Because various types of OCT systems are discussed in detail elsewhere in this handbook, they will not be discussed here. In general, LIF capability can be added to any type of OCT system, although the sharing of components is facilitated if the scan types are common (e.g., both one, two, or three dimensional scanning and full field). An appropriate method of signal separation (spectral, temporal, or spatial) is required.

Steady-state LIF measurements require a light source with a well-defined output spectrum, a conduit to transport the excitation light to the sample, illumination and collection optics, and a conduit to transport the emitted light back to a spectral analyzer for optical detection. Instrumentation for fluorescence measurements is well described by Lakowicz [39], and here we focus on spectral compatibility and common in vivo measurement limitations.

One of the major challenges in fluorescence detection is the signal to background ratio. Background signals are caused by instrument autofluorescence (see Sect. 50.5.4), ccc from the light source, and scattering within the spectral analyzer. Because endogenous signals can be weak, background signals need to be avoided. Out-of-band illumination suppression should reach 10⁻⁵ or 10⁻⁶ for endogenous tissue fluorophore measurements, especially in excitation ranges where the fluorescence efficiency is low (blue to green). Because the amount of collected instrument background is also dependent on the sample reflectivity, traditional subtraction of the sample signal by the instrument response to a negative standard (e.g., a measurement in water) has its limitations. As a rule of thumb, instrument background should be kept below 10–30 % of the sample signal. Traditionally, illumination and collection paths have been constructed separately [126] to reduce system autofluorescence. Irradiance levels on excitation optics are significantly higher than on collection optics making it necessary to incorporate autofluorescence reduction techniques in the illumination conduit and optical elements at the endoscope tip. In typical applications, the collection conduit will receive a significant amount of excitation light backscattered by the sample. If the stray light suppression capabilities of the detection system are exceeded (a spectrograph might have 10⁻⁴ suppression), the backscattered excitation light can also contribute to the signal background. This can be avoided by dampening excitation light in the detection path with long-pass filters before the emitted light is spectrally analyzed.

For LIF, short-arc lamps have been used as a standard low-cost excitation source when coupled to an excitation monochromator or dielectric band-pass filters. Example radiance data for mercury–xenon and xenon lamps are shown in Fig. 50.2, top panel. Their broad emission range allows wavelength tunability, but a common disadvantage of short-arc lamps is the limited ability to focus light into small optical fibers. A large variety of laser modules has been developed for confocal microscopy, and most of these lasers are compact and sufficiently transportable for endoscopic fluorescence measurements. Example sources are shown in Fig. 50.2 and abbreviated as XeCl, xenon chloride (308 nm emission); HeCd, helium–cadmium (325 nm); N2, nitrogen (337 nm); Ar, argon (351, 364, 458, 488, and 514 nm); YAG, neodymium YAG (1,064, 532 nm); and HeNe, helium–neon (543, 633 nm). Because of the small Stoke’s shift of most fluorophores, narrowband sources are preferred to enable efficient collection efficiency of emission light and rejection of excitation light. In contrast, OCT sources are by necessity broadband, with spectral widths of tens to hundreds of nanometers. Also, OCT sources are generally centered in the NIR, to enable greatest depth of imaging in highly scattering tissue. The light source emission spectra for some OCT sources (short-pulsed lasers, swept-wavelength lasers, and superluminescent diodes) are shown in Fig. 50.2. Also shown in Fig. 50.2 is the extinction spectrum of three tuned nanoshells (exogenous scattering agents) [127].

Epithelial tissues usually exhibit a strong fluorescence peak at 280 nm excitation and 350 nm emission, which is associated with tryptophan fluorescence (Fig. 50.2, bottom panel, cell emission at 280 nm excitation). Cells excited at UV-A wavelengths show strong fluorescence from NADH, which has a maximum at 350 nm excitation and 450 nm emission (Fig. 50.2, cell emission at 350 nm excitation), whereas the oxidized form NAD is not fluorescent [39]. Conversely, FAD fluoresces at a maximum at 450 nm excitation (blue) and 535 nm emission (Fig. 50.2, cell emission at 450 nm excitation), whereas the reduced form FADH₂ fluoresces minimally. Lipofuscin has a very broad excitation and emission spectra. The excitation spectrum is notable in particular because its peak at 470 nm is comfortably above the strong excitation bands of many other materials. Fluorescence from interstitial matrix is dominated by collagen; average fluorescence emission from polymerized collagen that was reconstituted from rat tail type I fragments is shown in Fig. 50.2 middle and bottom panel. Typical emission at 280 nm excitation/300 nm emission (consistent with tyrosine), 320 nm excitation/400 nm emission (consistent with the enzymatic cross-links lysyl hydroxypyridoxyline and hydroxylysyl hydroxypyridoxyline), and 370 nm excitation/450 nm emission (consistent with the nonenzymatic cross-links vesperlysine and crossline) can be observed. Unfortunately, at many excitation wavelengths, these collagen emission spectra overlap with the emission observed from cell suspensions. Porphyrins show an excitation maximum at 410 nm and two emission maxima at 630 and 690 nm, while the peak at 630 nm is substantially more intense (Fig. 50.2, protoporphyrin IX). Emission spectra from exogenous fluorophores fluorescein; cyanine dyes Cy3, Cy7, and ICG; and CdSe/InAs quantum dots are also shown in Fig. 50.2. Some of these agents have NIR excitation and/or emission.

Figure 50.2 illustrates that endogenous fluorescence, targeted fluorescence imaging, and optical coherence tomography can be spectrally incorporated because most endogenous components do not absorb or emit within the OCT illumination spectral range, and some exogenous agents are available outside that range. If spectral overlap occurs, LIF and OCT need to be time or spatially multiplexed. Spectral overlap has been used to advantage in one application [128], with simultaneous OCT/ICG imaging in the eye’s fundus. In this case, the same superluminescent diode (793 nm center wavelength, 22 nm spectral full width at half maximum) was used for OCT imaging and for excitation of ICG. The key in this implementation was proper selection of the dichroic filter separating backscattered source light (directed to the OCT detector) from fluorescence emission (directed to the LIF detector). This selection minimized distortion of the source correlation function and maximized collection efficiency of fluorescence emission. Excellent simultaneous OCT and ICG images were obtained. However, use of the same source for OCT and fluorescence excitation could become increasingly difficult as either the source bandwidth increases or fluorophore emission intensity decreases, unless very large Stoke’s shift NIR-excitable agents become available.

A strategy to provide greater spectral separation has been described [129]. This custom system included a mode-locked Ti:sapphire laser and a 0.9 numerical aperture water immersion objective to generate simultaneous en face optical coherence and two-photon-excited fluorescence images. An image of a green fluorescent protein expressing drosophila embryo was presented; presumably endogenous fluorophore distribution could be imaged as well. Because fluorescence occurred at a much shorter wavelength range, no difficulties were encountered separating the fluorescence from the backscattered light. However, implementation of this method in an endoscopic fashion would be technically challenging because of high peak-power needs, three-dimensional scanning requirements, and the need to control probe–tissue separation.

As described above, many interesting endogenous fluorophores are best excited at ultraviolet wavelengths that also cause autofluorescence in most optical materials. Many glasses and polymers exhibit autofluorescence that is similar in spectral shape and range to the fluorescence of endogenous tissue fluorophores [130]. Because probe autofluorescence can easily become significant compared to the tissue signal, each component of the OCT–LIF system should be considered for autofluorescence, particularly if the excitation wavelength lies in the UV. The location of the element within the system is equally important as the absorption and emission characteristics of the material. Autofluorescence from an optically thick piece of fiber is generally more significant than from the same material used in a thin window. Additionally, isotropically emitted fluorescence is more easily collected from an element located near focus than from a lens or window located far from focus or in a low-numerical-aperture portion of the beam path. Lenses and windows made of fused silica and CaF₂ have low fluorescence under UV-A excitation and good NIR transmission and are recommended throughout the beam path. Thick windows near the tissue should be restricted to lowest fluorescence glasses, although very thin windows of other glasses and even some low-fluorescence plastic films may be acceptable [131]. Usually, some fluorescent materials must be used despite the best attempts to avoid them. For example, the emission from even “low-fluorescence” ultraviolet transmitting optical epoxies is significant, and care should be taken to keep glue joints thin [132]. The jacketing materials available on many commercially available communications grade fibers ideal for OCT fluoresce strongly. Although fiber choice could be restricted only to low-fluorescence jacketed materials, it is often more practical to simply shield the fiber from UV-A excitation. This approach of hiding the fluorescent material behind absorbent material, a dichroic filter, or simply out of the beam path is the most flexible and commonly used. Care should be taken with items outside the device that may also autofluoresce, including tissue index matching gels, preservation media, culture plates, or wax mounting blocks.

The tip of a combined OCT–LIF endoscope is either in contact with or at a close distance to a tissue surface or body fluid and therefore should be analyzed for potential hazards. A thorough analysis of potential risks and protection against those risks are a requirement for human subject studies. Hazards include electrical shock hazards, clinical hazards, material toxicity hazards, and radiation hazards. Of unique importance to combined OCT–LIF systems is the fact that both systems may be operated concurrently, requiring cumulative radiation exposure analyses. Radiation exposure in the UV poses a different hazard than in the visible/NIR. UV exposure is a cumulative hazard due to the potentially ionizing energy of each photon (exposure limits can be calculated for repeated exposure over days), whereas visible and NIR light exposure is analyzed for potential thermal injuries. “Blue-light hazard” is unique to eye exposure, caused by free radical release due to interaction with visual photopigments [133]. Threshold limit values for radiation exposure are defined by several standardization organizations: The American National Standards Institute (ANSI) publishes maximal permissible exposure levels [134]; the American Conference of Governmental Industrial Hygienists (ACGIH) publishes similar threshold limit values (TLV) and biological exposure indices [135]. Applications which intentionally introduce light into the eye should verify that standards set for ophthalmic instruments ISO 15004-2 are met in addition to laser safety standards. In contrast to the above-listed standardization organizations, the International Commission on Non-Ionizing Radiation Protection (Oberschleissheim, Germany) makes their guidelines available online at no cost and publishes them in Health Physics. Usually the threshold limit values are expressed for laser beam [136, 137] and incoherent exposure [138, 139] of the eye and skin. For broadband UV exposure (<400 nm), the biologically effective radiation (device emission weighted by the biologic action spectrum, which is normalized at 270 nm) should not exceed the TLV for melano-compromised skin and eye (3 mJ/cm²) [138]. The TLV for the skin and eye has also been recommended for the cervix in a guidance document developed by the US Food and Drug Administration’s Center for Device and Radiological Health (CDRH) [140]. For the wavelength range of 0.38–1.4 μm, exposure limits are governed by thermal injury to the skin and eye. Unfortunately, the authors are not aware of any studies evaluating combinational effects of NIR radiation and UV exposure.

The optical design of a combined system is a result of an attempt to optimally control the photon paths in the tissue for each modality, while meeting material constraints in the optical path and separating the signals (chromatically, spatially, or temporally) to the appropriate detector. The useful signal in OCT depends on photons that undergo a single scattering event in a nearly backscattering direction. The confocal design of typical OCT systems helps reject photons that take any other path. Lateral resolution is dependent upon the beam cross section at a given depth. Usually, a balance between lateral resolution and depth of focus is struck in order to avoid axially scanning the objective [141]. Conversely, LIF photons are typically allowed to take a circuitous path: the excitation light may be scattered multiple times before it is absorbed. If fluorescence conversion occurs, the emission spectrum of the intrinsic fluorescence will be modified as the emitted photons wander randomly through the tissue until they are absorbed by chromophores or released at the tissue surface. The path taken, and volume subsequently probed by the LIF system, depends on the tissue absorption and scattering properties as well as factors determined by the probe geometry: the insertion angles of excitation photons, the distance along the surface the collection to point (source-detector separation distance – SDSD), and the angles over which the photons are collected. Fiber-optic probes have been proposed for preferential depth-sensitive fluorescence sensing with the use of variable illumination and collection apertures [142, 143], illumination collection distances [143–146], and illumination collection angles [147–149].

Monte Carlo analysis of specific fiber-optic probe implementations on media with tissue-like optical properties and their associated experimental verifications show trends that are useful in designing an LIF probe. First, it is critical to control the distance between the illumination/collection fibers and the tissue for quantitative tissue analysis [145, 150]. When the illumination and collection areas are superimposed at the tissue surface (SDSD = 0), either by means of single illumination–collection fiber in contact with the tissue, or a spacer distal to a group of fibers that allows the illumination–collection areas to diverge over the same tissue area, the probe will be primarily sensitive to superficial layers less than 400 μm deep [142, 144, 145]. Similarly shallow tissue sampling can be achieved when source and collection fibers are separated at the tissue surface but are inclined towards each other to cause an overlap of the fiber numerical apertures in the superficial tissue [148, 149, 151]. The average probed depth can be increased from 700 to 1,200 μm by placing normally incident illumination and collection fibers that are directly in contact with the tissue at increased separation [142, 144, 145]. Evidence with diffuse reflectance modeling suggests that the probed depth can be increased yet further by inclining the fibers away from each other [152]. In general, greater overlap of the numerical apertures of the illumination and collection fibers improves collection efficiency while the expected probing depth decreases. The lateral extent of a collected LIF photon’s path has been relatively unexplored, although lateral resolution somewhat larger than the SDSD could be expected. Increasing the fiber core diameter and numerical aperture increases the light gathering capacity of the collection fiber but may have effects on the depth of collected fluorescence that depends upon the configuration. In cases with fibers in contact with the tissue, the size of the illumination and collection area introduces a superposition of many SDSDs, while the numerical aperture of the fiber introduces a superposition of many angles of insertion and collection. In addition to the depth selectivity provided by SDSD, some geometries appear to be optimally sensitive to fluorophore distribution rather than scattering properties of the tissue [153, 154]. Knowledge of tissue layer thicknesses from OCT images could help in the selection of optimal LIF illumination/detection fiber configurations.

In the case where the majority of information from LIF derives from direct emission from a fluorophore in a relatively superficial tissue layer, a confocal arrangement delivers significant advantages. Confocal arrangements are common in fluorescent microscopes because they have the potential to deliver high-resolution imaging with high contrast. High resolution derives from the tightly limited photon paths allowed by the confocal arrangement. Contrast is enhanced by spatial filtering of out-of-plane, unwanted fluorescence, including that from the optics of the instrument itself. In combination with point-scanning OCT, the two confocal modalities will share similar design needs and constraints.

Because OCT systems generally utilize single-mode fibers and LIF systems utilize multimode fiber (to optimize the collected fluorescence emission signal), the most straightforward OCT–LIF endoscope design utilizes separate fibers for each modality. Maintaining separate conduits for OCT and LIF minimizes the cross talk between the systems and helps maintain high signal-to-background ratio. The proximal coupling optics between the fibers and sources/detectors are also simplified. However, other fiber configurations are possible. Dual-clad fibers consist of a central core with two layers of cladding. The core can be made single mode, whereas the inner cladding has a large radius and high-numerical aperture (multimode). OCT light is channeled to the sample through the core. LIF excitation light can be carried by the core (if compatible) or the inner cladding. Remitted OCT, excitation, and emission light are coupled into the core and inner cladding. In addition to coupling light into the dual-clad fiber, the proximal optics must spatially separate returned OCT light which has propagated through the core from that which has taken unwanted paths through the inner cladding, and spectrally separate the OCT and fluorescence emission. Dual-clad fibers have successfully been used in OCT–LIF systems [155–158], and one system is described in more detail below.