Emerging Tools for Single-Cell Analysis

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instrumentation are discussed. From there, recent advances in two-photon deep-tissue imaging are addressed: (1) the application of a blind deconvolution algorithm in further improving image quality and providing point spread function (psf) information in deep tissue, (2) the development of video-rate two-photon instrumentation to extend the imaging technology to potential clinical applications, and (3) the acquisition and analysis of two-photon spectroscopic data as complementary information to structural imaging.

Confocal versus Two-Photon Fluorescence Microscopy

Confocal microscopy has been traditionally the most promising technology for the study of three-dimensional (3D) deep-tissue structure with subcellular resolution (Raijadhyaksha et al., 1995; Corcuff et al., 1993; Masters et al., 1997c; Masters, 1996; Rummett et al., 1994). The basic concept behind confocal microscopy is straightforward (Wilson and Sheppard, 1984; Pawley, 1995). Excitation light deflected by a beam splitter is focused through the objective onto the sample. The excitation source generates scattered light (reflected mode) or fluorescence (fluorescence mode) throughout the hour-glass-shaped excitation volume. To obtain an optical section, the light collected by the objective is refocused onto a pinhole placed at a telecentric plane. At this plane, only the photons originating from the objective focal volume will be focused through the aperture whereas the offfocal photons are rejected. High-resolution 3D image stacks from human skin and eyes have been obtained in vivo (Raijadhyaksha et al., 1995; Corcuff et al., 1993; Masters et al., 1997b, 1997c; Masters, 1990, 1996; Piston et al., 1995). This technique provides spatial resolution with radial and depth resolution on the order of 0.2 and 0.5 m, respectively. The penetration depth of confocal techniques varies depending on the experimental conditions. In the reflected light mode, the penetration depth of a confocal microscope is about 200 m. However, in the fluorescence mode, the penetration distance is much shorter. With near-UV excitation, the penetration depth in human skin is reduced to only 30 m. Furthermore, the one-pho- ton excitation used in typical confocal microscopy results in signal generation throughout the excitation volume and can result in unnecessary photodamage in off-focal regions.

In the past decade, development in two-photon microscopy, initiated by Denk et al. (1990), bypasses many of the drawbacks of conventional confocal microscopy and emerges as a new and powerful technique in applications involving deep-tissue imaging. In this method, chromophores can be excited by the simultaneous absorption of two photons each having half the energy needed for the excitation transition (Denk et al., 1990; Göppert-Mayer, 1931). Unlike one-photon excitation, two-photon absorption occurs most significantly in regions with high excitation photon flux density. The high photon flux density can be easily achieved by focusing lasers with short pulse widths using microscope objectives with high numerical aperture (NA). In this manner, excitation photons arrive at the spatially confined focal volume within a narrow temporal window, thus creating the high instantaneous photon flux density necessary

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for efficient two-photon excitation of fluorescent molecules. While two-photon excitation can be achieved using continuous wave (CW) laser excitation, tissue thermal damage is a serious concern because of the high average power needed (Booth and Hell, 1998). The efficiency of two-photon excitation under different experimental conditions can be examined by calculating na, the number of photon pairs absorbed per laser pulse per chromophore. For a focused, mode-locked laser source with an average power p0, repetition rate fp, pulse width p, and wavelength , na can be estimated from (Denk et al., 1990)

na	p02	(NA)2		2
	f 2	hc

where NA is the numerical aperture of the focusing lens, c is the speed of light, h is Planck’s constant, and is the two-photon cross section, typically on the order of 10-50–10-49 cm4 photon-1 molecule-1.

There are a number of advantages to implementing two-photon-induced absorption in microscopy:

1.Since excitation occurs appreciably only at the focal region, excellent axial depth discrimination is achieved. For one-photon excitation in a spatially uniform fluorescent sample, significant fluorescence is generated from different axial sections above and below the focal plane. On the other hand, since two-photon excitation is nonlinear in nature (quadratic dependence on excitation power), over 80% of the total fluorescence intensity originates from a 1- m-thick region near the focal point for a NA 1.25 objective. The confinement of excitation volume allows 3D imaging with excellent depth discrimination to be achieved. An added benefit of the restricted excitation volume is the accompanying reduction in photodamage. Since only fluorophores near the focal volume are excited, sample photodamage is limited to the focal region, and this prolongs sample viability.

2.A second major advantage is the achievable greater penetration depth. The typical absorbancy of infrared photons is more than an order of magnitude less than that of near-UV or blue-green light. Furthermore, since Rayleigh scattering is inversely proportional to the fourth power of wavelength, the red photons used in two-photon microscopy scatter much less than the blue photons used in one-photon excitation (Jackson, 1975). Therefore, the red/near-IR excitation light source used in twophoton excitation can be focused to a greater depth, and optically thick samples can be examined without disruptive procedures.

3.Large-area photodetectors that can be used in two-photon microscopy allow more efficient photon collection than confocal detection. In confocal microscopy, the emission pinhole is used to reject out-of-focus light. However, inside deep tissue, scattering of signal photons is inevitable and can lead to significant loss of these photons at the confocal pinhole. The detection geometry for the fluorescence photons is less critical in the two-photon case, and most of the forward-scattered photons can be retained (Fig. 10.1).

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4. Two-photon excitation wavelengths are typically about twice that used in onephoton techniques. This wide separation between excitation and emission spectra ensures easy optical rejection of the excitation light and Raman scattering without significant attenuation of the fluorescence. The result is a more complete collection of the signal photons.

Tissue Autofluorescence

The importance of autofluorescence imaging has been recognized because of its use for noninvasive monitoring of tissue cellular metabolism by redox fluorometry (Masters, 1990; Chance and Thorell, 1959; Chance, 1976; Chance et al., 1979; Bennett et al., 1996; Masters and Chance, 1993). Intrinsic fluorescent probes include pyridine nucleotides and flavoproteins. The pyridine nucleotide NAD(P)H can be excited in the 365-nm region and fluorescence emission occurs in the 400–500 nm region. Spectroscopic properties of the flavoproteins are more red shifted with excitation

F i g . 10.1. A comparison between two-photon and confocal detection geometry. In a highly scattering specimen, a large fraction of emitted photons can be scattered before they are collected by the objective. In the two-photon system, most of these scattered photons can be collected by a large area detector. In the confocal system, these scattered photons can be blocked by the confocal pinhole aperture and escape detection. (Reprinted with permission from IEEE. IEEE Engineering in Medicine and Biology Magazine 18:23–30, 1999.)

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around 450 nm and fluorescence between 500 and 600 nm. Typically, the changes in fluorescence intensity follow changes in cell and tissue oxidative metabolism. Fluorescence imaging of cellular redox metabolism has been successfully performed in the cornea and in the skin.

Recent studies by Webb’s group at Cornell have demonstrated that two-photon microscopy is capable of imaging other tissue components such as collagen/elastin using their endogenous fluorescence (Maiti et al., 1997; Zipfel, 1997). Optical properties of these extracellular matrix components are known. Tropocollagen and collagen fibers typically have absorption in the range of 280–300 nm and emission ranges from 350 to 400 nm (LaBella and Gerald, 1965; Dabbous, 1966; Hoerman and Balekjan, 1966). Elastin can be excited by wavelengths of 340–370 nm and fluoresces from 400 to 450 nm with excitation (Thomas et al., 1963; LaBella, 1961; LaBella and Lindsay, 1963).

SKIN AND CORNEA AS DEEP-TISSUE IMAGING MODELS

Two important physiological models for deep-tissue studies are skin and cornea. Both are epithelial tissues and can be easily accessed by a two-photon microscope.

Anatomy of Skin

Typical mammalian skin (Fig. 10.2a) can be divided into two main layers: the epidermis and the dermis. The outer layer, the epidermis, is a squamous, multilayered structure composed primarily of keratinocytes. Epidermal tissue continuously renews itself by mitotic cell division from its deepest stratum, the basal layer. As the cells gradually migrate to the surface, they progressively mature by the process of keratinization, which results in the synthesis of the fibrous protein keratin. The basal monolayer is composed of cuboidal cells adhering to the basement membrane from which anchoring fibrils extend into the superficial dermis (papillary dermis). In a light microscope, the dermal-epidermal junction is observed as an undulating pattern of ridges. Cells produced by basal-cell division form the outer layers as they migrate toward the surface. Stratum corneum, the surface layer, is composed of hexagonalshaped, fully keratinized dead cells without nuclei and cytoplasmic organelles. The dermis is composed mainly of collagen/elastin fibers mixed with a sparse population of fibroblasts. In human, the epidermal thickness varies from 50 m on the eyelids to 1.5 mm on the palms and soles. In mouse skin, the epidermis is typically 30 m thick. In the case of a mouse’s ear, cartilage forms the structural support for the mouse ear below the skin.

Anatomy of Cornea

As an epithelial tissue, cornea has great structural similarity to skin (Fig. 10.2b). The corneal epithelium is about 50 m thick. The epithelium consists of a single layer of

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basal epithelial cells, an adjacent layer of wing cells, and additional poorly organized wing cell strata up to the surface. The cells in the uppermost layer are the superficial epithelial cells. The region below the epithelium is called the stroma, which is structurally similar to the dermal layer in the skin, with a typical thickness of about 350–450 m, and consists of collagen fibers interspersed with stromal keratocyte cells. The cell bodies of the stromal keratocytes are about 1 m thick and form long extended processes in the plane of the nuclei.

TWO-PHOTON FLUORESCENCE MICROSCOPE

Setting up a Two-Photon Fluorescence Microscope

A typical two-photon microscope previously presented is shown in Figure 10.3 (So et al., 1995; Masters et al., 1997a). Central to this microscope is an ultrafast, modelocked titanium–sapphire (Ti–sapphire) laser (Mira 900, Coherent, Santa Clara, CA). The typical pulse width of the Ti–sapphire laser is about 150 fs, which is ideal for inducing two-photon excitation. The excitation wavelength of the laser can be

F i g . 10.2. Anatomy of (a) skin and (b) cornea. (Reprinted with permission from IEEE. IEEE Engineering in Medicine and Biology Magazine 18:23–30, 1999.)

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tuned between 700 and 1000 nm. Optical components such as a Glan–Thomson polarizer can be used to control the excitation laser power. The beam-expanded laser light is directed into the microscope via a galvanometer-driven x–y scanner (Cambridge Technology, Watertown, MA). Images are generated by raster scanning the x–y mirrors. The excitation light enters Zeiss microscopes (Zeiss, Thornwood, NY) via a modified epiluminescence light path. The scan lens is positioned such that the x–y scanner is at its eye point while the field aperture plane is at its focal point. Since the objectives are infinity corrected, an excitation tube lens is positioned to recollimate the excitation light. The scan lens and the excitation tube lens function together as a beam expander that overfills the back aperture of the objective lens. The dichroic mirror reflects the excitation light to the objective. The dichroic mirrors are custom-made short-pass filters (Chroma Technology, Brattleboro, VT) that maximize reflection in the infrared and transmission in the blue-green region of the spectrum. Various high-NA objectives can be used for the experiments. High-NA objectives ensure tight focusing and result in highly efficient two-photon excita-

F i g . 10.3. A two-photon fluorescence microscope.

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tion. Two commonly used objectives are Zeiss c-Apochromat 40X (NA 1.2) and Zeiss Fluar 40X (NA 1.3). A piezo-driven device, which is interfaced to a computer, controls the objective axial position. The field of view is about 50–100 m on a side. Typical image acquisition time for a frame is about 1 s.

The fluorescence emission is collected by the same objective and transmitted through the dichroic mirror along the emission path. An additional barrier filter is needed to further attenuate the scattered excitation light because of the high excitation intensity used. Since two-photon excitation has the advantage that the excitation and emission wavelengths are well separated (by 300–400 nm), suitable short-pass filters such as 2 mm of BG39 Schott glass filter (CVI Laser, Livermore, CA) eliminate most of the residual scatter with a minimum attenuation of the fluorescence. A descan lens can be inserted after the tube lens. A single-photon counting signal detection scheme is commonly used to detect and process the fluorescence photons. Fluorescence signal at each pixel is first detected by an R5600-P photomultiplier tube (PMT; Hamamatsu, Bridgewater, NJ), which is a compact single-photon counting module with high quantum efficiency. Then the signal is conditioned by a low-noise preamplifier and a photon discriminator (Advanced Research Instrument, Boulder, CO) before being transferred to the memory of the data acquisition computer.

Imaging of Mouse Ear

To demonstrate the power of two-photon imaging in deep tissue, we use an ex vivo mouse ear as a model imaging system. Images at five layers of the tissue were acquired (Fig. 10.4), and distinct physiological features at each layer can be readily resolved. At the tissue surface, that is, the stratum corneum, large, hexagonal-shaped cells were observed. These highly fluorescent cells in the stratum corneum typically have lost their nuclei. Cells in the process of sloughing were also observed. Two layers of living cells were observed between the stratum corneum and the basal layer. Quantitatively, the fluorescence intensity originating from these layers was one-half to one-third that of the stratum corneum. The cells in these layers were fairly large (about 15 m across) and well organized. The fluorescence of these cells originates mainly from the cytoplasm. From previous studies (Albota et al., 1998), the major chromophore in the cytoplasm responsible for the observed fluorescence is NAD(P)H. The cellular nuclei with a negligible NAD(P)H concentration were clearly observed as large oblong dark regions. The basal layer lies above the epidermal– dermal junction and cellular autofluorescence from the basal layer is slightly more intense as compared to the upper cell layers. Cell packing in this layer was less well organized. These cells were smaller in cross section but were thicker as compared with cells in the upper epidermal layers. In the dermal layer, we observed fiberlike structures likely to be collagen or elastin fibers. Below the dermis, we observed honeycomblike structures with a loose hexagonal packing arrangement. Fluorescence originates from the walls of these honeycomb structure. The morphology of this layer

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F i g . 10.4. Sections of mouse ear structures obtained by two-photon deep-tissue microscopy. From left to right, the images are of stratum corneum, epidermal cell layer, basal cell layer, dermal structure, and cartilage. The depth of these images is indicated in the upper left corner of the each frame. The bar scale is 20 m. (Reprinted with permission from OSA. Optics Express 3:339–350, 1998.)

is consistent with the cartilage structure below the dermis, which is the primary structural support of the mouse ear.

APPLYING A BLIND DECONVOLUTION ALGORITHM TO IMPROVE TWO-PHOTON IMAGES

Blind Deconvolution Algorithm

Although two-photon microscopy is an excellent tool for obtaining structural information in deep tissues, greater image distortion in the axial axis is expected due to the larger dimension of the psf along that direction. This discrepancy in radial and axial psf means that out-of-plane noises can contribute significantly to distortion in image quality, resulting in the presence of “haze” in the acquired images. To rectify this problem, 3D deconvolution techniques can be used to further enhance spatial resolution and therefore the image quality of the specimen. In addition to improving image quality, deconvolution can also provide information about the psf in deep tissue. For example, a 3D blind deconvolution algorithm based on the principle of maximum likelihood may not require the psf to be known. In fact, the AutoDeblur software (AutoQuant, Watervliet, NY) we use actually returns the psf as a parameter. This software is based on the blind deconvolution technique developed by Holmes (1992) and is extremely important for several reasons:

1.Inside inhomogeneous samples such as tissues, it is very difficult to derive the psf theoretically. Since the psf in a given tissue specimen is generally unknown, a blind deconvolution algorithm that does not require prior knowledge of psf is a more general technique.

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2.The ability of the blind deconvolution algorithm to produce psf at different depths in the sample allows us to characterize optical properties inside the complex, inhomogeneous specimen and can lead to a better understanding of photonic processes in such samples.

F i g . 10.5. Two-photon human skin at different depths. Left: raw two-photon images. Right: deconvoluted results. The blind deconvolution algorithm also reveals the psf. FWHM’s of the psf in the order of stratum corneum, basal layer, and dermal layer are: radial 0.49, 0.45, and 0.43 m; axial: 2.5, 2.3, and 2.2 m.

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Applying Deconvolution to Deep-Tissue Images of Human Skin

To realize the power of deconvolution in improving image quality and derive information on the psf inside deep tissue, we applied the blind deconvolution software AutoDeblur to different layers inside human skin imaged by two-photon fluorescence microscopy. A 100x oil-immersion objective (Zeiss Fluar, NA 1.3) was used to image three distinct regions inside a sample of human skin: (1) stratum corneum near the surface, (2) basal layer at the epidermal–dermal junction, and (3) filamentous structures inside the dermal layer. These images along with the deconvoluted results are shown in Figure 10.5. First, the images clearly show a qualitative improvement in image quality. Finer details are revealed in the deconvoluted images. Furthermore, the psf’s determined from the deconvolution show that the full widths at half maximum (FWHMs) did not change significantly even though these three layers are located at different depths inside the skin. To be specific, the radial FWHMs varied from 0.49 m in the stratum corneum to 0.45 m in the basal layer and is 0.43 m in the dermal layer. Axial FWHMs showed the same trend. The FWHMs are 2.5, 2.3, and 2.2 m in the respective layers. This is consistent with results published by Centonze and White (1998) in which latex particles were imbedded in tissue phantoms.

VIDEO-RATE TWO-PHOTON IMAGING

Implementing Video-Rate Two-Photon Microscopy

Although two-photon methodology can be applied to in-depth imaging of thick tissue, challenges remain in implementing this technology in the clinical setting. A major difficulty is associated with the relatively slow imaging speed of typical twophoton microscopes. The typical frame rates of two-photon and confocal microscopes are between 0.5 to 10 s. In order to acquire a three-dimensional image of a 200- m- thick tissue with 500 optical sections, data acquisition time on the order of hours is needed. In the clinical setting, such long acquisition time is impractical for patient diagnosis. In addition, the slow data acquisition rate can present problems in image acquisition throughout the 3D image stack. For long acquisition times, the images acquired are more likely to be subject to motional artifacts caused by the patients. Two different approaches have been demonstrated in bringing two-photon imaging speed to video rates (about 20–30 frames per second) (Brakenhoff et al., 1996; Guild and Webb, 1995; Bewersdorf et al., 1998).

The first technique is based on the line scanning approach. A line scanning approach reduces image acquisition time by covering the image plane with a line instead of a point (Brakenhoff et al., 1996; Guild and Webb, 1995). Line focusing is typically achieved by using a cylindrical element in the excitation beam path. The resulting fluorescent line image is acquired with a spatially resolved detector such as a charge-coupled device (CCD) camera. The main drawback associated with line scanning is the degradation of the image psf, especially in the axial direction.