optical coherence tomography optical coherence tomography optical coherence tomography (oct) is an imaging technique that uses low-coherence light to capture micrometer-resolution, two- and three-dimensional images from within optical scattering media (e.g., biological tissue). it is used for medical imaging and industrial nondestructive testing (ndt). optical coherence tomography is based on low-coherence interferometry, typically employing near-infrared light. the use of relatively long wavelength light allows it to penetrate into the scattering medium. confocal microscopy, another optical technique, typically penetrates less deeply into the sample but with higher resolution. optical coherence tomographyoptical coherence tomography (oct) image of a sarcomameshd041623ops-301 code3-300 depending on the properties of the light source (superluminescent diodes, ultrashort pulsed lasers, and supercontinuum lasers have been employed), optical coherence tomography has achieved sub-micrometer resolution (with very wide-spectrum sources emitting over a ~100 nm wavelength range). optical coherence tomography is one of a class of optical tomographic techniques. commercially available optical coherence tomography systems are employed in diverse applications, including art conservation and diagnostic medicine, notably in ophthalmology and optometry where it can be used to obtain detailed images from within the retina. recently, it has also begun to be used in interventional cardiology to help diagnose coronary artery disease, and in dermatology to improve diagnosis. a relatively recent implementation of optical coherence tomography, frequency-domain optical coherence tomography, provides advantages in the signal-to-noise ratio provided, thus permitting faster signal acquisition. oct is not the same as optical coherence microscopy (ocm), which "is a microscopic incarnation of optical coherence tomography (oct)" that can be used for 3d imaging reconstruction through intrinsic contrasting of back-scattered (coherent) light. introduction optical coherence tomogram of a fingertip. it is possible to observe the sweat glands, having "corkscrew appearance" starting from adolf fercher and colleagues’ work on low-, partial coherence or white-light interferometry for in vivo ocular eye measurements in vienna in the 1980s, imaging of biological tissue, especially of the human eye, was investigated in parallel by multiple groups worldwide. a first two-dimensional in vivo depiction of a human eye fundus along a horizontal meridian based on white light interferometric depth scans was presented at the ico-15 sat conference in 1990. further developed in 1990 by naohiro tanno, then a professor at yamagata university it was referred to as heterodyne reflectance tomography, and in particular since 1991 by huang et al., in prof. james fujimoto laboratory at massachusetts institute of technology, who successfully coined the term "optical coherence tomography". since then, oct with micrometer resolution and cross-sectional imaging capabilities has become a prominent biomedical tissue-imaging technique that continuously picked up new technical capabilities starting from early electronic signal detection, via utilisation of broadband lasers and linear pixel arrays to ultrafast tuneable lasers to expand its performance and sensitivity envelope. it is particularly suited to ophthalmic applications and other tissue imaging requiring micrometer resolution and millimeter penetration depth. first in vivo oct images – displaying retinal structures – were published in 1993 and first endoscopic images in 1997. oct has also been used for various art conservation projects, where it is used to analyze different layers in a painting. oct has interesting advantages over other medical imaging systems. medical ultrasonography, magnetic resonance imaging (mri), confocal microscopy, and oct are differently suited to morphological tissue imaging: while the first two have whole body but low resolution imaging capability (typically a fraction of a millimeter), the third one can provide images with resolutions well below 1 micrometer (i.e. sub-cellular), between 0 and 100 micrometers in depth, and the fourth can probe as deep as 500 micrometers, but with a lower (i.e. architectural) resolution (around 10 micrometers in lateral and a few micrometers in depth in ophthalmology, for instance, and 20 micrometers in lateral in endoscopy). oct is based on low-coherence interferometry. in conventional interferometry with long coherence length (i.e., laser interferometry), interference of light occurs over a distance of meters. in oct, this interference is shortened to a distance of micrometers, owing to the use of broad-bandwidth light sources (i.e., sources that emit light over a broad range of frequencies). light with broad bandwidths can be generated by using superluminescent diodes or lasers with extremely short pulses (femtosecond lasers). white light is an example of a broadband source with lower power. light in an oct system is broken into two arms – a sample arm (containing the item of interest) and a reference arm (usually a mirror). the combination of reflected light from the sample arm and reference light from the reference arm gives rise to an interference pattern, but only if light from both arms have traveled the "same" optical distance ("same" meaning a difference of less than a coherence length). by scanning the mirror in the reference arm, a reflectivity profile of the sample can be obtained (this is time domain oct). areas of the sample that reflect back a lot of light will create greater interference than areas that don't. any light that is outside the short coherence length will not interfere. this reflectivity profile, called an a-scan, contains information about the spatial dimensions and location of structures within the item of interest. a cross-sectional tomogram (b-scan) may be achieved by laterally combining a series of these axial depth scans (a-scan). a face imaging at an acquired depth is possible depending on the imaging engine used. layperson's explanation ocular oct retinal thickness map, right eye time-domain oct of the macular area of a retina at 800 nm, axial resolution 3 µm spectral-domain oct macula cross-section scan. optical coherence tomography (oct) is a technique for obtaining sub-surface images of translucent or opaque materials at a resolution equivalent to a low-power microscope. it is effectively ‘optical ultrasound’, imaging reflections from within tissue to provide cross-sectional images. oct has attracted interest among the medical community because it provides tissue morphology imagery at much higher resolution (less than 10 μm axially and less than 20 μm laterally ) than other imaging modalities such as mri or ultrasound. the key benefits of oct are: live sub-surface images at near-microscopic resolution instant, direct imaging of tissue morphology no preparation of the sample or subject, no contact no ionizing radiation oct delivers high resolution because it is based on light, rather than sound or radio frequency. an optical beam is directed at the tissue, and a small portion of this light that reflects from sub-surface features is collected. note that most light is not reflected but, rather, scatters off at large angles. in conventional imaging, this diffusely scattered light contributes background that obscures an image. however, in oct, a technique called interferometry is used to record the optical path length of received photons allowing rejection of most photons that scatter multiple times before detection. thus oct can build up clear 3d images of thick samples by rejecting background signal while collecting light directly reflected from surfaces of interest. within the range of noninvasive three-dimensional imaging techniques that have been introduced to the medical research community, oct as an echo technique is similar to ultrasound imaging. other medical imaging techniques such as computerized axial tomography, magnetic resonance imaging, or positron emission tomography do not use the echo-location principle. the technique is limited to imaging 1 to 2 mm below the surface in biological tissue, because at greater depths the proportion of light that escapes without scattering is too small to be detected. no special preparation of a biological specimen is required, and images can be obtained ‘non-contact’ or through a transparent window or membrane. it is also important to note that the laser output from the instruments used is low – eye-safe near-infrared or visible-light – and no damage to the sample is therefore likely. theory the principle of oct is white light, or low coherence, interferometry. the optical setup typically consists of an interferometer (fig. 1, typically michelson type) with a low coherence, broad bandwidth light source. light is split into and recombined from reference and sample arms, respectively. fig. 2 typical optical setup of single point oct. scanning the light beam on the sample enables non-invasive cross-sectional imaging up to 3 mm in depth with micrometer resolution. fig. 1 full-field oct optical setup. components include: super-luminescent diode (sld), convex lens (l1), 50/50 beamsplitter (bs), camera objective (co), cmos-dsp camera (cam), reference (ref), and sample (smp). the camera functions as a two-dimensional detector array, and with the oct technique facilitating scanning in depth, a non-invasive three-dimensional imaging device is achieved. fig. 4 spectral discrimination by fourier-domain oct. components include: low coherence source (lcs), beamsplitter (bs), reference mirror (ref), sample (smp), diffraction grating (dg) and full-field detector (cam) acting as a spectrometer, and digital signal processing (dsp) fig. 3 spectral discrimination by swept-source oct. components include: swept source or tunable laser (ss), beamsplitter (bs), reference mirror (ref), sample (smp), photodetector (pd), and digital signal processing (dsp) time domain f o d t ( ν ) = 2 s 0 ( ν ) k r ( ν ) k s ( ν ) ( 3 ) dt\left(\nu \right)=2s_\left(\nu \right)k_\left(\nu \right)k_\left(\nu \right)\qquad \quad (3)} in time domain oct the path length of the reference arm is varied in time (the reference mirror is translated longitudinally). a property of low coherence interferometry is that interference, i.e. the series of dark and bright fringes, is only achieved when the path difference lies within the coherence length of the light source. this interference is called autocorrelation in a symmetric interferometer (both arms have the same reflectivity), or cross-correlation in the common case. the envelope of this modulation changes as path length difference is varied, where the peak of the envelope corresponds to path length matching. the interference of two partially coherent light beams can be expressed in terms of the source intensity, i s , as i = k 1 i s + k 2 i s + 2 ( k 1 i s ) ⋅ ( k 2 i s ) ⋅ r e [ γ ( τ ) ] ( 1 ) i_+k_i_+2i_\right)\cdot \left(k_i_\right)}}\cdot re\left\qquad (1)} where k 1 + k 2 < 1 +k_<1} represents the interferometer beam splitting ratio, and γ ( τ ) is called the complex degree of coherence, i.e. the interference envelope and carrier dependent on reference arm scan or time delay τ , and whose recovery is of interest in oct. due to the coherence gating effect of oct the complex degree of coherence is represented as a gaussian function expressed as γ ( τ ) = exp ⁡ [ − ( π δ ν τ 2 ln ⁡ 2 ) 2 ] ⋅ exp ⁡ ( − j 2 π ν 0 τ ) ( 2 ) \tau \right)\qquad \quad (2)} where δ ν represents the spectral width of the source in the optical frequency domain, and ν 0 is the centre optical frequency of the source. in equation (2), the gaussian envelope is amplitude modulated by an optical carrier. the peak of this envelope represents the location of the microstructure of the sample under test, with an amplitude dependent on the reflectivity of the surface. the optical carrier is due to the doppler effect resulting from scanning one arm of the interferometer, and the frequency of this modulation is controlled by the speed of scanning. therefore, translating one arm of the interferometer has two functions; depth scanning and a doppler-shifted optical carrier are accomplished by pathlength variation. in oct, the doppler-shifted optical carrier has a frequency expressed as f d o p p = 2 ⋅ ν 0 ⋅ v s c ( 3 ) =\cdot v_\qquad \qquad \qquad \qquad \qquad \qquad \qquad \quad (3)} where ν 0 is the central optical frequency of the source, v s is the scanning velocity of the pathlength variation, and c is the speed of light. interference signals in td vs. fd-oct the axial and lateral resolutions of oct are decoupled from one another; the former being an equivalent to the coherence length of the light source and the latter being a function of the optics. the axial resolution of oct is defined as l c = 2 ln ⁡ 2 π ⋅ λ 0 2 δ λ \cdot ^ ≈ 0.44 ⋅ λ 0 2 δ λ ( 4 ) ^\qquad \qquad \qquad \qquad \qquad \qquad \qquad \qquad (4)} where λ 0 and δ λ are respectively the central