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Optical Coherence Tomography and Its Application to Imaging of Skin and Retina
Optical Coherence Tomography and Its Application to Imaging of Skin and Retina Michael Pircher, Medical University of Vienna, Vienna, Austria r 2018 Elsevier Ltd. All rights reserved.
Abbreviations AMD Age related macular degeneration BCC Basal cell carcinoma CNV Choroidal neovascularization DOF Depth of focus
Nomenclature A-scan Depth profile recorded with OCT B-scan Cross-sectional image consisting of several laterally displaced A-scans
FWHM Full width at half maximum OCT Optical coherence tomography SLO Scanning laser ophthalmoscope UHR Ultra high resolution
C-scan scan
En-face image extracted from an OCT volume
Introduction Optical coherence tomography (OCT) provides cross-sectional images of tissue with high axial resolution. Thereby the sample is probed using light in the infrared range. Image acquisition is performed noninvasively and nondestructively. The axial resolution of OCT lies in the micrometer range and is decoupled from the transverse resolution. This overcomes limitations associated with the numerical aperture of the objective and allows imaging with high axial resolution despite a moderate transverse resolution. Thus layered structures, such as human skin or the human retina, are well suited to be investigated with OCT. The penetration depth in tissue depends on the corresponding scattering properties and typically lies in the order of 1–2 mm. Structures that lie deeper within tissue can be visualized either because of overlying transparent tissue (as is the case in the human eye) or because of endoscopic probes that can be moved inside the body. OCT can be regarded as bridge between microscopic imaging technologies such as confocal microscopy (with limited depth penetration) to macroscopic imaging technologies such as ultrasound or computer tomography. OCT is a highly sensitive technique and can achieve sensitivities ranging up to 110 dB or more. This allows detection of light even from weak backscattering samples. Image acquisition in OCT is done through recording of depth profiles (A-scans). Several laterally displaced A-scans form a cross-sectional image or B-scan. In a volume scan of the sample several laterally displaced Bscans are recorded. Another advantage of OCT is the high imaging speed. A-scan rates of several megahertz have been reported. Thus highly sampled volumes or several volumes per second can be recorded. The highest impact of OCT certainly lies in the field of ophthalmology. No other imaging modality is capable to provide threedimensional (3D) information of retinal tissue. Diagnosis and treatment control of retinal diseases are strongly depending on OCT images and the impact to the field can be compared with the invention of the direct ophthalmoscope more than 200 years ago. Many instruments are commercially available and OCT imaging has found a wide range of applications. Apart from ophthalmology OCT imaging can be found in dermatology, cardiovascular research, pulmonology, and many more. In addition to the biomedical field OCT has been applied to material sciences and art preservation (Stifter, 2007).
Basic Principle of OCT The technique evolved from low coherence interferometry and was initially intended to measure axial distances in the eye. The basic principle is similar to ultrasound technology. However, instead of sound waves light is used to generate an image. The speed of light is much higher than that of sound waves. Thus, an interferometric technique is needed in order to distinguish between light that has traveled different distances in the micrometer range. Fig. 1(A) shows a scheme of an OCT interferometer. Light is emitted from an OCT light source and is split by the beam splitter into two arms, the reference arm and the sample arm, respectively. In the reference arm, the light is back reflected by the reference arm mirror. In the sample arm the beam is deflected on a scanning unit (for x- and y-scanning) before the light is sent to the sample. The light is backscattered from the sample, de-scanned and brought to interference with light from the reference arm at the beam splitter. The interfering light is then detected by the detector. It is very common to use fiber optics for the interferometer part. This eliminates the need for interferometer alignment and enables the fabrication of very compact systems. There are several techniques available for OCT.
Encyclopedia of Modern Optics II, Volume 5
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Fig. 1 (A) Scheme of an optical coherence tomography (OCT) interferometer. BS denotes the beam splitter of the interferometer. For time domain optical coherence tomography (TD-OCT) the reference mirror is mounted on a translation stage. (B) Measured intensity I at the detector (interference pattern) when the mirror is moved in axial direction (reflectivity in sample and reference arm are equal) full width at half maximum (FWHM) denotes the full width at half maximum of the coherence function which is defined as axial resolution in OCT. (C) Measured interference pattern when a glass plate with thickness d is measured (n denotes the refractive index of the glass plate).
Timed Domain OCT Initially the so-called time domain OCT (TD-OCT) technique has been used. In this technique the interferometer is illuminated with a broadband light source. To generate an A-scan (depth profile) the reference mirror is moved axially. In the case of a mirror in the sample arm, the detector will record an intensity pattern that is shown in Fig. 1(B). Due to the broad bandwidth of the source, interference fringes will only be visible if sample and reference arm lengths are matched (within the coherence gate). Thus light originating from different depths can be separated. In the case of a glass plate the light will be backscattered (or back reflected) at the front and back surface of the plate. During axial scanning of the reference arm mirror two coherence functions (arising at the surface and back surface of the glass plate) as illustrated in Fig. 1(C) will be detected. The distance between the two signals corresponds to the optical thickness of the plate (i.e., the geometrical thickness multiplied with the refractive index of the plate). The full width at half maximum (FWHM) of the coherence function is defined as the axial resolution of OCT and equals the round trip coherence length ( ¼ 1/2 of the coherence length) of the light source. The A-scan rate of this technique is limited by the inertia of the mirror to a few 100 Hz. However, the method can be highly sensitive. Motion of the mirror will introduce a Doppler frequency shift for each wavelength. Thus, a carrier frequency for the OCT signal is generated which allows for band pass filtering of the signal. This procedure filters 1/f noise out and increases the sensitivity. Depending on the bandwidth of the light source and the speed of motion a certain electronic signal bandwidth needs to be detected. Faster speeds will require larger electronic bandwidths (of the detector and amplifying electronics) and thus lower system sensitivity. To overcome the limitation given by the inertia of the mirror, a galvanometer scanning mirror and a grating (rapid scanning optical delay line) can be placed in the reference arm. Instead of a translational motion in spatial domain, this configuration allows for a rotational motion in wavelength domain. The method has the advantage that the group delay (axial scanning of the sample) and the phase delay (modulation frequency of the OCT signal) can be adjusted independently. Therefore the imaging speed can be improved up to a few kilohertz. TD-OCT detects only light backscattered from within the coherence volume of the sample. The coherence volume is defined through the round trip coherence length and the transverse resolution of the system. Light from other depths (outside the coherence volume) is rejected which limits the sensitivity of the method. However, the depth range of the system depends on the travel range of the reference mirror. Thus, in principle very long axial distances without loss in sensitivity can be scanned with TD-OCT.
Fourier Domain OCT With the development of Fourier domain (FD) OCT the sensitivity could be greatly improved. The gain in sensitivity can be directly translated into imaging speed that allows A-scan rates an order of magnitude higher than in TD-OCT. Two different FDOCT schemes are possible. The spectral domain (SD) OCT technique uses a similar broadband light source as in TD-OCT. However, instead of a single detector at the interferometer exit the light is spectrally dispersed trough a grating and detected with a line scan camera (cf. Fig. 2(A)). Thereby, the reference arm mirror is kept stationary. In this way light from the entire sample depth is recorded simultaneously which results in a sensitivity advantage compared to TD-OCT. In case of a mirror in the sample arm (same light power returning from sample and reference arm) the spectrometer camera will detect an interference pattern that is displayed in Fig. 2(B). The modulation of the spectrum arises from constructive and destructive interference for each wavelength and depends on the path length difference between sample and reference arm beam. Thus the frequency of the modulation depends on the amount of path length difference. Smaller differences result in low frequencies, while larger differences cause higher frequencies. In addition, the spectrum (and the interference pattern) is detected in the wavelength domain. Thus, each modulation frequency that corresponds to a certain depth will have a chirp along the spectrum. To remove this chirp, the spectrum needs to be transformed into k-space ( ¼1/l-space). Associated with this chirp will
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Fig. 2 (A) Scheme of a spectral domain optical coherence tomography (SD-OCT) instrument using a spectrometer in the detection arm (BS, beam splitter). (B) Measured interference pattern when a mirror is placed in the sample arm (light power returning from sample and reference arm are equal). (C) Fast Fourier transformation of the spectrum measured in (B) corresponding to a depth profile.
be a drop of sensitivity (sensitivity fall off) with depth because of the finite width of the detector elements (camera pixels). Higher frequency components (constructive and destructive interference of neighboring wavelengths) will therefore be detected by a single detector element and thus not be visible. It should be noted that FD-OCT requires a minimum detection speed. Jitter within the interferometer or axial movement of the sample (in case of in vivo imaging) will cause a change (phase shift) in the interference pattern which degrades the visibility of the interference pattern when this change occurs during the exposure time of the camera. In order to avoid this “fringe washout” the exposure time should be shorter than 1 ms. In the case displayed in Fig. 2(B) one arm is only slightly longer than the other. At this point it should be noted that it is not possible to distinguish (without any other means) between positive (reference arm is longer) or negative (sample arm is longer) path length differences because both will result in the same frequency modulation of the spectrum. In order to generate an A-scan the interference pattern (modulated spectrum) needs to be Fourier transformed. The corresponding depth profile is plotted in Fig. 2(C). Because of the inability to distinguish between positive and negative frequencies, the transformation of the spectrum yields in fact two depth profiles that are mirrored around zero delay. Zero delay refers to the position where sample and reference arm lengths are exactly matched. The total number of sample points after Fourier transformation stays constant. However, this means that the number of sample points per A-scan is only half the number of camera pixels. Because the mirrored A-scan does not contain additional information it is very common in OCT to simply omit this part as is done in Fig. 2(C). The depth range of a SD-OCT system depends on the spectral resolution of the spectrometer as increasing depth is associated with higher modulation frequencies that need to be sampled according to the Nyquist theorem. This resolution is determined by the dispersion properties of the grating (number of lines that are illuminated on the grating), the spectrometer optics and the number of pixels that are available on the line scan camera. According to Leitgeb et al. (2003) the maximum depth range can be calculated via zmax ¼
l2 4δl
ð1Þ
where l and δl denote the wavelength and the spectral resolution, respectively. A typical OCT configuration uses a light source with a central wavelength of 840 nm, a bandwidth of 50 nm, and a 2048 pixel camera. Assuming that 150 nm bandwidth is imaged onto the camera and that the separation between spectral lines (corresponding to the resolution provided by the grating) is smaller than the size of a pixel an imaging depth of 2.4 mm is achieved. Back reflections from larger imaging depths will appear attenuated as aliasing frequencies and may cause unwanted artifacts inside an OCT image. Another technique for realizing FD-OCT is swept source (SS) OCT. SS-OCT uses fast (wavelength) tunable lasers as light source in combination with a single (fast) photo detector. During a laser sweep the interference pattern is recorded depending on wavelength and separated in time. The pattern will be similar to the one displayed in Fig. 2(B). It should be noted that the instantaneous power at each wavelength can be higher compared to SD-OCT which compensates for the nonparallel data acquisition scheme of SS-OCT compared to SD-OCT. Similar as in SD-OCT, light backscattered from the entire sample depth contributes to the interference pattern and the depth profile is reconstructed in a similar way via Fourier transformation of the recorded spectral information. The instantaneous linewidth of the source determines the instantaneous coherence length and defines the depth range of the system. The instantaneous linewidth thus plays an equivalent role as the spectral resolution of SDOCT. Within this range light from the reference arm and sample arm are coherent and can interfere. However, the instantaneous line width is a completely different measure as the round trip coherence length. The latter is defined by the FWHM of the laser sweep, while the former is the width of a laser line emitted at a certain wavelength within the sweep. Fig. 3(A) illustrates the emission of a SS laser. The output power of many swept laser sources fluctuates. In order to compensate for this fluctuations balanced photo detection is employed. Fig. 3(B) shows a schematic of a fiber-based SS-OCT instrument. The light emitted from the SS is already confined in a single mode fiber. Before the entrance of the interferometer the light traverses a fiberized circulator and is split thereafter by a 50:50 beam splitter into reference and sample beam, respectively. The light back reflected from both arms is recombined by the 50:50
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Fig. 3 (A) Illustration of a swept source (SS) laser sweep. The sweep range determines the axial resolution of the system. The instantaneous line width (drawn not to scale) determines the imaging depth provided by the light source. (B) Representative scheme of a SS optical coherence tomography (OCT) instrument based on fiber optics and balanced photo detection. BS 50:50, fiber beam splitter; GS, galvanometer scanner.
beam splitter and is detected with a balanced photo detector. The circulator ideally transmits nearly 100% of the light from the light source to the interferometer, while nearly 100% of the light coming from the interferometer is transmitted to the detector. There are several advantages of SS-OCT compared to SD-OCT. In SD-OCT the light is detected with a spectrometer. The efficiency of the spectrometer cannot be 100% which results in a lower sensitivity compared to SD-OCT. In addition the wavelength sweep can be performed in 1/l space which results in an elimination of the frequency chirp that is present in spectrometer-based systems and correspondingly into an elimination of the sensitivity drop with depth. Finally, the spectrometer needs to be perfectly aligned in order to achieve a high sensitivity which sets high demands, from a technical point of view, on the spectrometer design in commercial instruments. No such alignment is required for SS-OCT instruments.
Resolution The FWHM of the interference pattern (cf. Fig. 1(B)) defines the axial resolution of OCT and is inversely proportional to the bandwidth of the light source. In case of a Gaussian-shaped spectrum the resolution can be calculated via following equation (Drexler and Fujimoto, 2015): Dz ¼
2lnð2Þl npDl
2
ð2Þ
where l denotes the central wavelength, Dl the bandwidth, and n the refractive index of the medium. As can be seen from this equation, the resolution in OCT depends on the central wavelength of the light source and the corresponding bandwidth. Fig. 4 shows the axial resolutions for three wavelength regions that are commonly used in OCT. Standard OCT resolution refers to bandwidths around 50 nm corresponding to an axial resolution between 5 and 15 mm. The use of larger bandwidths (100 nm or more) is known as ultrahigh resolution OCT and requires ultra-broadband light sources, such as titanium sapphire laser or supercontinuum light sources. Important for achieving the resolution in OCT that is provided by the light source is a sufficient optical transparency of the optical components and matching of dispersion (see below) introduced in reference and sample arm, respectively. Especially fiberized interferometers may cause unwanted effects because single mode propagation (multimode propagation results in coherent image artifacts) depends on the size of the core diameter in respect to the wavelength. Thus larger wavelengths will be stronger attenuated if the core size is chosen to ensure single mode propagation for the shortest wavelengths. The transverse resolution mainly depends on the imaging optics in the sample arm. In order to achieve maximum fringe visibility it is common in OCT to use light with high spatial coherence. Light emitted from a single mode fiber has this property. Because of its directionality, divergence, and Gaussian intensity profile, the beam propagation can be described using Gaussian optics. In many OCT systems the objective lens is not the limiting aperture as the beam is scanned over this lens in order to achieve transverse scanning of the sample. Hence, the beam remains Gaussian and the resolution can be calculated via the FWHM of the Gaussian beam waist o0 and the angular spread yS of the Gaussian beam (Fercher, 2010): pffiffiffiffiffiffiffi pffiffiffiffiffiffiffi l Dx ¼ 2 ln2o0 ¼ 2 ln2 pyS
ð3Þ
l denotes the central wavelength of the beam. Strictly speaking, this resolution can be obtained only at the focal (or certain depth) plane. Outside this focal plane the resolution is degraded. A measure for the imaging depth where the resolution is degraded only by a factor of √2 is the Rayleigh range zR Twice this range defines the depth of focus (DOF), which can be calculated with 2zR ¼ DOF ¼ 2
l py2S
ð4Þ
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Fig. 4 Axial resolution of an optical coherence tomography (OCT) system depending on the bandwidth and the central wavelength for three commonly used wavelength regions. Standard OCT uses a bandwidth between 30 and 70 nm, which results in a resolution below 20 mm. Ultrahigh resolution (UHR) OCT refers to the use of larger bandwidths (4100 nm) and corresponding resolutions that are below 5 mm.
Thus for high-resolution systems only part of the imaged volume will be in focus and imaged sharply. For an OCT system at 1300 nm and 20 mm transverse resolution the DOF will be 2 mm. This DOF will be reduced to 20 mm in the case of 2 mm transverse resolution. Hence, the imaging optics needs to be carefully designed in order to optimize the OCT performance. Several methods have been proposed to maintain high transverse resolution throughout the imaging depth. These are based on a dynamic shift of focus with the coherence gate (only possible for TD-OCT) or on the use of special illumination techniques (such as Bessel beams).
Sensitivity Several noise sources influence the sensitivity in OCT. Receiver noise from the photo detectors, excess noise arising from fluctuations in light power, and shot noise that originates from the quantum nature (and random arrival) of photons. Ideally, an OCT system is limited by shot noise only. This limit can be reached through the coherent amplification process inherent to OCT as the light power returning from the reference arm greatly exceeds the power returning from the sample. In the case of sufficient light power from the light source, the amount of light returning from the reference arm can be optimized in order to reach the shot noise limit. In TD-OCT the shot noise limited sensitivity depends on the detection bandwidth and the power efficiency of the interferometer (Izatt et al., 2015): Sensitivity ¼
rPS 2eDBTD
ð5Þ
where r denotes the quantum efficiency of the detector, PS is the light power returning from the sample arm (with a 100% reflective mirror as sample), e is the electron charge (e ¼ 1.6 1019 1C), and DBTD is the electronic detection bandwidth of the TDOCT systems. The light power returning from the sample depends on the power incident on the sample, the interferometer configuration, and the optical losses within the system. Let us consider a fiber-based Michelson interferometer with a 50/50 splitting ratio and 4 mW of optical power (central wavelength at 1300 nm with a bandwidth of 50 nm) on the sample. The light is emitted from a single mode fiber. Thus the coupling efficiency (of the light returning from the sample) will be B50%. Another 50% of the light is lost due to the fiber beam splitter which results in an optical power returning from the sample arm (with a 100% reflective mirror as sample) of 1 mW. When the above mentioned TD-OCT instrument is operated at 1 kHz A-scan rate and 3 mm depth range a minimum electronic bandwidth of 800 kHz will be required. This results in a sensitivity of the OCT system (assuming 0.8 quantum efficiency of the detector) of 103 dB.
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In FD-OCT the sensitivity can be expressed in a similar way (Leitgeb et al., 2003): Sensitivity ¼
rPS M 4eDBFD
ð6Þ
with DBFD the electronic bandwidth of FD-OCT systems (inversely proportional to the exposure time of the camera in SD-OCT) and M the number of spectral channels of the FD-OCT systems. By comparing this equation with Eq. (5) we can see that the sensitivity of FD-OCT is a factor of M/2 larger than in TD-OCT. Thus, for a FD-OCT instrument operating with the same light source and interferometer as before will have an improved sensitivity. Nevertheless, following issues need to be considered. The light returning from the sample will be slightly attenuated by the efficiency of the spectrometer (this typically lies in the range of 60%) and the exposure time of the camera will be shorter than the time corresponding to the A-scan rate because the camera requires a certain readout time. For an A-scan rate of 70 kHz we assume that 1000 pixels on the camera are (equally) illuminated and that the camera requires 4 ms for readout. This results in an exposure time of B10 ms. The detection bandwidth B of FD-OCT is given by (Leitgeb et al., 2003) BFDOCT ¼
2zR lc t
ð7Þ
with zR denoting the depth range of the system and lc the double pass coherence length. t is the exposure time of the camera. In our case (zR ¼6 mm, lc ¼ 15 mm, and t¼10 ms) the detection bandwidth is 80 Mhz which results in a shot noise limited sensitivity of 108 dB. Thus, a higher sensitivity can be achieved with this SD-OCT system although the imaging speed is increased by a factor of 70.
Dispersion Dispersion describes the effect of a wavelength dependent refractive index of a medium. For example, the refractive index of water (cf. Fig. 5) is not constant and decreases with wavelength. For OCT dispersion mismatch between the interferometer arms results in a degradation of the axial point spread function. This is associated with a degradation of axial resolution and sensitivity. Thus the interferometer arms need to be balanced which is normally achieved by inserting corresponding materials that are present in the sample arm (such as lenses for scanning optics) into the reference arm of the interferometer. For compensating dispersion of the eye cuvettes that are filled with water are commonly used. FD-OCT records spectral data. Therefore it is possible to compensate for dispersion in post-processing.
Wavelength Range The choice of imaging wavelength in OCT depends on the investigated sample as well as on the availability of light sources. For a high penetration depth the total attenuation coefficient of tissue at the imaging wavelength should be minimal. Both, absorption and scattering contribute to the total attenuation coefficient. Water is the main constituent of tissue (apart from teeth
Fig. 5 The wavelength dependence of the refractive index of water in the visible and near-infrared (NIR) region. Data from Hale, G.M., Querry, M. R., 1973. Optical-constants of water in 200-nm to 200-mm wavelength region. Applied Optics 12(3), 555–563.
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Fig. 6 Absorption coefficient of water and scattering coefficient of dense tissue. The three main optical coherence tomography (OCT) imaging wavelengths (840, 1060, 1300 nm) are outlined in the graph. Data from Hale, G.M., Querry, M.R., 1973. Optical-constants of water in 200-nm to 200-mm wavelength region. Applied Optics 12(3), 555–563.
or bones) and its absorption properties will be dominating. As can be seen in Fig. 6 the absorption coefficient of water strongly increases from visible light to infrared light. On the other hand, scattering gradually decreases in this wavelength range. At the 1300 nm wavelength region the contribution of scattering and absorption are largely within the same range. Thus, this wavelength region is the preferred choice for imaging dense tissue such as human skin as for shorter wavelengths the penetration will be limited by scattering, while for longer wavelengths absorption will be dominating. Imaging at 1300 nm requires detector materials that are sensitive in this region. These are much more expensive than comparable materials that are sensitive in the visible range. As such the preferred OCT imaging technology will be based on SS-OCT because this technology requires a singlebalanced detector (instead of a detector array as in SD-OCT) and provides some benefits (as outlined above) compared to the SD-OCT counterpart. For imaging the retina, the light has to traverse the ocular media (eye length of B2 cm) twice which significantly changes the situation. The scattering properties of the retina are of less importance compared to the transmittance of the light through the eye. Ocular media are very weakly scattering (apart from pathologies such as cataract). Thus the imaging wavelength should be as close to the visible range as possible to minimize absorption by water. However, for highly sensitive measurements the quantum efficiency of the detector should be high. The used detector material, silicon, has a peak around 900 nm and levels off toward shorter wavelength. Thus imaging wavelength close to the maximum of the quantum efficiency of the detector is desirable. An additional aspect is that in the visible range the photoreceptors of the retina will be very sensitive to light (tails of the absorption spectra reach into the infrared region). Hence, the subjective perception of the light power will increase rapidly at wavelength regions that are close or within the visible range. This is associated with increasing subject discomfort during imaging, even though the light levels that are sent to the eye have to be below the limits for safe exposure. In addition dispersion is more pronounced for shorter wavelength regions. Because of these issues the wavelength region around 840 nm has been found as compromise and is widely used in OCT for retinal imaging. Since line scan cameras in this wavelength region are easily available and relatively cheap, the preferred imaging technique in this wavelength region is SD-OCT. Nevertheless, another wavelength region at 1060 nm which is around a local minimum of the water absorption spectrum (cf. Fig. 6) has gained interest in recent years. Due to the longer wavelength region the penetration depth within the retina can be increased and the influence of media opacities (such as cataract) on the image quality will be lower. Another advantage of this wavelength region is the availability of fast SSs. Thus high sensitivity can be maintained throughout the imaging depth which further improves the retinal image quality. Due to technical issues (dispersion introduced by fiber optics), SSs at 840 nm are relatively slow compared to their counterparts at 1060 or 1300 nm.
OCT Angiography In recent years, OCT angiography (OCTA), a method to visualize vessel structure within tissue, has gained increasing interest. There are many different approaches to realize OCTA. The underlying basic principle, however, is the same for all methods. The idea is that two (or more) measurements (A-scans or B-scans) are taken at the same location but separated in time. By comparing these measurements changes (intensity or phase or both) caused by blood flow can be detected. The time separation between these
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measurements needs to be short enough in order to minimize the influence of sample motion and long enough in order to ensure sufficient changes introduced by blood flow. With A-scan rates of 100 kHz or more the most common OCTA technique is based on recording several B-scans at the same location. This, however, increases the overall measurement time. The major benefit of OCTA is that it provides information on the microvascular perfusion of tissue noninvasively. In addition, due to the spatial separation of capillaries, these can be visualized even though the lateral extension lies below the transverse resolution of the system. Other techniques that are capable to provide information on the vascular structure require the application of contrast agents. Some subjects may show allergic reactions or other complications after administration of such contrast agents. Thus state of the art angiographic imaging is always associated with residual risk factors. Another benefit of OCTA lies in the fact that depth resolved information on the vasculature is provided. This allows for a separation between individual vascular beds and thus an improved image contrast. It should be noted, however, that OCTA is not able to detect vessel leakages because the technique relies on motion contrast. Within leakages motion occurring between individual B-scan recordings will be too slow compared to subject motion to be detected in vivo. However, in vitro experiments have shown that without bulk sample motion slow fluid motion (such as Brownian motion) can be visualize with OCTA techniques.
Imaging of Human Skin The human skin is easily accessible with optical methods. In clinical routine, lesions will be first inspected visually with high magnification using a microscope (dermatoscope). In the case of suspicious lesions a biopsy will be taken which will be sliced and prepared for histologic examination. The cellular details provided by this histologic examination can then be used to determine the malign or benign status of a lesion and finally determines the further clinical procedure. OCT may represent an alternative to the histologic examination as it provides in vivo and noninvasively cross-sectional images of skin. Current studies focus on improving the sensitivity and specificity of OCT methods for detecting malign tissue (Ulrich et al., 2016). Thereby, additional contrast mechanism such as polarization sensitive OCT or OCTA are used to gather additional information. Especially OCTA provides essential additional information as tumors, such as basal cell carcinoma (BCC) are known to present with increased vascularization. Using fiber-based or handheld systems practically every location can also be reached with OCT. In order to minimize scattering and absorption, the preferred imaging wavelength for human skin will be in the 1300 nm range. The penetration depth of OCT in skin is in the order of 1–2 mm. Fig. 7 shows a representative B-scan image of a finger recorded with a 1310 nm SD-OCT instrument. The A-scan rate of the instrument was 91 kHz. The image consists of 500 A-scans and was recorded in B5 ms. The bandwidth of the light source was 50 nm which results in an axial resolution of 10 mm in tissue (assuming a refractive index of 1.4). The power onto the skin was 8 mW. Individual layers of the skin can be clearly seen in Fig. 7 and are labeled in the image. Supplementary material related to this article can be found online at http://dx.doi.org/10.1016/B978-0-12-803581-8.09787-3. An interesting skin region is the nail fold of a finger. In this region capillaries are close to the surface (close to the nail some capillaries extend parallel to the surface) and can be investigated with optical methods. In order to assess the status of the capillary network, capillaroscopy is used in clinical routine. This method examines the skin with a microscope after applying a contrast agent (such as oil) to the skin region. Using OCTA this network can be visualized completely noninvasively and without the
Fig. 7 Representative B-scan image recorded at 1300 nm of the inner side of a finger. Superficial layers down to deeper layers of the dermis can be visualized. The multimedia file (media 1) shows a fly through of B-scans of the entire three-dimensional (3D) data set.
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Fig. 8 Representative optical coherence tomography (OCT) images recorded in the nail fold region of a healthy volunteer. (A) OCT-intensity B-scan (averaged over 5 B-scans). (B) Depth integrated en-face intensity image. The arrow indicates the location of the B-scan displayed in (A). (C) Depth integrated en-face intensity image and overlaid (red) vessel map.
application of a contrast agent. Fig. 8 shows representative image data recorded with the SD-OCT instrument described above. The intensity B-scan (Fig. 8(A)) clearly shows different structures such as the nail plate and different skin layers of the nail fold. The banding structure within the nail plate is an artifact that originates from birefringence of the plate. The sensitivity of the system is sufficient to penetrate through the nail plate as structures underneath the plate can be seen. The axial displacement of the nail plate beneath the nail fold (cf. right hand side in Fig. 8(A)) is caused by the increased optical path length that is introduced by light propagation through the overlying nail fold tissue. The en-face intensity projection clearly shows the different regions of the nail fold and nail plate. Fig. 8(C) shows a composite image where the depth integrated vessel structure is overlaid to the intensity image of Fig. 8(B). In this image the capillary network can be clearly seen. Close to the nail plate individual capillary loops can be seen which are typical for this skin region.
Imaging of Human Retina The human retina consists of several layers and represents an essential part for the vision process. All layers that can be seen in histology can be visualized with OCT. Fig. 9 shows a representative OCT B-scan image of the fovea region with a labeling of the different retinal layers. In healthy subjects the fovea is the location of best visual acuity. In this region the highest density of cone photoreceptors can be found. These cells are responsible for photopic vision and can be distinguished into three types (sensitive in the short, medium, and long wavelength region). In the fovea no rod photoreceptors (responsible for scotopic vision) are present. The light is incident from the top of this image and has to penetrate different retinal layers until it is detected by the photoreceptors. The pigments of the photoreceptors are located within the outer segments of cones and rods. The image was recorded with a research grade prototype SD-OCT instrument operating at 863 nm with a bandwidth of 60 nm which resulted in an axial resolution of B4 mm in tissue (assuming a refractive index of 1.4). The A-scan rate was 70 kHz and one B-scan consisting of 1024 A-scans was recorded in B15 ms. Although the axial resolution is sufficient to resolve individual retinal layers, individual cells (such as photoreceptors) cannot be visualized. In order to increase the transverse resolution, aberrations that are introduced by imperfections of the eye optics need to be compensated for. This can be done with the implementation of adaptive optics, a technology known from astronomy (Pircher and Zawadzki, 2007). Many commercial OCT devices for imaging the retina are available. Fig. 10 shows representative data recorded with a commercial SD-OCT instrument. The instrument provides a scanning laser ophthalmoscope (SLO) overview mode that is used for retinal tracking. The location of the recorded B-scan is marked with a green line in the SLO image. The image quality is very similar to the previous image. The entire 3D data set can be viewed in media 2. Supplementary material related to this article can be found online at http://dx.doi.org/10.1016/B978-0-12-803581-8.09787-3. In order to outline the difference of retinal images recorded at different wavelengths, Fig. 11 shows a retinal image recorded at 1040 nm. Compared to the previous images enhanced penetration into the choroid and even into the sclera can be seen. Otherwise all retinal layers can be seen with similar quality. The image was recorded with a research grade SS-OCT instrument operating at 100 kHz A-scan rate and 20 20 field of view. The axial resolution of the system was 5.6 mm in tissue. Through a change in the scanning pattern (recording of several B-scans at the same location) angiographic data can be retrieved. Fig. 12 shows representative data of a healthy volunteer recorded with a commercial instrument. The en-face projection images were generated through depth integration over the depth extension indicated with white arrows in Fig. 12(A)). Fig. 12(B)) shows cross-sectional angiographic data (red) overlaid to the intensity image. The contrast of retinal vessels is greatly increased using the angiographic data evaluation. The en-face projection of the intensity images (cf. Fig. 12(C)) shows larger vessels as shadows.
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Fig. 9 Optical coherence tomography (OCT) image of the retina with a labeling of the different layers (the image spans an angle of 28 degree which corresponds to B8 mm on the retina). COST, cone outer segment tips; ELM, external limiting membrane; GCL, ganglion cell layer; INL, inner nuclear layer; IPL, inner plexiform layer; IS/OS, junction between inner and outer segments of cone photoreceptors; ONL/HF, outer nuclear layer/Henle’s fiber layer; OPL, outer plexiform layer; PRL photoreceptor layer; RNFL, retinal nerve fiber layer; RPE, retinal pigment epithelium.
Fig. 10 Representative retinal image data of the fovea of a healthy volunteer recorded with a commercial instrument (imaging wavelength¼840 nm). Left: overview scanning laser ophthalmoscope (SLO) image. Right: optical coherence tomography (OCT) B-scan recorded at the location indicated in the left image with the green arrow. The multimedia file (media 2) shows a fly through of B-scans of the entire 3D data set.
However, the entire vascular structure can be seen in the corresponding OCTA projection map. This map clearly shows the avascular zone of the fovea region. From these maps deteriorations (such as non-perfused areas) from the normal vasculature structure can be easily determined. Fig. 13 shows representative image data of patients with age related macular degeneration (AMD). AMD is regarded as one of the major causes of blindness in the industrialized world. The disease mainly affects the macula region (containing the fovea) of subjects that are older than 60 years. AMD can be differentiated into several forms: dry AMD and wet AMD. The latter is the most severe form which results (untreated) into a rapid degradation of visual acuity. An early indicator of AMD is the presence of Drusen. Drusen are elevations of the retinal pigment epithelium (RPE)/photoreceptor bands because of accumulation of retinal debris underneath these layers. Fig. 13(A) shows a typical example of OCT images recorded in a patient with Drusen. The elevations of the photoreceptors cause image distortions and degradation in visual acuity of the patient. At locations of the Drusen an additional layer (Bruch’s membrane) can be seen. In the healthy case, this layer is obscured by the overlying RPE. Due to the highly scattering of the pigments within the RPE, this layer appears as broadened diffuse layer although it is a mono cellular layer. The 3D visualization of Drusen allows for a quantitative assessment of the size and volume of the Drusen. As such the effect of therapeutic interventions on the Drusen size can be precisely monitored. The most severe form of AMD is choroidal neovascularization (CNV). In this case new vasculature is generated within the retina and leakages of these lead to damage of the sensitive neural tissue of the retina. Fig. 13(B) shows a representative B-scan of the fovea region of a patient with CNV. Scarring of retinal tissue and associated disintegration of photoreceptors is very often accompanied by this disease. Using OCT the size of edema or cysts can be quantified. In addition vascular growth can be determined using OCTA. Therapeutic intervention (such as anti-vascular endothelial growth factor (anti-VEGF) therapy) can be easily monitored and the technology can be meanwhile regarded as gold standard for this kind of diagnosis and for treatment control.
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Fig. 11 Retinal image of the fovea of a healthy volunteer recorded with a swept source optical coherence tomography (SS-OCT) instrument at 1040 nm. Enhanced penetration into the choroid and sclera can be observed.
Fig. 12 Retinal images of the fovea of a healthy volunteer recorded with a commercial instrument. (A) Central B-scan of the volume scan. The arrows indicate the image depth that was used to generate the en-face images in (C) and (D). (B) Central B-scan of the volume scan and overlaid vessel structure (in red) retrieved with the angiography algorithm. (C) En-face intensity image (white ¼high intensity, black¼low intensity) generated through depth integration over the area indicated in (A). En-face vessel map generated through maximum intensity projection of the angiographic data within the area indicated in (A).
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Fig. 13 Representative optical coherence tomography (OCT) B-scans (averaged over 50 frames) at 840 nm of patients with age related macular degeneration (AMD). (A) OCT B-scan image of Drusen (some are indicated by white arrows). The black arrow indicates Bruch’s membrane. (B) OCT B-scan image of a choroidal neovascularization (CNV). Images are displayed on a logarithmic intensity scale.
There are many other retinal diseases where OCT plays a key role in diagnosis. A summary of the impact of OCT to all these diseases will be beyond the scope of this article. Nevertheless, one specific example should be mentioned: Glaucoma, a disease that is associated with visual field defects and one of the major causes for blindness worldwide. At an early stage of the disease, these defects occur in the periphery of the retina which is not easily noticeable by patients themselves because for reading etc. only the fovea region is used. Before changes in the visual field can be detected, a loss of retinal nerve fiber tissue (loss in layer thickness) will occur. The 3D imaging capabilities of OCT enable a precise measurement of the retinal nerve fiber layer thickness. Thus OCT provides an excellent indicator for early signs of Glaucoma and for monitoring therapeutic success.
Conclusion and Outlook OCT is a very powerful imaging technology. With the development of FD-OCT imaging speed could be greatly improved and Ascan rates in the megahertz range have become possible. This paved the way for many new applications of OCT. This article focused on skin and retinal imaging but many more applications of OCT are available. Associated with the improved imaging speed is the possibility to investigate dynamic processes or to increase the contrast through the introduction of angiographic evaluation of OCT data. The noninvasive visualization of capillaries in tissue, such as skin or retina, opens completely new perspectives for clinical research. The high clinical demand of this technology manifests itself through the rapid commercialization of OCTA for both, skin, and retinal imaging. Thereby, OCT certainly plays a key role for accurate diagnosis especially in ophthalmic routine. Nevertheless, there are new developments in the field. Currently new SSs are developed that provide higher speeds (A-scan rates) and very long coherence lengths. Meanwhile, imaging depth ranges of some meters have been demonstrated with OCT. This opens new possibilities such as imaging of the whole eye (from the cornea to the retina) with a single SS-OCT system. In addition further parallelization of OCT is investigated. Line field OCT, for example, records one B-scan in parallel. Thereby SDOCT (this requires an area detector in the spectrometer) or SS-OCT can be used. The latter uses a SS in combination with a line scan camera. Very high speeds (several volumes per second) can be achieved using a full-field illumination in combination with a SS and a very fast area camera. The demands on the camera are very high because the phase stability (to avoid “fringe washout”) requires the recording of a volume within a short time frame (ideally below 1 ms). In order to have sufficient sampling points in depth (determined by the number of spectral lines) the frame rate of the camera needs to be in the order of 50 kHz. Apart from the high imaging speeds that are possible and that have been demonstrated for in vivo retinal imaging, the full-field technique provides phase stability over the entire field of view (in lateral direction). OCT provides direct access to the phase of the electric light field and computational techniques can be applied in order to compensate for aberrations introduced, for example, by the eye. Thus high resolution can be achieved even without the use of hardware-based adaptive optics. In addition, the focal plane of the image can be adjusted in post-processing. As such, an entirely sharp volume can be obtained with high transverse resolution despite of a limited DOF. Apart from these developments, OCT has become a valuable tool in cardiovascular research. Using fiberized probes, OCT imaging can be done from inside the body, for example, within a coronary artery. As such plaques, deposits, or stents within the artery can be visualized in vivo. Key element for this development is the high imaging speed provided by OCT. Currently clinical trials are performed to demonstrate the benefit of cardiovascular OCT compared to state of the art examination tools. Another development in the field is the combination of OCT with other imaging technologies. For example, the combination of OCT with photo acoustics has gained increasing interest. The photo acoustic technique relies on absorption of the probing light beam and provides a higher penetration depth but lower resolution than OCT. Therefore vasculature that is located deeper within
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tissue can be visualized. The combination with OCTA enables the visualization of the entire vascular network as superficial capillaries are detected with OCT, while the deeper vascular structure is detected with photo acoustics.
Acknowledgments The author likes to thank following people from the Medical University of Vienna for their assistance: Christoph K. Hitzenberger, Matthias Salas, Martin Fürst, Mitsuro Sugita, Teresa Torzicky, and Ursula Schmidt-Erfurth.
See also: Interferometric Imaging. Overview: Coherence
References Drexler, W., Fujimoto, J., 2015. Optical Coherence Tomography, Heidelber. New York, NY: Springer. Fercher, A.F., 2010. Optical coherence tomography – Development, principles, applications. Zeitschrift Fur Medizinische Physik 20, 251–276. Izatt, J.A., Choma, M.A., Dhalla, A., 2015. Theory of optical coherence tomography. In: Drexler, W., Fujimoto, J.G. (Eds.), Optical Coherence Tomography. Heidelberg; New York, NY; Dordrecht; London: Springer. Leitgeb, R., Hitzenberger, C.K., Fercher, A.F., 2003. Performance of Fourier domain vs. time domain optical coherence tomography. Optics Express 11, 889–894. Pircher, M., Zawadzki, R., 2007. Combining adaptive optics with optical coherence tomography: Unveiling the cellular structure of the human retina in vivo. Expert Review of Ophthalmology 2, 16. Stifter, D., 2007. Beyond biomedicine: A review of alternative applications and developments for optical coherence tomography. Applied Physics B-Lasers and Optics 88, 337–357. Ulrich, M., Themstrup, L., De Carvalho, N., et al., 2016. Dynamic optical coherence tomography in dermatology. Dermatology 232, 298–311.
Abbreviations AMD Age related macular degeneration BCC Basal cell carcinoma CNV Choroidal neovascularization DOF Depth of focus
Nomenclature A-scan Depth profile recorded with OCT B-scan Cross-sectional image consisting of several laterally displaced A-scans
FWHM Full width at half maximum OCT Optical coherence tomography SLO Scanning laser ophthalmoscope UHR Ultra high resolution
C-scan scan
En-face image extracted from an OCT volume
Introduction Optical coherence tomography (OCT) provides cross-sectional images of tissue with high axial resolution. Thereby the sample is probed using light in the infrared range. Image acquisition is performed noninvasively and nondestructively. The axial resolution of OCT lies in the micrometer range and is decoupled from the transverse resolution. This overcomes limitations associated with the numerical aperture of the objective and allows imaging with high axial resolution despite a moderate transverse resolution. Thus layered structures, such as human skin or the human retina, are well suited to be investigated with OCT. The penetration depth in tissue depends on the corresponding scattering properties and typically lies in the order of 1–2 mm. Structures that lie deeper within tissue can be visualized either because of overlying transparent tissue (as is the case in the human eye) or because of endoscopic probes that can be moved inside the body. OCT can be regarded as bridge between microscopic imaging technologies such as confocal microscopy (with limited depth penetration) to macroscopic imaging technologies such as ultrasound or computer tomography. OCT is a highly sensitive technique and can achieve sensitivities ranging up to 110 dB or more. This allows detection of light even from weak backscattering samples. Image acquisition in OCT is done through recording of depth profiles (A-scans). Several laterally displaced A-scans form a cross-sectional image or B-scan. In a volume scan of the sample several laterally displaced Bscans are recorded. Another advantage of OCT is the high imaging speed. A-scan rates of several megahertz have been reported. Thus highly sampled volumes or several volumes per second can be recorded. The highest impact of OCT certainly lies in the field of ophthalmology. No other imaging modality is capable to provide threedimensional (3D) information of retinal tissue. Diagnosis and treatment control of retinal diseases are strongly depending on OCT images and the impact to the field can be compared with the invention of the direct ophthalmoscope more than 200 years ago. Many instruments are commercially available and OCT imaging has found a wide range of applications. Apart from ophthalmology OCT imaging can be found in dermatology, cardiovascular research, pulmonology, and many more. In addition to the biomedical field OCT has been applied to material sciences and art preservation (Stifter, 2007).
Basic Principle of OCT The technique evolved from low coherence interferometry and was initially intended to measure axial distances in the eye. The basic principle is similar to ultrasound technology. However, instead of sound waves light is used to generate an image. The speed of light is much higher than that of sound waves. Thus, an interferometric technique is needed in order to distinguish between light that has traveled different distances in the micrometer range. Fig. 1(A) shows a scheme of an OCT interferometer. Light is emitted from an OCT light source and is split by the beam splitter into two arms, the reference arm and the sample arm, respectively. In the reference arm, the light is back reflected by the reference arm mirror. In the sample arm the beam is deflected on a scanning unit (for x- and y-scanning) before the light is sent to the sample. The light is backscattered from the sample, de-scanned and brought to interference with light from the reference arm at the beam splitter. The interfering light is then detected by the detector. It is very common to use fiber optics for the interferometer part. This eliminates the need for interferometer alignment and enables the fabrication of very compact systems. There are several techniques available for OCT.
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Fig. 1 (A) Scheme of an optical coherence tomography (OCT) interferometer. BS denotes the beam splitter of the interferometer. For time domain optical coherence tomography (TD-OCT) the reference mirror is mounted on a translation stage. (B) Measured intensity I at the detector (interference pattern) when the mirror is moved in axial direction (reflectivity in sample and reference arm are equal) full width at half maximum (FWHM) denotes the full width at half maximum of the coherence function which is defined as axial resolution in OCT. (C) Measured interference pattern when a glass plate with thickness d is measured (n denotes the refractive index of the glass plate).
Timed Domain OCT Initially the so-called time domain OCT (TD-OCT) technique has been used. In this technique the interferometer is illuminated with a broadband light source. To generate an A-scan (depth profile) the reference mirror is moved axially. In the case of a mirror in the sample arm, the detector will record an intensity pattern that is shown in Fig. 1(B). Due to the broad bandwidth of the source, interference fringes will only be visible if sample and reference arm lengths are matched (within the coherence gate). Thus light originating from different depths can be separated. In the case of a glass plate the light will be backscattered (or back reflected) at the front and back surface of the plate. During axial scanning of the reference arm mirror two coherence functions (arising at the surface and back surface of the glass plate) as illustrated in Fig. 1(C) will be detected. The distance between the two signals corresponds to the optical thickness of the plate (i.e., the geometrical thickness multiplied with the refractive index of the plate). The full width at half maximum (FWHM) of the coherence function is defined as the axial resolution of OCT and equals the round trip coherence length ( ¼ 1/2 of the coherence length) of the light source. The A-scan rate of this technique is limited by the inertia of the mirror to a few 100 Hz. However, the method can be highly sensitive. Motion of the mirror will introduce a Doppler frequency shift for each wavelength. Thus, a carrier frequency for the OCT signal is generated which allows for band pass filtering of the signal. This procedure filters 1/f noise out and increases the sensitivity. Depending on the bandwidth of the light source and the speed of motion a certain electronic signal bandwidth needs to be detected. Faster speeds will require larger electronic bandwidths (of the detector and amplifying electronics) and thus lower system sensitivity. To overcome the limitation given by the inertia of the mirror, a galvanometer scanning mirror and a grating (rapid scanning optical delay line) can be placed in the reference arm. Instead of a translational motion in spatial domain, this configuration allows for a rotational motion in wavelength domain. The method has the advantage that the group delay (axial scanning of the sample) and the phase delay (modulation frequency of the OCT signal) can be adjusted independently. Therefore the imaging speed can be improved up to a few kilohertz. TD-OCT detects only light backscattered from within the coherence volume of the sample. The coherence volume is defined through the round trip coherence length and the transverse resolution of the system. Light from other depths (outside the coherence volume) is rejected which limits the sensitivity of the method. However, the depth range of the system depends on the travel range of the reference mirror. Thus, in principle very long axial distances without loss in sensitivity can be scanned with TD-OCT.
Fourier Domain OCT With the development of Fourier domain (FD) OCT the sensitivity could be greatly improved. The gain in sensitivity can be directly translated into imaging speed that allows A-scan rates an order of magnitude higher than in TD-OCT. Two different FDOCT schemes are possible. The spectral domain (SD) OCT technique uses a similar broadband light source as in TD-OCT. However, instead of a single detector at the interferometer exit the light is spectrally dispersed trough a grating and detected with a line scan camera (cf. Fig. 2(A)). Thereby, the reference arm mirror is kept stationary. In this way light from the entire sample depth is recorded simultaneously which results in a sensitivity advantage compared to TD-OCT. In case of a mirror in the sample arm (same light power returning from sample and reference arm) the spectrometer camera will detect an interference pattern that is displayed in Fig. 2(B). The modulation of the spectrum arises from constructive and destructive interference for each wavelength and depends on the path length difference between sample and reference arm beam. Thus the frequency of the modulation depends on the amount of path length difference. Smaller differences result in low frequencies, while larger differences cause higher frequencies. In addition, the spectrum (and the interference pattern) is detected in the wavelength domain. Thus, each modulation frequency that corresponds to a certain depth will have a chirp along the spectrum. To remove this chirp, the spectrum needs to be transformed into k-space ( ¼1/l-space). Associated with this chirp will
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Fig. 2 (A) Scheme of a spectral domain optical coherence tomography (SD-OCT) instrument using a spectrometer in the detection arm (BS, beam splitter). (B) Measured interference pattern when a mirror is placed in the sample arm (light power returning from sample and reference arm are equal). (C) Fast Fourier transformation of the spectrum measured in (B) corresponding to a depth profile.
be a drop of sensitivity (sensitivity fall off) with depth because of the finite width of the detector elements (camera pixels). Higher frequency components (constructive and destructive interference of neighboring wavelengths) will therefore be detected by a single detector element and thus not be visible. It should be noted that FD-OCT requires a minimum detection speed. Jitter within the interferometer or axial movement of the sample (in case of in vivo imaging) will cause a change (phase shift) in the interference pattern which degrades the visibility of the interference pattern when this change occurs during the exposure time of the camera. In order to avoid this “fringe washout” the exposure time should be shorter than 1 ms. In the case displayed in Fig. 2(B) one arm is only slightly longer than the other. At this point it should be noted that it is not possible to distinguish (without any other means) between positive (reference arm is longer) or negative (sample arm is longer) path length differences because both will result in the same frequency modulation of the spectrum. In order to generate an A-scan the interference pattern (modulated spectrum) needs to be Fourier transformed. The corresponding depth profile is plotted in Fig. 2(C). Because of the inability to distinguish between positive and negative frequencies, the transformation of the spectrum yields in fact two depth profiles that are mirrored around zero delay. Zero delay refers to the position where sample and reference arm lengths are exactly matched. The total number of sample points after Fourier transformation stays constant. However, this means that the number of sample points per A-scan is only half the number of camera pixels. Because the mirrored A-scan does not contain additional information it is very common in OCT to simply omit this part as is done in Fig. 2(C). The depth range of a SD-OCT system depends on the spectral resolution of the spectrometer as increasing depth is associated with higher modulation frequencies that need to be sampled according to the Nyquist theorem. This resolution is determined by the dispersion properties of the grating (number of lines that are illuminated on the grating), the spectrometer optics and the number of pixels that are available on the line scan camera. According to Leitgeb et al. (2003) the maximum depth range can be calculated via zmax ¼
l2 4δl
ð1Þ
where l and δl denote the wavelength and the spectral resolution, respectively. A typical OCT configuration uses a light source with a central wavelength of 840 nm, a bandwidth of 50 nm, and a 2048 pixel camera. Assuming that 150 nm bandwidth is imaged onto the camera and that the separation between spectral lines (corresponding to the resolution provided by the grating) is smaller than the size of a pixel an imaging depth of 2.4 mm is achieved. Back reflections from larger imaging depths will appear attenuated as aliasing frequencies and may cause unwanted artifacts inside an OCT image. Another technique for realizing FD-OCT is swept source (SS) OCT. SS-OCT uses fast (wavelength) tunable lasers as light source in combination with a single (fast) photo detector. During a laser sweep the interference pattern is recorded depending on wavelength and separated in time. The pattern will be similar to the one displayed in Fig. 2(B). It should be noted that the instantaneous power at each wavelength can be higher compared to SD-OCT which compensates for the nonparallel data acquisition scheme of SS-OCT compared to SD-OCT. Similar as in SD-OCT, light backscattered from the entire sample depth contributes to the interference pattern and the depth profile is reconstructed in a similar way via Fourier transformation of the recorded spectral information. The instantaneous linewidth of the source determines the instantaneous coherence length and defines the depth range of the system. The instantaneous linewidth thus plays an equivalent role as the spectral resolution of SDOCT. Within this range light from the reference arm and sample arm are coherent and can interfere. However, the instantaneous line width is a completely different measure as the round trip coherence length. The latter is defined by the FWHM of the laser sweep, while the former is the width of a laser line emitted at a certain wavelength within the sweep. Fig. 3(A) illustrates the emission of a SS laser. The output power of many swept laser sources fluctuates. In order to compensate for this fluctuations balanced photo detection is employed. Fig. 3(B) shows a schematic of a fiber-based SS-OCT instrument. The light emitted from the SS is already confined in a single mode fiber. Before the entrance of the interferometer the light traverses a fiberized circulator and is split thereafter by a 50:50 beam splitter into reference and sample beam, respectively. The light back reflected from both arms is recombined by the 50:50
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Fig. 3 (A) Illustration of a swept source (SS) laser sweep. The sweep range determines the axial resolution of the system. The instantaneous line width (drawn not to scale) determines the imaging depth provided by the light source. (B) Representative scheme of a SS optical coherence tomography (OCT) instrument based on fiber optics and balanced photo detection. BS 50:50, fiber beam splitter; GS, galvanometer scanner.
beam splitter and is detected with a balanced photo detector. The circulator ideally transmits nearly 100% of the light from the light source to the interferometer, while nearly 100% of the light coming from the interferometer is transmitted to the detector. There are several advantages of SS-OCT compared to SD-OCT. In SD-OCT the light is detected with a spectrometer. The efficiency of the spectrometer cannot be 100% which results in a lower sensitivity compared to SD-OCT. In addition the wavelength sweep can be performed in 1/l space which results in an elimination of the frequency chirp that is present in spectrometer-based systems and correspondingly into an elimination of the sensitivity drop with depth. Finally, the spectrometer needs to be perfectly aligned in order to achieve a high sensitivity which sets high demands, from a technical point of view, on the spectrometer design in commercial instruments. No such alignment is required for SS-OCT instruments.
Resolution The FWHM of the interference pattern (cf. Fig. 1(B)) defines the axial resolution of OCT and is inversely proportional to the bandwidth of the light source. In case of a Gaussian-shaped spectrum the resolution can be calculated via following equation (Drexler and Fujimoto, 2015): Dz ¼
2lnð2Þl npDl
2
ð2Þ
where l denotes the central wavelength, Dl the bandwidth, and n the refractive index of the medium. As can be seen from this equation, the resolution in OCT depends on the central wavelength of the light source and the corresponding bandwidth. Fig. 4 shows the axial resolutions for three wavelength regions that are commonly used in OCT. Standard OCT resolution refers to bandwidths around 50 nm corresponding to an axial resolution between 5 and 15 mm. The use of larger bandwidths (100 nm or more) is known as ultrahigh resolution OCT and requires ultra-broadband light sources, such as titanium sapphire laser or supercontinuum light sources. Important for achieving the resolution in OCT that is provided by the light source is a sufficient optical transparency of the optical components and matching of dispersion (see below) introduced in reference and sample arm, respectively. Especially fiberized interferometers may cause unwanted effects because single mode propagation (multimode propagation results in coherent image artifacts) depends on the size of the core diameter in respect to the wavelength. Thus larger wavelengths will be stronger attenuated if the core size is chosen to ensure single mode propagation for the shortest wavelengths. The transverse resolution mainly depends on the imaging optics in the sample arm. In order to achieve maximum fringe visibility it is common in OCT to use light with high spatial coherence. Light emitted from a single mode fiber has this property. Because of its directionality, divergence, and Gaussian intensity profile, the beam propagation can be described using Gaussian optics. In many OCT systems the objective lens is not the limiting aperture as the beam is scanned over this lens in order to achieve transverse scanning of the sample. Hence, the beam remains Gaussian and the resolution can be calculated via the FWHM of the Gaussian beam waist o0 and the angular spread yS of the Gaussian beam (Fercher, 2010): pffiffiffiffiffiffiffi pffiffiffiffiffiffiffi l Dx ¼ 2 ln2o0 ¼ 2 ln2 pyS
ð3Þ
l denotes the central wavelength of the beam. Strictly speaking, this resolution can be obtained only at the focal (or certain depth) plane. Outside this focal plane the resolution is degraded. A measure for the imaging depth where the resolution is degraded only by a factor of √2 is the Rayleigh range zR Twice this range defines the depth of focus (DOF), which can be calculated with 2zR ¼ DOF ¼ 2
l py2S
ð4Þ
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Fig. 4 Axial resolution of an optical coherence tomography (OCT) system depending on the bandwidth and the central wavelength for three commonly used wavelength regions. Standard OCT uses a bandwidth between 30 and 70 nm, which results in a resolution below 20 mm. Ultrahigh resolution (UHR) OCT refers to the use of larger bandwidths (4100 nm) and corresponding resolutions that are below 5 mm.
Thus for high-resolution systems only part of the imaged volume will be in focus and imaged sharply. For an OCT system at 1300 nm and 20 mm transverse resolution the DOF will be 2 mm. This DOF will be reduced to 20 mm in the case of 2 mm transverse resolution. Hence, the imaging optics needs to be carefully designed in order to optimize the OCT performance. Several methods have been proposed to maintain high transverse resolution throughout the imaging depth. These are based on a dynamic shift of focus with the coherence gate (only possible for TD-OCT) or on the use of special illumination techniques (such as Bessel beams).
Sensitivity Several noise sources influence the sensitivity in OCT. Receiver noise from the photo detectors, excess noise arising from fluctuations in light power, and shot noise that originates from the quantum nature (and random arrival) of photons. Ideally, an OCT system is limited by shot noise only. This limit can be reached through the coherent amplification process inherent to OCT as the light power returning from the reference arm greatly exceeds the power returning from the sample. In the case of sufficient light power from the light source, the amount of light returning from the reference arm can be optimized in order to reach the shot noise limit. In TD-OCT the shot noise limited sensitivity depends on the detection bandwidth and the power efficiency of the interferometer (Izatt et al., 2015): Sensitivity ¼
rPS 2eDBTD
ð5Þ
where r denotes the quantum efficiency of the detector, PS is the light power returning from the sample arm (with a 100% reflective mirror as sample), e is the electron charge (e ¼ 1.6 1019 1C), and DBTD is the electronic detection bandwidth of the TDOCT systems. The light power returning from the sample depends on the power incident on the sample, the interferometer configuration, and the optical losses within the system. Let us consider a fiber-based Michelson interferometer with a 50/50 splitting ratio and 4 mW of optical power (central wavelength at 1300 nm with a bandwidth of 50 nm) on the sample. The light is emitted from a single mode fiber. Thus the coupling efficiency (of the light returning from the sample) will be B50%. Another 50% of the light is lost due to the fiber beam splitter which results in an optical power returning from the sample arm (with a 100% reflective mirror as sample) of 1 mW. When the above mentioned TD-OCT instrument is operated at 1 kHz A-scan rate and 3 mm depth range a minimum electronic bandwidth of 800 kHz will be required. This results in a sensitivity of the OCT system (assuming 0.8 quantum efficiency of the detector) of 103 dB.
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In FD-OCT the sensitivity can be expressed in a similar way (Leitgeb et al., 2003): Sensitivity ¼
rPS M 4eDBFD
ð6Þ
with DBFD the electronic bandwidth of FD-OCT systems (inversely proportional to the exposure time of the camera in SD-OCT) and M the number of spectral channels of the FD-OCT systems. By comparing this equation with Eq. (5) we can see that the sensitivity of FD-OCT is a factor of M/2 larger than in TD-OCT. Thus, for a FD-OCT instrument operating with the same light source and interferometer as before will have an improved sensitivity. Nevertheless, following issues need to be considered. The light returning from the sample will be slightly attenuated by the efficiency of the spectrometer (this typically lies in the range of 60%) and the exposure time of the camera will be shorter than the time corresponding to the A-scan rate because the camera requires a certain readout time. For an A-scan rate of 70 kHz we assume that 1000 pixels on the camera are (equally) illuminated and that the camera requires 4 ms for readout. This results in an exposure time of B10 ms. The detection bandwidth B of FD-OCT is given by (Leitgeb et al., 2003) BFDOCT ¼
2zR lc t
ð7Þ
with zR denoting the depth range of the system and lc the double pass coherence length. t is the exposure time of the camera. In our case (zR ¼6 mm, lc ¼ 15 mm, and t¼10 ms) the detection bandwidth is 80 Mhz which results in a shot noise limited sensitivity of 108 dB. Thus, a higher sensitivity can be achieved with this SD-OCT system although the imaging speed is increased by a factor of 70.
Dispersion Dispersion describes the effect of a wavelength dependent refractive index of a medium. For example, the refractive index of water (cf. Fig. 5) is not constant and decreases with wavelength. For OCT dispersion mismatch between the interferometer arms results in a degradation of the axial point spread function. This is associated with a degradation of axial resolution and sensitivity. Thus the interferometer arms need to be balanced which is normally achieved by inserting corresponding materials that are present in the sample arm (such as lenses for scanning optics) into the reference arm of the interferometer. For compensating dispersion of the eye cuvettes that are filled with water are commonly used. FD-OCT records spectral data. Therefore it is possible to compensate for dispersion in post-processing.
Wavelength Range The choice of imaging wavelength in OCT depends on the investigated sample as well as on the availability of light sources. For a high penetration depth the total attenuation coefficient of tissue at the imaging wavelength should be minimal. Both, absorption and scattering contribute to the total attenuation coefficient. Water is the main constituent of tissue (apart from teeth
Fig. 5 The wavelength dependence of the refractive index of water in the visible and near-infrared (NIR) region. Data from Hale, G.M., Querry, M. R., 1973. Optical-constants of water in 200-nm to 200-mm wavelength region. Applied Optics 12(3), 555–563.
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Fig. 6 Absorption coefficient of water and scattering coefficient of dense tissue. The three main optical coherence tomography (OCT) imaging wavelengths (840, 1060, 1300 nm) are outlined in the graph. Data from Hale, G.M., Querry, M.R., 1973. Optical-constants of water in 200-nm to 200-mm wavelength region. Applied Optics 12(3), 555–563.
or bones) and its absorption properties will be dominating. As can be seen in Fig. 6 the absorption coefficient of water strongly increases from visible light to infrared light. On the other hand, scattering gradually decreases in this wavelength range. At the 1300 nm wavelength region the contribution of scattering and absorption are largely within the same range. Thus, this wavelength region is the preferred choice for imaging dense tissue such as human skin as for shorter wavelengths the penetration will be limited by scattering, while for longer wavelengths absorption will be dominating. Imaging at 1300 nm requires detector materials that are sensitive in this region. These are much more expensive than comparable materials that are sensitive in the visible range. As such the preferred OCT imaging technology will be based on SS-OCT because this technology requires a singlebalanced detector (instead of a detector array as in SD-OCT) and provides some benefits (as outlined above) compared to the SD-OCT counterpart. For imaging the retina, the light has to traverse the ocular media (eye length of B2 cm) twice which significantly changes the situation. The scattering properties of the retina are of less importance compared to the transmittance of the light through the eye. Ocular media are very weakly scattering (apart from pathologies such as cataract). Thus the imaging wavelength should be as close to the visible range as possible to minimize absorption by water. However, for highly sensitive measurements the quantum efficiency of the detector should be high. The used detector material, silicon, has a peak around 900 nm and levels off toward shorter wavelength. Thus imaging wavelength close to the maximum of the quantum efficiency of the detector is desirable. An additional aspect is that in the visible range the photoreceptors of the retina will be very sensitive to light (tails of the absorption spectra reach into the infrared region). Hence, the subjective perception of the light power will increase rapidly at wavelength regions that are close or within the visible range. This is associated with increasing subject discomfort during imaging, even though the light levels that are sent to the eye have to be below the limits for safe exposure. In addition dispersion is more pronounced for shorter wavelength regions. Because of these issues the wavelength region around 840 nm has been found as compromise and is widely used in OCT for retinal imaging. Since line scan cameras in this wavelength region are easily available and relatively cheap, the preferred imaging technique in this wavelength region is SD-OCT. Nevertheless, another wavelength region at 1060 nm which is around a local minimum of the water absorption spectrum (cf. Fig. 6) has gained interest in recent years. Due to the longer wavelength region the penetration depth within the retina can be increased and the influence of media opacities (such as cataract) on the image quality will be lower. Another advantage of this wavelength region is the availability of fast SSs. Thus high sensitivity can be maintained throughout the imaging depth which further improves the retinal image quality. Due to technical issues (dispersion introduced by fiber optics), SSs at 840 nm are relatively slow compared to their counterparts at 1060 or 1300 nm.
OCT Angiography In recent years, OCT angiography (OCTA), a method to visualize vessel structure within tissue, has gained increasing interest. There are many different approaches to realize OCTA. The underlying basic principle, however, is the same for all methods. The idea is that two (or more) measurements (A-scans or B-scans) are taken at the same location but separated in time. By comparing these measurements changes (intensity or phase or both) caused by blood flow can be detected. The time separation between these
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measurements needs to be short enough in order to minimize the influence of sample motion and long enough in order to ensure sufficient changes introduced by blood flow. With A-scan rates of 100 kHz or more the most common OCTA technique is based on recording several B-scans at the same location. This, however, increases the overall measurement time. The major benefit of OCTA is that it provides information on the microvascular perfusion of tissue noninvasively. In addition, due to the spatial separation of capillaries, these can be visualized even though the lateral extension lies below the transverse resolution of the system. Other techniques that are capable to provide information on the vascular structure require the application of contrast agents. Some subjects may show allergic reactions or other complications after administration of such contrast agents. Thus state of the art angiographic imaging is always associated with residual risk factors. Another benefit of OCTA lies in the fact that depth resolved information on the vasculature is provided. This allows for a separation between individual vascular beds and thus an improved image contrast. It should be noted, however, that OCTA is not able to detect vessel leakages because the technique relies on motion contrast. Within leakages motion occurring between individual B-scan recordings will be too slow compared to subject motion to be detected in vivo. However, in vitro experiments have shown that without bulk sample motion slow fluid motion (such as Brownian motion) can be visualize with OCTA techniques.
Imaging of Human Skin The human skin is easily accessible with optical methods. In clinical routine, lesions will be first inspected visually with high magnification using a microscope (dermatoscope). In the case of suspicious lesions a biopsy will be taken which will be sliced and prepared for histologic examination. The cellular details provided by this histologic examination can then be used to determine the malign or benign status of a lesion and finally determines the further clinical procedure. OCT may represent an alternative to the histologic examination as it provides in vivo and noninvasively cross-sectional images of skin. Current studies focus on improving the sensitivity and specificity of OCT methods for detecting malign tissue (Ulrich et al., 2016). Thereby, additional contrast mechanism such as polarization sensitive OCT or OCTA are used to gather additional information. Especially OCTA provides essential additional information as tumors, such as basal cell carcinoma (BCC) are known to present with increased vascularization. Using fiber-based or handheld systems practically every location can also be reached with OCT. In order to minimize scattering and absorption, the preferred imaging wavelength for human skin will be in the 1300 nm range. The penetration depth of OCT in skin is in the order of 1–2 mm. Fig. 7 shows a representative B-scan image of a finger recorded with a 1310 nm SD-OCT instrument. The A-scan rate of the instrument was 91 kHz. The image consists of 500 A-scans and was recorded in B5 ms. The bandwidth of the light source was 50 nm which results in an axial resolution of 10 mm in tissue (assuming a refractive index of 1.4). The power onto the skin was 8 mW. Individual layers of the skin can be clearly seen in Fig. 7 and are labeled in the image. Supplementary material related to this article can be found online at http://dx.doi.org/10.1016/B978-0-12-803581-8.09787-3. An interesting skin region is the nail fold of a finger. In this region capillaries are close to the surface (close to the nail some capillaries extend parallel to the surface) and can be investigated with optical methods. In order to assess the status of the capillary network, capillaroscopy is used in clinical routine. This method examines the skin with a microscope after applying a contrast agent (such as oil) to the skin region. Using OCTA this network can be visualized completely noninvasively and without the
Fig. 7 Representative B-scan image recorded at 1300 nm of the inner side of a finger. Superficial layers down to deeper layers of the dermis can be visualized. The multimedia file (media 1) shows a fly through of B-scans of the entire three-dimensional (3D) data set.
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Fig. 8 Representative optical coherence tomography (OCT) images recorded in the nail fold region of a healthy volunteer. (A) OCT-intensity B-scan (averaged over 5 B-scans). (B) Depth integrated en-face intensity image. The arrow indicates the location of the B-scan displayed in (A). (C) Depth integrated en-face intensity image and overlaid (red) vessel map.
application of a contrast agent. Fig. 8 shows representative image data recorded with the SD-OCT instrument described above. The intensity B-scan (Fig. 8(A)) clearly shows different structures such as the nail plate and different skin layers of the nail fold. The banding structure within the nail plate is an artifact that originates from birefringence of the plate. The sensitivity of the system is sufficient to penetrate through the nail plate as structures underneath the plate can be seen. The axial displacement of the nail plate beneath the nail fold (cf. right hand side in Fig. 8(A)) is caused by the increased optical path length that is introduced by light propagation through the overlying nail fold tissue. The en-face intensity projection clearly shows the different regions of the nail fold and nail plate. Fig. 8(C) shows a composite image where the depth integrated vessel structure is overlaid to the intensity image of Fig. 8(B). In this image the capillary network can be clearly seen. Close to the nail plate individual capillary loops can be seen which are typical for this skin region.
Imaging of Human Retina The human retina consists of several layers and represents an essential part for the vision process. All layers that can be seen in histology can be visualized with OCT. Fig. 9 shows a representative OCT B-scan image of the fovea region with a labeling of the different retinal layers. In healthy subjects the fovea is the location of best visual acuity. In this region the highest density of cone photoreceptors can be found. These cells are responsible for photopic vision and can be distinguished into three types (sensitive in the short, medium, and long wavelength region). In the fovea no rod photoreceptors (responsible for scotopic vision) are present. The light is incident from the top of this image and has to penetrate different retinal layers until it is detected by the photoreceptors. The pigments of the photoreceptors are located within the outer segments of cones and rods. The image was recorded with a research grade prototype SD-OCT instrument operating at 863 nm with a bandwidth of 60 nm which resulted in an axial resolution of B4 mm in tissue (assuming a refractive index of 1.4). The A-scan rate was 70 kHz and one B-scan consisting of 1024 A-scans was recorded in B15 ms. Although the axial resolution is sufficient to resolve individual retinal layers, individual cells (such as photoreceptors) cannot be visualized. In order to increase the transverse resolution, aberrations that are introduced by imperfections of the eye optics need to be compensated for. This can be done with the implementation of adaptive optics, a technology known from astronomy (Pircher and Zawadzki, 2007). Many commercial OCT devices for imaging the retina are available. Fig. 10 shows representative data recorded with a commercial SD-OCT instrument. The instrument provides a scanning laser ophthalmoscope (SLO) overview mode that is used for retinal tracking. The location of the recorded B-scan is marked with a green line in the SLO image. The image quality is very similar to the previous image. The entire 3D data set can be viewed in media 2. Supplementary material related to this article can be found online at http://dx.doi.org/10.1016/B978-0-12-803581-8.09787-3. In order to outline the difference of retinal images recorded at different wavelengths, Fig. 11 shows a retinal image recorded at 1040 nm. Compared to the previous images enhanced penetration into the choroid and even into the sclera can be seen. Otherwise all retinal layers can be seen with similar quality. The image was recorded with a research grade SS-OCT instrument operating at 100 kHz A-scan rate and 20 20 field of view. The axial resolution of the system was 5.6 mm in tissue. Through a change in the scanning pattern (recording of several B-scans at the same location) angiographic data can be retrieved. Fig. 12 shows representative data of a healthy volunteer recorded with a commercial instrument. The en-face projection images were generated through depth integration over the depth extension indicated with white arrows in Fig. 12(A)). Fig. 12(B)) shows cross-sectional angiographic data (red) overlaid to the intensity image. The contrast of retinal vessels is greatly increased using the angiographic data evaluation. The en-face projection of the intensity images (cf. Fig. 12(C)) shows larger vessels as shadows.
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Fig. 9 Optical coherence tomography (OCT) image of the retina with a labeling of the different layers (the image spans an angle of 28 degree which corresponds to B8 mm on the retina). COST, cone outer segment tips; ELM, external limiting membrane; GCL, ganglion cell layer; INL, inner nuclear layer; IPL, inner plexiform layer; IS/OS, junction between inner and outer segments of cone photoreceptors; ONL/HF, outer nuclear layer/Henle’s fiber layer; OPL, outer plexiform layer; PRL photoreceptor layer; RNFL, retinal nerve fiber layer; RPE, retinal pigment epithelium.
Fig. 10 Representative retinal image data of the fovea of a healthy volunteer recorded with a commercial instrument (imaging wavelength¼840 nm). Left: overview scanning laser ophthalmoscope (SLO) image. Right: optical coherence tomography (OCT) B-scan recorded at the location indicated in the left image with the green arrow. The multimedia file (media 2) shows a fly through of B-scans of the entire 3D data set.
However, the entire vascular structure can be seen in the corresponding OCTA projection map. This map clearly shows the avascular zone of the fovea region. From these maps deteriorations (such as non-perfused areas) from the normal vasculature structure can be easily determined. Fig. 13 shows representative image data of patients with age related macular degeneration (AMD). AMD is regarded as one of the major causes of blindness in the industrialized world. The disease mainly affects the macula region (containing the fovea) of subjects that are older than 60 years. AMD can be differentiated into several forms: dry AMD and wet AMD. The latter is the most severe form which results (untreated) into a rapid degradation of visual acuity. An early indicator of AMD is the presence of Drusen. Drusen are elevations of the retinal pigment epithelium (RPE)/photoreceptor bands because of accumulation of retinal debris underneath these layers. Fig. 13(A) shows a typical example of OCT images recorded in a patient with Drusen. The elevations of the photoreceptors cause image distortions and degradation in visual acuity of the patient. At locations of the Drusen an additional layer (Bruch’s membrane) can be seen. In the healthy case, this layer is obscured by the overlying RPE. Due to the highly scattering of the pigments within the RPE, this layer appears as broadened diffuse layer although it is a mono cellular layer. The 3D visualization of Drusen allows for a quantitative assessment of the size and volume of the Drusen. As such the effect of therapeutic interventions on the Drusen size can be precisely monitored. The most severe form of AMD is choroidal neovascularization (CNV). In this case new vasculature is generated within the retina and leakages of these lead to damage of the sensitive neural tissue of the retina. Fig. 13(B) shows a representative B-scan of the fovea region of a patient with CNV. Scarring of retinal tissue and associated disintegration of photoreceptors is very often accompanied by this disease. Using OCT the size of edema or cysts can be quantified. In addition vascular growth can be determined using OCTA. Therapeutic intervention (such as anti-vascular endothelial growth factor (anti-VEGF) therapy) can be easily monitored and the technology can be meanwhile regarded as gold standard for this kind of diagnosis and for treatment control.
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Fig. 11 Retinal image of the fovea of a healthy volunteer recorded with a swept source optical coherence tomography (SS-OCT) instrument at 1040 nm. Enhanced penetration into the choroid and sclera can be observed.
Fig. 12 Retinal images of the fovea of a healthy volunteer recorded with a commercial instrument. (A) Central B-scan of the volume scan. The arrows indicate the image depth that was used to generate the en-face images in (C) and (D). (B) Central B-scan of the volume scan and overlaid vessel structure (in red) retrieved with the angiography algorithm. (C) En-face intensity image (white ¼high intensity, black¼low intensity) generated through depth integration over the area indicated in (A). En-face vessel map generated through maximum intensity projection of the angiographic data within the area indicated in (A).
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Fig. 13 Representative optical coherence tomography (OCT) B-scans (averaged over 50 frames) at 840 nm of patients with age related macular degeneration (AMD). (A) OCT B-scan image of Drusen (some are indicated by white arrows). The black arrow indicates Bruch’s membrane. (B) OCT B-scan image of a choroidal neovascularization (CNV). Images are displayed on a logarithmic intensity scale.
There are many other retinal diseases where OCT plays a key role in diagnosis. A summary of the impact of OCT to all these diseases will be beyond the scope of this article. Nevertheless, one specific example should be mentioned: Glaucoma, a disease that is associated with visual field defects and one of the major causes for blindness worldwide. At an early stage of the disease, these defects occur in the periphery of the retina which is not easily noticeable by patients themselves because for reading etc. only the fovea region is used. Before changes in the visual field can be detected, a loss of retinal nerve fiber tissue (loss in layer thickness) will occur. The 3D imaging capabilities of OCT enable a precise measurement of the retinal nerve fiber layer thickness. Thus OCT provides an excellent indicator for early signs of Glaucoma and for monitoring therapeutic success.
Conclusion and Outlook OCT is a very powerful imaging technology. With the development of FD-OCT imaging speed could be greatly improved and Ascan rates in the megahertz range have become possible. This paved the way for many new applications of OCT. This article focused on skin and retinal imaging but many more applications of OCT are available. Associated with the improved imaging speed is the possibility to investigate dynamic processes or to increase the contrast through the introduction of angiographic evaluation of OCT data. The noninvasive visualization of capillaries in tissue, such as skin or retina, opens completely new perspectives for clinical research. The high clinical demand of this technology manifests itself through the rapid commercialization of OCTA for both, skin, and retinal imaging. Thereby, OCT certainly plays a key role for accurate diagnosis especially in ophthalmic routine. Nevertheless, there are new developments in the field. Currently new SSs are developed that provide higher speeds (A-scan rates) and very long coherence lengths. Meanwhile, imaging depth ranges of some meters have been demonstrated with OCT. This opens new possibilities such as imaging of the whole eye (from the cornea to the retina) with a single SS-OCT system. In addition further parallelization of OCT is investigated. Line field OCT, for example, records one B-scan in parallel. Thereby SDOCT (this requires an area detector in the spectrometer) or SS-OCT can be used. The latter uses a SS in combination with a line scan camera. Very high speeds (several volumes per second) can be achieved using a full-field illumination in combination with a SS and a very fast area camera. The demands on the camera are very high because the phase stability (to avoid “fringe washout”) requires the recording of a volume within a short time frame (ideally below 1 ms). In order to have sufficient sampling points in depth (determined by the number of spectral lines) the frame rate of the camera needs to be in the order of 50 kHz. Apart from the high imaging speeds that are possible and that have been demonstrated for in vivo retinal imaging, the full-field technique provides phase stability over the entire field of view (in lateral direction). OCT provides direct access to the phase of the electric light field and computational techniques can be applied in order to compensate for aberrations introduced, for example, by the eye. Thus high resolution can be achieved even without the use of hardware-based adaptive optics. In addition, the focal plane of the image can be adjusted in post-processing. As such, an entirely sharp volume can be obtained with high transverse resolution despite of a limited DOF. Apart from these developments, OCT has become a valuable tool in cardiovascular research. Using fiberized probes, OCT imaging can be done from inside the body, for example, within a coronary artery. As such plaques, deposits, or stents within the artery can be visualized in vivo. Key element for this development is the high imaging speed provided by OCT. Currently clinical trials are performed to demonstrate the benefit of cardiovascular OCT compared to state of the art examination tools. Another development in the field is the combination of OCT with other imaging technologies. For example, the combination of OCT with photo acoustics has gained increasing interest. The photo acoustic technique relies on absorption of the probing light beam and provides a higher penetration depth but lower resolution than OCT. Therefore vasculature that is located deeper within
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tissue can be visualized. The combination with OCTA enables the visualization of the entire vascular network as superficial capillaries are detected with OCT, while the deeper vascular structure is detected with photo acoustics.
Acknowledgments The author likes to thank following people from the Medical University of Vienna for their assistance: Christoph K. Hitzenberger, Matthias Salas, Martin Fürst, Mitsuro Sugita, Teresa Torzicky, and Ursula Schmidt-Erfurth.
See also: Interferometric Imaging. Overview: Coherence
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