Ординатура / Офтальмология / Английские материалы / Wavefront Customized Visual Correction The Quest for Super Vision II_Krueger, Applegate, MacRae_2003

important for learning about the fundamental properties of vision as well as for diagnosing retinal disease. Most effort on the Rochester AO ophthalmoscope to date has been concentrated on imaging cone photoreceptors, and that effort has produced the best pictures ever of the cone mosaic in the living human eye.

AO noninvasive imaging has produced the best images of the cone mosaic in the living human eye.

bright spots in the image are cone photoreceptors, which at this retinal location are about 5 µm in diameter. Nearly all photoreceptors are resolved in a single image. Registration of multiple images, as shown on the rightmost image, improves the signal- to-noise ratio in the image to an extent where all photoreceptors are resolved. Even the smallest cones across the very center of the fovea have been resolved, as shown in Figure 5-5, which is a composite of many frames taken around the central 2.2 degrees of one subject.

Signal-to-noise is the ratio of the signal (what you want to see—the cone mosaic) to the noise (information that degrades the signal).

Examples of the types of images that can be obtained are shown in the three panels on Figure 5-4. The image on the left is a single snapshot taken after only defocus and astigmatism have been corrected in the eye.6,16 Prior to the implementation of AO, sphere and cylinder were the only corrections that were applied to improve image quality. In the case shown here, the aberrations for the subject were sufficiently low so that some photoreceptor structure could be seen in the uncompensated images. This ability to resolve photoreceptors without using AO had already been documented by other groups who showed similar results.17-20 Nonetheless, the improvement in image quality obtained after AO compensation (see the middle panel in Figure 5-4) is striking. In the compensated image, the fine structures are better resolved and have higher contrast than in the uncompensated image. The

Arrangement of S, M, and L Cones

Over 200 years ago, Thomas Young proposed that the human retina was comprised of three cone types,21 but it was not possible to determine the spatial arrangement of those cones prior to the development of high-resolution imaging with AO. While retinal densitometry has been used for years to measure the pigment concentration in cone photoreceptors,22 the advantage of using AO is that it allows one to perform the same measurements on individual photoreceptors. Full details of this experiment are described elsewhere16 but a brief description follows. All photoreceptor images were taken with 550 nm light; a wavelength chosen to maximize the absorptance by L and M cone photopigments. Individual cones were classified by comparing images when the photopigment was fully bleached with those taken when it was either dark-adapted or exposed to a light that selectively bleached one photopigment. Images of fully bleached retinas were obtained following exposure to 550 nm light. Images of dark-adapted retina were taken following 5 minutes of dark adaptation. The S cones were distinguished from M and L cones by comparing fully bleached and dark-adapted images. Since the S cones absorb negligibly while the M and L cones absorb strong-

Figure 5-4. Images before and after adaptive compensation for the right eye of a living human subject. All three images are of the same retinal area located 1 degree from the central fovea. Images were taken with 550 nm light (25 nm bandwidth) through a 6 mm pupil. The dark, vertical band down the center of each image is an out-of-focus shadow of a blood vessel. The image to the far left shows a single snapshot taken after defocus and astigmatism have been corrected. The middle image is a snapshot after additional aberrations have been corrected with adaptive optics. The image to the far right shows the benefits in image quality obtained by registering and averaging multiple frames, 61 in this example.

Figure 5-6. Images of the cone mosaics of eight subjects with normal color vision, obtained with the combined methods of adaptive optics imaging and retinal densitometry developed by Roorda and Williams.16 The images are false colored so that blue, green, and red are used to represent the S, M, and L cones, respectively. (The true colors of these cones are yellow, purple, and bluish-purple). The mosaics illustrate the enormous variability in L/M cone fractions. The L:M cone ratios are A, 0.33; B, 1.04; C, 1.12; D, 1.5; E, 2.00; F, 2.33; G, 4.00; and H, 19.00. Images were taken either 1 or 1.25 degree from the foveal center.

ly at the imaging wavelength of 550 nm, the S cones reflect relatively more light than the M and L cones, which absorb the light and appear dimmer. Once the sparse population of S cones was identified, they were removed from analysis so that the M and L cones could be distinguished. To distinguish L from M cones, images were taken immediately following either of two bleaching conditions. In the first bleaching condition, dark-adapted retina was exposed to a 650 nm light that selectively bleached the L pigment. In the second, a 470 nm light selectively bleached the M pigment. The image following the 650 nm bleach revealed relatively brighter, low absorptance L cones that had been heavily bleached and darker, high absorbing M cones spared from bleaching. The absorptance images for the 470 nm bleach showed the opposite. Densitometric measurements were repeated in this way until the signal-to-noise in the data was sufficiently high to confidently identify the individual cones. Pseudocolor images of

Figure 5-5. Image of the center of the living human fovea obtained with the Rochester Adaptive Optics Ophthalmoscope. Because of the 1 degree field of view of the instrument, this larger image was constructed by merging a number of overlapping images centered at different locations. The height of the image is 0.98 degrees high and its width is 2.2 degrees wide, extending from 0.7 degrees temporal to 1.5 degrees nasal retina. Note the increase in cone spacing with increasing distance from the foveal center.

the arrangement of the cones near the fovea for eight human retinas are shown in Figure 5-6. The figure shows L:M ratios ranging from 0.33:1 to 19:1.

S, M, and L cones: S is short wavelength sensitive cones, M is medium wavelength sensitive cones, and L is long wavelength sensitive cones.

Angular Tuning

The angular tuning of the cones is often referred to as the Stiles-Crawford effect after the individuals who discovered the fact that in most eyes light entering the edge of the pupil is less effective in eliciting a visual response than light entering more centrally in the pupil.

It is well known that human photoreceptors act as waveguides. High resolution AO imaging can be brought to bear on some remaining questions about these waveguiding properties by looking at how each individual photoreceptor contributes to the overall directional properties. For example, it is suggested, based on psychophysics23 and reflectometric measurements from an ensemble of cones,24 that all the cones are narrowly tuned. This question can be answered directly now that individual cones can be resolved in living eyes. In the AO ophthalmoscope, the directional properties of individual cones were measured by determining how efficiently light is coupled into the cones as a function of illumination angle.

If less light gets into a cone because it is illuminated away from the cone’s optical axis, then the reflected light from that cone will also be less.25,26 Images were taken of the same cone mosaic under identical conditions except that the illumination angle was controlled by translating a 2 mm entrance pupil beam to different locations in the pupil. The images in Figure 5-7 are of

Figure 5-7. Composite of seven images of the same patch of retina taken with different entrance beam locations. Each image is a registered sum of 10 images. The number in the upper right corner of each image shows the position of the entrance beam relative to the central illumination beam location. The central location was centered with the best estimation of the Stiles-Crawford peak. The symbols S, I, N, and T indicate the superior, inferior, nasal, and temporal directions in retinal space. The reflectance of each cone changes as the illumination angle moves off the axis of the photoreceptor.

the same retinal patch taken with seven different entrance pupil positions. An angular tuning function was fit to over 200 contiguous cones in the images. The main result was that some disarray can be seen in the individual cones, but it is very small, accounting for less than 1% of the breadth of the overall tuning of an ensemble of cones.27 Thus, the angular tuning of ensembles of cones is a good measure of the angular tuning of a single cone photoreceptor.

The angular tuning of ensembles of cones is a good measure of the angular tuning of a single cone photoreceptor.

Adaptive Optics Scanning Laser Ophthalmoscopy

Optical Sectioning

The retina is a thin, mostly transparent, multilayered tissue. Its thickness is typically 300 µm from the nerve fibers on the surface to the retinal pigment epithelium. Beneath the retina is the choroid, which is primarily composed of blood vessels that nourish the retinal pigment epithelium and the back of the eye. Because of its complex structure, it is desirable to be able to image the retina in three dimensions. The improvement in optical sectioning in the AOSLO is important because current instruments report a maximum axial resolution (full width at half maximum [FWHM] of the axial point spread function [PSF]) of greater than 300 µm, which is about the thickness of the neural retina.

Figure 5-8. The three images are of a retinal location that is about 1.5 mm from the fovea. Each panel shows a different optical section of the retina, starting with the nerve fiber layer, and a blood vessel that runs across the retinal surface. The second panel reveals a second blood vessel and other capillaries that run beneath the nerve fibers. The final panel shows the mosaic of cone photoreceptors, which lie about 300 µm below the retinal surface. The scale bar on the middle figure represents about 100 µm.

AOSLO allows noninvasive optical sectioning in real time of the microscopic structure of the living human retina.

Improved axial sectioning has allowed for direct imaging of specific layers in the retina, such as nerve fibers, capillary layers, and photoreceptors. Figure 5-8 shows a series of optical sections of an area in a living human retina located about 4 degrees (1.2 mm) from the fovea. The sequence of images reveals the three dimensional structure of the retina.

Blood Flow

The AOSLO images at 30 frames per second. This allows us to track dynamic changes in the retina, such as blood flow. The movement of single white blood cells has been observed in the smallest capillaries. Moreover, this blood flow can be observed directly without the use of fluorescein. The movement is impossible to observe in static frames, but digital videos can be found at www.opt.uh.edu/research/aroorda/aoslo.htm.

THE FUTURE OF

ADAPTIVE OPTICS OPHTHALMOSCOPY

Clinical Applications

Early diagnosis and treatment of retinal disorders have been hampered by the inability to resolve microscopic structures in the living human eye. In many cases, the retinal disease is detected only after significant and irreversible retinal damage has occurred. Since early detection and appropriate treatment is the best way to maintain good vision, it is important that we develop instruments that are sensitive to the specific changes, like photoreceptor loss, that are known to occur with these disorders.

Rods are difficult to see because they are barely inside the resolution limits of the eye and because of low contrast due to the fact that, unlike cones, they do not direct their reflected light toward the pupil.

Figure 5-9. The two images are taken from different locations of the retina of a diabetic patient. The left image is focused on the photoreceptor layer, and shows shadows of small microaneurysms in the capillaries just nasal to the fovea. The right image shows a detail of a hard exudate that is about 2.5 degrees temporal to the fovea. The area outside the hard exudate appears dark because the photoreceptor layer is beyond the focal plane. The scale bar on the figure represents about 100 µm.

The increased contrast and resolving power offered with AO imaging will provide this sensitivity. More importantly, this increased sensitivity will open the possibility of testing the effectiveness of treatment interventions and also learning more about the mechanisms of the retinal disease.

To date, little AO imaging effort has been spent on the study of features other than the cone photoreceptors. To develop the instrument for clinical applications, it is desirable to expand the range of features that can be imaged (eg, rods, nerve fibers, and retinal pigment epithelium cells). Rods outnumber the cones by 20:1 in the human eye.28 They are not important for central vision, but they have important roles in peripheral and night vision and rods are also the first photoreceptors to be affected in common retinal diseases like retinitis pigmentosa. Rods have never been resolved with the AO ophthalmoscope, but it remains a real possibility and a future challenge. Nerve fibers have diameters of about 3 µm and the striated patterns they produce are readily seen in conventional fundus images. With AO, we should be able to resolve individual nerve fibers and even measure their diameters.

AO imaging will soon allow individual nerve fibers to be resolved, providing a powerful new tool for cutting edge research in glaucoma, as well as the monitoring and evaluation of therapy designed to alter the natural history of glaucoma.

Figure 5-9 shows two AOSLO pictures from preliminary investigations of the retina of a diabetic patient. The two images show microaneurysms and a detailed view of hard exudates.

Resolution, however, is not the only challenge to overcome for microscopic retinal imaging. Contrast is also a major limiting factor. Figure 5-1 illustrates that the contrast of features are improved in an image by improving the MTF, but successful imaging still relies on the object having its own intrinsic contrast. For example, RPE cells, whose health is implicated in several important retinal diseases, are difficult to see. The problem is not that they are too small, but rather because they have such low contrast. The low contrast is mainly due to the fact that to see the RPE, one must look through the strong reflection from the photoreceptor layer. Likewise, ganglion cells have low contrast, but

Retinal Imaging Using Adaptive Optics 49

in this case it is because they are transparent, a necessary requirement for tissue that lies anterior to the photoreceptors. Successful imaging of the broad range of structures in the retina will likely demand the marriage of AO ophthalmoscopy (either conventional or SLO) with other imaging and detection modalities, such as optical coherence tomography, differential interference contrast microscopy, or fluorescein angiography.

ALTERNATE TECHNOLOGIES

Developing and expanding the scope of the use of AO for basic and clinical investigations will require that AO technology become simpler to use, more compact, and less expensive.

The DM is the most effective technology for AO, but these devices are large and expensive and an economical, smaller version of DM technology is not likely. However, alternate technologies are on the horizon. For more details on these alternate technologies, the reader is referred to some recent publica- tions.29-31 For vision science, the best choice for an alternate wavefront corrector may depend on the particular imaging application. For example, the requirement of using polarized light for liquid crystal wavefront correcting devices might be a drawback for some applications but an advantage for others. Economical wavefront correctors such as phase plates32 or membrane mirrors,33 which have limitations, might be sufficient for many applications. On the other hand, for the best imaging possible, high temporal bandwidth to compensate for small eye movements, tear film changes, and accommodative fluctuations will be necessary.34

The most promising new technologies are in Micro Electrical Mechanical Systems, or MEMS. The manufacturing process, similar to that used for computer chips, can be done inexpensively and in mass quantities once a suitable system has been designed. The first MEMS DM for ophthalmic applications was demonstrated in 2002 by Doble et al,35 where they corrected an eye over a 6 mm pupil down to 0.12 µm. Although the level of improvement did not reach performance of the Xinetics DM, the results were promising and efforts to make MEMS mirrors that are suitable for vision applications are continuing in several laboratories.

BEYOND IMAGING

Since AO can be used to image the retina with high spatial resolution, it follows that AO can be used to deliver light to the retina with the same precision. This opens a number of possibilities that range from studying the perception of aberration-free retinal images to realizing the potential for pinpoint laser treatment of the retina.

Prior to the development of AO for vision applications, studies of the perception of high spatial frequency retinal images could only be done by producing interference fringes on the retina.36,37 This was and remains a very useful and productive technique, but the complexity of the retinal image is limited to sinusoidal gratings. Using AO, any complex near aberration-free image can be projected on the retina. Projection of such images is already being used to test the potential benefits of aberration-cor-

recting refractive surgical techniques.38 While there are some obvious benefits to vision, such as an improvement in contrast sensitivity, it remains to be seen whether or not hyperacuity tasks might be compromised by improving retinal image quality. Delivering small spots of color to the single photoreceptor cells might also be used to learn about the early stages of color processing in the human retina.

For clinical applications, laser systems can be equipped with AO to potentially pinpoint the treatment of features as small as individual capillaries or single cells. Using this technology, laser treatments like photocoagulation or photodynamic therapy can be restricted to a localized region, thereby preserving neighboring functional tissue in the retina. Finally, a real-time high resolution image offers the opportunity to do eye tracking measurements with unprecedented resolution.

CONCLUSION

With new technologies on the horizon and a host of new scientists and companies developing their own AO programs, the future promises to be exciting and productive for years to come.

AO in ophthalmoscopy is still a young field. Both the technology and the ideas of how to apply it are still developing. With new technologies on the horizon and a host of new scientists and companies developing their own AO programs, the future promises to be exciting and productive for years to come.

ACKNOWLEDGMENTS

Preparation of this chapter was supported by the following grants: NSF "Center for Adaptive Optics" a Science and Technology Center managed by the University of California at Santa Cruz under cooperative agreement number AST-9876783 to AR and DRW, NIH/NEI grant RO1 EY13299 to AR and NIH/NEI grant RO1 EY04367.

REFERENCES

1.Young T. On the mechanism of the eye. Phil Trans Roy Soc London. 1801;91:23-88.

2.Helmholtz H. Helmholtz’s Treatise on Physiological Optics.

Washington, DC: Optical Society of America; 1924.

3.Snyder AW, Miller WH. Photoreceptor diameter and spacing for highest resolving power. J Opt Soc Am. 1977;67:696-698.

4.Liang J, Grimm B, Goelz S, Bille JF. Objective measurement of wave aberrations of the human eye with use of a Hartmann-Shack wavefront sensor. J Opt Soc Am A. 1994;11:1949-1957.

5.Liang J, Williams DR. Aberrations and retinal image quality of the normal human eye. J Opt Soc Am A. 1997;14:2873-2883.

6.Liang J, Williams DR, Miller D. Supernormal vision and high-reso- lution retinal imaging through adaptive optics. J Opt Soc Am A. 1997;14:2884-2892.

7.Babcock HW. The possibility of compensating astronomical seeing. Pub Astr Soc Pac. 1953;65:229-236.

8.Hardy JW, Lefebvre JE, Koliopoulis CL. Real-time atmospheric compensation. J Opt Soc Am. 1977;67:360-369.

9.Benedict R, Breckinridge JB, Fried DL. Atmospheric-compensation technology: introduction. J Opt Soc Am A. 1994;11:257-262.

10.Dreher AW, Bille JF, Weinreb RN. Active optical depth resolution improvement of the laser tomographic scanner. Appl Opt. 1989;28: 804-808.

11.Bille JF, Dreher AW, Zinser G. Scanning laser tomography of the living human eye. In: Masters BR, ed. Noninvasive Diagnostic Techniques in Ophthalmology. New York: Springer-Verlag; 1990:528-547.

12.Campbell FW, Gubisch RW. Optical quality of the human eye. J Physiol. 1966;186:558-578.

13.Hofer H, Chen L, Yoon G, Singer B, Yamauchi Y, Williams DR. Improvement in retinal image quality with dynamic correction of the eye’s aberrations. Optics Express. 2001;8:631-643.

14.Webb RH, Hughes GW, Pomerantzeff O. Flying spot TV ophthalmoscope. Appl Opt. 1980;19:2991-2997.

15.Roorda A, Romero-Borja F, Donnelly WJ, Queener H, Hebert TJ, Campbell MCW. Adaptive optics scanning laser ophthalmoscopy. Optics Express. 2002;10:405-412.

16.Roorda A, Williams DR. The arrangement of the three cone classes in the living human eye. Nature. 1999;397:520-522.

17.Miller D, Williams DR, Morris GM, Liang J. Images of cone photoreceptors in the living human eye. Vision Res. 1996;36:1067-1079.

18.Roorda A, Campbell MCW. Confocal scanning laser ophthalmoscope for real-time photoreceptor imaging in the human eye. Vision Science and its Applications: Technical Digest. 1997;1:90-93.

19.Wade AR, Fitzke FW. In vivo imaging of the human cone-photore- ceptor mosaic using a confocal laser scanning ophthalmoscope.

Lasers and Light in Ophthalmology. 1998;8:129-136.

20.Marcos S, Navarro R, Artal P. Coherent imaging of the cone mosaic in the living human eye. J Opt Soc Am A. 1996;13:897-905.

21.Young T. On the theory of light and colours. Phil Trans Roy Soc London. 1802;91:12-48.

22.Rushton WAH, Baker HD. Red/green sensitivity in normal vision. Vision Res. 1964;4:75-85.

23.MacLeod DIA. Directionally selective light adaptation: a visual consequence of receptor disarray? Vision Res. 1974;14:369-378.

24.Marcos S, Burns SA. Cone spacing and waveguide properties from cone directionality measurements. J Opt Soc Am A. 1999;16:995-1004.

25.van Blokland GJ. Directionality and alignment of the foveal receptors, assessed with light scattered from the human fundus in vivo. Vision Res. 1986;26:495-500.

26.Burns SA, Wu S, Chang He J, Elsner AE. Variations in photoreceptor directionality across the central retina. J Opt Soc Am A. 1996;14:20332040.

27.Roorda A, Williams DR. Optical fiber properties of individual human cones. Journal of Vision. 2002;2:404-412.

28.Curcio CA, Sloan KR, Kalina RE, Hendrickson AE. Human photoreceptor topography. J Comp Neurol. 1990;292:497-523.

29.Tyson RK. Principle of Adaptive Optics. 2nd ed. San Diego, Calif: Academic Press; 1998.

30.Tyson RK, ed. Adaptive Optics Engineering Handbook. New York: Marcel Dekker; 2000.

31.Séchaud M. Wave-front compensation devices. In: Roddier F, ed. Adaptive Optics in Astronomy. Cambridge: Cambridge University Press; 1999:57-91.

32.Burns SA, Marcos S, Elsner AE, Bará S. Contrast improvement for confocal retinal imaging using phase correcting plates. Optics Letters. 2002;27:400-402.

33.Fernández EJ, Iglesias I, Artal P. Closed-loop adaptive optics in the human eye. Optics Letters. 2001;26:746-748.

34.Hofer H, Artal P, Aragon JL, Williams DR. Dynamics of the eye’s wave aberration. J Opt Soc Am A. 2001;18:497-506.

35.Doble N, Yoon G, Bierden P, Chen L, Olivier S, Williams DR. Use of a microelectromechanical mirror for adaptive optics in the human eye. Optics Letters. 2002;27:1537-1539.

36.Westheimer G. Modulation thresholds for sinusoidal light distributions on the retina. J Physiol. 1960;152:67-74.

37.Arnulf MA, Dupuy MO. La transmission des contrastes par le système optique de l’oeil et les seuils de contrastes retinines. C R Acad Sci (Paris). 1960;250:2757-2759.

38.Yoon G, Cox I, Williams DR. The visual benefit of static correction of the monochromatic wave aberration. Invest Ophthalmol Vis Sci. 1999;40(Suppl):40-40.

39.Curcio CA, Hendrickson A. Organization and development of the primate photoreceptor mosaic. In: Osborne NN, Chader GJ, eds. Progress in Retinal Research. Oxford: Pergamon; 1991:89-120.

40.Kolb H, Linberg KA, Fisher SK. Neurons of the human retina: a Golgi study. J Comp Neurol. 1992;318:147-187.

41.Zinn KM, Benjamin-Henkind JV. Anatomy of the human retinal pigment epithelium. In: Zinn KM, Marmor MF, eds. The Retinal Pigment Epithelium. Cambridge: Harvard University Press; 1979:3-31.

42.Snodderly DM, Weinhaus RS, Choi JC. Neural-vascular relationships in central retina of Macaque monkeys (Macaca fascicularis). J Neuroscience. 1992;12:1169-1193.

Contemporary visual optics research is changing our mindset, our way of thinking about the optical system of the eye, and in the process is redefining the field of visual optics. In the past, optical imperfections of the eye were conceived as simple refractive errors—defocus, astigmatism, and perhaps a bit of prism. Although clinical students learned about other kinds of optical imperfections, such as spherical aberration, coma, oblique astigmatism, and the other Seidel aberrations, those concepts were confined to courses in optical theory, not to clinical practice. This is for good reason: these higher-order aberrations of the eye could not be measured routinely in the clinic, and even if they could, we did not have the means to correct them optically at a reasonable cost to patients. Furthermore, since the effects of such aberrations on visual function were largely unknown, there was little reason to suppose that correcting them would do any good for the patient’s vision. However, the introduction of laser refractive surgery, with its potential for removing as well as introducing unwanted optical aberrations into the eye, demands changes in established ways of thinking and answers to these unresolved issues.

Today, optical imperfections of the eye are being re-examined within a comprehensive theoretical framework that expresses the combined effect of all the eye’s aberrations in a two-dimensional aberration map of the pupil plane. An aberration map is similar in concept to corneal topographic maps used to describe the corneal surface. The major difference is that a corneal map describes the curvature of a physical surface, whereas an aberration map describes the difference between a wavefront of light and a reference wavefront. By concentrating our attention on light instead of the refracting surface, we gain an ability to compute image quality on the retina for simple points of light, for clinical test targets, or for any complex object in the real world. For example, Figure 6-1 shows a wavefront aberration map for a defocused eye from which the retinal image of an acuity chart may be computed. Such computations are poised to become routine clinical tools of the future for predicting the visual benefit of aberration correction to the patient, and for explaining the risks and visual consequences of unintended increases in optical aberrations following refractive surgery or other forms of treatment.

An aberration map is similar in concept to corneal topographic maps used to describe the corneal surface.

Customized corneal ablation is a surgical procedure designed to improve the optical quality of the eye, thereby improving vision. To assess the outcome of this procedure requires measures of the direct effect on retinal image quality and secondary effects on visual performance and the quality of visual experience. A variety of methods for specifying optical quality are well established in the field of optics and may be readily applied to the optical image on the retina. Similarly, a variety of visual performance measures are sensitive to the optical quality of the retinal image and therefore may be used to assess the effect of refractive surgery on vision. However, optical limits normally imposed by the eye’s optical aberrations may recede in the near future if refractive surgery, contact lenses, or intraocular lenses (IOLs) improve retinal image quality beyond limits imposed by the neural component of the visual system. If this occurs, then common measures of visual performance such as letter acuity, which are traditionally regarded as good measures of optical quality of the retinal image, may no longer be optically limited. When this happens, visual performance will be limited instead by the neural architecture and physiology of the retina and visual brain, thereby generating a demand for new measures of vision that are sensitive to even the smallest departures from perfect retinal image quality.

MEASURES OF OPTICAL QUALITY

The quality of an optical system may be specified in three different, but related, ways. The first method is to describe the detailed shape of the image for a simple geometrical object such as a point of light, or a line. The distribution of light in the image plane is called a point spread function (PSF) for a point object or a line spread function (LSF) for a line object. Simple measurements derived from these functions, such as the width (blur circle diameter) or height (Strehl ratio) of the intensity distribution, are taken as figures of merit that capture the blurring effects of optical imperfections.

The second method is a description of the loss of contrast suffered when an image of a sinusoidal grating object is cast. The sinusoidal grating is a very special object in optics because it has the unique property of producing images of the same form. In other words, a sinusoidal grating object forms sinusoidal images with the same spatial frequency (expressed in cycles/degree [c/deg]) and the same orientation.