Adaptive optics for precision_

Document Sample
Adaptive optics for precision_ Powered By Docstoc
					                                Adaptive optics for precision,
                                   guided, laser surgery
                                        D. G. Cole and A. Melamud
                                              Duke University

Research Plan
The overall goal of this project is to define the clinical, histologic, angiographic, and functional effects of an
adaptive optics, guided laser system versus conventional laser treatment. To meet this goal, we will develop an
adaptive optics, guided, precision laser surgery system. This system will incorporate adaptive optics in order to
better visualize retinal structures, more precisely deliver laser energy to target tissue, and improve the surgeon's
control of the laser delivery system. Aberations of the eye can be compensated for using adaptive optics
{Unknown}. This permits better visualization of the retina and associated tissues providing the diagnostician
and surgeon improved information about retina pathology and health. Furthermore, we hypothesize that by
accounting for these aberations the dimensions and characteristics of the laser spot can be more accurately
controlled, allowing for increase precision in the application of laser energy.

A.      Specific Aims
Diabetic retinopathy and exudative age related macular degeneration are leading causes of irreversible vision
loss in the United States. In both of these diseases, the main pathological entity is abnormal vascular tissue that
can proliferate and lead to bleeding, fibrosis, scarring and permanent loss of vision. Laser treatment of
neovascular lesions is the standard of care in both of these diseases. The goal of laser treatment is to destroy the
neovascular tissue and promote healing. This is accomplished by photocoagulating the abnormal vessels with
the goal of sparing as much surrounding tissue as possible. The current state of the art in retinal laser surgery is
argon laser photocoagulation. In this technique, laser energy is focused by the surgeon on the patient’s retina
through a dilated pupil using conventional slit lamp optics. Added magnification and some stabilization of the
eye are provided by a contact lens that the surgeon holds apposed to the patient’s eye during the procedure.
There are a number of limitations of current laser technology. Slit lamp optics, with contact lens magnification
does not allow the surgeon to visually distinguish the exact boundaries of the abnormal vascular lesion from the
surrounding healthy tissue. This is true for identifying both the width and depth of the pathologic lesion.
Moreover, when applying laser energy, the surgeon is limited by an ablation width of 50 um, a limit imposed by
the surgical optics and aberations of the eye. This creates a spot size that is larger than many targeted lesions.
Thermal coagulation, applied at an energy level that is considered therapeutic, penetrates through multiple layers
of the retina above and below the targeted tissue, causing destruction of viable tissue. Additional imprecision in
the surgical system is caused by distortion in the optics of eye caused by the force of the contact lens held by the
surgeon in apposition against the patient’s cornea. Finally, even though some stability may be achieved by
anesthetizing the extra-ocular muscles with a retrobulbar injection (a procedure that carries significant risks) it is
usually impossible to completely neutralize all movement of the patient’s eye during the procedure.
A number of complications of laser therapy have been described in the literature {Lovestam-Adrian2000,
Shartz1991, Rutledge 93, Lewis90, Varley 1988}. These include loss of central visual acuity, diminished
contrast sensitivity, decreased color vision, loss of peripheral visual field and others. Because the retina is a
delicate neural tissue, packed with multiple photoreceptors per square micron, any thermal coagulation of viable
structures outside the target area results in measurable vision loss. While accomplishing the goal of destroying

abnormal vascular tissue, laser photocoagulation therapy itself causes loss of normal photoreceptors and other
retinal elements responsible for vision. This side effect of treatment can have devastating effects on the
patient’s quality of life.
        There is clearly a need for a more precise, better controlled laser delivery system; a system that
        would enable the surgeon to identify the exact cellular boundaries of the target tissue and apply
        laser energy to coagulate only the abnormal vessel and not the surrounding healthy tissue.
        Such a system would achieve the destruction of pathologic tissue with maximum precision and
The PI’s in this study bring together expertise in clinical, biological, and engineering research to develop an
adaptive optics guided laser delivery system that will achieve the goal of ablating retinal tissue with minimal
damage to surrounding structures. The proposed system will enable the surgeon to visualize retinal structures at
the cellular level. Using adaptive optics technology to correct for the eye's aberations, focused laser energy will
be more easily directed and controlled. It will be possible to minimize the width of the ablation zone to the
diameter of XXX um to an ablation depth of XXX um. This system will also be equipped with the ability to
track eye movements enhancing laser precision, and potentially obviating the need to administer retrobulbar
anesthesia to paralyze the extra-ocular muscles. There is no laser delivery system reported in the literature that
is able to focus laser energy and enhance surgical precision to the degree proposed in this study.
For the successful completion of this research, we propose the following specific aims:

Develop an Adaptive Optics, Guided Laser, Surgical System.
To develop an automated, adaptive optics, surgical laser system that exceeds the state of the art in retinal laser
surgery by:
        a) enabling the surgeon to visualize and focus laser energy on retinal structures XXX um in diameter;
        b) enabling the surgeon to contain thermal energy to a depth of XXX um;
        c) enhancing surgical precision by utilizing gaze tracking to stabilize the laser delivery system.

Determine the Effectiveness of the Surgical Laser System for Precision Surgery
      We propose to conduct studies to evaluate the efficacy, and safety of our surgical system in an animal
   model. The pigmented rabbit is a commonly used animal model in ophthalmic research. The rabbit eye
   offers a large size and optical characteristics that make it a suitable model for laser surgery. (Blankenship
       In this study we hope to accomplish the goal of understanding the capabilities of an adaptive optics
   system in targeting specific retinal structures. The structures of particular interest in clinical ophthalmology
   are micro-vascular elements within the retina and choroid. To this effect, we will investigate the ability of
   the surgical laser system to ablate targeted micro-vascular structures such as arterioles, venules and
   capillaries. Our goals will be twofold: to characterize the tissue response of the micro-vascular structures at
   different laser energy levels and wavelengths, and to evaluate tissue damage to retinal elements surrounding
   the targeted structure. We plan to characterize the short and long term effects of applied laser energy on
   retinal tissue.
      To accomplish our goals we will perform pre-operative assessment of retinal structure and function using
   standard clinical techniques such as elecroretinography, optical coherence tomography, fluorescein
   angiography and photography. Laser surgery to select structures in the rabbit retina and choroid will be
   performed. Animals will then be re-evaluated at specific post-operative time points with the same diagnostic

     techniques used pre-operatively in addition to careful histological analysis of enucleated eyes.

Application of the Novel Laser System in Treatment of Diabetic Macular Edema in
  Human Patients.
To test in human diabetic patients:
         a) the effect of photocoagulation of micro-aneurysms and ablation of selected areas of the retinal
            pigment epithelium on decreasing focal diabetic macular edema;
         b) the effect of grid photocoagulation on decreasing diffuse diabetic macular edema;
         c) to evaluate the effectiveness of the novel laser system, as compared to conventional treatment, to
            improve or stabilize visual acuity in patients with macular edema.

Application of the Novel Laser System in Treatment of Choroidal Neovascular
  Membranes in Age Related Macular Degeneration.
To test in human patients with wet age related macular degeneration:
         a) the effectiveness of focused laser treatment to destroy and prevent further growth of choroidal
            neovascular membranes;
         b) to evaluate the effectiveness of the novel laser system, as compared to conventional treatment, to
            improve or stabilize visual acuity in patients with choroidal neovascular membranes.

B.       Background and Significance

Historical perspective
Laser photocoagulation surgery was introduced to ophthalmology over 40 years ago. From its inception, the
goal of focusing laser energy on the retina was to cause thermal damage to selected ocular structures. Retinal
photocoagulation works by producing heat generated by the absorption of light by ocular pigments. Heat is
generated in the tissue that absorbs light directly and in the surrounding tissue by thermal conduction. Thermal
damage results from the biochemical changes that take place when ocular tissues are heated to temperatures high
enough to denature proteins. Usually a temperature increase of 10 to 20 C is sufficient to produce the desired
biochemical changes.
Laser surgery to the retina is the standard of care in the treatment of numerous ophthalmic diseases. Diseases
treated by laser photocoagulation include proliferative diabetic retinopathy, diabetic macular edema {ETDRS-
85, ETDRS-87}, cystoid macular edema, retinal vein occlusion, choroidal neovascularization, central serous
chorioretinopathy, retinal tears, and other lesions. In using photocoagulation therapy, the clinician attempts to
visualize the pathologic lesion and apply treatment directly over the lesion or adjacent to the lesion. Imaging
modalities such as fundus photography, fluorescein angiography and idiocyanide green angiography are often
used to help evaluate the retina prior to surgery and to plan the procedure.

Laser Surgery
Ophthalmologists have several parameters at their control when photocoagulating. These parameters include
exposure time, power, spot size, and sometimes wavelength (color). Different settings are used for
photocoagulation of different lesions. The most important principle in laser surgery is to achieve
photocoagulation of the desired lesion with minimal damage to surrounding structures.

Great precision and accuracy is needed in any laser surgery but particularly in the treatment of pathology located
in the macular area, the area responsible for central vision in humans. Laser photocoagulation in the macula,
close to the fovea, can produce irreversible vision loss that is debilitating to patients. The precision that can be
achieved in laser retinal surgery depends, in part, on the skill of the operator but also, in great proportion, on the
technology that the surgeon has at his disposal. First and foremost, the surgeon must be able to clearly visualize
the lesion. In practical terms this often means visualizing structures such as micro-aneurysms, capillaries, and
vascular membranes that are microns in diameter {Stitt95}. The optics of slit-lamp biomicroscope in
conjunction with contact lenses designed for laser surgery provide a limited level of magnification for
visualizing retinal structures {Dieckert84}. Structures that cannot be visualized with conventional optics
technology include individual cells within different layers of the retina. Pathological structures such as micro-
aneurysms are visualized as small red lesions during photocoagulation surgery. The exact cellular boundaries of
the micro-aneurysm separating it from the surrounding tissue cannot be distinguished with conventional optics.
A second limitation to the precision achievable in conventional photocoagulation techniques is the spot size, or
diameter, of the laser beam itself. Probably the smallest diameter spot size used in clinical practice was
evaluated by the Early Treatment of Diabetic Retinopathy study (ETDRS) {ETDRS-87}. This study applied
focal laser photocoagulation to macular micro-aneurysms using a 50–200 um spot size. It must be kept in mind
that most commercially available laser surgery systems make use of a contact lens with a flat front surface that
eliminates the focusing power of the first corneal surface of the eye. If the lens is not held with the flat surface
perpendicular to the laser beam, the laser spot size on the retina may be even larger than that indicated by the
photocoagulation system. However, even an ideal 50 um spot size will likely be larger than many diabetic
micro-aneurysms. Histopathological studies of micro-aneurysms prove that these lesions may be 30–40 um in
diameter or less {Venecia76, Engerman84}. Moreover a micro aneurysm is never a perfectly round lesion.
Conventional macular photocoagulation applied with the goal of ablating the micro-aneurysm will inevitable
destroy surrounding healthy tissue.
The depth of the photocoagulation effect is similarly a factor that influences the precision of laser treatment.
Laser energy applied directly to a vessel in the inner retina creates a pyramid shaped lesion {Apple73}. The
apex of this lesion is located at the entry site of the beam into the vessel, the base of the pyramid is situated
along the plane defined by the retinal pigment epithelium (RPE) and choriocapillaris. Current laser technology
does not allow the operator to focus the laser beam within one particular cell layer of the retina. In fact,
conventional laser treatment applied with the goal of cauterizing retinal vessels, achieves greater tissue
disruption at the level of the choroid and the RPE than the retinal vessel itself {Tso77}. This limits the amount
of energy that can be used during photocoagulation surgery of microvascular structures. High levels of energy
can create full-thickness retinal necrosis, chorioretinal anastamosis, and localized areas of retinal detachment.
Severe lesions can weaken the retinochoroidal interface, promoting the growth of chorioretinal or
chorioretinovitreal neovascularization {Wallow85}.
Finally, precision laser surgery is affected by eye movement and operator stability. Focal macular
photocoagulation is usually performed with the aid of a contact lens that the surgeon holds against the patient’s
eye throughout the procedure. This minimizes most eye movement, but small movements are still possible.
Likewise, operator stability in guiding the laser beam to the target area plays a role in the process. The force of
the contact lens held against the eye by the surgeon can create distortions in the eye and additional aberations in
the eye's optics. These aberations further affect the precision of the laser surgery process.
Some of the most important causes of blindness in the United States, such as diabetes mellitus and age related
macular degeneration (AMD), present with vascular lesions located in close proximity to the fovea. In the case
of AMD, the pathological lesion is a choroidal neovascular membrane. These membranes are classified based
on their proximity to the foveal center. Thus lesions located more than 200 um from the foveal center are
termed extrafoveal; those 1–199 um from the center are juxtafoveal; and CNV (???) lesions extending under the
center of the foveal avascular zone are termed subfoveal. Laser treatment to neovascular membranes in AMD
was found to be superior to observation by the Macular Photocoagulation Study trials. Among these trials was a

study looking at the use of krypton red laser (647 nm, 676 nm) treatment to juxtafoveal neovascular membranes
{Guyer86}. In this study, the persistence of neovascular membranes in eyes having 10% or more of the foveal
side of the lesion not covered by treatment was twice as high as in the eyes with more extensive treatment
coverage {MPSG-95}. Reluctance to treat within such a close proximity to the foveal avascular zone and
reduced visibility of retinal landmarks probably both contribute to the limited success of laser treatment.
The potential for devising a system combining the elements of laser delivery with adaptive optics has been
suggested elsewhere {Roorda00}. To our knowledge however, no one has published any studies looking at the
application of such a devise to retinal photocoagulation surgery. We believe that a combination of adaptive
optics technology with laser technology can lead to several advantages in retinal laser surgery. This includes
better visualization of retinal structures (normal and pathological); more precise delivery of laser energy to
target tissue; decreased laser damage to surrounding tissue; and increased operator control of laser delivery

Adaptive Optics, Imaging, and Laser Delivery
The limit for both imaging and laser delivery is determined by diffraction. The Rayleigh criterion defines this
limit by

                                                  d 0.61
For the a typical dilated eye (0.25 NA, wavelength 550 nm) this limit is ~1.3 um. This is the smallest laser spot
size that can be achieved, and using conventional imaging systems objects smaller than this limit cannot be
discriminated. However, the eye is far from being a diffraction limited system {Navarro98}. Accordingly, the
limits of retina imaging and laser delivery systems are defined by the aberations of the eye. In addition to
defocus and astigmatism normal eyes are known to suffer from spherical aberations, coma, and as well as other
higher-order aberations {Unknown}. Furthermore, these aberations are more pronounced when the pupil is
large (> 3 mm) {Liang97}. Clearly, for increased resolution and laser surgery precision a larger pupil is desired
in order to maximize the numerical aperture, but this comes with the cost of increased aberation {Unknown}.

Aberations & Adaptive Optics
Liang et al. {Liang94} demonstrated a technique for measuring the eye's aberations using a Shack-Hartmann
sensor. The Shack-Hartmann sensor is the state of the art in wavefront sensing, but other wavefront sensor
designs exist {Unknown} and may be more attractive based on cost and ease of implementation. The Shack-
Hartmann measurements are used to determine the appropriate contrubtions of the various types of aberations,
described by Zernike modes, using a modal method{Unknown}. Using a double-pass technique {Unknown}, a
laser beam is focused onto the retina and reflected light from the retina forms an aberrated wavefront at the exit
pupil of the eye. The wavefront distortion is measured by the wavefront sensor, and the wavefront is
reconstructed as a linear combination of Zernike modes.

Figure XXX shows grayscale images of the various Zernike modes, and the mode's order corresponds to the
radial power of the corresponding Zernike polynomial. Orders one through four correspond to the the so called
primary aberrations. The first-order modes are linear corresponding to tilt and are often ignored. The second-
order modes are quadratic and characterize the common aberrations of defocus and astigmatism. The third-
order modes correspond to coma and trefoil aberrations. The fourth-order modes contain primary sperical,
secondary astigmatism, and tetrafoil aberrations. Fifth and higher-order modes contain so called irregular
aberrations. These irregular aberrations typically play little role in vision for pupil diameters less than ~3 mm
but are very significant for large diameter pupils ~7 mm, which is the case for precision retina imaging and laser

     Figure 2 source {Liang97}

                                   Liang et al. {Liang97} evaluated the aberrations and retinal image quality of
                                   normal human eyes for 3.4 mm and 7.3 mm diameter pupils for multiple
                                   subjects (12 and 14 respectively). Figure XXX shows the contribution to the
                                   wavefront error for each modal order (a given order has multiple modes)
                                   averaged over all subjects for the small and large diameter pupil cases. The
RMS error for the small pupil is ~3–4 times lower than the large diameter
case illustrating the fact that aberrations increase with pupil size. The         Figure 1 source {Liang97}
Maréchal criteria provides a commonly used specification for diffraction
limited optics, requiring the RMS wavefront error to be less than       /14
(0.038 um at 533 nm (green)); this criteria corresponds to a Strehl ratio of
0.8. Figure XXX shows, for small and larger diameter pupils, the Strehl
ratio that results if all aberrations upto and including a certian order are
removed (compensated for). The key result is that for small diameter
pupils compensation of second and third order aberrations is sufficient for
diffraction limited vision, making conventional correction of sphere and
cylinder sufficient for most vision requirements. However, for large
diameter pupils as many as eight orders (corresponding to 44 modes)
must be accounted for before achieving diffraction limited optics. More

importantly, the higher-order, irregular aberrations reduce retinal contrast when the pupil is large, e.g.,
aberrations beyond defocus and astigmatism can reduce retinal image contrast by as much as a factor of 7 at 20
cycles/deg, an effect that is particularly limits the ability to image the retina at high resolution {Liang97}. In
order to achieve diffraction limited performance for large pupil diamters, techniques such as adaptive optics can
be used to correct for both primary and irregular aberrations.
The Shack-Hartmann has been combined with active optical elements (AOE's), e.g., deformable mirror
{Unknown} or spatial light modulator {Unknown}, to create an adaptive optics system that can correct for the
total wave aberation of the eye. Correcting AOE's modify the wavefront according to the measured aberations.
Such systems can easily compensate for the large variations in aberations from eye to eye. Liang et al.
{Liang97b} investigated super-normal vision and high-resolution using adaptive optics. They showed that in
natural daylight conditions for normal eyes the pupil diameter is small, ~3 mm, and aberrations beyond defocus
and astigmatism are negligible. However, correction of high-order, irregular aberrations using a deformable
mirror provides the greatest benefit when the pupil is large. They showed that eyes with adaptive compensation
could resolve fine gratings (55 cycles/deg) that were invisible under normal viewing conditions. This result is
significant for imaging of the retina. Imaging systems will be able to resolve fine features of the retina, but only
for large pupil diameters and only if the irregular aberrations are removed. In particular, Liang et al.
{Liang97b} point out that with adaptive compensation the width (FWHH) of the point spread function (PSF) is
often smaller than a single foveal cone. Their adaptive optics fundus camera provided unprecedented resolution,
so that the living retina could be imaged at a scale previously only possible for excised tissue. For an 8 mm
pupil and 555 nm (green) light, one could in principle achieve a 1.18 um PSF (FWHH) and a 3.2-fold increase
in transverse resolution over that for a 2.5 mm pupil, a common pupil diameter for current fundus imaging.
Furthermore, the axial resolution, critical for optical sectioning and discriminating retina layers, increases as the
square of the pupil diameter. For an AO corrected 8 mm pupil, the depth of field of the PSF would be ~30 um
(FWHH) approaching OCT but with high transverse resolution and rapid imaging. The application of AO to
confocal scanning laser ophthalmoscopy (CSLO) is obvious and could provide a necessary tool for diagnosing
retina pathology and health. This level of precision applies to laser delivery as well, making precise application
of laser energy possible with accuracy an order of magnitude greater than currently possible.
Deformable mirrors (DM) have been the standard in adaptive optics systems. An array of actuators, typically
50–100 channels, is used to deform a flexible mirror, the shape of which is chosen to compensate for the
aberations introduced by the eye. The cost of DM's is on the order of $1000 per channel, limiting such devices
to astronomy and military applications. Another drawback with DM's is that they are often larger than the pupil
requiring large optical path lengths to match their size to the eye's pupil. Spatial light modulators (SLM's), on
the other hand, provide a cheaper alternative to DM's. They are lower cost ($5k–20k), provide more channels
(~1,000,000), and better match the pupil size.
The eye's aberations and point spread function are not static and fluctuate with time at frequencies ~1–10 Hz.
For example, during steady-state accomodation, the eye's focus fluctuates about its mean ~0.3–0.5 D at
frequencies up to 5 Hz {Hofer01, Charman88}. While these and higher-order fluctuations are large enough to
affect retinal image quality, they probably do not produce a perceptible effect on vision under normal
circumstances {Charman88}. However, these fluctuations could adversely affect high quality images of the
retina and would cause unwanted spreading of focuses laser energy during surgery. Using feedback control, an
adaptive optics system can track aberations in real-time. The process of measurement and correction is repeated
at video frequecies (~30}{\hertz}) enabling tracking of these fluctuation {Hofer01}.

Improving the eye's optical quality will also improve the resolution of fundus images. These improvements
could be invaluable for clinical diagnosis and the timely and accurate identification of pathologic lesions.
Current fundus cameras including scanning laser ophthalmoscopes, while well suited to capture macroscopic
retinal features, do not resolve retinal structures as small as single cells. The smallest cones resolved using a

high resolution fundus camera had a spacing of 3.0 um corresponding to a cone mosaic sampling frequency of
~100 c/deg {Miller96}. Scanning laser ophthalmoscopes (SLO) {Plesch87} provide advantages for retinal
imaging because of it ability to improve contrast {Elsner96}, and confocal versions {Webb87} eliminate of out-
of-focus information, and provide the ability to image thick specimens in sections; these sections can be
combined to reconstruct a 3D image with remarkable quality. However, fundus imaging techniques use a pupil
diameter ~2–3 mm. In a tradeoff between diffraction resolution and high-order aberations {Unknown}, this
diameter provides the best images. This limitation also restricts the axial resolution which varies quadratically
with the transverse resolution. In this case, a factor of two improvement in the scanning spot size will improve
the axial resolution by a factor of four. Increased resolution could be achieved for larger pupil diameters, but
higher order aberations, such as coma and spherical aberation, become the limiting factor {Unknown}.
Adaptive optics provides the best technique to compensate for these aberations. Roorda et al. {Roorda02}
demonstrated an adaptive optics scanning laser ophthalmoscope. This instrument increased resolution in lateral
and axial directions, and could visualise photoreceptors, nerve fibers, and the flow of blood cells in retinal

Laser Delivery and Surgery Systems
Researchers at the University of Texas at Austin and the U.S. Air Force Academy worked for the development
of a robotic retinal laser surgery system {Barrett95, Barrett96, Barrett97, Oberg96, Oberg97}. Their system
contained several features that would be requisite for any laser surgery system. Imaging in this system was
achieved using a CCD camera in a conventional fundus photography setup; the digitized were used for retina
(eye) tracking. lesion placement subsystem tracked a specific lesion coordinates on the fundus using blood
vessel tracking templates. In order to control the laser energy delivered for photocoagulation, a lesion depth
control subsystem was developed. This subsystem used lesion reflectance during photocoagulation as a measure
of lesion size (depth), stopping treatment when a pre-defined level of reflectance was reached {Jerath94,
Eye movements have a small effect on the wavefront measurements, because the shift of pupil position is small
for the fixating eye. For example, a relatively large fixational eye movement of 10 arc min produces a pupilary
displacement less than 35 um, a small fraction of the spacing on the wavefront sensor. The important aberations
correspond to gradual enough variations in phase across the pupil that normal fixational eye movements are not
a problem {Liang97}.

C.      Preliminary Studies

Research Outline and Work Tasks

Adaptive Optics, Precision, Guided Laser System

Adaptive Optics Subsystem

Scanning Confocal Imaging Subsystem

Beam Steering and Lesion Placement Subsystem
Markow93 developed a retinal tracking system used in a lesion placement system.

Eye/Retina Tracking Subsystem
W. Kosnik, J. Fikre, and R. Sekuler, Visual fixation stability in older adults, Invest. Ophthalmol. Vis. Sci., 27,
1720-1725, 1986.

Wright et al. (1996) use a quadrature technique similar to that used for single molecule fluorescence tracking.

Research Design and Methods:

To define the clinical, histologic, angiographic and functional effects of adaptive optics guided laser (AOGL)
versus conventional argon laser treatment on the rabbit retina.
Pigmented rabbits will be used in this study. (Is this the best animal model) Each animal will undergo retinal
laser photocoagulation surgery using the AOGL system to both eyes.
AOGL will be used to deliver focused laser to discrete retinal and sub-retinal cell layers of the rabbit right eye.
Spot size, burn duration, power and wavelength will be adjusted as necessary. We will photographically record
the location and clinical appearance of the laser burns at different energy levels.
Laser settings will be varied in a stepwise fashion by adjusting parameters of spot size, power, wavelength and
duration. how to vary? Treatment spots will be applied to the retina, retinal vessels, RPE and choroid.
A detailed map of the areas where laser was applied will be kept for each animal.
Each rabbit will be evaluated using a battery of tests on two occasions: baseline measurements obtained before
the surgical procedure will be compared to those obtained at a specific post-operative time point. At baseline,
rabbits will be anesthetized with ketamine (50 mg/kg; 10 mg/kg/hr thereafter) and xylazine (5 mg/kg; 0.5
mg/kg/hr thereafter) and the pupils will be dilated with eyedrops (1% tropicamide; 2.5% phenylephrine), and the
corneal surface anesthetized with another eyedrop (0.5% proparacaine). After full pupil dilation, the status of
the retina and eye will be examined by several methods. Two of these are noninvasive (indirect
ophthalmoscopy followed by fundus photography). In addition, intraocular pressure (IOP) will be measured
using a Tonopen, a device that is used clinically and which makes minimal contact with the corneal surface.
The architecture of the retinal blood vessels will then be documented with fluorescein angiography. Fluorescein
angiography will consist of an intravenous bolus injection of 0.5 ml of 10% fluorescein sodium into an ear vein.
Using a standard fundus camera with appropriate excitation and barrier filters, the fluorescein pattern of the
rabbit fundus will be documented and stored on film.
At specific post-operative time points (immediately after treatment, 1 day, 1 week, 2 weeks, 1 month, 3 months),
5 rabbits will be anesthetized with ketamine (50 mg/kg; 10 mg/kg/hr thereafter) and xylazine (5 mg/kg; 0.5
mg/kg/hr thereafter) and the procedures performed at baseline will be repeated. At the completion of these tests,
and while still under anesthesia, the rabbit will be killed by using an intravenous dose of Beuthanasia D Special
(1ml/5 kg). Both eyes will then be enucleated and fixed in 10% buffered formalin for 24 hours, for histological
Histologic evaluation will include light microscopy and electron microscopy. (What is the best way to perform
histological analysis)

D.      Human Subjects
Not applicable.

E.      Vertebrate Animals

F.      Literature Cited

Apple, D.J.; Goldberg, M.F. & Wyhinny, G.
Histopathology and ultrastructure of the argon laser lesion in human retinal and choroidal
Am. J. Ophthalmol., 1973 , 75 , 595-609
Barrett, S.F.; Wright, C.H.G.; Jerath, M.R.; R. S. Lewis, I.; Dillard, B.C.; H. G. Rylander, I. & Welch,
Automated retinal robotic laser system
Biomed. Sci. Instrum., 1995 , 31 , 89-93
Barrett, S.F.; Wright, C.H.G.; Oberg, E.D.; Rockwell, B.A.; Cain, C.; H. G. Rylander, I. & Welch, A.J.
Digital integrated retinal surgical laser system
Biomed. Sci. Instrum., 1997 , 33 , 354-359
Barrett, S.F.; Wright, C.H.G.; Oberg, E.D.; Rockwell, B.A.; Cain, C.; H. G. Rylander, I. & Welch, A.J.
Development of an integrated automated retinal surgical laser system
Biomed. Sci. Instrum., 1996 , 32 , 215-224
Burns, S.A.; Marcos, S.; Elsner, A.E. & Bara, S.
Contrast improvement of confocal retinal imaging by use of phase-correcting plates
Optics Letters, 2002 , 27 , 400-402
Charman, W.N. & Heron, G.
Fluctuations in accommodation: a review
Ophthalmic Physiol. Opt., 1988 , 8 , 153-163
Dieckert, J.P.; Mainster, M.A. & Ho, P.C.
Contact lenses for laser applications
Ophthalmology, 1984 , Instrument and Book suppl. , 79-87
Elsner, A.E.; Burns, S.A.; Weiter, J.J. & Delori, F.C.
Infrared imaging of sub-retinal structures in the human ocular fundus
Vision Res., 1996 , 36 , 191-205
Engerman, R.L. & Kern, T.S.
Experimental galactosemia produces diabetic-like retinopathy
Diabetes, 1984 , 33 , 97-100
Guyer, D.R.; Fine, S.L.; Murphy, R.P. & Green, W.R.
Clinicopathologic correlation of krypton and argon laser photocoagulation in a patient with a subfoveal
choroidal neovascular membrane
Retina, 1986 , 6 , 157-163
Hofer, H.; Artal, P.; Singer, B.; Aragón, J.L. & Williams, D.R.
Dynamics of the eye's wave aberrations

J. Opt. Soc. Am. A, 2001 , 18 , 497-506
Hofer, H.; Chen, L.; Yoon, G.Y.; Singer, B.; Yamauchi, Y. & Williams, D.R.
Improvement in retinal image quality with dynamic correction of the eye's aberrations
Optics Express, 2001 , 8 , 631-643
Jerath, M.R.; Chundru, R.; Barrett, S.F.; H. G. Rylander, I. & Welch, A.J.
Preliminary results on reflectance feedback control of photocoagulation in vivo
IEEE Trans. Biomed. Eng., 1994 , 41 , 201-203
Liang, J.; Grimm, B.; Goelz, S. & Bille, J.F.
Objective measurement of wave aberations of the human eye with the use of a Hartmann-Shack
wave-front sensor
J. Opt. Soc. Am. A, 1994 , 11 , 1949-1957
Liang, J. & Williams, D.R.
Aberrations and retinal image quality of the normal human eye
J. Opt. Soc. Am. A, 1997 , 14 , 2873-2883
Liang, J.; Williams, D.R. & D. T. Miller
Supernormal vision and high-resolution retinal imaging through adaptive optics
J. Opt. Soc. Am. A, 1997 , 14 , 2884-2892
Ludwig, D.A.; Barrett, S.F. & Kubichek, R.F.
Laser dosimetry control for retinal surgery
Biomed. Sci. Instrum., 2001 , 37 , 479-484
The influence of treatment coverage on the visual acuity of eyes treated with krypton laser for
juxtafoveal choroidal newovasculariation
Arch. Ophthalmol., 1995 , 113 , 190-194
Miller, D.T.; Williams, D.R.; Morris, G.M. & Liang, J.
Images of cone photoreceptors in the living human eye
Vision Res., 1996 , 36 , 1067-1079
Navarro, R.; Moreno, E. & Dorronsoro, C.
Monochromatic aberations and point-spread functions of the human eye across the visual field
J. Opt. Soc. Am. A, 1998 , 15 , 2522-2529
Oberg, E.D.; Barrett, S.F. & Wright, C.H.G.
The development of a hybrid analog/digital retinal surgical laser system
Biomed. Sci. Instrum., 1997 , 34 , 224-228

Oberg, E.D.; Barrett, S.F. & Wright, C.H.G.
Development of an integrated automated retinal surgical laser system
Biomed. Sci. Instrum., 1996 , 32 , 215-224
Plesch, A.; Klingbeil, U. & Bille, J.
Digital laser scanning fundus camera
Appl. Optics, 1987 , 26 , 1480-1486
Roorda, A.
Adaptive optics ophthalmoscopy
J. Refractive Surgery, 2000 , 16 , S602-S607
Roorda, A.; Romero-Borja, F.; W. J. Donnelly, I. & Queener, H.
Adaptive optics scanning laser ophthalmoscopy
Optics Express, 2002 , 10 , 405-412
Stitt, A.W.; Gardiner, T.A. & Archer, D.B.
Histological and ultrastructural investigation of retinal microaneurysm development in diabetic patients
Br. J. Ophthalmol., 1995 , 79 , 362-367
Tso, M.O.M.; Wallow, I.H.L. & Elgin, S.
Experimental photocoagulation of the human retina I: correlation of physical, clinical, and pathologic
Arch. Ophthalmol., 1977 , 95 , 1035-1040
Venecia, G.; Davis, M. & Engerman, R.
Clinicopathologic correlations in diabetic retinopathy
Arch. Ophthalmol., 1976 , 94 , 1766-173
Wallow, I.H.L.; Johns, K.; Barry, P.; Chandra, S. & Bindley, C.
Chorioretinal and choriovitreal neovascularization after photocoagulation for proliferative diabetic
retinopathy: a clinicopathologic correlation
Ophthalmology, 1985 , 92 , 523-532
Webb, R.H.; Hughes, G.W. & Delori, F.C.
Confocal scanning laser opthamoscope
Appl. Optics, 1987 , 26 , 1492-1499
Treatemtn techniques and clinical guidelines for photocagulation of diabetic macular edema: : Early
Treatment Diabetic Retinopathy Study report No. 2
Ophthalmology, 1987 , 94 , 761-774
Photocagulation for dabetic macular edema: Early Treatment Diabetic Retinopathy Study report No. 1

Arch. Ophthalmol., Early Treatment Diabetic Retinopathy Study Research Group, 1985 , 103 , 1796-
(Blankenship 1986)
Blankenship, W., Red Krypton and Blue-Green Argon Panretinal Laser Photocoagulation for
Proliferative Diabetic Retinopathy: A Laboratory and Clinical Comparison. Tr. Am. Ophth. Soc., 1986,
vol.LXXXIV, 967-1003.
(Lewis 1990)
Lewis H., Schachat AP, Haimann MH, et al. Choroidal neovascularization after laser photocoagulation
for diabetic macular edema. Ophthalmology 1990;97:503-10.
(Varley 1988)
Varley MP, Frank E, Purnell EW, Subretinal neovascularization after focal argon laser for diabetic
macular edema. Ophthalmology 1988;95:567-73.
(Rutledge BK)
Rutledge BK, Wallow IH, Paulsen GL. Sub-pigment epithelial membranes after photocoagulation for
diabetic macular edema. Arch Ophthalmol 1993;111:608-13.
(Schatz 1991)
Schatz H, Madeira D, McDonald, et al. Progressive enlargement of laser scars following grid laser
photocoagulation for diffuse diabetic macular edema. Arch Ophthalmol 1991;109:1549-51.
(Lovestam-Adrian 2000)
Lovestam-Adrian M, Agardh E. Photocoagulation of diabetic macular oedema-complications and
visual outcome. Acta Ophthalmol Scand 2000;78:667-71.

G.     Consortium/Contractual Arrangements
Not applicable

   H. Letters of Support