Two-Photon Scanning Laser Ophthalmoscope
Two-photon excitation (TPE) fluorescence imaging is an important tool in biological sciences which provides high penetration depth, inherent three-dimensional (3-D) sectioning and good detection sensitivity at very low molecular concentrations. Due to the high transparency of the preretinal tissue, TPE fluorescence imaging can be used to image single neurons optically and non-invasively. The combination of scanning laser ophthalmoscopy (SLO) which provides reflectance images of the retina with TPE fluorescence imaging represents an imaging modality well suited for non-invasive retinal imaging. Here we present the implementation, application and future prospective of a novel compact two photon laser scanning ophthalmoscope capable of acquiring simultaneous confocal reflectance and two-photon images.
KeywordsTwo-photon imaging Confocal reflectance Fluorescence lifetime imaging Ganglion cells In-vivo imaging Non-invasive imaging Hollow core fiber Femtosecond laser Retinal imaging
Two-photon excitation (TPE) fluorescence imaging is a powerful emerging tool in biomedical applications, providing high penetration depth and inherent three-dimensional (3-D) sectioning at the subcellular level [1, 2, 3]. The retina is the only tissue in which single neurons can be imaged optically and noninvasively due to the high transparency of the preretinal tissues [4, 5]. TPE with infrared light (IR) is particularly well suited for in vivo retinal imaging. Reporter molecules in cell bodies can be excited with IR light to allow differential activation of rod and cone photoreceptors by wavelengths in the visually sensitive range to evoke responses in the retina [6, 7, 8, 9]. An additional advantage of TPE is reduced phototoxicity . TPE fluorescence imaging enables the study of functional physiological processes, which, in combination with in vivo ophthalmoscopy, represent powerful imaging techniques that are well suited for noninvasive in vivo retinal imaging. Applications include longitudinal tracking of disease progression, for example, in optic neuropathies in which retinal ganglion cells (RGCs ), the output neurons from the eye to the brain, are lost.
9.1.1 Retinal Signaling
9.1.2 Imaging Retinal Neurons
Exogenous fluorophores commonly used as reporters to image retinal neurons
Single photon excitation (nm)
Two photon excitation (nm)
Small molecule dyes
Fura-2 (Ca2+ bound/Ca2+ unbound)
Studies have developed TPE imaging systems to examine functional responses of RGCs in vivo [9, 21]. In TPE imaging, pulses of long wavelength light (typically in the IR range), approximately double the single photon excitation wavelength (Table 9.1) , are used to create conditions where two photons simultaneously stimulate the fluorescent reporter molecule to create energy similar to that of a single short wavelength photon (detailed theory in Sect. 9.2). Yin et al.  used adaptive optics (AO, described in Chap. 16) in conjunction with TPE to image RGC activity, reported by GCaMP5, in non-human primate retina. Similarly, but without the use of AO, Bar-Noam et al.  measured in vivo changes in GCaMP6 fluorescence in RGCs, in response to stimulation of the visual pathway. This study demonstrated in vivo calcium transients in RGCs similar to those reported previously in ex vivo retinal preparations . Together, these studies provide a foundation to further develop TPE as an imaging platform to study RGC function in vivo; however, clinical translation requires further refinement.
9.1.3 Imaging Other Retinal Cell Types In Vivo
Endogenous fluorophores expressed in retinal cells
Single photon excitation (nm)
Two photon excitationa (nm)
Fluorescence lifetime (ns)
NAD(P)H (free/protein bound)
FAD (free/protein bound)
9.2 Theoretical Background
9.2.1 Luminescence, SPA and TPA
Electronic excitation of molecules can be performed either by a physical (absorption of light), chemical or mechanical process. Luminescence describes the effect of a molecule emitting light after de-excitation into the ground state. If this excitation was created by the absorption of photons, then it is called photoluminescence. Photoluminescence of molecules can be divided into two groups, fluorescence and phosphorescence, differing in electronic configuration in the excited state and emission pathways .
Fluorescence is the ability of some atoms and molecules to absorb photons with a particular energy and to re-emit photons with reduced energy (red shifted) after a short time interval in the nanosecond time scale, referred to as fluorescence life time. Phosphorescence differs from fluorescence by the electronic transition pathway (intersystem crossing) into the excited triplet state resulting in a much longer excited state lifetime in the range of milliseconds to hundreds of seconds .
Molecules can transit from the ground state (lower energy) to the excited state (higher energy) by absorbing photons with an energy being at least equal to the energy difference between the excited and ground state. This electronic excitation can be achieved either by linear or nonlinear photon absorption. Linear excitation of the molecule is achieved by single photon absorption (SPA) whereas nonlinear absorption describes the case when two or more photons with less energy (as compared to the single photon excitation) combine to bridge the energy gap needed to excite the atom or molecule. Most widely used nonlinear excitation in biomedical research is two-photon absorption (TPA) , where two photons with half the energy of SPA combine for electronic excitation of the molecule .
Combination of confocal scanning laser microscopy (CSLM) with SPA or TPA allows one to generate two-dimensional fluorescence images of the specimen under investigation. Within the next sections, CSLM with SPA will be termed linear fluorescence imaging (LFI) and CSLM with TPA will be termed two-photon excitation fluorescence imaging (TPEFI).
9.2.2 TPA Probability and Dependencies
The probability for a two-photon absorption (TPA) process to occur, is dependent on the physical properties of the molecule (termed as TPA cross-section (σ2)) and the spatiotemporal properties of the excitation light. TPA requires the “simultaneous” arrival of two photons (time interval within 10−18 s) wherefore it has a quadratic dependency on the average incident light power (Pavg), making it a nonlinear process. Since the TPA cross sections are usually very low as compared to SPA cross sections [40, 41], temporal and spatial confinement is crucial for increasing TPA signal generation.
Temporal confinement is achieved by the use of pulsed laser sources (fp) with pulse durations (τp) below a few pico-seconds, resulting in high peak powers. The use of pulsed laser sources as compared to continous wave lasers enhances the signal by a factor of ~1/τpfp, which would lead to a signal enhancement of 105 by using commonly available femtosecond laser sources with pulse durations of 100 fs and pulse repetition rates of 100 MHz.
Spatial concentration is dependent on the excitation wavelength (λexc) and high numerical aperture (NA) objectives, leading to a small focal volume resulting in high peak intensities.
h is Planck’s constant
c is the speed of light
As can be seen from the formula above the laser pulse duration is inversely proportional to the two-photon signal generation, resulting in increased two-photon signal generation when reducing the laser pulse duration. In practice however, short pulses lead to higher dispersion which have to be compensated. Dispersion is a phenomenon in which the phase velocity of a wave is coupled to its frequency resulting in wavelength-dependent refractive indices in optical media. The temporal profile of the laser pulse is directly related to its spectral bandwidth and therefore shorter pulses are more vulnerable to dispersion effects leading to broadened pulses and less efficient two-photon excitation if not compensated [4, 43].
9.2.3 Optical Resolution
In linear fluorescence imaging, fluorescence photons are also generated above and below the focal plane, wherefore axial sectioning can be enhanced by spatially filtering the emitted fluorescence signal at the detection plane. Even though axial resolution is improved by reducing the pinhole diameter, fluorescence yield is also reduced, since fluorescence photons suffering from chromatic aberrations and strong scattering are blocked and cannot reach the detector . Axial and lateral confinement in TPEFI on the other hand is an intrinsic property of the nonlinear excitation process neglecting the need for spatially filtering the signal by a pinhole.
At first sight the theoretical lateral and axial resolution of TPEFI seems worse than the case of LFI due to the longer wavelength being used for excitation. In practice however, the spatial resolution of LFI and TPEFI that can be achieved are similar due to the use of finite-sized pinholes which broaden the theoretical PSF in LFI [22, 42].
9.2.4 Linear SPA vs. Nonlinear TPA Imaging
In nonlinear imaging, in particular TPEFI two excitation, photons combine their quantum energies and generate a photon with higher quantum energy which leads to a “bluer” emission as compared to the excitation. This is different from the red shifted emission occurring in LFI, and allows the use of excitation light in the near-infrared (NIR) wavelength range (700–1000 nm) for commonly used fluorescent markers emitting in the visible spectral range [32, 39].
The use of longer wavelengths in TPEFI allows the excitation light to penetrate deeper into scattering tissue, since longer wavelengths exhibit less scattering, and also phototoxic effects are reduced, since less one-photon endogenous absorbers are available at this wavelength [22, 44]. Another major advantage of nonlinear imaging is the nonlinear dependence of the signal intensity (S) to the excitation light intensity (I) which is quadratic for TPEFI (S ∝ In). This quadratic dependency allows the TPA to occur only at the focal volume and its close vicinity (spatially confined excitation) when focusing the laser beam through a microscope. As can be seen in Fig. 9.7b, no TPA fluorescence is created in planes above or below the focal volume which differs from SPA fluorescence, where fluorescence is created over the entire depth of the excitation light cone. The lack of out of focus TPA fluorescence provides advantages for long term in vivo imaging of biological tissue, since tissue viability is enhanced by reduced photo damage [22, 44, 45]. TPEFI further provides inherent three-dimensional sectioning capabilities without the need of spatially filtering the emitted light by the use of a confocal pinhole as it is the case for LFI.
Even in strongly scattering media this inherent sectioning capability is maintained, because the density of scattered excitation photons are usually too low for nonlinear signal generation which is important for deep imaging, since all the signal photons reaching the detector originated from the focal volume and its vicinity carrying useful information [2, 22, 42]. However, special care has to be taken by choosing the collection optics for optimizing the fluorescence for deep imaging, because of the increased spatio-angular space due to scattering.
9.3 Experimental Setup and Results
From  Reproduced under Creative Commons Attribution License (CC BY; https://creativecommons.org/licenses/by/3.0/legalcode ).
After dispersion compensation, the laser beam was coupled to a 2-m HCF (GLOphotonics, Limoges, France) with a 45-mm focal length achromatic lens (AC254-045-B, Thorlabs) with which a coupling efficiency of ~89% was achieved. The coupling lens was mounted on a 25-mm XYZ translation stage (PT3, Thorlabs) and the fiber on a three-axis microblock stage (MBT616D, Thorlabs).
The fiber output was coupled via FC connector to the fiber adapter plate that was mounted to the camera head of the modified Spectralis unit. The divergence of the output beam from the fiber was increased with a −6.0-mm focal length, biconcave, negative lens (LD2746-B, Thorlabs) to avoid the use of a longer focal length collimator before it was collimated with an achromatic doublet lens with a focal length of 25 mm (AC 127-025-B, Thorlabs). Two customized lens pairs of equal focal length, f = 20 mm (L4, L5 in Fig. 9.8), were integrated to enable finer focus adjustment in the axial plane. Horizontal and vertical beam scanning was performed with the standard Spectralis scan unit . In combination with a customized 50-mm focal length scan lens, an intermediate image field of 5 × 5 mm2 was produced. A customized 16-mm focal length tube lens translated the intermediate image field to a field of view of ~17.5° while achieving a beam size of ~2.2 mm (overfilling the dilated mouse pupil). Both reflectance and two-photon fluorescence signal were repassed through the scanning unit, resulting in a stationary, descanned light beam. A dichroic mirror (FF735-Di02, Semrock, Rochester, New York) was used to couple the fluorescence and reflectance signal into the detection branch, consisting of an 40-mm focal length achromatic lens and a 100-μm multimode fiber, which guides the signal light to the external detection unit. The fiber output was collimated with a 12-mm focal length lens and a second dichroic mirror (FF735-Di02, Semrock) separated the visible fluorescence light from the near infrared (NIR) reflectance light. The reflectance light was further attenuated with a neutral density (ND) filter and focused on an avalanche photodiode (RCA, New York City) with a 20-mm focal length customized achromatic lens. In the fluorescence signal path, a short-pass filter (FF01-720/SP-25, Semrock) was used to remove any leakage from the excitation light before being focused by a 100-mm focal length achromatic lens (47-972, Edmund Optics, Barrington, New Jersey) onto the photon counting detector (HPM-100-50, Becker-Hickl, Berlin, Germany) connected to a time-correlated single photon counting (TCSPC) module (SPC-150, Becker-Hickl). All imaging was performed with a horizontal line scan rate of 8 kHz and a pixel clock of 10 MHz. Confocal reflectance images were digitized at a resolution of 768 × 768 pixels with a frame rate of ~9 Hz. Fluorescence images were digitized at a 256 × 256 pixel resolution (by binning the corresponding signal to superpixels) and were averaged over 2–3 min. Customized software for real-time eye tracking (Heidelberg Engineering) was used for imaging. The confocal reflectance image served as a reference for the two-photon fluorescence signals, whose acquisition took place after correct positioning and localization. In this manner, each detected fluorescence photon could be assigned to the corresponding pixel from the reflectance image.
Two mouse strains, Thy1-YFP-16 [B6.Cg-Tg(Thy1-YFP)16Jrs/J; 6 months old male, 35–40 g, The Jackson Laboratory, Bar Harbor, Maine] and Thy1-GCaMP3 [B6.Cg-Tg(Thy1-GCaMP3)6Gfng/J; 6 months old male, 35–40 g, The Jackson Laboratory], were used for imaging. Mice were anesthetized with ketamine (100 mg/kg) and xylazine (10 mg/kg) by intraperitoneal injection. Pupils were dilated with one drop of 1% tropicamide (Mydriacyl, Alcon Laboratories, Mississauga, Ontario, Canada) and one drop of 2.5% phenylephrine (Mydfrin, Alcon Laboratories). After dilation, a 3.2-mm plano contact lens (Cantor and Nissel, Brackley, United Kingdom) was placed on the cornea to maintain corneal hydration and compensate for most of the nonspherical refractive errors arising from the corneal surface. During imaging, the mouse was placed on a custom-built translation stage and a bite bar was used to stabilize the head for camera alignment. All experimental procedures followed the guidelines of the Canadian Council on Animal Care, and protocols were approved by the Dalhousie University Committee on Laboratory Animals.
Fluorescence lifetime maps, obtained with fluorescence lifetime imaging (FLIM), deliver additional information about cell health [48, 49, 50]. FLIM is mainly concentration independent and measures the average duration a molecule remains in an excited state. This duration is unique, providing a molecular fingerprint . Changes in fluorescence lifetime reflect changes in cellular environment, such as temperature, pH, ion, and oxygen concentration [45, 52].
Focused at the level of the retinal nerve fiber layer where axons of RGCs are located, the confocal reflectance images visualize the mouse fundus (Fig. 9.10a), whereas the TPE images show RGCs (Fig. 9.10b). For each TPE pixel, the mean fluorescence lifetime (tm) was determined and displayed in a color-coded FLIM map with a decay time range from 2.1 ns (red) to 3.2 ns (blue) (Fig. 9.10c). The same imaging was performed ~20-μm deeper but at the same lateral position. This axial position was approximately at the level of the RGC somas (Fig. 9.10e, f). Although RGCs are clearly visualized in the Thy1-YFP-16 mouse strain (Fig. 9.10b, e), the acquisition of static fluorescence intensity measurements of RGCs does not deliver sufficient information to discriminate functional from nonfunctional RGCs . Previous work from our group has shown that, after experimental optic nerve injury, some RGCs expressing GCaMP3, a calcium indicator whose dynamics are related to changing calcium levels during neuronal action potentials, do not respond to a stimulus . Dynamic fluorescence intensity imaging of these markers enables probing of cellular function of individual RGCs in response to physiologic stimuli.
FLIM measurements as presented in (Fig. 9.10c, f) have the potential to provide critical and complementary information about cell status [31, 53] and differentiate among RGC subtypes with different levels of vulnerability to damage. This work is currently in progress.
9.4 Future Application of Two-Photon Scanning Laser Ophthalmoscopy
Currently, most experimental models of diseases leading to RGC loss, such as glaucoma and ischemic optic neuropathy rely on qualitative or quantitative assessment of RGC loss after tissues have been processed after termination of the experiment. Therefore, a longitudinal assessment of the degree or rate of RGC loss is not typically made in vivo, resulting in a limitation in the assessment of damage in chronic disease models.
The availability of transgenic mice that express fluorophores under the control of promoters thought to be expressed by RGCs, such as Thy1 , have greatly advanced the field. Several examples of characterization of RGC loss in vivo after experimental optic nerve diseases have been published [14, 15]. While readily available, these transgenic strains pose some limitations given that the expression profile of the promoters change after injury and encapsulation of the fluorophore by phagocytosing cells after RGC death. Hence, the specificity and accuracy of RGC loss with these strains can limit the accuracy of experiments. Even if attempts to move away from transgenic animals to make these applications more translatable, these problems exist even if the fluorophore is introduced exogenously via techniques such as virus-based transfection .
However, as recent evidence shows , structural fluorescent markers are likely inadequate markers of RGC integrity since the presence of fluorescence does not necessary indicate functional viability. The ability to dynamically image fluorescence after a light stimulus represents a significant advance. As this chapter has indicated, two-photon scanning laser ophthalmoscopy offers a real potential to scientists to study the functional impact of diseases causing RGC loss in experimental disease models. This technique also offers a powerful assay to study the impact of therapeutics.
Ultimately, if fluorescent indicators of functional activity can be safely introduced into RGCs in humans and safely imaged after visual stimulation with two-photon scanning laser ophthalmoscopy, remarkable progress would be made in the diagnosis and treatment of many ocular diseases. Indeed, single cell functional imaging of RGCs could represent one of the single-most important imagining innovations.
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