Optical coherence tomography (OCT)




Optical Coherence Tomography, or 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.

It provides images of tissue morphology at a far higher resolution (better than 10 µm) than is possible with other imaging modalities such as MRI or ultrasound.

The key benefits of OCT to the user are:

*  live sub-surface images at near-microscopic resolution;

*  enables instant, direct imaging of tissue morphology;

*  no preparation of the sample or subject is required;

*  no ionizing radiation – enabling safe use in office, laboratory or clinic.


OCT can deliver much higher resolution because it is based on light and optics, rather than sound or radio frequency radiation. It works as follows: an optical beam is projected into the subject, and light is reflected from the layers and sub-surface artefacts as the beam penetrates. Most of this light is scattered on its way back to the surface. Scattered light has lost its original direction and therefore cannot be used for imaging: this is why scattering material such as tissue appears opaque to the human eye. However, a very small proportion of the light is not scattered but reflected (as with a mirror). This light is detected and used in an OCT microscope for optical imaging.

The reflected light has the unique property that it is coherent, so that it can be detected in an OCT instrument using a device called an optical interferometer (see Interferometry). Essentially, the interferometer is used to separate the useless, incoherent, scattered light from the valuable, coherent, reflected light. It also provides the depth and intensity information from the light reflected from a sub-surface feature, enabling an image of it to be built up, rather like an echo-sounder. In this respect, OCT is more similar to ultrasound than to MRI or to PET.

The technique is limited to imaging some below the surface in tissue, because at greater depths the proportion of light that escapes without scattering is vanishingly small. No special preparation of the 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 is low – eye-safe near-infra-red light is used – and no damage to the sample is therefore likely.


Basic working principle

Optical coherence tomography (OCT) is an interferometric, (the technique of superimposing or interfering two or more waves, to detect differences between them, see also Huygens' principle) non-invasive optical tomographic imaging technique offering mm-penetration (mostly 2-3 mm, sometimes more) with μm-scale axial and lateral resolution and cross-sectional imaging capabilities.

OCT works through low-coherence Interferometry. In conventional interferometry with long coherence length (laser interferometry), interference of light occurs over a distance of meters. In OCT, this interference is shortened to a distance of μms, thanks to the use of broadband light sources (broad range of wavelengths, e.g. white light). Light with broad bandwidths can be generated by using super luminescent diodes (super bright LEDs) or lasers with extremely short pulses (femtosecond lasers, since this pulse comprises many frequencies as follows from a Fourier transform).

Light in an OCT system is broken into two pathways (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 travelled 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 (e.g. determination of the eye length for calculation of intraocular lens power). A cross-sectional tomograph (B-scan)) may be achieved by laterally combining a series of these axial depth scans (A-scan), e.g.  to produce a two-dimensional, cross-sectional view of the vitreous and retina. En face imaging (C-scan) at an acquired depth is possible depending on the imaging engine used.


The basic optical setup typically consists of an interferometer (Fig. 1, typically Michelson type) with a low coherence, broad bandwidth (white) light source. The horizontal black arrow in Fig. 1 off the sample represents the scanning movements in the sample plane (x,y). The scattered beam of the sample comes from a range of depths (z axis). The black arrow off the reference indicates the scanning in the z direction. (The scattered beam from the reference has no depth info; it comes from the surface). The scattered beam of each sample depth interferes with the scattered beam of the reference when the both path lengths are identical. The camera functions as a 2D detector array, and by performing a sequence of depth (z) scans a non-invasive 3D imaging device is achieved.

The classical type of OCT is Time Domain OCT (see More Info).




Fig. 1   Full-field OCT optical design.  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).





OCT is widely applied, especially in ophthalmology and other tissue imaging requiring μm resolution and mm penetration depth.

OCT has important advantages over e.g. medical ultrasonography (see Echography), MRI, (see MRI: general)) and confocal microscopy (see Optical microscopy: confocal laser scanning). They are not suited to morphological tissue imaging at μm-scale. The former two having poor resolution and the latter are lacking mm-penetration depth.




Fig. 2  Typical optical design 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.



More info


Time domain OCT

The path length of the reference arm is translated longitudinally in time. 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 (see Interferometry) of the light source. This interference is called autocorrelation in a symmetric interferometer (both arms have the same reflectivity), or crosscorrelation 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, Is, as

  I = k1Is  + k2 Is + 2(k1Isk2 Is)0.5∙Re [γ(τ)],        (1)

where k1 + k2 < 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 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:

,                             (2)

where Δν represents the spectral width of the source in the optical frequency domain, and ν0 is the centre optical frequency of the source. In equation (2), the Gaussian envelope is amplitude (1st term) modulated by an optical carrier (2nd term). The peak of this envelope represents the location of sample under test microstructure, with an amplitude dependent on the reflectivity of the surface. The optical carrier is, due to the Doppler effect (see Doppler principle) 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 path length variation. In OCT, the Doppler-shifted optical carrier has a frequency expressed as:

  fDoppler = 2f0vs/c,                                                   (3)

where f0 is the central optical frequency of the source, vs is the scanning velocity of the path length variation, and c is the speed of light. 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 coherence length of a source and hence the axial resolution of OCT is defined as

  lc = (2ln2/π)(λ02/Δλ)                                            (4)


Frequency Domain OCT (FD-OCT)

In frequency domain OCT the broadband interference is acquired with spectrally separated detectors (either by encoding the optical frequency in time with a spectrally scanning source or with a dispersive detector, like a grating and a linear detector array). Since the Fourier transform of the autocovariance function is the power spectral density the depth scan can be immediately calculated by a Fourier-transform from the acquired spectra, without movement of the reference arm. This feature improves imaging speed dramatically, while the reduced losses during a single scan improve the signal to noise proportional to the number of detection elements. The parallel detection at multiple wavelength ranges limits the scanning range, while the full spectral bandwidth sets the axial resolution.


Spatially Encoded Frequency Domain OCT (also Fourier Domain OCT)

SEFD-OCT extracts spectral information by distributing different optical frequencies onto a detector stripe (line-array CCD or CMOS) via a dispersive element (see Fig. 4). Thereby the information of the full depth scan can be acquired within a single exposure. However, the large signal to noise advantage of FD-OCT is reduced due the lower dynamic range of stripe detectors in respect to single photosensitive diodes, resulting in an SNR (signal to noise ratio) advantage of ~10 dB at much higher speeds. The drawbacks of this technology are found in a strong fall-off of the

Fig. 3 Spectral discrimination by swept-source OCT. Components include: swept source or tunable laser (SS), beam splitter (BS), reference mirror (REF), sample (SMP), photodetector (PD), digital signal processing (DSP).


SNR, which is proportional to the distance from the zero delay and a sin(z)/z type reduction of the depth dependent sensitivity because of limited detection line width. (One pixel detects a quasi-rectangular portion of an optical frequency range instead of a single frequency, the Fourier-transform leads to the sin(z)/z behavior). Additionally the dispersive elements in the spectroscopic detector usually do not distribute the light equally spaced in frequency on the detector, but mostly have an inverse dependence. Therefore the signal has to be re-sampled before processing, which can not take care of the difference in local (pixel wise) bandwidth, which results in further reduction of the signal quality.


Fig. 4 Spectral discrimination by SEFD-OCT. Components include: low coherence source (LCS), beam splitter (BS) (Light: beam splitter), reference mirror (REF), sample (SMP), diffraction grating (DG) and full-field detector (CAM) act as a spectrometer, and digital signal processing (DSP).


Time Encoded Frequency Domain OCT (also called swept source OCT)

TEFD-OCT tries to combine some of the advantages of standard TD and SEFD-OCT. Here the spectral components are not encoded by spatial separation, but they are encoded in time. The spectrum either filtered or generated in single successive frequency steps and reconstructed before Fourier-transformation. By accommodation of a frequency scanning light source (i.e. frequency scanning laser) the optical setup (see Fig. 5) becomes simpler than SEFD, but the problem of scanning is essentially translated from the TD-OCT reference-arm into the TEFD-OCT light source. Here the advantage lies in the proven high SNR detection technology, while swept laser sources achieve very small instantaneous bandwidths (is line width) at very high frequencies (20-200 kHz). Drawbacks are the nonlinearities in the wavelength, especially at high scanning frequencies, the broadening of the line width at high frequencies and a high sensitivity to movements of the scanning geometry or the sample (below the range of nanometers within successive frequency steps).


Full-field OCT

Focusing the light beam to a point on the surface of the sample under test, and recombining the reflected light with the reference will yield an interferogram with sample information corresponding to a single A-scan (Z axis only). Scanning of the sample can be accomplished by either scanning the light on the sample, or by moving the sample under test. A linear scan will yield a 2D data set corresponding to a cross-sectional image (X-Z axes scan), whereas an area scan achieves a 3D data set corresponding to a volumetric image (X-Y-Z axes scan), also called.


Single point (confocal) OCT

Systems based on single point, or flying-spot time domain OCT, must scan the sample in two lateral dimensions and reconstruct a three-dimensional image using depth information obtained by coherence-gating through an axially scanning reference arm (Fig. 2). Two-dimensional lateral scanning has been electromechanically implemented by moving the sample using a translation stage, and using a microelectro-mechanical system scanner.


Parallel OCT

Parallel OCT using a CCD camera has been used in which the sample is full-field illuminated and en face imaged with the CCD, hence eliminating the electromechanical lateral scan. By stepping the reference mirror and recording successive en face images, a 3D representation can be reconstructed.