An interference effect fast modulations in intensity are seen at the detector only if the time travelled by light in the reference and sample arms is nearly equal. Thus the presence of interference serves as a relative measure of distance travelled by light. Optical coherence tomography leverages this concept by replacing the mirror in the sample arm with the sample to be imaged Figure 2.
The reference arm is then scanned in a controlled manner and the resulting light intensity is recorded on the detector. The rapid modulation interference pattern occurs when the mirror is nearly equidistant to one of the reflecting structures in the sample, and can be processed to register the presence of that structure. The distance between two mirror locations where interference occurs corresponds to the optical distance between two reflecting structures of the sample in the path of the beam.
Even though the light beam passes through different structures in the sample, the low-coherence interferometry described above helps to distinguish the amount of reflection from each unique structure in the path of the beam. In doing the so, the material scattering, and hence structure, can be measured as a function of depth. Exploring the transverse or x-y localization of the sample structure is simpler.
The broadband light source beam that is used in OCT is focused to a small spot on the order of a few micrometers and scanned over the sample. By mapping the sample in x and y with a scanning arm while collecting depth information using interferometry, a complete 3D picture of the sample can be constructed. Instead of recording intensity at different locations of the reference mirror, the intensity is recorded as a function of the wavelengths or frequencies of the light.
The intensity modulations when measured as a function of frequency are called spectral interference. The rate of variation of intensity over different frequencies is indicative of the location of the different reflecting layers in the samples. It can be shown that a Fourier transform of spectral interference data provides information equivalent to that which would be obtained by moving the reference mirror Figure 3.
There are two common methods of measuring spectral interference in OCT: spectral-domain and swept-source. In spectral-domain OCT SD-OCT , a broadband light source delivers many wavelengths to the sample, and all are measured simultaneously using a spectrometer as the detector.
In swept-source OCT SS-OCT , the light source is swept through a range of wavelengths and the temporal output of the detector is converted to spectral interference. Fourier-domain OCT allows for much faster imaging than scanning of the sample arm mirror in the interferometer, as all the back reflections from the sample are being measured simultaneously. This speed increment introduced by Fourier-domain OCT has opened a whole new arena of applications for the technology. Live video, in-vivo OCT imaging can be easily obtained using commercial systems, allowing it to be used for process monitoring and guided surgery.
The customer should seek devices that incorporate other capabilities such as microperimetry. Showing 4 of 4 products. Select All. Select up to 5 products from below to compare or request more information. Featured products. Compare for selected. These sequentially acquired OCT-B scans are compared; the decorrelation signal used to generate the image of the vasculature corresponds to regions of erythrocyte movement. Though currently available OCTA devices use SD-OCT systems, prototype swept-source devices have been used in the research setting to visualize changes in the choriocapillaris in diabetes and age-related macular degeneration.
Both spectral-domain and swept-source devices have been used to qualitatively and quantitatively describe the microvascular morphology of choroidal neovascularization.
In the United States, swept-source devices will soon become commercially available, expanding the number of choices available to ophthalmologists when considering OCT systems. Swept-source OCT devices are able to operate at higher scanning speeds than spectral-domain systems and, though their clinical superiority is still unclear, remain promising prospects for future imaging development. She can be reached at emilycole10 gmail. His email is jduker tuftsmedicalcenter.
Cole has no financial disclosures. Selected References 1. Fujimoto J, Swanson E. The development, commercialization, and impact of optical coherence tomography. Visualizing the choriocapillaris under drusen: Comparing nm swept-source versus nm spectral-domain optical coherence tomography angiography. OCT using a frequency-tunable optical source. Opt Lett ; Opt Express ; Three-dimensional and high-speed swept-source optical coherence tomography for in vivo investigation of human anterior eye segments.
Sensitivity advantage of swept source and Fourier domain optical coherence tomography. Investigating the choriocapillaris and choroidal vasculature with new OCT technologies.
Progress in retinal and eye research ; Cross-sectional and en face optical coherence tomographic features of polypoidal choroidal vasculopathy.
Retina ; Correlation of indocyanine green angiography and optical coherence tomography findings after intravitreal ranibizumab for polypoidal choroidal vasculopathy. Am J Ophthalmol ; e Direct comparison of spectral-domain and swept-source OCT in the measurement of choroidal thickness in normal eyes.
In the lower left inset of Fig. OCT axial resolution depends on the spectral bandwidth of the light source and on the center wavelength. The exemplary plots of identical axial resolution in the eye show that the bandwidth needs to be increased for longer center wavelengths to maintain the same resolution. As indicated by the dotted curve of the absorption coefficient of water [ 12 ], not all wavelengths are equally suitable.
For greater wavelengths, the eye is considerably less transparent. Figure 3. It is evident that, for a longer wavelength, the bandwidth of the light source needs to be increased to achieve the same axial resolution. The water absorption curve dashed red line shows that absorption is increased for nm compared to nm. The spectral width of the absorption dip limits the maximum achievable resolution, e. For SS-OCT it is given by the number of readouts of the photo diode during one sweep of the light source.
The maximum imaging range, divided by 0. This number characterizes how many micrometer per pixels are imaged and provide the axial scaling of the scan.
It is often mistaken with the axial resolution, which defines the minimal distance of structures, which can still be distinguished in the OCT-B-Scan. Left: Lateral image parameters of retinal OCT depend on the focusing of the probing beam by the human eye. Right: Schematic of the sampling of an OCT volume.
Outside the focal volume the intensity coming back from the sample is reduced considerably. Therefore a compromise of focal depth and lateral resolution needs to be found with the optical design of the OCT system.
As OCT measures optical delays, all axial distances are optical distances. To achieve scaling in geometrical distances to allow for instance thickness measurements, the refractive index n of the medium needs to be known and axial distances measured in OCT scans are divided by the refractive index n.
To record a 3D data set, the sample beam is stepped in the second lateral direction after each B-Scan, as shown in the right part of Fig.
The recorded B-Scan series is stacked together. From this volume, a transversal image can be calculated, referred to as enface OCT image. The step width of the scanner defines the lateral sampling in both directions. Usually a B-Scan is sampled more densely than the slow direction y of a volume.
The OCT A-scan presents a profile of backscattered light intensity over tissue depth. The height of a signal compared to the image noise floor is called signal to noise ratio SNR. The SNR is different for each individual structure, because the signal strength is determined by the backscattering properties, often referred to as reflectivity. Backscattering originates from local changes in refractive index within the tissue due to alterations in the microscopic structure or in the density of scattering particles.
The detection of reflectivity enables OCT to reveal the internal structure of an object and is particularly useful to visualize its layer architecture. However, without elaborate modelling, the OCT signal does not provide an absolute quantitative measure of local reflectivity.
Due to absorption and scattering in the upper layers less light will reach the lower layers and backscattered light from lower layers is attenuated on its return path again.
Sensitivity has been established as a useful figure of merit to characterize or compare the performance of an OCT system. It is defined by the minimum sample reflectance the system can detect by achieving a SNR of 1. An OCT signal which is generated by specular reflection of an ideal mirror i. SNR and sensitivity are commonly specified in units of power decibel dB denoting a logarithmic scaling of the OCT power values.
Single A-scans sometimes can be affected by specular reflection on the ILM or the center of the macula. The maximum signal level and the noise floor span a range of about 40 dB for healthy retinal tissue and clear media.
In a linear scale, the OCT power values exceed the limited number of distinct grey values of common display devices and the perception of the human eye. Therefore, the power signal needs to be mapped to grey scale in a meaningful way. Usually, a logarithmic transformation or a comparable mathematical operation is first applied to the data, compressing the distribution of power values to approach a more Gaussian-like shape.
The resulting data is then mapped to 8 bit grey values. The mapping can be further adapted by applying different curves for gamma-correction. This allows to variably assign a range of signal power levels within an OCT B-Scan to a range of grey values and thus increase the contrast for distinct regions of interest. This means that compared to the inherent and unavoidable characteristic noise of photons, other noise sources can be neglected.
Therefore, it depends linearly on the incident optical power, the efficiency of photon detection and the sensor integration time. Consequently, there is a principal tradeoff between acquisition speed and system sensitivity. Every FD-OCT system has a characteristic decrease in sensitivity with imaging depth, also called roll-off.
It is related to the finite spectral resolution of the system component providing spectral separation. As shown in Fig. Two main contributions are therefore responsible for the characteristic decrease in sensitivity: the finite pixel size of the line detector and the finite spot size created by the spectrometer optics.
Its spectral resolution is determined by the instantaneous line width of the swept laser source and may be impacted by the bandwidth of the analog-to-digital conversion. SD-OCT is assumed to have a pronounced roll-off. The interferometric principle of OCT gives rise to a granular intensity pattern called speckle, which inherently exists due to the coherent detection scheme of OCT. Within the coherence volume or resolution element which is given essentially by the optical lateral and axial resolution of the system, mutual interference from multiple scattering events can occur.
As a result, the OCT signal from a single resolution element can vary to a large amount and is sensitive to variations in scan geometry. Homogenously scattering tissue manifests in a speckle pattern with a typical speckle size corresponding to the size of the resolution element and the spatial average brightness reflecting the backscattering properties of the tissue.
Structural OCT images suffer from speckle noise because it might obscure small image features or hamper the recognition of layer boundaries. A common way to reduce speckle and thereby improving the visibility of structures is achieved by signal averaging. The intrinsic variation in scan geometry together with patient movement serves the purpose to induce the necessary variation in the speckle pattern.
Averaging not only reduces the speckle noise but also reduces fluctuations in background noise. The SNR constantly increases with the square root of the number of acquisitions. However, changes in speckle pattern reflect changes in the distribution of scattering particles within the resolution element. This is used to distinguish steady tissue from moving particles for blood flow imaging see Chap.
Some selected applications are presented in Chap. OCT is usually combined with IR confocal imaging, though other combinations are possible as well.
Confocal imaging creates a transversal image of the retina corresponding to the en-face plane of OCT. Live images are presented throughout the imaging procedure to control image acquisition and quality.
Based on laser safety guidelines, the optical output power is limited to 1. Essentially, the scan angle determines the field of view FOV of the imaging area on the retina, the diameter of the scan pupil aperture defines the diffraction limited optical lateral resolution.
The OCT scanning unit is comprised of two linear scanners, which are driven synchronously with the read-out of the line scan camera in the spectrometer. The OCT frame rate is therefore determined by the scan density i.
The OCT2 module supports a line rate of 85 kHz, resulting in a frame rate of about Hz for the fastest scan pattern. As discussed in the technical section, the spectral resolution of the spectrometer determines the characteristic roll-off in sensitivity with imaging depth. There is a tradeoff between acquisition speed and sensitivity: The higher the line rate, the faster the image acquisition but the less that photons can be detected. Acquisition speed therefore is inherently coupled to the sensitivity of the system.
For retinal imaging, the maximum laser power is set by the exposure limit according to the laser safety guidelines. Therefore, to compensate for shorter integration time, the power can only be increased up to this limit. At the same time, eye motion, heart beat and any motion in general requires accelerated acquisition. Some eye motion occurs at frequencies faster than the OCT frame rate and requires software algorithms to ensure precise and reliable positioning of the OCT scan pattern.
As a result, the OCT image is precisely aligned even in cases with eye movement during image acquisition. In addition, the co-registration of OCT and cSLO images allows for follow-up examinations at exactly the same position and at any later point in time. The algorithm to combine multiple images which have been captured in the same location is called ART mean automatic real time mean. While ART is active, the SNR of the image is continuously increasing with approximately the square root of the number of averaged single B-Scans, to a maximum selected by the user.
As a result, faint signals elevate from the noise floor and the contrast between single retinal layers is increased. Follow-up series for treatment control of wet AMD: two follow-up images 2, 3 are co-registered to the baseline image 1.
The exact same scan position allows for identifying changes. The red lines indicate identical scan locations. Without APS, the influence of head tilt and eye rotation can impede the sectorial analysis of RNFL thickness and assessment of progression c, lower row.
Without APS, differences in patient alignment may impede the sectorial analysis of RNFL thickness and thereby impact the assessment of progression, which is presented in Fig. The segmentation of retinal layers is a basic prerequisite for many subsequent visualization and analysis features, such as the display of retinal thickness profiles or the definition and visualization of retinal slabs between any retinal boundaries. Multi-layer segmentation of a retinal OCT scan with the naming convention used throughout the software.
If the retina is affected by pathology or by poor image quality, the automatic segmentation may fail. The segmentation editor tool therefore supports manual segmentation of some scans and corrects the rest of the volume dataset accordingly. Enface OCT images can be calculated from OCT volume scans, which have been segmented accordingly: For each slab—vitreoretinal, RPE and choroid—maximum intensity projection along the depth direction is used to generate transversal images.
The contrast of choroidal vascular detail and the visibility of the choroidal-scleral interface CSI may be important in assessment of choroidal pathologies, e. Imaging of the lamina cribrosa benefits from the setting as well. For EDI, the characteristic roll-off is reversed in depth. The optimum imaging position, also called sweet spot, is moved to the lower part of the displayed OCT image. Technically, the EDI mode is realized by shifting the position of the reference mirror. Deeper layers then have smaller differences in optical path length and are therefore encoded in interference fringes of lower spatial frequency: their OCT signal gains an additional SNR of 2—3 dB as it is not affected by the roll-off anymore.
However, EDI cannot account for the losses induced by scattering, which may be enhanced for several pathologies and affects all layers below. An emerging area of interest is widefield OCT. While widefield technology in other imaging modalities, such as angiography and autofluorescence is already widely used, the application of widefield OCT is still currently being adopted into clinical practice. Widefield OCT imaging may provide significant benefit in the visualization of multifocal macular disorders or in the understanding of peripheral vitreoretinal diseases.
Nerve fiber layer thickness analysis: Three peripapillary circular sans are placed at the optic nerve head with a fixed starting point relative to the macula position top left inset and in each circle scan the RNFL and ILM are segmented top right inset. Standardized measurements include thickness in predefined segments bottom left and comparison of the thickness with a normative database bottom right inset.
The black line indicates the measurement of the individual patient in comparison with the average thickness for this age and population green line and in comparison with margins of normative data base green—normal, yellow—borderline and red—out of normal range. Standardized measurements include thickness in predefined segments bottom left and the BMO minimal rim width according to the previously found landmarks in the OCT B-Scans bottom right inset.
The black line indicates the measurement of the individual patient in comparison to margins of normative data base green—normal, yellow—borderline and red—out of normal range. Segmentation of the ganglion cell layer GCL and resulting color-coded thickness maps. Since more than a decade structural OCT measurements have been used very successfully in clinical routine for diagnosis of retinal and neurodegenerative diseases see Chaps.
OCT technology was also established for measuring and assessing structural parameters within the eye bulbus e. Furthermore, during the last years research work investigating additional or complementary contrast mechanism based on OCT technology has been published continuously.
In the following section, a short overview of the most important contrast mechanisms is given. OCT signal originating from blood vessels shows a much larger variance compared to the OCT signal of stationary tissue. The signal alterations in the vessels are due to the flow of the back-scattering particles mainly erythrocytes. Increased variance is observed for both, for the intensity and for the phase of the complex valued OCT signal and is used to compute OCTA images.
For OCTA images B-Scans at the same position are repetitively acquired, and sophisticated mathematical and statistical algorithms were developed to discriminate vascular structures from stationary tissue based on the variance of the OCT signal. These algorithms face several challenges: The presence of fast eye movements bulk motion causes a signal variance also for stationary tissue, which needs to be separated from the variance caused by the retinal blood flow.
Also the blood flow in larger vessels within the inner retina can cause so-called projection artefacts in the deeper vascular plexus. See also reference [ 15 ] for an overview and Chap. Whereas in OCTA the blood flow and thus the geometry of the different vascular plexus are visualized, no quantification of the blood flow e. For such a quantitative assessment of the ocular blood flow, also the knowledge of the velocity profile within the vessel is required.
If the blood flow contains a velocity component in z-direction—as it is the case for the large vessels at the rim of the optical nerve head—this velocity component can be extracted from OCT phase measurements of consecutive A-scans.
0コメント