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Chapter 13 – Ophthalmic Instrumentation

Chapter 13 – Ophthalmic Instrumentation

DAVID MILLER
EDMOND H. THALL
NEAL H. ATEBARA

KEY FEATURES
• The ability of a transparent medium to bend a ray of light is the basis for most of the instruments used in ophthalmology today.
• Spherical lenses, prisms, mirrors, slit-shaped illumination, astronomical and Galilean telescopes, and a multitude of other optical components—both simple and complex—have been devised and manufactured for over two centuries in order to study the human eye and its function.

INTRODUCTION
In this chapter, the basic principles that underlie some of the more common instruments used in ophthalmology will be reviewed, including the following:

1.
Keratometer and corneal topographer

2.
Slit-lamp biomicroscope

3.
Slit-lamp funduscopic lenses

4.
Applanation tonometer

5.
Optical corneal pachymeter

6.
Specular microscope

7.
Operating microscope

8.
Retinoscope

9.
Automated objective refractometer

10.
Lensometer

11.
Binocular indirect ophthalmoscope

12.
Direct ophthalmoscope

13.
Fundus camera

14.
Magnifying devices
KERATOMETER AND CORNEAL TOPOGRAPHER
The keratometer and its cousin, the Placido disc corneal topographer, are among the most widely used, yet misunderstood, instruments in clinical practice today. The principle of the keratometer is simple enough. An illuminated target, usually a ring, is placed near the patient’s eye. The cornea (with its overlying tear film) acts as a convex mirror to produce a virtual, erect, image of the ring. The size and position of the virtual image are measured ( Fig. 13-1 ). The corneal radius of curvature may be estimated from this, provided several assumptions are valid. They are: (1) the position of the illuminated ring and image must be known to high precision, (2) the cornea is assumed to be spherical, (3) paraxial optics is assumed, and (4) the power of the back corneal surface is estimated. However, the information gained is, in and of itself, very little.
Because the image is erect and smaller than the object, it can be stated with certainty that the cornea is a convex surface. However, any smooth convex surface can produce an image identical to the mires produced by the cornea. The corneal shape

Figure 13-1 Keratometer principle. An illuminated ring is placed in front of the cornea, which acts as a convex mirror and produces a virtual image of the ring approximately 4?mm behind the cornea.
could be spherical, paraboloid, hyperboloid, ellipsoid, or any other convex aspherical shape and would still produce images identical to those observed using the keratometer. A very fundamental problem in the design of the instrument is that it does not provide enough information to determine corneal shape accurately. The same is true for the Placido disc.
Perhaps the easiest way to understand the shortcomings of these instruments is to compare their characteristics with those of a technique commonly used in optical engineering to measure optical surfaces accurately, the Hartman test. [1] In the Hartman test a narrow beam of light from a precisely known position and direction is reflected from the surface to be measured. The position and direction of the reflected beam are measured precisely ( Fig. 13-2 ). The intersection of the two beams precisely locates the position of one point on the mirror. The process is repeated point by point, until enough points have been measured to determine the overall shape of the surface.
The problem with the keratometer and the Placido disc is that the characteristics of only one of the two beams are known with certainty. Those of the reflected beam are known, but those of the incident beam could be any of an infinite number of possibilities depending on the corneal shape ( Fig. 13-3 ). To determine the corneal shape that produces the image of the mire is not possible. Adding additional rings or data points does not solve the problem. No matter how many data points are sampled, the direction of the incident beam for each point is unknown.
Why not simply perform a Hartman test on the cornea? Researchers have tried this, but it is difficult because of the steepness

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Figure 13-2 Hartman test. The direction and location of the incident ray are known to high precision, and the same characteristics of the reflected ray are measured to high precision. The intersection of the two rays precisely locates a point on the optical surface.
of the cornea. To construct a clinical instrument is prohibitively expensive.
The above is not to say that these devices are useless, but rather that the data produced by these devices must be interpreted carefully. Before keratometry or corneal topography is undertaken, the clinician must have a definite purpose for the test and a clear understanding of the assumptions implicit in the data.
A fundamental difference between manual keratometers and Placido disc corneal topographers is that the former use a Scheiner double-pinhole focusing system so that the position of the virtual image is known precisely. In Placido disc corneal topographers, each ring produces an image in a slightly different plane and the instrument cannot assess the position of each image ( Fig. 13-4 ).
In keratometry, the assumption is that the cornea is spherical or toric. In general, the cornea is not spherical, so this assumption is known to be false. Nevertheless, this is the underlying assumption, in part because a sphere has a much simpler geometry than the true corneal surface. A sphere may be described completely by a single parameter (the radius), and a toroid may be described by three parameters (the two radii and an axis). The keratometer does not reveal the corneal shape, but it does describe the shape of a toric surface that would produce the same mires.
Clinical Use of Keratometric Information
For fitting rigid contact lenses this information is accurate enough, especially because keratometry is used to provide a starting point that is refined empirically. Multiple attempts over nearly half a century to automate the fitting process completely by relying on either keratometry or topography have failed.
Since the advent of IOLs, the keratometer has been used to measure corneal power, which requires additional assumptions. The keratometer and Placido disc topographers assess only the anterior corneal surface, but corneal power depends on both anterior and posterior surfaces and, to a much lesser degree, on thickness. The keratometer makes an assumption about the posterior corneal surface. Because the posterior surface has minus power, most keratometers compensate by using a smaller refractive index in the lens maker’s formula. The typical, but not universal, formula is given in Equation 13-1 , in which P is the

Figure 13-3 Fundamental flaw of the Placido disc. For the Placido disc and its many variants, the direction and location of the reflected ray are measured precisely, but the direction of the incident ray is unknown.

Figure 13-4 Placido rings are imaged in different planes. The virtual images of Placido rings do not lie in the same plane.
corneal power (D, diopters) and r is the radius (m). [2] This formula assumes a corneal refractive index of 1.3375 instead of Gullstrand’s value of 1.376.

Reliability of Keratometry for the Measurement of Corneal Power
The reliability depends to some extent on the formula used to calculate implant power. Many formulas are derived empirically or contain “fudge factors” that may compensate for erroneous assumptions inherent in the keratometer. Whatever the reason, keratometry appears to work reasonably well for the calculation of intraocular lens (IOL) power.

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Figure 13-5 The slit lamp. A, Some slit lamps first bring the light to a sharp focus within the slit aperture, and the light within the slit is focused by the condensing lens on to the patient’s eye. The observation system of a modern slit lamp has many potential reflecting surfaces; antireflection coatings on these surfaces help reduce loss of light. B, Slit-lamp apparatus. (A, Modified from Spalton DJ, Hitchings RA, Hunter PA. Atlas of clinical ophthalmology. New York: Gower Medical; 1984:10.)
The interpretation of Placido disc corneal topography data is much more complicated. First, Placido disc topography does not use the Scheiner double pinhole, so even more assumptions are involved, which is why corneal topography companies constantly discuss improved algorithms. Fundamentally, every different algorithm is simply a different set of assumptions. Whether one set of assumptions is uniformly superior to another is difficult to establish. For this reason, corneal topography must be regarded as a developing technology.
SLIT-LAMP BIOMICROSCOPE
The slit lamp is the piece of equipment most frequently used by the ophthalmologist. With the addition of auxiliary lenses, it can give unique, magnified views of every part of the eye. With the use of auxiliary devices it can be used to make quantitative measurements, including intraocular pressure, endothelial cell counts, pupil size, corneal thickness, anterior chamber depth, and others, and to take photographs. Illumination and observation are discussed in this section.
Illumination
Although Purkinje, in 1823, attempted to develop a type of slit lamp by using one handheld lens to magnify and another handheld lens to focus strong oblique illumination,[3] it was not until almost 100 years later that a version of the slit lamp appeared that is recognizable today. By 1916, the slit lamp was composed of the newly developed bright Nernst lamp; the Gullstrand illumination system, which condensed the light onto a slit aperture and then projected the slit onto the eye; the rotatable Henker arm; and the Azanski stereomicroscope, which slid across a glass-topped table.[3] Because the transparent cornea only backscatters about 10% of incident light, [4] development of the bright lamp and a powerful condensing system was essential before the faint nuances of the cornea could be seen.
In the terminology of the visual physiologist, the brighter illumination allowed the observer to move farther along the contrast-versus-background intensity curve and exploit the heightened sensitivity of the arrays of retinal cones.[5] At present, the physiological limits of illumination may have been reached for both the doctor and the patient. Most patients are unable to endure bright lights; illumination beyond a certain limit simply produces a noisy frenzy in the retinal circuitry and does not improve resolution. It is possible that a dim illumination system coupled with electronic light amplification may be used in the future.
Specifically, the modern slit lamp produces an intensity of about 200?mW/cm2 . When operated at the rated voltage, halogen lamps have a higher luminance and color temperature than do conventional incandescent lamps. For slit-lamp work, a high color temperature (e.g., a greater amount of blue light) is useful. Because many of the ocular structures are seen via light scatter, and because the shorter wavelengths are scattered most, a light with a high blue component illuminates the structures best. The light is first brought to a focus at the slit aperture ( Fig. 13-5 ),[6] and the light within the slit is focused by the condensing lens onto the patient’s eye.
Improving Tissue Contrast
One of the great strengths of the modern slit lamp is the way in which contrast can be improved by various maneuvers:
• Optical sectioning: as the beam is narrowed, the scattered light of adjacent tissue is removed and greater detail of the optical section is seen.
• Tangential illumination: when the light is brought in from the side, highlights and shadows become stronger, and the texture (i.e., elevations and depressions) is seen better.
• Pinpoint illumination: the cells and flare in the anterior chamber in a patient who has iritis are best seen using a narrow beam focused into the aqueous, so that the black pupil becomes the background. The combination of the narrow beam and the dark pupillary background eliminates any extraneous light, which would reduce contrast. The same principle holds when the examiner pushes the lower lid up to examine the tear meniscus. For example, the stagnant cell pattern of an obstructed tear duct is best seen using a narrow beam, with the dark iris in the background.
• Specular reflection: in this technique the angle of observation is set to equal the angle of illumination. In this way, the structure of the front surfaces of the cornea (i.e., ulcers, dry areas) and the rear surfaces (endothelial pattern) may be assessed.
• Proximal indirect illumination: in this technique a moderately wide beam is directed to the areas adjacent to the area

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of interest. Against a dark background, the backscattered light from the lesion yields a higher contrast, which often allows the observer to see the borders of the lesion more precisely. For example, using this technique subtle corneal edema, with its minute pools of fluid, stands out more distinctly against a dark pupil.
• Sclerotic scatter: with the slit illuminator offset from its isocentric position, light is directed to the limbus. The light then follows the cornea as if it were a fiberoptic element and reaches the other side of the limbus. However, if a lesion or particles within the cornea exist, the backscattered light from the lesion or particles is seen clearly against the dark pupillary background.
• Retroillumination from the fundus: light sent through the pupil to the fundus is reflected and yields an orange background. Holes in the iris or subtle wrinkles in the cornea become silhouetted and much easier to see.
Observation System
The observation system has many glass-to-air surfaces. Theoretically, 4% of the incident light may be lost at each surface by reflection, which results in a substantial light loss. If the elements are given antireflection coatings (as they are in all modern slit lamps), the total gain in brightness rises by 20% compared with that of an uncoated system.
Another feature of the slit lamp is that the observation system is really a microscope, but with a long working distance (i.e., about 3.9?in [10?cm]). Prisms take the divergent rays from the patient’s eye and force them to emerge as parallel pencils from each eyepiece. Thus, a stereoscopic appreciation of the patient’s eye is achieved without convergence of the observer’s visual axis. Most slit-lamp microscopes offer magnifications between ×5 and ×50, with ×10, ×16, and ×25 being the most popular. The issue of resolution ultimately becomes an issue of diffraction limit. The working distance of the slit lamp (for the ×10 objective) is at least 100 times longer than that of the laboratory microscope. Therefore, the cone of light that emanates from the patient’s eye and is captured by the slit-lamp objective is small relative to that of the microscope. This narrower cone yields a wider diffraction pattern. Because the ability to resolve two cells implies a distinct space visible between each cell, and because the image of each cell border has a diffraction fringe, tighter diffraction fringes allow the viewer to distinguish two separate cells that are closer together. Abbe, an optical physicist, was able to combine the factors of aperture size and working distance (i.e., the focal length for a microscope) into an index of resolution called the numerical aperture (NA).[5] The NA not only includes aperture size and working distance, but also index of refraction of the media (i.e., oil, water, or air) and the wavelength of illuminating light. However, these latter factors are relatively constant for most systems.
The NA of the slit lamp is substantially lower because the laboratory microscope’s objective may be larger, its working distance smaller, or both. If the NA is known, the optimal magnification of an optical system may be calculated. The term optimal magnification is used here to define the limit beyond which further magnification yields no more information.
In the calculation of the magnification needed to see individual endothelial cells using slit-lamp specular reflection, the limitations are set by the resolution limit of the observer’s eye. For example, if the 20/20 (6/6) Snellen letter represents the accepted resolution limit, and each bar of a 20/20 (6/6) “E” subtends 1 minute of arc, then the details of each endothelial cell must be magnified to subtend 1 minute of arc on the observer’s retina. Although 1 minute of arc is the threshold, the lens designer often uses 3 minutes of arc to minimize fatigue. The two adjacent endothelial cells may be thought of as two outer cell borders that contain a common double cell border, a Snellen “E” with the open side closed. To calculate the magnification needed, convention assumes that the magnified image is positioned at the comfortable observation distance of 9.8?in (25?cm) from the observer’s eye. Thus, if the border-to-border dimension of an endothelial cell is assumed to be 10?m, then this must be magnified by a certain amount to allow it to subtend 3 minutes of arc at 9.8?in (25?cm). If tan 3 minutes of arc = 0.0009, then the cell must be enlarged to 225?m, a magnification of about ×22.5. The ultimate theoretical microscope has an index of resolution (NA) of 1.0 and can magnify objects by about 31,000. If a microscope with an NA of 1.0 has a magnification of ×1000, then a slit lamp with an NA of 0.05 can magnify up to ×50 effectively. Therefore, it is reasonable to expect good slit-lamp resolution at ×22.5, but only if other variables such as involuntary eye movement can be controlled.
Unfortunately, a magnification factor of ×50 also magnifies fine eye movements. Therefore, ×50 produces a movement smear, so magnification of more than ×25 rarely gives more information unless the eye is immobilized or the view frozen by photography or video.
SLIT-LAMP FUNDUS LENSES
Because the cornea has such a high refractive power, the slit-lamp microscope can view only the first one third of the eye. Special lenses, in conjunction with the slit-lamp microscope, can be used to view the posterior vitreous and the posterior pole retina. The two ways to overcome the high corneal refractive power are: (1) nullify the corneal power, or (2) utilize the power of the cornea as a component of an astronomical telescope, in a manner similar to that exploited by the indirect ophthalmoscope.
The Goldmann contact lens ( Fig. 13-6 ) and other similar lenses work in conjunction with the slit-lamp microscope to nullify the dioptric power produced by the corneal curvature and to bring the retina into the focal range of the slit-lamp microscope. These plano-concave contact lenses are placed on the cornea, forming virtual, erect, and diminished images of the illuminated retina near the pupillary plane, within the focal range of the slit-lamp microscope.
The Hruby lens is a powerful plano-concave lens, minus 58.6 diopters in power. It is held immediately in front of the cornea, forming a virtual, erect, and diminished image of the illuminated retina, near the pupillary plane, bringing it within focal range ( Fig. 13-7 ).
The 60D, 78D, and 90D funduscopic lenses ( Fig. 13-8 ) use a different approach to view the posterior vitreous and posterior pole retina. These lenses act as high-powered, biconvex, condensing lenses, projecting an inverted, real image in front of the lens, within focal range. This is the same optical principle used by the indirect ophthalmoscope; the higher the power of the lens, the lower the magnification of the image.
The Goldmann three-mirror contact lens ( Fig. 13-9 ), as its name implies, incorporates three internal mirrors. The contact lens nullifies the refractive power of the patient’s cornea, and the three mirrors then reflect light from the patient’s midperipheral retina, peripheral retina, and the iridocorneal angle, respectively. The posterior pole of the fundus can be visualized, also, in a manner similar to that of the Goldmann posterior pole contact lens.
The panfundoscope contact lens and the Rodenstock contact lens are high-powered, spherical, condensing lenses, as are corneal contact lenses ( Fig. 13-10 ). A real, inverted image of the fundus is formed within the spherical glass element, which is within the focal range of the slit-lamp microscope. Because the condensing lens is so close to the eye and has such a high power, the field of view is very wide, making these lenses specially suited for a wide-angle view of the posterior pole and midperipheral fundus.
The applanation tonometer ( Fig. 13-11 ) is used to measure intraocular pressure. It relies on an interesting physical principle: for an ideal, dry, thin-walled sphere, the pressure inside a sphere

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Figure 13-6 The Goldmann fundus contact lens, or any similar plano-concave contact lens, nullifies the refractive power of the cornea, thereby moving the retinal image close to the pupillary plane and into the focal range of the slit-lamp microscope. The image formed is virtual, erect, and diminished in size.

Figure 13-7 The concave Hruby lens, when placed close in front of the patient’s eye, forms a virtual, erect image of the illuminated retina that lies within the focal range of the slit-lamp microscope.

Figure 13-8 The 60D, 78D, and 90D lenses produce inverted, real images of the retina within the focal range of the slit-lamp microscope in a fashion similar to that employed by the indirect ophthalmoscope.

Figure 13-9 The contact lens of the Goldmann lens nullifies the refractive power of the patient’s cornea, while the three mirrors then reflect light from the patient’s peripheral retina (orange ray) and iridocorneal angle (green ray). The posterior pole of the fundus also can be visualized in a manner similar to that of the Goldmann posterior pole contact lens (blue ray).

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Figure 13-10 The panfundoscope lens consists of a corneal contact lens and a high-powered, spherical condensing lens. A real, inverted image of the fundus is formed within the spherical glass element, which is within the focal range of the slit-lamp microscope.
is proportional to the force applied to its surface. Unlike an ideal sphere, however, the human eye is not thin walled and it is not dry, producing two confounding forces: (1) a force produced by the eye’s scleral rigidity (because the eye is not thin walled), directed away from the globe, and (2) a force produced by the surface tension of the tear film (because the eye is not dry), directed toward the globe ( Fig. 13-12 ). Goldmann determined that when a flat surface is applied to the cornea with enough force to produce a circular area of flattening 3.06?mm in diameter, then the force caused by scleral rigidity exactly cancels out the force caused by surface tension. Therefore, the applanating force required to flatten a circular area of cornea exactly 3.06?mm in diameter is directly proportional to the intraocular pressure. Specifically, the force (measured in dynes) multiplied by 10 is equal to the intraocular pressure (measured in millimeters of mercury).
How does the observer know when the area of applanation is exactly 3.06?mm in diameter so that the intraocular pressure can be measured? The applanation tonometer is mounted on a biomicroscope to produce a magnified image. When the cornea is applanated, the tear film, which rims the circular area of applanated cornea, appears as a circle to the observer. The tear film often is stained with fluorescein dye and viewed under a cobalt-blue light in order to enhance the visibility of the tear film ring. Higher pressure from the tonometer head causes the circle to have a wider diameter, because a larger area of cornea becomes applanated ( Fig. 13-13 ). Split prisms, each mounted with their bases in opposite directions, are mounted in the applanation head, creating two images offset by exactly 3.06?mm. The clinician looks through the applanation head and adjusts the pressure until the half circles just overlap one another ( Fig. 13-14 ). At this point, the circle is exactly 3.06?mm in diameter, and the reading on the tonometer (multiplied by a factor of 10) represents the intraocular pressure in millimeters of mercury ( Fig. 13-15 ).
OPTICAL PACHYMETER
A review of the literature concerning corneal thickness[7] reveals that the average thickness of the central 3?mm varies in the range 0.50–0.57?mm, with a standard deviation of about 0.04. This means that a change of thickness of two standard deviations, or 0.08?mm, is considered a statistically significant change. Such a statement indicates the precision required of corneal pachometry. Electronic corneal pachometry is probably more accurate than optical pachometry, and precision is important prior to refractive surgery procedures.
The value of optical corneal pachometry is manifold. Aside from providing information about the normal physiological
GOLDMANN APPLANATION TONOMETER

Figure 13-11 Photograph of a Goldmann applanation tonometer in working position on a slit-lamp microscope.
hydration dynamics of the cornea, pachometry has had a significant clinical impact. In the field of contact lens wear, corneal pachometry was the first quantitative objective parameter which differentiated a contact lens that fits well from a poorly fitting one.[8] Because corneal thickness is directly related to the health of the corneal endothelium, its measurement after surgery is an excellent indication of the amount of endothelial trauma sustained during surgery.[9] For example, corneal pachometry was used to demonstrate that an anterior chamber filled by the viscoelastic substance hyaluronic acid during IOL implantation protected the corneal endothelium from the injurious touch of the IOL.[10] Corneal pachometry also is a sensitive method with which to follow the health of a transplanted cornea.[11] Experience with post-operative LASIK patients has taught us that corneal thinning lowers the applanation tonometer reading. Thus, in the future, corneal thickness measurements may be used to arrive at a correct IOP measurement.
Modern optical pachometry depends upon the measurement of the thickness of the optical cross-section of the cornea, as seen in the slit lamp. However, a simple direct measurement using a ruler (a measuring reticule in the eyepiece of the slit lamp)

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Figure 13-12 A, When a flat surface is applied to the cornea with enough force (w) to produce a circular area of flattening greater than 3.06?mm in diameter, the force caused by scleral rigidity (r) is greater than that caused by the tear film surface tension (s). B, When the force of the flat surface produces a circular area of flattening exactly 3.06?mm in diameter, the confounding forces caused by scleral rigidity and tear film surface tension cancel each other. The applied force (w) then becomes directly proportional to the intraocular pressure (p).

Figure 13-13 When viewed through a transparent applanation head, the fluorescein-stained tear film appears as a circular ring (A). Greater applanation pressure causes the ring to increase in diameter (B).

Figure 13-14 The split prism in the applanation head creates two images offset by 3.06?mm, allowing greater ease in determining when the circular ring is exactly 3.06?mm in diameter. When the area of applanation is smaller than 3.06?mm, the arms of the semicircles do not reach each other (A). When the area of applanation is greater than 3.06?mm, the arms of the semicircles reach past each other (B). When the area of applanation is exactly 3.06?mm, the arms of the semicircles touch each other (C). This is the end point at which the intraocular pressure can be measured.

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Figure 13-15 When the applanation pressure is too low (1.0 dynes in this illustration) the circular ring is smaller than 3.06?mm in diameter, and the arms of the ring do not reach each other in the split image (A). When the applanation pressure is too high (3.0 dynes in the illustration) the circular ring is larger than 3.06?mm in diameter, and the arms of the ring stretch past each other in the split image (B). When the applanation pressure creates a circular ring exactly 3.06?mm in diameter, the arms of the ring just reach each other in the split image (C). In this illustration, the end point is reached at 2.0 dynes of applanation pressure, which corresponds to an intraocular pressure of 20?mmHg.
across the corneal cross-section gives only the apparent value. Figure 13-16 shows that the slit lamp, with its oblique illumination, views an oblique slice of the cornea. If the observation microscope makes a 40° angle with the illumination system, then the real thickness is the hypotenuse of the triangle of which the oblique optical section is the base. Thus, the true thickness equals the oblique apparent section divided by the sine of 40°.

Figure 13-16 The oblique beam of the slit lamp gives an apparent corneal thickness. To calculate the real thickness, the length of the hypotenuse of the triangle, of which the oblique optical section is the base, must be calculated.

Figure 13-17 The observed view in the pachometer. The back of the top corneal section is aligned with the front of the bottom section. When this configuration is reached, the exact corneal thickness is read from the scale.
The apparent thickness is a physical measurement made, in theory, by the movement of a marker from the front to the back of the cornea. In reality, the modern pachometer produces two images of the apparent oblique section. Rotation of the top image so that its endothelial surface aligns with the epithelial surface of the bottom image essentially moves a marker from the front to the back of the cornea. Figure 13-17 shows the appearance of the alignment when a measurement is taken.
The endothelial surface, itself, is not seen. Instead, an image of that surface produced by the optics of the front of the cornea is seen. The effect of front-surface power and corneal index of refraction on the position of the image of the back surface is shown in Equation 13-2 , in which F is the power of front surface (42D), n is the index of refraction of the cornea (1.3375), u is the corneal thickness (i.e., the distance of the endothelial surface from the refracting surface [the front corneal surface]; 0.5?mm), and v is distance of the image of the endothelial surface from the

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Figure 13-18 Endothelial layer as seen in slit-lamp microscopy. Note the wide angle between the illumination beam and the observation path needed to remove the bothersome surface reflection and stromal scatter from the view of the endothelial mosaic.
refracting surface. Thus, for corneas of average curvature, the image of the endothelial surface is within 0.01?mm of the real endothelial surface.

SPECULAR MICROSCOPE
It seems remarkable that just 8 years after Gullstrand unveiled his first model of the slit lamp in 1911,[12] Vogt described the endothelium in the living eye using a modified Gullstrand slit lamp.[13] [14] The optics of the visualization of the endothelial mosaic produced by slit-lamp microscopy are shown in Figure 13-18 . [14] A photograph of the endothelium in the enucleated rabbit eye was first produced by Maurice in 1968.[15] In 1975, 56 years after Vogt presented a painting of the endothelium in vivo, Laing et al. developed a camera to photograph the layer in the living, human subject.[16]
Optics of Endothelial Microscopy
A number of significant obstacles stand in the way of easy microscopic observation of the living corneal endothelium. First, the reflection from the front corneal surface interferes with a sharp view of the endothelium. Second, the intervening stromal layers backscatter light, which decreases the contrast of the endothelial details. In addition, the thicker and more edematous the stroma, the hazier the views of the endothelium. Finally, because of the small difference in index of refraction between the cornea (1.376) and the aqueous (1.336), only 0.02% of the incident light (for most angles of incidence) is reflected from the interface between corneal endothelium and aqueous.[17]
To eliminate the bothersome reflection from the front corneal surface, two approaches are used. An increase in the angle of incidence

Figure 13-19 Optics of endothelial specular microscopy using a contact lens. (Adapted from Bigar F. Specular microscopy of the corneal endothelium. In: Straub W, ed. Developments in ophthalmology, vol 6. Basel: Karger; 1982:1–88.)
moves the anterior reflection to the side, so it covers less of the specular reflection from the endothelium. This approach alone is used in the noncontact technique. If the cornea could be thickened artificially (without an increase in light scatter), this also would move the surface reflection further to the side. By using a contact lens that has a coupling fluid of index of refraction similar to that of the cornea, the surface reflection is eliminated and the corneal thickness may be assumed to include the contact lens thickness. The reflection from the surface of the contact lens replaces that of the corneal surface. However, because of the thickness of the contact lens, the surface reflection is moved well over to the side ( Fig. 13-19 ).

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