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Record and pay accounts payable. Issue invoices to and collect from customers. View all Transbio Incorporated jobs - Manila jobs. The minor tick on the scale is 1 mm. Fluorescence excitation light blue passes through a filter and dichroic mirror, and is coupled into the doublet microendoscope probe using a conventional microscope objective.

The probe delivers the excitation light to the specimen, and fluorescence emissions green return back through the probe, the microscope objective lens, and an emission filter. The fluorescence image can be viewed through eyepieces or projected onto a camera. White arrowheads point to individual red blood cells. An ultra-short pulsed laser beam is focused just external to the relay lens of the microendoscope probe. The focal spot is raster scanned in two lateral dimensions, and the probe projects and demagnifies the scan pattern to the specimen plane.

Two-photon excited fluorescence originates from the laser focus in the specimen, returns through the endoscope probe, is separated from the laser light, and is detected by a photomultiplier tube PMT. A computer controls the scanning process and constructs the digital image. Cylindrically shaped GRIN lenses act like optical fibers in the sense that they transmit light between distal locations using total internal reflection. Although this modality offers fast frame-rate acquisition, it lacks optical sectioning.

Confocal FME that combines micro-lenses with a fiber bundle provides a flexible hand-held probe, but in doing so sacrifices focusing ability [ 7 , 13 ]. Although FME remains in infancy, it seems well poised to achieve widespread application because of its ability to provide high-resolution in vivo imaging of deep brain areas. Further studies must address this issue carefully, and optimal insertion methods, perhaps involving cannulation, need to be developed.

Other likely advances include improved resolution for visualizing fine cellular structures, development of chronic imaging preparations, and creation of FME devices for use in freely moving animals. Several fiber optic modalities are being developed to exploit the intrinsic optical properties of brain tissue and optical correlates of brain activity that have traditionally been studied without fibers Table 1.

Metabolic responses to brain activity that can be detected optically include changes in blood flow [ 19 , 20 ], hemoglobin oxygenation [ 20 — 22 ], oxygen tension [ 23 ], and redox states of intracellular NADH [ 24 , 25 ] and flavoproteins [ 26 , 27 ]. Such effects generally occur within 0. These faster phenomena alter light scattering by 0. Fiber optics are being harnessed to examine both slow and fast signals in vivo.

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Researchers are increasingly performing near infrared absorption spectroscopy NIRS studies of brain activation using assemblies of fiber optic emitters and detectors, typically spaced about 5—40 mm apart on the head, to deliver and collect light through the intact skull Figure 2a,d ; [ 34 — 40 ]. Because scattering lengths of near infrared light in the brain are at most hundreds of microns but absorption lengths are in the millimeter to centimeter range, photons undergo multiple scattering events and migrate diffusively between emitters and detectors.

This enables diffuse optical imaging DOI , in which changes in optical fluence between pairs of emitter and detector fibers Figure 2 are usually ascribed to changes in light absorption within tissue.

The inverse problem of taking these measured pairwise changes in fluence and working backwards to reconstruct the spatiotemporal images of absorption changes within tissue is usually mathematically underdetermined and complicated by cardiac and respiratory artifacts. Ongoing improvement of heuristics for image reconstruction and for artifact removal continues to be key to the development of DOI. Each source fiber emitted light from a pair of laser diodes with wavelengths of and nm. Each of the 18 laser diodes was intensity modulated in a square wave at a unique frequency between 4 and 8 kHz, allowing the signals from distinct light sources to be distinguished at each detector.

The pattern of activation, as assessed using the change in absorbance at nm averaged over 32 stimulus presentations, is superimposed on the fiber array and the image of the rat, with warmer colors indicating greater magnitude of activation. Although there is good agreement regarding the lateral separation of the two identified regions, the activated regions found by DOT are larger than those found by fMRI and also extend into the scalp and skull.

The latter effect is a common artifact with DOT that occurs when image reconstruction is performed without mathematically constraining absorbance changes to the brain. Areas of activation are then biased somewhat towards the scalp. Each curve is an average of 27 trials. Most DOI studies of the brain examine slow absorption changes arising from activity-dependent variations in blood flow and hemoglobin oxygenation. However, use of two or more wavelengths for NIRS permits changes in total blood content and in both oxy- and deoxy-hemoglobin concentrations to be distinguished, which conventional fMRI does not allow.

Although diffuse photon migration limits the penetration depth and spatial resolution of DOI to the centimeter-scale in humans, such non-invasive measurements provide a low-cost means of identifying regions of brain activation [ 46 ], for example with subjects who cannot undergo MRI [ 35 , 37 , 47 ]. Commercial DOI systems for use in humans are already available.

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A key facet of DOT is that each detector should receive light from emitters beyond its nearest neighbors. In small animals, in which the distance from scalp to cortex is short, close packing of fibers yields superior resolution than can be obtained in adult humans, for which the skull thickness is a key limitation. Thus, a broadly applicable advance that appears within reach is development of DOT for use in freely behaving animals.

Other advances might come from improved analysis of blood chromophore concentrations [ 56 , 57 ], optical assemblies that can incorporate large numbers of precisely oriented channels using a non-contact mode displaced from tissue [ 54 , 55 ], and improved algorithms for reducing physiological noise artifacts [ 38 , 58 ]. Eventually, dense arrays of integrated semiconductor emitters and photodetectors [ 59 ] might obviate the need for optical fibers and significantly enhance resolution.

Optical coherence tomography OCT uses light that is backscattered off tissue to provide image contrast. Because the slight differences in travel times from different points in tissue are too short to measure directly, low coherence interferometry LCI is used to compare the optical path lengths of light from the sample and the interferometer reference arm Figure 3a ; [ 60 , 61 ].

Although diffraction determines the lateral resolution of OCT, the coherence length of the optical source, which is inversely related to the spectral bandwidth, determines the axial resolution of optical sectioning. Although it is not strictly necessary to implement OCT within fiber optic hardware [ 64 ], this is most commonly done because of the benefits regarding instrument portability, ability to fabricate miniaturized OCT probes [ 65 — 68 ], and availability of superior fiber optic components for telecom wavelengths, such as and nm.

Ultrahigh-resolution in vivo optical coherence tomography OCT of the human retina. The main optical paths are shown in red, electronic components in green. The laser light is coupled into a fiber optic Michelson interferometer in which the sample arm comprises the ophthalmologic system.

In the reference arm the light is coupled from optical fiber into free space, reflects off a translating reference mirror, and returns into the optical fiber. The position of the translating mirror sets the optical path length of the reference arm. Light in the sample arm that travels an equivalent optical path length, to within a coherence length, recombines interferometrically with light in the reference arm.

The amplitude of this interferometric signal represents the contrast parameter in conventional OCT. Scanning the position of the reference mirror varies the depth of the optical section acquired within retinal tissue, and the scanning speed sets the Doppler frequency shift of light in the reference arm. This Doppler frequency becomes the heterodyne detection frequency from which the interferometric amplitude is extracted.

A fraction of the light focused onto the retina backscatters and re-enters the interferometer. To optimize the interferometric signal, fiber optic polarization controllers in both the sample and reference arms are adjusted to match the polarization of light within the two arms. Chromatic dispersion must also be matched, or else optical sectioning degrades.

Chromatic dispersion is fine-tuned using two translatable prism pairs, one of fused silica FS , the other BK7 high-index glass BK7 , which allow the thickness of glass in the reference arm to be varied. This detection approach cancels amplitude fluctuations common to the two arms, such as laser power noise, and increases the detection sensitivity to nearly the physical limit set by photon shot noise.

An amplified version of the interferometer signal is band-pass filtered around the Doppler frequency of the reference arm and then demodulated. A computer reconstructs the retinal tomogram. Because the dynamic range of backscattering signals from tissue typically extends over several orders of magnitude, OCT images are usually represented on a logarithmic scale.

Panel c shows a comparison between conventional and spectroscopic OCT retinal imaging. The mean wavelength of backscattered light is encoded by the color hue, with shorter wavelengths assigned to greener hues and longer wavelengths to redder hues, as indicated by the color bar.

Results and Discussion

The intensity of backscattered light is encoded by the color saturation. A vein shown in green and an artery shown in red that are not easily identifiable by conventional OCT are apparent in the spectroscopic OCT image. The blood vessels attenuate light and so the underlying areas appear darkened in ci and cii. This illustrates the confound between attenuation and backscattering that can arise in OCT image interpretation.

The two vessels backscatter light of different spectra, in part because their opposing directions of blood flow induce Doppler frequency shifts of opposite sign. The distinct absorption spectra of oxy- and deoxy-hemoglobin across the laser emission band are additional influences. This difference helps the artery and vein to be distinguished on the basis of the wavelength dependence of signal attenuation in the respective underlying tissues.