for diagnosing conditions like macular degeneration.[9] By the early 2000s, limitations in imaging speed prompted a pivotal transition to Fourier-domain OCT (FD-OCT), with Wojtkowski et al. demonstrating the first in vivo human retinal tomograms using this approach in 2002, achieving up to 100 times faster scan rates than time-domain methods.[10] Swept-source OCT, a variant of FD-OCT using wavelength-swept lasers, was introduced in 2005, offering deeper tissue penetration and higher speeds for enhanced choroidal imaging.[11] The mid-2000s saw the commercialization of spectral-domain OCT (SD-OCT), with the first FDA-approved system from Optovue in 2006, revolutionizing routine retinal diagnostics through volumetric imaging.[12] In the 2010s, optical coherence tomography angiography (OCTA) emerged as a major advancement, with commercial systems like those from Zeiss and Optovue gaining FDA clearance around 2015-2016, enabling noninvasive visualization of retinal blood flow without dye injection.[13] This period also witnessed OCT's expansion beyond ophthalmology into cardiology and dermatology, driven by improved resolution and integration with other modalities. By the early 2020s, OCT evolved from bulky lab prototypes to portable, handheld devices suitable for point-of-care screening, such as the Nidek RS-3000 Advance2 and emerging smartphone-compatible systems.[14] Integration with artificial intelligence for automated analysis became prominent by 2024, with AI algorithms enhancing image segmentation and disease detection in tools like Altris AI and home-monitoring OCT devices.[15] The global OCT market reached approximately $2.5 billion in 2025, reflecting widespread clinical adoption and ongoing innovations.[16]Principles of Operation
Interferometry Basics
Optical coherence tomography (OCT) relies on low-coherence interferometry, typically implemented using a Michelson interferometer configuration. In this setup, light from a broadband source is directed to a beam splitter, which divides it into two arms: a reference arm containing a movable mirror and a sample arm directed toward the tissue of interest. The light backscattered from the tissue in the sample arm recombines with the light reflected from the reference mirror at the beam splitter, and the resulting interference is detected by a photodetector.[17] Interference fringes are produced only when the optical path lengths of the reference and sample arms match within the coherence length of the light source, enabling depth-resolved detection of tissue reflectivity. The low temporal coherence of the broadband source confines the interference to a narrow axial range, acting as a "coherence gate" that localizes the signal to specific depths in the sample. This condition ensures that backscattered light from different tissue layers interferes selectively with the reference beam, allowing for high-resolution ranging without mechanical scanning in the sample depth direction in some implementations.[17][18] The interference intensity at the detector is given byI = I_r + I_s + 2\sqrt{I_r I_s} \cos(\phi),
where I_r and I_s are the intensities from the reference and sample arms, respectively, and \phi is the phase difference between the arms, which depends on the optical path length mismatch. The cosine term oscillates rapidly, but its envelope corresponds to the axial position of scatterers in the tissue, providing the depth profile.[17][18] Axial scanning to form an A-scan—a one-dimensional reflectivity profile versus depth—is achieved by varying the reference arm length, such as through mechanical movement of the reference mirror, or by spectral analysis of the interference signal. The low temporal coherence plays a critical role in determining the axial resolution, with the coherence length l_c given by
l_c = \frac{2 \ln 2}{\pi} \frac{\lambda_0^2}{\Delta \lambda},
where \lambda_0 is the central wavelength and \Delta \lambda is the full width at half maximum of the source spectrum; broader bandwidths yield shorter coherence lengths and thus finer resolution.[18] To extract the tissue reflectivity profile, the raw interferogram undergoes signal processing, typically involving envelope detection using the Hilbert transform, which computes the analytic signal and yields the magnitude of the interference envelope for direct visualization of depth-resolved scatterers. This process isolates the slowly varying amplitude from the carrier fringe pattern, enhancing the signal-to-noise ratio and enabling accurate A-scan formation.[18]