It has never been easy to peer inside the living eye. Though several investigators in the 19^th century were aware of the illumination conditions necessary to cause the otherwise impenetrable pupil to glow red, they failed to obtain a clear view of the living retina until Helmholtz invented the ophthalmoscope (Helmholtz, 1851). Since then, the evolution of this instrument has been punctuated by important, if occasional, technical advances such as adding a camera that could record retinal images on film (Jackman and Webster, 1886). Surprisingly, before the late 1980's there were few improvements in optical resolution since Helmholtz's landmark invention. A rare example is the incorporation of the electronic flash, invented ~30 years earlier by Edgerton, that allowed exposures brief enough to avoid eye motion blur (Ogle and Rucker, 1953). In contrast, the last quarter century has witnessed gains in spatial resolution that have completely transformed our capacity to recover information from the retinal image.

These experiments lead to psychophysical methods to characterize the arrangement and spacing of cones that relied on aliasing effects caused by imaging interference fringes on the retina (Williams, 1985; 1988). Artal and Navarro (1989) wondered whether another variant of interferometry, based on the interference of light returning from the retina rather than light entering the eye, could be used to estimate the spacing of cones in the living eye. Their approach was derived from a technique used in astronomy, stellar speckle interferometry, that had been developed to recover high spatial frequency information from speckled images, such as whether an object in the sky was a single or a binary star (Labeyrie, 1970), They collected images of the speckle patterns generated by illuminating small patches of living human retina with coherent light. The average power spectrum of multiple images revealed a ring corresponding to spatial frequency components of the cone mosaic, providing the first evidence about the granularity of the cone mosaic obtained from images of the living human eye.

Unfortunately, Artal and Navarro's method did not allow for the direct observation of the cone mosaic because the cones are obscured by the interference of the coherent light reflected from multiple layers of the retina. Experiments in animals had previously demonstrated direct imaging of receptors. In 1985, Land and Snyder (1985) showed in the snake and Jagger (1985) showed in the cane toad that it is possible to image photoreceptor cells through the natural optics of the intact eye. While these observations were made on animal eyes with good optics and unusually large receptors, they encouraged the possibility that this could also be achieved in the human eye. Cidecyian and Jacobson (1994) noticed small bright spots in fundus images from patients who were carriers of x-linked retinitis pigmentosa. They speculated that these spots were cones back-illuminated by the highly-reflective tapetum-like fundus in these patients.

The idea had been around since the early 1970's that one might recover the depth profile of thick tissue from the variation in the time it takes for light to return from different depths. But the speed of light is so high that this is not practical to measure, and the early pioneers of OCT seized instead on low coherence interferometry. The historical roots of low coherence interferometry lay in the observation by Robert Hooke in 1664 that interference could be observed between two reflective surfaces illuminated in white light. This phenomenon was later analyzed by Isaac Newton, and became known as “Newton's rings”. The basic principle of marshalling low coherence to measure depth profiles in the eye is shown in Fig. 4a, in the context of what is now called time domain OCT. A light source is used that has very short coherence length, ideally small compared with the retinal layers one wishes to distinguish. This short coherence length implies that when the beam is split in two by a fiber coupler and one beam is directed at the retina and the second one to a reference mirror, the two beams, when brought back together, can only interfere if the two beams traveled the same optical path length to within the coherence length of the source. Shifting the axial position of the reference mirror changes the depth in the retina where interference can be produced, and the magnitude of the interference signal at that depth increases with the intensity of the reflection. By measuring the magnitude of the interference signal as a function of the axial position of the reference mirror, it is possible to identify the light reflected from different depths, without the need to measure the transit time of light. The earliest application of these ideas to the eye produced a depth profile along a line passing through a single point on the retina, allowing, for example, very precise measurements of the axial length of the eye (eg. Fercher et al.,1988). By transverse scanning, sequentially acquiring axial depth profiles along a linear series of locations across the retina, one can construct a 2 dimensional tomograph or retinal cross-section in the living eye. OCT is able to detect much weaker signals from the retina than conventional fundus imaging because it takes advantage of the increased gain provided by heterodyne detection. In heterodyne detection, the amplitude of the interference signal is proportional to the product of the amplitudes of the two interfering beams so that the use of a large amplitude reference beam amplifies the weak interference signal. David Huang and colleagues, working in Jim Fujimoto's laboratory at MIT, made the first in vitro optical coherence tomograph of the retina in 1991 (Huang, 1991). Adolph Fercher's group made the first in vivo optical coherence tomography of the optic disc in 1993 (Fercher et al., 1993).

The next major step in the evolution of OCT was to improve the axial resolution even further. The first systems had a fairly long coherence length, which provided a theoretical resolution of about 10-15 microns. In 1999, Wolfgang Drexler, working in Jim Fujimoto's lab, pioneered the use of broader band, shorter coherence length, sources that could in principle achieve an axial resolution of 3 microns in the eye, two orders of magnitude better than the confocal SLO. Though the resolution improvement is limited to a single spatial dimension, ultrahigh resolution OCT could in theory produce a point spread function in the axial dimension smaller than the size of many cells in the retina.

A major limitation of time-domain OCT is the speed at which the retina can be scanned. This is especially problematic because large movements of the eyes could occur even with the fastest scanning speeds available. The main limitation on the speed of image acquisition in time domain OCT is the time required to shift the optical path length in the reference arm of the interferometer, either through mechanical displacement of the mirror or a variable delay line. Avoiding this delay, Fercher introduced the basic principle of spectral domain OCT in 1995. A typical layout for a spectral domain OCT is shown in Fig. 4b. It is similar to the time domain OCT with the exception of the detection arm of the system, which contains a spectrometer instead of a simple light detector. The spectrometer consists of a diffraction grating and a one-dimensional CCD that can acquire the entire spectrum of the light returning from a single retinal location in parallel. Spectral domain OCT obviates the need for scanning in the reference arm to acquire the depth profile at each retinal location. Instead, the reflectance at each depth is obtained from variations in the pattern of interference across the spectrum of the reflected light. The spectrum contains depth information because the light reflected from a single depth at one retinal location will generally produce a sinusoidal variation in the interference pattern across wavelength. The amplitude of this sinusoid is proportional to the square root of the amplitude of the electric field reflected at that point. The frequency of this sinusoid increases as the optical path length difference between the reference and sample increases. Therefore, the Fourier transform of the spectrum of the light returning from reflection will reveal the amount of light reflected at each axial depth at that retinal location. The first demonstration of its application in retinal imaging was published in 2002 (Wojtkowski et al., 2002a, b). The introduction of spectral domain OCT (sometimes called Fourier or frequency domain OCT) increased the sensitivity of OCT by roughly 2 orders of magnitude. The reason for this improvement is because, in time-domain OCT, at any instant, all the light is lost that is illuminating other retinal layers than the one that happens to be interfering with the reference arm. In spectral domain OCT, at any instant, light from all retinal layers contribute to defining the axial profile. An example of a spectral domain OCT scan from Jack Werner's group at UC, Davis is shown in Fig. 5.

In conventional fundus cameras designed for wide field clinical imaging, no attention is usually paid to correcting aberrations other than defocus. But if the goal is to maximize the lateral resolution of microscopes for imaging the living retina at a cellular spatial scale, it is necessary to compensate not only for defocus, but also astigmatism and a host of other aberrations that we now refer to as higher order aberrations. A quarter of a century ago, higher order aberrations were collectively called “irregular astigmatism”. Only a few devoted disciples of physiological optics had succeeded in characterizing some of the individual aberrations that make up irregular astigmatism, and because no method existed to correct them, the importance of their work went largely unrecognized. Smirnov (1961) used a subjective vernier task to measure the third- and fourth-order aberrations of the eye. He lamented the lengthy calculations required to compute the wave aberration and stated that it was unlikely this method would have practical significance, not foreseeing the rapid increase in computing speed that would make it possible within 40 years to measure and compute the eye's wave aberration in a few tens of milliseconds. Walsh, Charman, and Howland (1984) demonstrated the first objective method to measure the eye's wave aberration. This was a significant breakthrough in the history of aberration measurement, but the method was never widely adopted because it was never automated.

In 1989, Andreas Dreher, working in Joseph Bille's laboratory at the University of Heidelberg, used a deformable mirror to improve retinal images in a scanning laser ophthalmoscope (Dreher et al, 1989), correcting the astigmatism in one subject's eye based on a conventional spectacle prescription. At the time, Bille's group had not yet developed the Shack-Hartmann wavefront sensor, so that they were not able to correct additional aberrations beyond astigmatism.

The signal to noise ratio of single images of the cone mosaic was too poor to reliably identify the pigment contained in each cone. This made it necessary to add many images. Eye movements that occurred between frames are typically many times larger than the diameter of a cone, so that it was necessary to register multiple frames with cross-correlation before adding them. This approach turned out to be effective at increasing signal-to-noise ratio without compromising image quality through registration error. Putnam et al. (2005) later showed that it is possible to record the retinal location of a fixation target on discrete trials with an error at least 5 times smaller than the diameter of the smallest foveal cones. These results showed how accurate this process could be, and that registration error with high quality AO images can be smaller than the point spread function of the aberration corrected image. The ability to register and add such high spatial resolution images turns out to be key to microscopic retinal imaging because safety limits often preclude delivering all the light in a single exposure.

Austin Roorda, then at the University of Houston, and his colleagues built the first instrument that combined the benefits of adaptive optics with those of the confocal SLO (Roorda et al., 2002). The optical layout of an AOSLO is shown in Fig. 10. This new instrument had a number of important advantages. First, adaptive optics increased confocality by increasing the quality of the retinal image at the pinhole, allowing the use of smaller pinholes and therefore better rejection of light from unwanted planes away from the image plane. The sectioning capability of the AOSLO is shown in Fig. 11 (see also, Fig. 3). In the most recent AOSLOs at the University of Rochester, this has improved the axial resolution of the first confocal SLOs by about an order of magnitude. Second, the AOSLO provides microscopic resolution with a real-time view of the retina. This has enabled the video-rate imaging of the motion of individual blood cells in the smallest capillaries in the retina (Martin and Roorda, 2005; Martin and Roorda, 2009; Zhong et al., 2008; Zhong et al., 2011). Tam et al., 2010 has improved the imaging of retinal vasculature without the need for fluorescein by using motion contrast to identify vessels, setting the stage for automated measurements of leukocyte velocity (Tam et al., in press). In vivo studies of blood flow in these smallest of vessels may provide valuable information about early vascular changes in diseases such as diabetic retinopathy.

The clinical utility of AO imaging is only now beginning to be explored (Baraas et al., 2007; Bhatt et al., 2010; Carroll et al., 2004, 2009, 2010; Choi et al., 2006; Chen et al., 2010; Hammer et al., 2008; McAllister et al., 2010; Talcott et al., 2010; Wolfing et al., 2006; Yoon et al., 2009). An exciting recent paper has also shown the potential of AO in drug development (Talcott et al., 2011). They have shown that the AOSLO can detect photoreceptor rescue by a drug, CNTF, in retinitis pigmentosa patients. The value of AO is highlighted by the fact that other standard clinical measures such as visual acuity failed to show a benefit. It seems likely that the microscopic views of cells provided by AO will find their place in the clinic as an especially sensitive indicator of the progression of retinal disease.

The developers of new technology for retinal imaging may be unaware that they have something in common with cave divers. Both enjoy an extreme sport that is more often than not an exercise in the management of claustrophobia. The cave diver's view of his cramped world is of almost entirely impenetrable rock. The developer of retinal imaging technology is similarly boxed in by the fundamental limits of physics and biology. Both grope along, driven by the hope of discovering a previously-missed passage that, if they can just squeeze through, will open up into whole new possibilities for exploration. The history of retinal imaging would suggest, however, that fundamental limits are rarely if ever fundamental limits of the natural world. They are almost always fundamental limits of the conceptual framework we have chosen to think in. The conventional wisdom in 1990 held that axial resolution was fundamentally constrained by geometrical optics and depth of focus. But by adding the wave properties of light to the conceptual framework, practitioners of low coherence interferometry in the eye surprised everyone with a 100-fold, and now recently a 1000-fold, improvement in axial resolution over what was thought possible before.

Though it is true that the speed of light has stood its ground as a fundamental limit for a very long time, can we really say with confidence that any limits are truly impenetrable? Progress in high-resolution retinal imaging demands that we treat these limits as invitations to overcome them. It may seem that the spectral transmittance of the cornea and lens would pose a fundamental limit on noninvasive optical imaging of the retina, forever preventing us from interrogating molecules that absorb light only in spectral regions outside the ocular transmittance window. But as shown in Fig. 19, Jennifer Hunter and colleagues have recently shown that two-photon imaging can excite autofluorescent molecules in the living retina that have excitation spectra in the near UV, outside the spectral pass band of the eye's optics (Hunter et al, 2011). If this method can be made more efficient, it could open up an entirely new way to study retinal structure and function in the living eye.

Another major constraint on microscopic retinal imaging is the maximum permissible light exposure that can be delivered without damaging the eye. This constraint is especially troublesome given that the typical yield in reflectance imaging is one photon back for every 10,000 that enter the pupil. The problem is compounded by the high magnification that microscopic retinal imaging demands, as well as the low efficiency of many potentially informative light-tissue interactions. But the dramatic improvements in eye tracking described earlier make it possible to maintain high spatial resolution and avoid thermal light damage by harvesting light at reduced power over longer times. Moreover, compounds have been discovered that protect the eye from photochemical light damage (eg. Maeda et al., 2006), and the application of these prior to imaging could make it possible to observe faint retinal signals that are presently invisible. In addition, we are at the beginning of a revolution in the availability of contrast enhancing fluorophores that one can direct to particular classes of retinal neurons with increasing specificity, functional imaging capabilities, and improved quantum efficiency.