U.S. patent application number 16/163175 was filed with the patent office on 2019-02-14 for method for modifying the refractive index of ocular tissues.
The applicant listed for this patent is University of Rochester. Invention is credited to Li Ding, Krystel R. Huxlin, Wayne H. Knox, Jay F. Kunzler.
Application Number | 20190046357 16/163175 |
Document ID | / |
Family ID | 39739651 |
Filed Date | 2019-02-14 |
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United States Patent
Application |
20190046357 |
Kind Code |
A1 |
Knox; Wayne H. ; et
al. |
February 14, 2019 |
Method for Modifying the Refractive Index of Ocular Tissues
Abstract
A laser system for changing the index of refraction of cornea
tissue in a living eye. The laser system comprises a laser that
provides laser pulses with a wavelength from 400 nm to 900 nm and a
pulse energy from 0.01 nJ to 10 nJ, and a control device for
setting the operating parameters of the laser below an optical
breakdown threshold of the tissue to avoid photo-disruption and
tissue destruction of the tissue, and to direct the laser pulses at
the cornea tissue resulting in a change in the index of refraction
of the tissue within regions irradiated by the laser pulses.
Inventors: |
Knox; Wayne H.; (Rochester,
NY) ; Huxlin; Krystel R.; (Rush, NY) ;
Kunzler; Jay F.; (Canandaigua, NY) ; Ding; Li;
(Pleasanton, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
University of Rochester |
Rochester |
NY |
US |
|
|
Family ID: |
39739651 |
Appl. No.: |
16/163175 |
Filed: |
October 17, 2018 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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13789762 |
Mar 8, 2013 |
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16163175 |
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12146976 |
Jun 26, 2008 |
8486055 |
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13789762 |
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61026890 |
Feb 7, 2008 |
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60929397 |
Jun 26, 2007 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61F 9/00829 20130101;
A61F 9/008 20130101; A61F 2009/00872 20130101; A61F 2009/0088
20130101; A61F 2009/0087 20130101; A61F 9/00804 20130101; A61F
2009/00848 20130101; A61F 2009/00842 20130101 |
International
Class: |
A61F 9/008 20060101
A61F009/008 |
Goverment Interests
STATEMENT OF GOVERNMENT INTEREST
[0002] The U.S. Government has a paid-up license in this invention
and the right in limited circumstances to require the patent owner
to license others on reasonable terms as provided for by the terms
of Contract No. R01 EY015836 and Grant No. 08P0EY01319F awarded by
the National Institutes of Health.
Claims
1. A laser system for changing the index of refraction of cornea
tissue in a living eye comprising: a laser that provides laser
pulses with a wavelength from 400 nm to 900 nm and a pulse energy
from 0.01 nJ to 10 nJ; and a control device for setting the
operating parameters of the laser below an optical breakdown
threshold of the tissue to avoid photo-disruption and tissue
destruction of the tissue, and to direct the laser pulses at the
cornea tissue resulting in a change in the index of refraction of
the tissue within regions irradiated by the laser pulses.
2. The laser system of claim 1, wherein the regions irradiated by
the laser pulses are in the form of optical elements selected from
the group consisting of Bragg gratings, microlens arrays, zone
plates, Fresnel lenses, and combinations thereof.
3. The laser system of claim 1, further comprising a wavefront
sensor, in communication with the control device, for detecting
wavefront aberrations of the eye.
4. The laser system of claim 1, wherein the control device includes
the identification and measurement of aberrations resulting from a
surgical procedure, and a computer to process aberration
information and provide the requisite vision correction to a
patient.
5. The laser system of claim 3, wherein the wavefront sensor is
selected from the group consisting of Schemer disk, Shack Hartmann,
Hartmann screen, Fizeau interferometer and Twymann-Green
interferometer.
6. The laser system of claim 1, wherein an optical path of the
laser pulse comprises a compensation scheme to provide negative
dispersion that compensates for the positive dispersion of the
laser pulse due to focusing objectives of the laser.
7. The laser system of claim 1, wherein the laser pulse energy is
from 0.01 nJ to 2 nJ.
8. The laser system of claim 1, wherein the laser has a repetition
rate from 10 MHz to 500 MHz.
9. The laser system of claim 1, wherein the laser has a pulse width
from 10 fs to 100 fs.
10. The laser system of claim 1, wherein the laser has an average
power from 10 mW to 100 mW.
11. The laser system of claim 7, wherein the laser has an average
power from 10 mW to 100 mW.
12. The laser system of claim 11, wherein the operating parameters
are set to induce a 0.001 to 0.03 change in the index of refraction
of the cornea tissue within the irradiated region.
13. The laser system of claim 1, wherein the operating parameters
are set to induce a 0.001 to 0.03 change in the index of refraction
of the cornea tissue within the irradiated region.
Description
REFERENCE TO RELATED APPLICATIONS
[0001] This patent application is a divisional application that
claims the benefit under 35 U.S.C. .sctn. 120 of U.S. patent
application Ser. No. 12/146,976 filed Jun. 26, 2008, which in turn
claims the benefit under 35 U.S.C. .sctn. 119(e) to U.S.
provisional application Ser. No. 60/929,397 filed Jun. 26, 2007,
and U.S. provisional application Ser. No. 61/026,890 filed Feb. 7,
2008, whose disclosures are hereby incorporated by reference in
their entireties into the present disclosure.
FIELD OF THE INVENTION
[0003] The present invention relates to a method of using a laser
to modify the refractive index of ocular tissues, e.g., the corneal
stromal layer or lens cortex, for vision correction.
BACKGROUND OF THE INVENTION
[0004] Conventional ultraviolet nanosecond excimer lasers have been
very successfully used for corneal refractive surgery such as
photorefractive keratectomy (PRK), laser-assisted in situ
keratomileusis (LASIK) and laser sub-epithelial keratomileusis
(LASEK). By ablating corneal tissue through direct, one-photon
absorption of ultraviolet light, these lasers are able to alter the
curvature and thickness of corneas, ultimately altering their
optical power.
[0005] The rapid development of femtosecond laser technology has
provided an additional tool for corneal refractive surgery. In
contrast to the photo-ablative ultraviolet lasers, femtosecond
laser pulses in the near infrared or visible range can pass through
transparent corneal tissue without significant one-photon
absorption. Only when pulses are focused inside the cornea, is the
intensity of the beam sufficient to cause nonlinear, typically,
multi-photon absorption. Because the absorption is nonlinear, the
laser-affected region tends to be highly localized, leaving the
surrounding region unaffected, or minimally affected. See, Vogel A,
Noack J, Huttman G, Paltauf G, Mechanisms of femtosecond laser
nanosurgery of cells and tissues. Applied Physics B 2005, 81,
1015-47; Loesel F H, Niemz M H, Bille J F, Juhasz T, Laser-induced
optical breakdown on hard and soft tissue and its dependence on the
pulse duration: experiment and model. IEEE Journal of Quantum
Electronics 1996, 32, 1717-22; and Giguere D, Olivie G, Vidal F, et
al., Laser ablation threshold dependence on pulse duration for
fused silica and corneal tissues: experiments and modeling, Journal
of the Optical Society of America A 2007, 24, 1562-68.
[0006] In the past two decades, extensive experimental and
theoretical work has been done to characterize laser-induced
optical breakdown thresholds in different materials, including the
cornea and the lens. Most of this work, however, centered on the
use of continuous wave (CW) lasers or on single pulses from
low-repetition-rate lasers in which thermal diffusion time is much
shorter than the time interval between adjacent pulses. Thus, each
pulse is responsible for a change in the material. Indeed, it has
been established that for pulses longer than 10 ps, the optical
breakdown threshold fluence scales as the square root of the pulse
duration. To date, most femtosecond lasers used to cut corneas in
clinical practice use microJoule (.mu.J) femtosecond laser pulses
with a low-repetition-rate (Hz-kHz range) and spot diameters of
more than 5 microns (.mu.m). See, Kurtz R M, Horvath C, Liu H H,
Krueger R R, Juhasz T, Lamellar refractive surgery with scanned
intrastromal picosecond and femtosecond laser pulses in animal
eyes, Journal of Refractive Surgery 1998, 14, 541-48; and Juhasz T,
Loesel C, Horvath C, Kurtz R M, Mourou G, Corneal refractive
surgery with femtosecond lasers, IEEE Journal of Quantum
Electronics 1999, 5, 902-09.
[0007] This contrasts with the range of femtosecond laser
parameters that have been established for biomedical applications.
See, Loesel F H, Niemz M H, Bille J F, Juhasz T, Laser-induced
optical breakdown on hard and soft tissue and its dependence on the
pulse duration: experiment and model, IEEE Journal of Quantum
Electronics 1996, 32, 1717-22. Compared with the
low-repetition-rate femtosecond lasers with .mu.J or milliJoule
(mJ) pulse energies, high-repetition-rate (>1 MHz) femtosecond
laser oscillators usually have pulse energies on the order of
nanoJoule (nJ). Such low-pulse-energy femtosecond lasers have been
used for both micromachining and nanosurgery. See, Konig K, Krauss
O, Riemann I, Intratissue surgery with 80 MHz nanojoule femtosecond
laser pulses in the near infrared, Optics Express 2002, 10,
171-76.
[0008] While most femtosecond laser surgical procedures involve (by
definition) some sort of disruption, either affecting membranes,
organelles or other cellular components, they can be performed with
such precision and selectivity so as not to kill the cells.
Recently, research within our group on both silicone and
non-silicone-based hydrogels, demonstrates that femtosecond
micromachining works by inducing a significant change in refractive
index of the materials without visible plasma luminescence or
bubble formation, and without the generation of undesirable
scattering or absorbing centers. See, U.S. patent application Ser.
No. 11/745,746, filed May 8, 2007, and Ser. No. 11/948,298 filed
Nov. 30, 2007. Our success with creating refractive structures in
hydrogel materials led us to explore whether similar type
structures could be created in ocular tissues.
[0009] There exists an ongoing need for ways to improve or correct
vision. Changing the refractive index of ocular tissues, e.g., the
corneal stromal layer or lens cortex, using a femtosecond laser,
without tissue destruction or wound healing response would
represent a major advance in the field of laser refractive
correction or vision correction generally.
SUMMARY OF THE INVENTION
[0010] The invention is directed to a laser system for changing the
index of refraction of cornea tissue in a living eye. The laser
system comprises a laser that provides laser pulses with a
wavelength from 400 nm to 900 nm and a pulse energy from 0.01 nJ to
10 nJ, and a control device for setting the operating parameters of
the laser below an optical breakdown threshold of the tissue to
avoid photo-disruption and tissue destruction of the tissue, and to
direct the laser pulses at the cornea tissue resulting in a change
in the index of refraction of the tissue within regions irradiated
by the laser pulses.
BRIEF DESCRIPTION OF THE DRAWINGS
[0011] The invention will be better understood from the following
description and in consideration with the accompanying figures. It
is to be expressly understood, however, that each of the figures is
provided to further illustrate and describe the invention and is
not intended to further limit the invention claimed.
[0012] FIG. 1 is a schematic representation of a three-dimensional
structure in the corneal stromal layer that can be produced by
laser irradiation;
[0013] FIG. 2 is a schematic representation of creating a convex,
plano or concave structure in corneal stromal layer to yield a
positive or negative correction;
[0014] FIG. 3 is a schematic representation of the laser and
optical system used to provide the refractive structures;
[0015] FIGS. 4A and 4C are Differential Interference Contrast (DIC)
photographic images of a line grating in lightly-fixed cat corneal
stroma at or near the tissue breakdown threshold;
[0016] FIGS. 4B and 4D are Bright Field (BF) photographic images of
a line grating in lightly-fixed cat corneal stroma at or near the
tissue breakdown threshold;
[0017] FIG. 5A is a DIC photographic image of a line grating in
lightly-fixed cat lens cortex at or near the tissue breakdown
threshold;
[0018] FIG. 5B is a BF photographic image of a line grating in
lightly-fixed cat lens cortex at or near the tissue breakdown
threshold;
[0019] FIG. 6A is a DIC photographic image of a line grating in
lightly-fixed cat corneal stroma below the tissue breakdown
threshold;
[0020] FIG. 6B is a zoomed-in DIC image of the line grating
refractive structure shown in FIG. 6A;
[0021] FIG. 6C is a BF photographic image of a line grating in
lightly-fixed cat corneal stroma below the tissue breakdown
threshold;
[0022] FIG. 7A is a DIC photographic image of a line grating in
lightly-fixed cat lens cortex below the tissue breakdown
threshold;
[0023] FIG. 7B is a BF photographic image of a line grating in
lightly-fixed cat lens cortex below the tissue breakdown
threshold;
[0024] FIG. 8A is a graph plotting the 2nd and 3rd order
diffraction efficiencies and the corresponding laser-induced
refractive index changes of eight gratings micromachined in
different corneal samples;
[0025] FIG. 8B is a photograph of the diffraction pattern obtained
with a 632.8 nm He--Ne laser when illuminating the grating of FIG.
6A
[0026] FIG. 9A is a graph plotting the 2nd and 3rd order
diffraction efficiencies of eight gratings micromachined in
different corneal samples after one month of storage;
[0027] FIG. 9B is a photograph of the diffraction pattern obtained
with a 632.8 nm He--Ne laser when illuminating the grating of FIG.
6A after one month;
[0028] FIG. 10A is a graph plotting the 1st and 2nd order
diffraction efficiencies and the corresponding laser-induced
refractive index changes of eight gratings micromachined in
different lens cortex samples;
[0029] FIG. 10B is a photograph of the diffraction pattern obtained
with a 632.8 nm He--Ne laser when illuminating the grating of FIG.
7A;
[0030] FIG. 11A is a DIC photograph showing the line grating of
FIG. 7A after one month of storage;
[0031] FIG. 11B is a graph plotting the 1st and 2nd order
diffraction efficiencies of eight gratings micromachined in
different lens cortex samples after one month of storage;
[0032] FIG. 11C is a photograph of the diffraction pattern obtained
with a 632.8 nm He--Ne laser when illuminating the grating of FIG.
7A after one month;
[0033] FIG. 12 is a schematic representation of a use of the
preferred or another embodiment in providing ficucial marks in the
cornea; and
[0034] FIGS. 13A-13C are schematic diagrams of a device in which
the preferred or another embodiment can be implemented.
DETAILED DESCRIPTION OF THE INVENTION
[0035] The use of high-repetition-rate, femtosecond laser pulses
can result in cumulative, free-electron-mediated chemical effects,
photochemical bond breaking and thermal effects, which can
contribute to the laser-tissue interaction. As a result, the
breakdown threshold fluence may be quite different from that
predicted by current models. See, Vogel et al., App. Physics B,
2005, 81, 1015-47. Several studies on the effects of
high-repetition-rate femtosecond lasers on fused silica and
borosilicate glass have found that laser pulses greatly increased
the temperature of the materials at the laser focus. See, Eaton et
al, Optics Express 2005, 13, 4708-16. Vogel calculated the
temperature change in water would be >10.degree. K with a 0.6 NA
focusing lens and 100 fs laser pulses assuming that with each
pulse, an energy density of 1 J/cm.sup.3 at the center of the
initial temperature distribution is deposited.
[0036] Using very high-repetition-rate (93 MHz), ultra-short laser
pulses (27 fs), we determined that the optical breakdown threshold
for the 0.70 NA focusing condition in lightly-fixed corneal stroma
and lens cortex to be 55 mW and 75 mW average laser power (0.6 nJ
and 0.8 nJ pulse energies), respectively. For both corneal stromal
and lens cortex both values are lower than the optical breakdown
power reported by Konig and colleagues using 1 nJ pulse energy, 170
fs pulse duration and 1.30 NA focusing in porcine corneas. See,
Konig et al, Optics Express 2002, 10(3), 171-76. By using 30 mW and
45 mW average laser power (0.3 nJ and 0.5 nJ pulses), we discovered
that one can induce Intra-tissue Refractive Index Shaping (IRIS),
without accompanying photo-disruption and tissue destruction.
[0037] We adapted our femtosecond micromachining approach with
hydrogel materials to carry out IRIS in biological tissues. We
initially measured the optical breakdown thresholds of
lightly-fixed cat corneas and lenses. We then reduced femtosecond
laser pulse energies below these optical breakdown thresholds to
create grating patterns that are associated with a change in the
refractive index of the tissue. Our investigation has led to the
development of a process to modify the refractive index of ocular
tissue, e.g., corneal stromal layer and lens cortex, without
apparent tissue destruction. Accordingly, a determination of the
appropriate laser parameters is important for achieving IRIS in
biological tissues. Not only does the femtosecond laser fluence at
the objective focus have to be below the optical breakdown
threshold of the tissue, the laser fluence must be strong enough to
induce nonlinear changes in the tissues. Moreover, the scan speed
must be set within a specified range.
[0038] The process involves irradiating the ocular tissue with a
low-pulse-energy, MHz femtosecond laser. If very short laser pulses
having a very select energy are tightly focused on ocular tissue,
the high intensity of light leads to a change in the refractive
index of the ocular tissue at the focus point. Moreover, the region
of the ocular tissue just outside the focal region is minimally
affected by the laser light. As a result, select regions of ocular
tissue can be modified resulting in a change in the refractive
index in these regions. Moreover, the long-term stability of the
observed change in refractive index suggests permanent molecular
and/or structural changes to the ocular tissue.
[0039] Accordingly, the invention is directed to a method for
modifying the refractive index of ocular tissue. The method
comprises irradiating select regions of biological tissue with a
focused, visible or near-IR laser. The irradiated regions exhibit
little or no scattering loss, which means that the structures
formed in the irradiated regions are not clearly visible under
appropriate magnification without contrast enhancement.
[0040] To date, the determined change in refractive index induced
in cornea and lens tissue using the described process is relatively
small, but very significant. Based on published values for the
power (39D) and native refractive index (1.376) of the cat cornea,
IRIS should generate a change in corneal power ranging between
0.14D and 0.56D (assuming an index change between 0.005 and 0.02).
Similarly, for the cat lens (power=53D, refractive index of the
homogeneous lens=1.554), the refractive index changes induced by
micromachining should theoretically alter lenticular power by
between 0.5D and 0.7D. The laser process described could completely
alter the approach to laser refractive surgery and to vision
correction.
[0041] In addition, the preservation of tissue clarity during the
treatment allows the application of IRIS for the creation of
corneal fiducial markings that could be used to align eye trackers
during LASIK, and for refractive corrections in a closed-loop
approach, e.g. with specific benefit for the correction of
higher-order aberrations, as well as for "touch-up corrections" of
ocular surface defects. Various types of refractive structures can
be created in biological tissues. Examples include high refractive
index structures such as Bragg gratings, microlens arrays, optical
zone plates, and Fresnel lenses.
[0042] As stated, the determination and selection of the laser
operating parameters are particularly important in implementing
IRIS. The inventors have found that various ranges of parameters
are particularly useful in implementing the present invention. In
treatment of the eye, the laser wavelength should be such that the
tissues through which laser pulses pass are transparent to the
pulses. There should also be no damage to the retina; any change
should be confined to the tissue located at the spot of focus of
the pulses. Also, for non-destructive alteration of ocular tissue,
a CO.sub.2 laser or excimer laser should not be used, since there
should be no ablation or removal of the tissue.
[0043] A laser pulse frequency from 1 MHz to 10 GHz, and preferably
from 10 to 500 MHz, should be used. For example, our work used a
laser pulse frequency (repetition rate) of 93 MHz.
[0044] Linked to the pulse frequency is the average laser power. A
preferable average laser power is from 1 mW to 1,000 mW, preferably
10 mW to 100 mW, and more preferably from 30 mW to 70 mW.
[0045] The energy of each pulse should be in a range from 0.01 nJ
to 10 nJ, preferably from 0.1 nJ to 2 nJ, and more preferably be
less than 1 nJ. For example, we have determined that a pulse energy
from 0.1 nJ to 0.5 nJ, is particularly preferred.
[0046] Although we used a laser pulse width of 30 fs, one can
likely operate within a range from 5 fs to 1 ps, and more
preferably from 10 fs to 100 fs. Also, the scanning speed of the
laser is preferably 0.4 .mu.m/s, though, depending on the
equipment, types of structures to be written and type of tissue,
scan speeds in a range from 0.1 .mu.m/s to 10 mm/s can be used. The
laser pulses were focused to a spot size of about 1 .mu.m, though a
range of spot size from 0.5 .mu.m to 50 .mu.m, from 0.5 .mu.m to 10
.mu.m, from 0.5 .mu.m to 5 .mu.m, could be used.
[0047] The pulse energy of the focused laser used in the method
will in-part depend on the type of structures to be written into
the ocular tissue, the type of ocular tissue and how much of a
change in refractive index is desired. The selected pulse energy
will also depend upon the scan rate at which the structures are
written into the ocular tissue. Typically, greater pulse energies
will be needed for greater scan rates.
[0048] The pulse width must be preserved so that the pulse peak
power is strong enough to exceed the nonlinear absorption threshold
of the ocular tissue. However, the glass of the focusing
objective(s) significantly increases the pulse width due to the
positive dispersion of the glass. A compensation scheme is used to
provide a corresponding negative dispersion that can compensate for
the positive dispersion introduced by the focusing objective(s).
Accordingly, the term "focused" in this application refers to the
focusing of light from a laser within ocular tissue using a
compensation scheme to correct for the positive dispersion
introduced by the focusing objective(s). The compensation scheme
can include an optical arrangement selected from the group
consisting of at least two prisms and at least one mirror, at least
two diffraction gratings, a chirped mirror and dispersion
compensating mirrors to compensate for the positive dispersion
introduced by the focus objective.
[0049] The wavelength should, as noted above, be one to which the
tissues through which the laser pulses must pass are transparent.
It should also preferably be just barely within the visible range
for the patient, or outside of the visible range (e.g.,
near-infrared), so as not to bother the patient. A wavelength of
800 nm is useful; preferable ranges include 600-1,000 nm (and more
preferably 700-900 nm) and above 1,000 nm.
[0050] In one embodiment, the compensation scheme comprises at
least one prism, in many cases at least two prisms, and at least
one mirror to compensate for the positive dispersion of the
focusing objective. In another embodiment, the compensation scheme
comprises at least two gratings to compensate for the positive
dispersion of the focusing objective. Any combination of prisms,
gratings and/or mirrors can be used for the compensation scheme in
accordance with optical principles known by those of ordinary skill
in the art.
[0051] The laser will generate light with a wavelength in the range
from violet to near-infrared radiation. In various embodiments, the
wavelength of the laser is in the range from 400 nm to 1500 nm,
from 400 nm to 1200 nm or from 600 nm to 900 nm.
[0052] In one particular embodiment, the laser is a pumped
Ti:sapphire laser with an average power of 10 mW to 1000 mW. Such a
laser system will generate light with a wavelength of approximately
800 nm. In another embodiment, an amplified fiber laser that can
generate light with a wavelength from 1000 nm to 1600 nm can be
used
[0053] The laser will have a peak intensity at focus of greater
than 10.sup.13 W/cm.sup.2. At times, it may be advantageous to
provide a laser with a peak intensity at focus of greater than
10.sup.14 W/cm.sup.2, or greater than 10.sup.15 W/cm.sup.2.
[0054] The irradiated regions of the ocular tissue are defined by
two- or three-dimensional structures. The two- or three-dimensional
structures can comprise an array of discrete cylinders.
Alternatively, the two- or three-dimensional structures can
comprise a series of lines (a grating) or a combination of an array
of cylinders and a series of lines. Moreover, the two- or
three-dimensional structures can comprise area or volume filled
structures, respectively. These area or volume filled structures
are formed by continuously scanning the laser over a select plane
or volume of the ocular tissue, respectively. As stated, various
types of refractive structures can be created in biological
tissues. Examples include high refractive index structures such as
lenses, prisms, Bragg gratings, microlens arrays, optical zone
plates, and Fresnel lenses.
[0055] The area-filled or volume-filled two- or three-dimensional
structures can be formed by continuously scanning the laser over
selected regions of the ocular tissue. Refractive-type optical
devices can be micro-machined inside the volume of ocular tissue by
repeatedly scanning a tightly focused beam of femtosecond pulses in
an area segment. The area of the segment can be changed
correspondingly with the depth of the scan, so as to produce
three-dimensionally shaped lenses with spheric, aspheric, toroidal
or cylindrical shapes as shown in FIG. 1. Alternatively, refractive
corrective lenses can be made in various combinations of convex,
plano- or concave to yield a positive correction, or negative
correction, as shown in FIG. 2. The refractive optical devices can
be stacked vertically, written separately in different planes, so
as to act as a single lens. Additional corrective layers can be
written as desired.
[0056] In one embodiment, the irradiated regions of the ocular
tissue are defined by a series of lines in an approximately two
dimensional plane having a width from 0.2 .mu.m to 3 .mu.m,
preferably a width from 0.6 .mu.m to 1.5 .mu.m and a height from
0.4 .mu.m to 8 .mu.m, preferably a height from 1.0 .mu.m to 4 .mu.m
(height is measured in the z direction, which is parallel to
direction of the laser light). For example, one can generate a line
grating comprising a plurality of lines with each line of any
desired length, about 0.8 .mu.m to about 1.5 .mu.m in width and
about 2 .mu.m to 5 .mu.m in height. The lines can be separated by
as little as 1.0 .mu.m (0.5 .mu.m spacing), and any number of lines
can be incorporated into the ocular tissue. Moreover, the grating
can be positioned at any selected depth (z-direction), and any
number of line gratings can be generated at various depths into the
ocular tissue.
[0057] In one embodiment, the irradiated portions of corneal
stromal layer will exhibit a change in refractive index of about
0.001 to about 0.02, and typically about 0.01 or more. Based on
published values for the power (39D) and native refractive index
(1.376) of the cat cornea, the refractive index changes induced by
micromachining should generate a change in corneal power ranging
between 0.14D and 0.28D (assuming that refractive index change
affects the thickness of the cornea uniformly).
[0058] In another embodiment, the irradiated portions of lens
cortex will exhibit a change in refractive index of about 0.005 to
about 0.03, and typically about 0.01 or more.
[0059] Our work demonstrates that it is possible to cause
low-scattering-loss, refractive index modifications in
lightly-fixed cat cornea and lens using 93 MHz repetition rate, 27
fs laser pulses with 0.3 nJ and 0.5 nJ pulse energies. These
modifications were visible only using DIC microscopy and were not
associated with apparent tissue damage. The resulting modifications
correspond to refractive index changes between 0.05.+-.0.001 and
0.021.+-.0.001. Preservation of IRIS over a month of refrigerated
storage suggested that the femtosecond laser-induced modifications
were likely to involve relatively long-term molecular/structural
alterations. In related experiments involving silicone hydrogels,
the micromachined gratings (and associated Raman spectra) are
stable for at least one year, even after drying and rehydration of
the hydrogel.
A Laser and Optical Configuration for Modifying Ocular Tissue
[0060] A non-limiting embodiment of a laser system 10 for
irradiating ocular tissue with a laser to modify the refractive
index of the tissue in select regions is represented in FIG. 3. A
laser source comprises a Kerr-lens mode-locked Ti:Sapphire laser 12
(Kapteyn-Murnane Labs, Boulder, Colo.) pumped by 4 W of green light
from a frequency-doubled Nd:YVO.sub.4 laser 14. The laser generates
pulses of 300 mW average power, 30 fs pulse width and 93 MHz
repetition rate at wavelength of 800 nm. Because there is a
reflective power loss from the mirrors and prisms in the optical
path, and in particular, from the power loss of the objective 20,
the measured average laser power at the objective focus on the
material is about 120 mW, which indicates the pulse energy for the
femtosecond laser is about 1.3 nJ.
[0061] Due to the limited laser pulse energy at the objective
focus, the pulse width must be preserved so that the pulse peak
power is strong enough to exceed the nonlinear absorption threshold
of the ocular tissue. Because the glass inside the focusing
objective significantly increases the pulse width due to the
positive dispersion inside of the glass, an extra-cavity,
compensation scheme is used to provide the negative dispersion that
compensates for the positive dispersion introduced by the focusing
objective. Two SF10 prisms 24 and 28 and one ending mirror 32 form
a two-pass one-prism-pair configuration. We used a 37.5 cm
separation distance between the prisms to compensate the dispersion
of the microscope objective and other optics within the optical
path. A collinear autocorrelator 40 using third-order harmonic
generation is used to measure the pulse width at the objective
focus. Both 2nd and 3rd harmonic generation have been used in
autocorrelation measurements for low NA or high NA objectives. We
selected third order surface harmonic generation (THG)
autocorrelation to characterize the pulse width at the focus of the
high-numerical-aperture objectives because of its simplicity, high
signal to noise ratio and lack of material dispersion that second
harmonic generation (SHG) crystals usually introduce. The THG
signal is generated at the interface of air and an ordinary cover
slip 42 (Corning No. 0211 Zinc Titania glass), and measured with a
photomultiplier 44 and a lock-in amplifier 46. After using a set of
different high-numerical-aperture objectives and carefully
adjusting the separation distance between the two prisms and the
amount of glass inserted, we selected a transform-limited 27-fs
duration pulse, which is focused by a 60.times.0.70 NA Olympus
LUCPlanFLN long-working-distance objective 48.
[0062] Because the laser beam will spatially diverge after it comes
out of the laser cavity, a concave mirror pair 50 and 52 is added
into the optical path in order to adjust the dimension of the laser
beam so that the laser beam can optimally fills the objective
aperture. A 3D 100 nm resolution DC servo motor stage 54 (Newport
VP-25XA linear stage) and a 2D 0.7 nm resolution piezo
nanopositioning stage (PI P-622.2CD piezo stage) are controlled and
programmed by a computer 56 as a scanning platform to support and
locate the samples. The servo stages have a DC servo-motor so they
can move smoothly between adjacent steps. An optical shutter
controlled by the computer with 1 ms time resolution is installed
in the system to precisely control the laser exposure time. With
customized computer programs, the optical shutter could be operated
with the scanning stages to micro-machine different patterns in the
materials with different scanning speed at different position and
depth and different laser exposure time. In addition, a CCD camera
58 along with a monitor 62 is used beside the objective 20 to
monitor the process in real time.
[0063] The method and optical apparatus described above can be used
to modify the refractive index of ocular tissue as follows. The
first step in our micromachining experiment was to establish
thresholds for the optical breakdown of lightly fixed feline cornea
and lens cortex. The neutral density filter was first adjusted to
minimize the focused incident laser power on the cornea and the
lens below their breakdown thresholds. The incident laser power was
then progressively increased by adjusting the neutral density
filter. The breakdown threshold power was considered to be reached
when visible plasma luminescence suddenly appeared and strong
scattering light as well as laser-induced damage became visible,
see FIGS. 4A to 4D and FIGS. 5A and 5B. Using the 0.70 NA
long-working-distance objective in our system, the measured
breakdown thresholds for cat cornea and lens was about 55 mW and 75
mW average laser power, respectively, which corresponds to a pulse
energy of 0.6 nJ and 0.8 nJ, respectively.
[0064] FIGS. 4A to 4D are microscopic photographs of line gratings
micromachined in lightly-fixed, cat corneal stroma using
femtosecond laser conditions at or near the tissue breakdown
threshold. FIGS. 4A and 4C are Differential Interference Contrast
(DIC) images of lines created in the stroma of two different,
lightly-fixed cat corneas with 0.6 nJ pulses and a scanning speed
of 10 .mu.m/s. Note, the spots of tissue destruction or "bubbles"
(arrowed) along the micromachined lines (the clear, horizontal
lines within stromal tissue). FIGS. 4B and 4D are Bright Field (BF)
images of the same line gratings of FIGS. 4A and 4C, respectively.
The BF images illustrate the visibility of tissue destruction
(arrowed) and the relative invisibility of the rest of the lines
that are clearly seen under DIC conditions.
[0065] Once tissue breakdown thresholds were established, the
focused laser power was lowered gradually by carefully adjusting
the neutral density filter until lines could be micromachined
without the induction of bubbles or burns. We determined an average
laser power setting of 30 mW for the cornea, which corresponds to a
pulse energy of about 0.3 nJ.
[0066] The gratings were micromachined in the horizontal plane
within the stromal layer of each corneal piece at a constant speed
of 0.7 .mu.m/s. The gratings consisted of 20-40 parallel lines, 100
.mu.m long, 1 .mu.m linewidth, 5 .mu.m apart and about 100 .mu.m
beneath the corneal epithelium. Likewise, gratings were
micromachined in the horizontal plane within the cortex of each
lens at a constant speed of 1.0 .mu.m/s. The gratings again
consisted of 20-40 parallel lines, 100 .mu.m long, 1 .mu.m
linewidth, 5 .mu.m apart and about 100 .mu.m beneath the lenticular
surface. The spherical aberration at the laser focus induced by
refractive index mismatch was compensated by an adjustable cover
slip correction of the focusing microscope objective in order to
achieve the smallest possible laser-affected region along the laser
propagation direction.
Observation and Measurement of Refractive Index Change.
[0067] After writing the observed structures in both corneal stroma
and lens cortex we assessed whether the micromachined gratings are
associated with a change in refractive index of the two different
tissues. Immediately after micromachining, the slide containing the
corneal piece and lens cortex was examined under an Olympus BX51
optical microscope. Bright field, phase contrast (PC) and
differential interference contrast (DIC) were used to view the
gratings. The slide was then moved to another setup where a low
power 632.8 nm He--Ne laser was used to irradiate the gratings. The
diffraction pattern from each grating was captured by a digital
camera. The refractive index changes attained were calculated as
described previously. See, Ding L, Blackwell R, Kunzler J F, Knox W
H, Large refractive index change in silicone-based and
non-silicone-based hydrogel polymers induced by femtosecond laser
micro-machining, Optics Express 2006, 14, 11901-909.
[0068] In brief, the intensity of 0th order to 3rd order of
diffracted light from the gratings was measured by a power meter.
The different order diffraction efficiencies was obtained by
calculating the ratios between the intensity of 1st, 2nd and 3rd to
0th order diffraction light. Because only one particular value of
the refractive index change matches one particular diffraction
efficiency value, one could calculate the index change within the
femtosecond laser micromachined regions. We note that several
factors could affect the results, such as the accuracy of
measurement for the different diffraction order intensities, and
the measurements of grating linewidth and thickness. To reduce
measurement error of the diffraction order intensities, we took
five measurements on each grating and calculated the average value
and the standard deviation of the results. In principle, the
spatial distribution of the refractive index change within the
micromachined region was a small-scale gradient-index structure.
For the purpose of this investigation, however, we presumed the
index profile to be uniform within the grating lines, which were
only 3 .mu.m deep because the spherical aberration at the focal
point was corrected.
[0069] The micromachined cat cornea and lens pieces were then
removed from the glass slides after discarding the cover slips, and
stored in the ethylene glycol/sucrose solution at 4.degree. C.
After one month, each corneal piece and lens piece was mounted onto
a new glass slide for imaging and the diffraction light intensity
measurement was repeated. This allowed us to assess whether the
refractive index change initially observed had been maintained
during storage.
[0070] Exposure of lightly-fixed cat corneal to 0.3 nJ femtosecond
laser pulses (30 mW average laser power) resulted in the reliable
creation of grating patterns about 100 .mu.m below the epithelial
surface in all test samples, even when they were obtained from
different cats. When imaged immediately after micromachining,
individual grating lines could be clearly observed and
distinguished with DIC microscopy (FIGS. 6A and 6B), but they were
practically invisible when viewed under bright field transmission
microscopy (FIG. 6C). This could be interpreted as the grating
lines having very low scattering properties, which is in great
contrast to the destructive tissue changes observed when laser
energy was increased above the optical breakdown threshold levels
(spots in FIG. 4). Using the knife-edge method, we ascertained that
the laser focus diameter was 2.5 .mu.m in air, which was much
bigger than the micromachined line-widths. Therefore, it appears
that only the central part of the laser focal area had sufficient
intensity to modify the refractive index of corneal tissue.
[0071] Likewise, exposure of lightly-fixed cat lens cortex to 0.5
nJ femtosecond laser pulses (45 mW average laser power) resulted in
the reliable creation of grating patterns about 100 .mu.m below the
lenticular surface in all test samples, even when they were
obtained from different cats. When imaged immediately after
micromachining, individual grating lines could be clearly observed
and distinguished with DIC microscopy (FIG. 7A), but they were
practically invisible when viewed under bright field transmission
microscopy (FIG. 7B). Again, this is interpreted as the grating
lines having very low scattering properties, which is in great
contrast to the destructive tissue changes observed when laser
energy was increased above the optical breakdown threshold levels
(spots in FIG. 5). Also, it appears that only the central part of
the laser focal area had sufficient intensity to modify the
refractive index of lens cortex.
[0072] In order to further assess the optical consequences of
low-pulse-energy femtosecond laser micromachining on corneal
stromal layer and lens cortex, we immediately irradiated the
micromachined gratings with low power 632.8 nm He--Ne laser light.
Because displacement of the stromal collagen lamellae as a result
of post-mortem corneal swelling could not be completely avoided,
scattering effect from the 0th order diffraction light was very
strong, obscuring the 1st order diffraction light. Thus, only the
2nd and 3rd order diffraction efficiencies of each grating could be
measured and used to calculate an approximate refractive index
change within the femtosecond laser micromachined regions, FIG. 8A.
In contrast, tissue swelling and opacification were minimal in
slices of lens cortex, the 0th through 3rd order diffraction light
could be measured clearly, and 1st and 2nd order diffraction
efficiencies were used to calculate the induced change in
refractive index (FIG. 10A).
[0073] Although a single diffraction efficiency is usually
sufficient to calculate refractive index, we measured 1st/2nd or
2nd/3rd combinations to confirm that the refractive indices
calculated were consistent through different diffraction orders.
For these calculations, the average refractive indices of cat
corneal stroma and lens were assumed to be 1.376 and 1.400,
respectively. For corneal stroma, the calculated range of
refractive index changes induced by the laser micromachining was
from 0.005.+-.0.001 to 0.01.+-.0.001. For lens cortex, the
calculated range of refractive index changes induced by the laser
micromachining was from 0.005.+-.0.001 to 0.03.+-.0.001.
[0074] After undergoing low-pulse-energy femtosecond laser
micromachining, each cornea piece was returned to the storage
solution in a -20.degree. C. freezer for one month in order to
determine if the micromachined structures could be maintained over
such a period of time. After one month, the cornea pieces were
removed from storage and re-examined. The storage solution
significantly slowed corneal swelling and opacification (relative
to conventional storage in 0.1M PBS, for example), but was not able
to completely prevent these events. In spite of a moderate loss of
corneal transparency, DIC microscopy did reveal that the grating
structures initially micromachined into the corneal stroma were
still present one month after they were originally created as
demonstrated by the diffraction pattern observed in FIGS. 7B and
9B. The edges of the lens slices became opaque following one month
storage, but the centers remained largely transparent and the
micromachined gratings were still clearly visible in a DIC image,
FIG. 11A.
[0075] The diffraction light distribution of one-month old gratings
in corneal stromal layer (FIG. 9B) was again measured and found to
be no different than that obtained right after the gratings'
creation (FIG. 8B). Also, the diffraction light distribution of
one-month old gratings in lens cortex (FIG. 11C) was again measured
and found to be no different than that obtained right after the
gratings' creation (FIG. 10B). In the corneal pieces, the
scattering light from the 0th order diffraction still obscured the
1st order diffraction. However, the 2nd, 3rd, and even 4th order
diffractions were still visible and easy to measure. The measured
refractive index change after one month of storage remained from
0.005.+-.0.001 to 0.01.+-.0.001 for the corneal pieces and from
0.005.+-.0.001 to 0.03.+-.0.001 for the lens pieces.
[0076] Applications in ophthalmic surgery will now be described. As
shown in FIGS. 4A to 4C, it is possible to write micron-scale
features into the stromal layer with minimal scattering loss by
carefully controlling the laser and scan parameters such as pulse
width, average power, repetition rate, scan rate and focusing
conditions. This result, which is significantly different than the
results in corneal surgery that have been previously reported using
femtosecond, focused pulses, suggests to us certain
applications.
[0077] One such application is in writing fiducial marks in the
corneal stroma. More particularly, in one application involving
excimer laser ablation of the cornea for vision correction--laser
in situ keratomileusis or LASIK--it is first necessary to cut
across the cornea with a `flap cutting` device. Typically, a
rapidly vibrating razor blade or microkeratome is used for this
purpose. This method generally produces acceptable results, however
the depth of the final cut is not precise, and sometimes the degree
of accommodation that can be achieved with excimer laser ablation
is compromised. A competing form of corneal flap-cutting involves
the use of a high-power, femtosecond laser. Femtosecond flap
cutting has not been widely adopted yet in clinical refractive
surgery practices, in part because of uncertainty about the
long-term photochemical, mechanical and biological effects of this
technique (Stonecipher et al., 2006; Wilson et al., 2007).
Recently, there have been reports about negative effects of this
technique, particularly in terms of tissue destruction, which
appears significantly stronger than that obtained following
microkeratome cutting (Stonecipher et al., 2006; Wilson et al.,
2007).
[0078] The micromachining process described provides a possible
solution to the problem of being able to make a precise cut in the
corneal stroma without additional tissue destruction. FIG. 12 shows
a situation in which fiducial marks 101 has been machined into the
stromal layer of the cornea, for example, at a specific location
and depth. Low-energy femtosecond laser pulses can be used to write
fiducial marks 101 in the stromal layer of the cornea at a specific
depth and location. The fiducial mark 101 would not be visible to a
human, as indicated by FIGS. 4b and 4D, however it is detectable by
specialized optical techniques such as Optical Coherence Tomography
(OCT) or Differential Interference Contrast (DIC) microscopy (FIGS.
4A and 4C).
[0079] The fiducial marks 101 could be used to `lock` the depth of
the cutting blade by using an imaging technique such as Optical
Coherence Tomography (OCT). OCT has been well developed for both
retinal and corneal imaging. This would ensure that the resulting
depth of the blade cut would be significantly better regulated than
is currently possible, even if a femtosecond laser is used to cut
the corneal flap. The location and depth of a fiducial mark 101 is
obtained using optical coherence tomography (OCT) interfaced with
blade control, and the depth of the blade 105 is `locked` to this
depth and `guided` along a specified cutting path, which can ensure
accurate flap cutting.
[0080] Another application is in altering the optical power of the
cornea. Currently, laser refractive surgery achieves changes in the
optical power of the cornea by destroying/removing corneal tissue.
Tissue destruction causes (1) a change in the surface profile (and
curvature) of the cornea, (2) a change in corneal biomechanical
properties (usually a flattening of the corneal surface), and (3) a
wound healing response. A change in surface shape of the cornea as
a result of points (1) and (2) is sufficient to correct large
optical aberrations such as defocus and astigmatism. However, as
mentioned earlier, the wound healing response that results from
corneal tissue destruction limits current laser refractive
procedures by decreasing their ultimate optical benefit. Exemplary
femtosecond micromachining patterns that could be written into the
corneal stroma include a continuous circular area, an annulus
pattern, or a segmented annulus pattern.
[0081] The use of femtosecond laser pulses as described to modify
the optical power of the cornea can be accomplished as follows: (1)
by changing the refractive index of the corneal stroma, and (2) by
altering corneal biomechanics without inducing a significant
corneal wound healing response. Because of the femtosecond laser's
ability to be focused non-invasively, in a non-contact manner, to
effect at any chosen depth within the corneal stroma, this
procedure would not require removal of the corneal epithelium or
creation of a corneal flap. Epithelial manipulations are one of the
major stimuli causing the wound healing response since such
manipulations destroy the normally close interaction (both
physically and biologically) of the corneal epithelium with its
underlying stroma. The femtosecond micromachining (i.e. use of
low-energy femtosecond pulses to alter tissue properties
non-destructively) could be applied over a continuous area, 6-8 mm
in diameter, in the center of the cornea or at particular locations
in the corneal periphery as mentioned above, depending on the
optical or biomechanical changes desired.
[0082] Yet another application is in altering the optical power of
the intraocular lens of the eye. Presbyopia, or the loss of
accommodation ability as a function of age, is currently of
epidemic proportions in the developed world. The most common
treatment approaches for this condition include reading glasses,
bifocal glasses, contact lenses (including bifocals), multifocal
laser refractive treatments, monovision laser refractive treatments
and the use of accommodative intraocular lens implants into the
eye. Other approaches involve the use of surgical or destructive
laser treatments to punch holes in the patient's intraocular lens,
thus decreasing its rigidity, and restoring some limited
accommodative power. However, any invasive or destructive
procedures induce a wound healing response in the lens, with
increased risk of opacification or cataract formation.
[0083] The micromachining process described also provides an
opportunity for an ocular surgeon to modify the refractive index of
the corneal stromal layer of a patient having gone cataract
surgery. The method allows the ocular surgeon to correct any
aberrations as a result of the surgery. For example, starting from
a lens of selected power, the power of which will vary according to
the ocular requirements of the patient, the surgeon can
subsequently adjust the refractive properties of the corneal
stromal layer to correct a patients vision based upon the
individual needs of the patient. In essence, an intraocular lens
would essentially function as a fixed power lens to correct for the
refractive error of a patient's eye. The patient's vision can then
be further adjusted post-implantation by modifying the refractive
index of select regions of the patient's corneal stromal layer. As
a result, post-operative refractive errors resulting from
pre-operative measurement errors, variable lens positioning during
implantation and wound healing (aberrations) can be corrected.
[0084] For instance, cataract surgery typically requires that the
natural lens of each eye be replaced with an intraocular lens
(IOL). Following insertion of the IOL the surgeon or eye specialist
can correct for aberrations resulting from the surgery or correct
for slight misplacement of the IOL. Following surgery, and after
allowing time for the wound to heal, the patient would return to
the surgeon to have select regions of his or her corneal stromal
layer irradiated. These irradiated regions would experience a
change in refractive index, which would correct for the aberrations
as well as the patients needs for vision correction.
[0085] Accordingly, the invention is directed to a method
comprising identifying and measuring the aberrations resulting from
the surgical procedure. Once the aberrations are identified and
quantified using methods well known in the art of ophthalmology,
this information is processed by a computer. Of course, information
related to the requisite vision correction for each patient can
also be identified and determined, and this information can also be
processed by a computer. There are a number of commercially
available diagnostic systems that are used to measure the
aberrations. For example, common wavefront sensors used today are
based on the Schemer disk, the Shack Hartmann wavefront sensor, the
Hartmann screen, and the Fizeau and Twymann-Green interferometers.
The Shack-Hartmann wavefront measurement system is known in the art
and is described in-part by U.S. Pat. Nos. 5,849,006; 6,261,220;
6,271,914 and 6,270,221. Such systems operate by illuminating a
retina of the eye and measuring the reflected wavefront.
[0086] Once the aberrations are identified and quantified, the
computer programs determine the position and shape of the optical
structures to be written into the corneal stromal layer to correct
for those aberrations. These computer programs are well known to
those of ordinary skill in the art. The computer than communicates
with the laser-optical system and select regions of corneal stromal
layer are irradiated with a focused, visible or near-IR laser
having a pulse energy from 0.01 nJ to 1.0 nJ.
[0087] The described micromachining process can also be used for
custom vision correction of higher order wavefront aberration in
the optical path of the eye. The basic technology for detecting and
correcting aberrations of at least third-, fifth-, and tenth orders
is taught in U.S. Pat. No. 5,777,719, whose disclosure is hereby
incorporated by reference in its entirety into the present
disclosure. Given that the region of refractive index change
generated by femtosecond laser micromachining can be as small as 1
.mu.m in diameter, this will make it possible to correct small,
localized optical wavefront aberrations (higher order aberrations)
in the optical path of the eye. Such aberrations exist both
naturally, or can be induced by ocular surgeries, such as laser
refractive surgery, corneal transplantation and wound healing
following trauma to the eye.
[0088] FIGS. 13A to 13C show a schematic diagram of a device 1 used
to carry out the preferred embodiment or another embodiment. The
device 1 includes a laser 3 for emitting femtosecond laser pulses,
a shutter 5, a focusing lens 7, a dichroic mirror 9, a wavefront
sensor 11 having a lenslet array 13 and a detector array 15, and a
control system 17 for controlling the operations described
herein.
[0089] As illustrated in FIGS. 13A to 13C, the process we propose
would include the following steps: (1) using a wavefront sensor to
detect and measure the lower and higher order aberrations along the
optical path of a given eye, (2) calculating the topography and
magnitude of refractive index changes required to achieve the
necessary aberration correction, (3) focusing the femtosecond laser
pulses either into the cornea or intraocular lens in order to carry
out the micromachining necessary to induce the required refractive
index change. Once the micromachining is complete, the wavefront
sensor would be used once again to check the correction of the
ocular wavefront. Since the resolution of the femtosecond laser
micromachining is about 1 .mu.m, this noninvasive method could be
used as a complement or an alternative method for current
customized wavefront correction methods.
[0090] In FIG. 13A, the shutter 5 is closed for detection of
wavefront aberration from the optical path through the wavefront
sensor 11, using aberrated light reflected from the retina of the
eye. In FIG. 13B, the shutter 5 is open, and light pulses from the
femtosecond laser 3 are used to correct the aberration by locally
changing the index in the cornea or the lens of the eye. In FIG.
13C, after femtosecond laser 3 micromachining, the wavefront
correction is verified once again using the wavefront sensor
11.
Calculation of Change in Refractive Index.
[0091] As mentioned. these gratings were investigated by focusing
an unpolarized He--Ne laser beam with a wavelength of 632.8 nm on
these gratings and monitoring the diffraction pattern. The
diffraction angles showed good agreement with the diffraction
equation
m.lamda.=d sin .theta. (1)
[0092] where m is the diffraction order, .lamda. is the wavelength
of the incident laser beam which here is 632.8 nm, and d is the
grating period.
[0093] The diffraction efficiency of the grating can be measured,
and since the efficiency is a function of the refractive index
change, it may be used to calculate the refractive index change in
the laser irradiation region. Consider the grating as a phase
grating, its transmittance function could be written as
t ( x 0 , y 0 ) = ( e i .phi. 2 - e i .phi. 1 ) rect ( x 0 a ) * 1
d comb ( x 0 d ) + e i .phi. 1 ( 2 ) ##EQU00001##
[0094] where a is the grating line width, d is the groove spacing,
.PHI..sub.2 and .PHI..sub.1 are the phase delays through the lines
and ambient region respectively,
.phi. 2 = 2 .pi. .times. ( n + .DELTA. n ) .times. b .lamda. and
.phi. 1 = 2 .pi. .times. n .times. b .lamda. , ##EQU00002##
b is the thickness of the grating line, n is the average refractive
index of the material, .DELTA.n is the average refractive index
change in the grating lines, and .lamda. is the incident light
wavelength of the measurement (632.8 nm). Here, the grating line
width is 1 .mu.m and the thickness is 3 .mu.m. The index change
within the laser effect region can be approximated to be uniform.
The convolution theorem can be used to calculate the spectrum of
the grating such as
T(f.sub.x,f.sub.y)=F{t(x.sub.0,y.sub.0)}=(e.sup.i.PHI..sup.2-e.sup.i.PHI-
..sup.1)a sin
c(af.sub.x)comb(df.sub.x).delta.(f.sub.y)+e.sup.i.PHI..sup.1(f.sub.x,f.su-
b.y) (3)
[0095] Then, the intensity distribution of the grating diffraction
pattern is:
I ( x , y ) = ( 1 .lamda. z ) 2 .times. [ ( e i .phi. 2 - e i .phi.
1 ) a d n = - .infin. .infin. sin c ( an d ) .delta. ( x .lamda. z
- n d y .lamda. z ) + e i .phi. 1 .delta. ( x .zeta. z , y .lamda.
z ) ] 2 ( 4 ) ##EQU00003##
[0096] From this formula, the intensity of the 0th (I0), 1st (I1),
and 2nd (I2) order diffraction light is
I 0 = ( 1 .lamda. z ) 2 .times. [ ( e i 2 .pi. .times. ( n +
.DELTA. n ) .times. b .lamda. - e i 2 .pi. .times. n .times. b
.lamda. ) a d + e i 2 .pi. .times. n .times. b .lamda. ] 2 ( 5 ) I
1 = ( 1 .lamda. z ) 2 .times. [ ( e i 2 .pi. .times. ( n + .DELTA.
n ) .times. b .lamda. - e i 2 .pi. .times. n .times. b .lamda. ) a
d sin c ( a d ) ] 2 and ( 6 ) I 2 = ( 1 .lamda. z ) 2 .times. [ ( e
i 2 .pi. .times. ( n + .DELTA. n ) .times. b .lamda. - e i 2 .pi.
.times. n .times. b .lamda. ) a d sin c ( 2 a d ) ] 2 ( 7 )
##EQU00004##
[0097] By comparing the light intensities of 1.sup.st, 2.sup.nd and
0.sup.th diffraction orders, the refractive index change within the
grating lines can be determined.
[0098] While specific embodiments of the present invention have
been described in the foregoing, it will be appreciated by those
skilled in the art that many equivalents, modifications,
substitutions, and variations may be made thereto without departing
from the spirit and scope of the invention as defined in the
appended claims.
EXAMPLES
Extraction and Preparation of Cat Corneas
[0099] Eight corneas and eight lenses were extracted under surgical
anesthesia from five normal, adult domestic short-hair cats (felis
cattus). All animal procedures were conducted in accordance with
the guidelines of the University of Rochester Committee on Animal
Research, the ARVO Statement for the Use of Animals in Ophthalmic
and Vision Research, and the NIH Guide for the Care and Use of
Laboratory Animals. Feline corneas and lenses were chosen because
of their similarity to human corneas and lenses in terms of
histological structure, molecular composition and optical
properties. See, Hughes A. The topography of vision in mammals of
contrasting life style: comparative optics and retinal
organization. Handbook of Sensory Physiology, VII/5. Berlin:
Springer Verlag; 1977. Also, in contrast with the problems
associated with obtaining post-mortem human eyes, using cat corneas
and lenses allowed us to precisely control post-mortem extraction
time and tissue processing parameters. This was important to avoid
degradation and opacification of the corneas and lenses prior to
femtosecond laser micromachining. Extracted feline tissues were
immediately drop-fixed for 10 minutes (corneas) in a solution
consisting of 1% paraformaldehyde in 0.1M phosphate buffered saline
(PBS), pH 7.4. Lenses were cut into 500 .mu.m thick slices using a
vibratome. The lens slices and whole corneas (.about.500 .mu.m
thick) were immersed in a mixture of 30% ethylene glycol+30%
sucrose in 0.1M PBS, pH 7.4 at 4.degree. C. The ocular tissues were
stored in this solution at all times in order to minimize tissue
swelling and loss of transparency.
Femtosecond Laser Micromachining
[0100] For laser micromachining, the corneas were trimmed to
generate small, flat pieces of tissue, averaging .about.1 cm.sup.2.
Each piece of cornea was then flattened onto a clear glass slide
(1.times.3 inches, 1 mm thick, Surgipath Medical Industries Inc.,
IL) with the epithelium facing up and the endothelium facing down.
A glass coverslip (Corning No. 0211 Zinc Titania glass) was placed
on the top of each piece of cornea or lens, stabilizing it for the
duration of the experiment. The ethylene glycol/sucrose storage
solution was used as mounting medium to prevent or at least
minimize dehydration of the cornea and lens since these effects are
known to alter the refractive index and transparency of both these
tissues.
[0101] Femtosecond laser micro-machining was conducted as
previously described in U.S. patent application Ser. No.
11/745,746, filed May 8, 2007 and U.S. patent application Ser. No.
11/948,298, filed Nov. 30, 2007. The laser source was a Kerr-lens
mode-locked Ti:Sapphire laser (K-M Labs). This laser oscillator
generates pulses averaging 300 mW, pulse duration of 27 fs and a 93
MHz repetition rate at 800 nm wavelength. A continuously variable,
metallic, neutral density filter was inserted into the optical path
and used to adjust the incident laser power onto each cat cornea
piece. The femtosecond laser pulses were focused 100 .mu.m below
the tissue surface using a 60.times., 0.70 NA Olympus LUCPlanFLN
long-working-distance microscope objective. Because the large
amount of glass within the microscope objective induces significant
chromatic dispersion into the femtosecond laser pulses, greatly
broadening the femtosecond pulse durations, we used a standard
extra-cavity-prism double-pass configuration to compensate for the
dispersion and maintain the ultrashort pulse duration. By carefully
adjusting this dispersion compensator, we obtained nearly
transform-limited 27 fs duration pulses at the focal point of the
focusing objective which were measured by a collinear
autocorrelator using 3rd order surface harmonic generation (THG).
During femtosecond laser micromachining, the slide containing the
biological tissue samples was mounted on a 3D scanning platform
consisting of a Physik Instrumente (PI) P-622.2CD XY scanning stage
with 250 .mu.m travel range and 0.7 nm close-loop resolution, and a
Newport VP-25XA linear servo Z-axis scanning stage with 25 mm
travel range and 100 nm resolution. An infrared CCD camera was used
to monitor the micromachining process and the generation of visible
plasma luminescence in real-time.
[0102] Our experiments were conducted at room temperature
(.about.25.degree. C.). It took about 40 minutes to create a (100
.mu.m.times.50 .mu.m) grating and conduct the immediate
post-micromachining measurements. Corneal trimming and mounting did
not exceed 10 minutes in duration, and the corneal tissue was
exposed to ambient air during the trimming process for at most 2
minutes. Application of the Ti:Sapphire femtosecond as described
resulted in the formation of micromachined gratings having 20 to 40
lines into the stromal layer of the corneas; each line
approximately 1 .mu.m wide, 100 .mu.m long and 5 .mu.m apart.
Refractive index changes in the micromachined regions were
calculated immediately and after one month of further storage by
measuring the intensity distribution of diffracted light when the
gratings were irradiated by 632.8 nm wavelength He--Ne laser light.
Because we observed no significant changes in cornea and lens
transparency or thickness at the end of our micromachining
experiments, we conclude that the described micromachining process
did not cause significant corneal or lenticular dehydration or
swelling.
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