U.S. patent application number 11/820773 was filed with the patent office on 2007-12-20 for spectral domain optical coherence tomography system.
Invention is credited to Kabir M. Arianta, Christopher J.R.V. Baker, Matthew J. Everett, Claus Flachenecker, James P. Foley, Martin Hacker, Jochen M.M. Horn, David Landhuis, Scott A. Meyer, Yue Qiu, Jochen Straub, Clement Louis-Rene Viard.
Application Number | 20070291277 11/820773 |
Document ID | / |
Family ID | 38861216 |
Filed Date | 2007-12-20 |
United States Patent
Application |
20070291277 |
Kind Code |
A1 |
Everett; Matthew J. ; et
al. |
December 20, 2007 |
Spectral domain optical coherence tomography system
Abstract
An ophthalmic imaging device for improved ophthalmic imaging
including: an optical coherence scanning device, a fundus imaging
device: an iris viewer; a motorized chin rest; an internal test
target, and a fixation target device wherein the optical coherence
scanning device, the ophthalmic scanning device, the iris viewer,
and the fixation target device all share at least one common
optical element. The optical coherence device preferably employs a
Mach-Zehnder interferometer with an all fiber reference path;
monitoring and attenuating power within the reference path. The
multiple devices are separately and in combination aligned with the
eye. The system includes internal and external calibration and
improved image formats.
Inventors: |
Everett; Matthew J.;
(Livermore, CA) ; Meyer; Scott A.; (Livermore,
CA) ; Hacker; Martin; (Jena, DE) ; Horn;
Jochen M.M.; (San Francisco, CA) ; Baker; Christopher
J.R.V.; (Moraga, CA) ; Arianta; Kabir M.;
(Livermore, CA) ; Foley; James P.; (Fremont,
CA) ; Straub; Jochen; (Pleasanton, CA) ; Qiu;
Yue; (Pleasanton, CA) ; Landhuis; David;
(Dublin, CA) ; Flachenecker; Claus; (Hayward,
CA) ; Viard; Clement Louis-Rene; (San Francisco,
CA) |
Correspondence
Address: |
STALLMAN & POLLOCK LLP
353 SACRAMENTO STREET
SUITE 2200
SAN FRANCISCO
CA
94111
US
|
Family ID: |
38861216 |
Appl. No.: |
11/820773 |
Filed: |
June 20, 2007 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60815107 |
Jun 20, 2006 |
|
|
|
60925104 |
Apr 18, 2007 |
|
|
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Current U.S.
Class: |
356/497 ;
356/479 |
Current CPC
Class: |
G01N 21/4795 20130101;
G01B 9/02044 20130101; G01B 9/02091 20130101; G01B 9/02077
20130101; G01B 2290/65 20130101; G01B 2290/70 20130101; G01B 9/0203
20130101; A61B 3/102 20130101 |
Class at
Publication: |
356/497 ;
356/479 |
International
Class: |
G01B 11/02 20060101
G01B011/02; G01B 9/02 20060101 G01B009/02 |
Claims
1. An ophthalmic imaging device comprising: an optical coherence
tomography (OCT) system including a first light source for
generating a first radiation beam; a fundus imaging system
including a second light source generating a second radiation beam;
optics for combining the first and second radiation beams and
directing the combined beams into the eye of a patient in a manner
such that the OCT system and the fundus imaging system are
confocal; and an iris viewing system including an imaging device
and optics for obtaining an image of the iris along an axis common
with the combined first and second radiation beams whereby the OCT
system and the fundus imaging system can be aligned with the
patient's eye based on images generated by the iris viewing system
and wherein the images generated by the OCT system and the fundus
imaging system are aligned.
2. An imaging device as recited in claim 1, further including a
fixation system for generating an image of a target and including
optics for projecting the image of the target into the eye along
axis common with the combined first and second radiation beams to
aid the patient in rotating the eye relative to the OCT and fundus
imaging systems.
3. The imaging device as recited in claim 1, wherein the OCT system
is a frequency domain optical coherence tomography scanner.
4. The imaging device as recited in claim 1, wherein the fundus
imaging system is a line scanning ophthalmoscope.
5. The imaging device as recited in claim 1, wherein the OCT system
contains a Mach-Zehnder interferometer.
6. The imaging device as recited in claim 1, wherein the OCT system
contains an optical coupler in which one tap is used to monitor
optical power.
7. The imaging device as recited in claim 1, for imaging an eye
further comprising: the OCT system having a component conjugate to
an image plane in a first region of the eye; and the a fundus
imaging system having a component conjugate to an image plane in a
second region of the eye; wherein the first and second regions of
the eye are at different depths.
8. An ophthalmic imaging device comprising: an optical coherence
scanning device; a fundus imaging device; an iris viewer; a
motorized chin rest; and a fixation target device wherein the
optical coherence scanning device, the ophthalmic scanning device,
the iris viewer, and the fixation target device all share at least
one common optical element.
9. The ophthalmic imaging device as recited in claim 8, wherein at
least one common optical element is a lens.
10. The ophthalmic imaging device as recited in claim 8, further
comprising an ocular lens, said ocular lens being the optical
element closest to the patient, wherein the motorized chin rest
positions the patient with respect to the imaging device; and the
ocular lens is arranged to be disposed to be adjusted in position
wherein the position of the ocular lens is separately adjustable
relative to the chin rest and jointly adjustable with the chin rest
relative to the remainder of the optical elements of the imaging
device.
11. A method for ophthalmic imaging comprising: moving the subject
to align focus of a first optical device; aligning a plurality of
optical devices while retaining focus alignment of the first
optical device; creating an image of the interior of an eye with an
optical coherence scanning device; creating a fundus image of the
posterior of the eye; creating an image of the iris of the eye; and
projecting a fixation target to be viewed by the eye wherein the
optical coherence scanning device, the ophthalmic scanning device,
the iris viewer, and the fixation target device are all optical
devices and all operate essentially simultaneously.
12. The method of claim 11, wherein the first optical device is an
iris viewer.
13. The method of claim 11, wherein the optical coherence scanning
device is a spectral domain Optical Coherence Tomography
scanner.
14. The method of claim 11, wherein the fundus imaging device is a
line scanning ophthalmoscope.
15. The method of claim 11, wherein least one common optical
element is a lens.
16. The method of claim 11, wherein the interior portion of the eye
scanned by the optical coherence scanning device is the retina.
17. The method of claim 11, wherein the interior portion of the eye
scanned by the optical coherence scanning device is the cornea.
18. The method of claim 17, wherein the optical coherence scanning
device is switched from scanning the retina to scanning the cornea
by inserting a lens in the OCT beam path.
19. The method of claim 11, wherein the subject is moved using a
motorized chinrest to center the image of the iris.
20. The method of claim 18, wherein the ophthalmic alignment
focuses the fundus image by moving the subject and an optical
component of the ophthalmic device concurrently.
21. The method of claim 20, wherein the optical component of the
ophthalmic device is an ocular lens.
22. The method of claim 21, wherein the OCT image of the eye is
created by: setting an OCT depth range; adjusting an OCT
polarization compensation; selecting an OCT scan pattern; adjusting
an OCT transverse scanning region; and acquiring an OCT image of
the eye.
23. The method of claim 22, further comprising the step of setting
the patient's prescription.
24. The method of claim 22, further comprising the step of
adjusting a lens within the ophthalmic scanning device to optimize
image brightness.
25. An ophthalmic imaging device comprising: an optical coherence
scanning device; a fundus imaging device; an iris viewer; an
internal test target; and a fixation target device wherein during
an optical examination, the optical coherence scanning device, the
ophthalmic scanning device, the iris viewer, and the fixation
target device all share at least one common optical element.
26. The method of claim 25, wherein the internal test target
comprises a series of horizontal and vertical strips.
27. An optical coherence tomography scanner comprising a
longitudinal delay device, an interferometer, a corner cube, and a
transverse scanner, wherein the longitudinal delay device lies
between the interferometer and the transverse scanner.
28. An optical coherence tomography scanner as recited in claim 27,
wherein the optical coherence source beam is in free space between
the longitudinal delay and the transverse scanner.
29. An optical coherence tomography scanner as recited in claim 28,
wherein the longitudinal delay device includes the corner cube.
30. An optical coherence tomography scanner as recited in claim 29,
wherein the corner cube is mounted on a rail and the rail is
mounted on a plate where the plate design reduces thermal variation
misalignment of the corner cube as the corner cube traverses the
rail.
31. An optical coherence tomography scanner as recited in claim 29,
wherein at least one surface of the corner cube is coated to reduce
reflections.
32. An optical coherence tomography scanner as recited in claim 27,
wherein the optical coherence scanning device further includes a
scanner adjustment device for the purpose of adjusting the center
of the beam on a central axis of the scanner.
33. An optical coherence tomography scanner as recited in claim 32,
wherein the adjustment device includes a mirror.
34. An optical coherence tomography scanner as recited in claim 32,
wherein the adjustment device adjust the optical coherence beam to
maintain a position error below 10 microns at the entrance to the
scanner while the longitudinal delay is varied by more than 30
mm.
35. An optical coherence tomography scanner as recited in claim 32,
wherein the adjustment reduces the phase shift associated with
scanning.
36. An optical coherence tomography scanner as recited in claim 35,
wherein the adjustment device maintains a phase shift below the
phase shift corresponding to 1 mm/s axial motion of the sample,
while the scanner scans the OCT beam over more than 2 mm on the
subject and while the longitudinal delay is varied by more than 30
mm.
37. An optical coherence tomography scanner as recited in claim 27,
wherein the optical coherence scanning device further includes an
alignment mechanism capable of aligning the beam parallel to the
axis of the longitudinal delay.
38. An optical coherence tomography scanner as recited in claim 37,
wherein the alignment mechanism reduces the transverse motion of
the optical coherence beam at the scanner associated with changes
in longitudinal delay.
39. An optical coherence tomography scanner comprising a reference
path in fiber and a reference power device within the reference
path for setting the reference power.
40. An optical coherence tomography scanner as recited in claim 39,
wherein the reference power device is a fiber tap and the reference
power is set by the coupling ratio in the fiber tap.
41. An optical coherence tomography scanner as recited in claim 40,
wherein the fiber tap removes power from the reference path, and
some of the removed power is received by an optical detector.
42. An optical coherence tomography scanner as recited in claim 39,
wherein the reference power device is a length of fiber with
optical loss.
43. An optical coherence tomography scanner as recited in claim 39,
wherein the optical coherence tomography scanner is a frequency
domain optical coherence tomography scanner.
44. An optical coherence tomography scanner as recited in claim 39,
wherein the reference power device includes a fiber-based device
through which optical transmission depends on polarization state of
the light, and further includes a fiber-based polarization
controller.
45. An optical coherence tomography scanner comprising a reference
path in fiber and a sample path and operating at a wavelength
having significant chromatic dispersion mismatch between the
reference path and the sample path.
46. An optical coherence tomography scanner as recited in claim 45,
wherein the sample path includes materials of high chromatic
dispersion.
47. An optical coherence tomography scanner as recited in claim 45,
where the reference and sample paths include loops of substantially
equal path length, and in which the loops are individually
compensated to have insignificant polarization mode dispersion.
48. An optical coherence tomography scanner as recited in claim 45,
where the reference path fiber is routed with bends in different
planes, the radii of these bends being chosen to substantially
cancel the bending-induced polarization mode dispersion.
49. An optical coherence tomography scanner as recited in claim 45,
including at least one path in optical fiber along which the
optical fiber is routed with bends in different planes, the radii
of these bends being chosen to substantially cancel the
bending-induced polarization mode dispersion.
50. An ophthalmic imaging device comprising: a first imaging system
having a component conjugate to an image plane in a first region of
the eye; a second imaging system having a component conjugate to an
image plane in a second region of the eye; wherein first and second
regions of the eye are at different depths; and wherein first and
second imaging systems share a common portion of an optical
path.
51. A device as recited in claim 50, wherein the first imaging
system is a retinal imager and the associated image plane
corresponds to a layer posterior to the RPE and the second imaging
system is a scanning retinal imager and the associated image plane
is located anterior to the image plane of the first imaging
system.
52. A method of generating images of a eye using an optical
coherence tomography (OCT) system, said system including a light
source for generating a radiation beam and a scanning mechanism for
moving the beam to a plurality of positions within an X/Y plane and
wherein the OCT system obtains a measurement of a reflectance
distribution within the eye as a function of depth (Z) at each X
and Y position, said method comprising the steps of: obtaining a
first set of depth scans a spaced positions along an X axis at a
first Y position; obtaining a second set of depth scans at spaced
positions along an X axis at a second Y position near the first Y
position and wherein the X positions of the second set of depth
scans are offset from the X positions in the first set of depth
scans; and generating a two dimensional slice image along the X
axis as a function of depth by treating the measurements as if they
were obtained along a common Y position.
53. A method of generating images of a eye using an optical
coherence tomography (OCT) system, said system including a light
source for generating a radiation beam and a scanning mechanism for
moving the beam to a plurality of positions within an X/Y plane and
wherein the OCT system obtains a measurement of a reflectance
distribution within the eye as a function of depth (Z) at each X
and Y position, said method comprising the steps of: scanning the
beam generally along the X direction at a particular Y axis
position, said Y axis position defining a centerline of the scan,
and while scanning the beam, varying the position of the beam on
the Y axis about the centerline while obtaining depth scans at
spaced X positions; and generating a two dimensional slice image
along the X axis as a function of depth by treating the
measurements as if they were obtained along a common Y
position.
54. A method of generating images of a eye using an optical
coherence tomography (OCT) system, said system including a light
source for generating a radiation beam and a scanning mechanism for
moving the beam to a plurality of positions within an X/Y plane and
wherein the OCT system obtains a measurement of a reflectance
distribution within the eye as a function of depth (Z) at each X
and Y position, said method comprising the steps of: scanning the
beam generally along the line and while scanning the beam,
dithering the lateral position of the beam with respect to the line
while obtaining depth scans at a plurality of positions; and
generating a two dimensional slice image along the line as a
function of depth by treating all the measurements as if they were
obtained along said line.
55. A device as recited in claim 54, wherein the line is
curved.
56. A method of obtaining B-scan data with reduced speckle in
optical coherence tomography (OCT) comprising: acquiring a
plurality of OCT A-scans; and forming a B-scan from said A-scans,
wherein the A-scans within one resolution cell of the B-scan
contains a subset of A-scans which are speckle diverse both tangent
to and orthogonal to the B-scan at that cell.
57. A device as recited in claim 56, wherein the B-scan is
substantially planar.
58. A device as recited in claim 56, wherein the B-scan is
substantially cylindrical.
59. A device as recited in claim 56, wherein the subset of A-scans
are compounded to represent the resolution cell of the B-scan.
60. A device as recited in claim 59, wherein the subset of A-scans
are sequentially acquired.
61. A method for improving the long term performance of an
ophthalmic scanning device in which a galvanometer is positioned by
a motor comprising: generating a command to direct the motor
driving the galvanometer to a desired position; characterizing the
closed-loop response of the galvanometer to the generating command;
and adjusting one of the command or the closed loop response to
achieve a desired galvanometer position.
62. A method as recited in claim 61, wherein the adjusting step
modifies the position command.
63. A method as recited in claim 61, wherein the closed-loop
response of the galvanometer includes a loop filter controlling a
servo and the adjusting step adapts the loop filter to keep the
closed-loop response constant.
Description
PRIORITY
[0001] This application claims the benefit of the filing date under
35 U.S.C..sctn.119(e) of Provisional U.S. Patent Application Ser.
No. 60/815,107, filed on Jun. 20, 2006, and Provisional U.S. Patent
Application Ser. No. 60/925,104, filed on Apr. 18, 2007, which are
hereby incorporated by reference in their entirety.
TECHNICAL FIELD OF THE INVENTION
[0002] The subject invention relates to diagnostic and measurement
devices for evaluating a patient's eye. In particular, a spectral
domain optical coherence tomography system is disclosed.
BACKGROUND OF THE INVENTION
[0003] Optical Coherence Tomography (OCT) is a technology for
performing high-resolution cross sectional imaging that can provide
images of tissue structure on the micron scale in situ and in real
time. OCT is a method of interferometry that uses light containing
a range of optical frequencies to determine the scattering profile
of a sample. Optical coherence tomography (OCT) as a tool for
evaluating biological materials was first disclosed in the early
1990's (see U.S. Pat. No. 5,321,501 for fundus imaging.). Since
that time, a number of manufacturers have released products based
on this technology. For example, the assignee herein markets a
device called the StratusOCT. This device is used for diagnostic
imaging and provides direct cross sectional images of the retina
for objective measurement and subjective clinical evaluation in the
detection of glaucoma and retinal diseases. The device can generate
images of macular thickness, the retinal nerve fiber layer, the
optic disc, the cornea, and other parts of the eye. This device is
based on a version of OCT known as time domain OCT.
[0004] In recent years, it has been demonstrated that frequency
domain OCT has significant advantages in speed and signal to noise
ratio as compared to time domain OCT (Leitgeb, R. A., et al.,
Optics Express 11:889-894; de Boer, J. F. et al., Optics Letters
28: 2067-2069; Choma, M. A., and M. V. Sarunic, Optics Express 11:
2183-2189).
[0005] In frequency domain OCT, a light source capable of emitting
a range of optical frequencies excites an interferometer, the
interferometer combines the light returned from a sample with a
reference beam of light from the same source, and the intensity of
the combined light is recorded as a function of optical frequency
to form an interference spectrum. A Fourier transform of the
interference spectrum provides the reflectance distribution along
the depth within the sample.
[0006] Several methods of Frequency domain OCT have been described
in the literature. In spectral-domain OCT (SD-OCT), also sometimes
called "Spectral Radar" (Optics letters, Vol. 21, No. 14 (1996)
1087-1089), a grating or prism or other means is used to disperse
the output of the interferometer into its optical frequency
components. The intensities of these separated components are
measured using an array of optical detectors, each detector
receiving an optical frequency or a fractional range of optical
frequencies. The set of measurements from these optical detectors
forms an interference spectrum (Smith, L. M. and C. C. Dobson,
Applied Optics 28: 3339-3342), wherein the distance to a scatterer
is determined by the wavelength dependent fringe spacing within the
power spectrum. SD-OCT has enabled the determination of distance
and scattering intensity of multiple scatters lying along the
illumination axis by analyzing a single the exposure of an array of
optical detectors so that no scanning in depth is necessary.
Typically the light source emits a broad range of optical
frequencies simultaneously. Alternatively, in swept-source OCT, the
interference spectrum is recorded by using a source with adjustable
optical frequency, with the optical frequency of the source swept
through a range of optical frequencies, and recording the
interfered light intensity as a function of time during the sweep
(U.S. Pat. No. 5,321,501).
[0007] The commercial OCT systems typically include some form of
scanning mirror configuration to scan the light beam across the eye
in a plane perpendicular to the propagation axis of the beam. The
most common interferometer configuration for OCT is the Michelson
interferometer [FIG. 1a of U.S. Pat. No. 5,321,501]. Michelson
interferometers return some reference arm light to the source,
causing a conflict between the desire to set the reference level
for best performance of the detector, and to set the reference
level low enough to be below the back-reflection tolerance. Some
alternative interferometer topologies allow the reference path to
be completely in fiber, allowing simple construction. If the
reference path is completely in fiber then the sample path length
can be varied instead (U.S. Pat. No. 5,321,501).
[0008] Non-reciprocal optical elements in the source arm [U.S. Pat.
No. 6,657,727 issued to Izatt, et al.] have been used to divert the
reflected light that would otherwise return to the source to a
detector. While this protects the source and increases its
longevity, non-reciprocal optical elements in the source arm add
significant costs to the interferometer manufacture.
[0009] Interferometers with topology different from the common
Michelson topology have been proposed for OCT (U.S. Pat. No.
5,321,501 FIG. 10, U.S. Pat. No. 6,201,608 issued to Mandella, et
al., and U.S. Pat. No. 6,992,776 issued to Feldchtein, et al.).
Some of these designs route the reference light without
retro-reflecting or otherwise reversing the reference light back
toward the source. In such interferometer designs some light
returned from the sample can reach the source, but this is less of
a concern because in many applications only a small fraction
(10.sup.-4 to 10.sup.-10) of the incident light is scattered from
the sample and returned to the interferometer.
[0010] There has been a continuing effort in the industry to
improve the existing OCT systems. For example, when measuring
living tissue such as an eye, movement during the measurement
period can cause a wide variety of difficulties. Efforts have been
made to increase the speed of data collection to reduce the effects
of motion of the subject. In addition, various approaches have been
suggested to measure sample motion and then compensate for that
motion.
BRIEF DESCRIPTION OF THE DRAWINGS
[0011] FIG. 1 is an optical path system diagram including the OCT
path, the LSO path, the Fixation path, the Iris Viewer path, and
the internal test target path.
[0012] FIG. 2 is a system block diagram including the OCT system,
the LSO system, the Fixation system, the Iris Viewer system, and
the motorized chin rest.
[0013] FIG. 3 is an electrical block diagram of the PC and system
peripherals.
[0014] FIG. 4 is a polarization paddle design.
[0015] FIG. 5 is a flow chart of scan states during the operation
of the instrument.
[0016] FIG. 6 is an example internal test target.
[0017] FIG. 7 illustrates analysis of the internal test target.
[0018] FIG. 8 illustrates one topology of an interferometer for
OCT.
[0019] FIG. 9 illustrates another topology of an interferometer for
OCT.
[0020] FIG. 10a illustrates the delay line alignment.
[0021] FIG. 10b illustrates a consequence of delay line
misalignment.
[0022] FIG. 11 illustrates a mounting block for the corner cube
rail.
[0023] FIG. 12 illustrates an embodiment for bending fiber as a
method of birefringence compensation.
[0024] FIG. 13 illustrates a system with a large number of optical
components working off the left side of a retinal conjugate, with
the ocular lens and eye on the right hand side of the retinal
conjugate.
[0025] FIG. 14a is a table of steps used for alignment of an OCT
imaging system.
[0026] FIG. 14b is a table of optional steps used for alignment of
an OCT imaging system.
[0027] FIG. 15a illustrates a grid of A-scans for acquiring a 3-D
OCT volume.
[0028] FIG. 15b illustrates a thick B-scan achieved by combining
2B-scans.
[0029] FIG. 15c illustrates a thick B-scan achieved by combining 3
horizontal B-scans.
[0030] FIG. 15d illustrates a thick B-scan achieved by combining 3
vertical B-scans.
[0031] FIG. 16a illustrates a wiggle pattern for acquiring speckle
reduced B-scans.
[0032] FIG. 16b illustrates how the A-scans of a traditional B-scan
relate to speckle at a given depth.
[0033] FIG. 16c illustrates how the A-scans acquired for speckle
reduced B-scans relate to speckle at a given depth.
[0034] FIG. 17 illustrates the Optical lay out of a test eye.
[0035] FIG. 18 illustrates the cross section of a retina.
[0036] FIG. 19a illustrates the traditional optical alignment of
two optical devices without chromatic aberration alignment.
[0037] FIG. 19b illustrates the traditional optical alignment of
two optical devices with chromatic aberration alignment.
[0038] FIG. 19c illustrates the preferred optical alignment of two
optical devices with chromatic aberration alignment.
[0039] FIG. 20 shows, in pictorial form, a conventional scanning
optical system and fundus camera.
[0040] FIG. 21 shows, in pictorial form, one embodiment of the
current invention.
[0041] FIG. 22 shows, in pictorial form, another embodiment of the
current invention.
[0042] FIG. 23 shows, in pictorial form, yet another embodiment of
the current invention.
DISCLOSURE OF THE INVENTION
[0043] This document is intended to describe a new OCT system being
developed by the assignee herein. The principals and applications
of the invention are set forth in part in the description which
follows, and, in part, will be obvious to those skilled in the art
from the description provided herein. Further advantages may be
learned by practice of the invention. The scope of the invention is
defined by the claims, which includes known equivalents and
unforeseeable equivalents at the time of filing of this
application. This new system is a spectral domain optical coherence
tomographer including a spectrometer. The system also includes a
line scanning ophthalmoscope (LSO) and an iris viewing system.
Certain aspects of the individual sub-systems are unique. In
addition, the combination of these sub-systems is also unique.
[0044] Some of the inventive concepts being employed in the subject
OCT system have been described in earlier filed patent applications
which will be referenced herein all of which are incorporated by
reference. This disclosure is intended to describe the overall
system.
[0045] FIG. 1 is a schematic of the principal optical components of
the system 10 specially designed to generate diagnostic images of
an eye 30. The system 10 includes four primary sub-systems, an
optical coherent tomography (OCT) system 20, a line scanning
ophthalmoscope (LSO) 40, a fixation system 50, and an iris viewer
60.
[0046] The OCT system is a spectral domain system generally of the
type described in the above-cited articles. The OCT system 20
includes a low coherence light source 101 which in this case is a
super luminescent diode that has bandwidth of about 800 to 900 nm
with a center wavelength of 840 nm. One choice for this device is
the SD371 manufactured by Superlum Diodes in Moscow. The output
from the SLD 101 is directed into a fiber based interferometer 100.
An input fiber 110 is connected to a first port I of a 70/30
optical coupler 111. The coupler 34 directs thirty percent of the
light out of port III to the sample arm fiber 112 and seventy
percent of the light out of port IV to the reference arm fiber 113.
The reference arm fiber 113 is optionally connected to an optical
attenuator 119. An optical attenuator is useful for attenuating
excess signal passing through the reference arm. A variable optical
attenuator can also compensate for the variability of other
components. In particular, a variable optical attenuator can be
optimized during system manufacture to account for attenuation
differences between parts and ensure that the signal transmitted to
port I of a 99/1 coupler 131 is sufficient. This configuration
defines a transmissive reference path, which has an advantage over
Michelson interferometers in that the reference light is not
returned back to the source 101.
[0047] A photodetector 132 in the interferometer monitors light
coupled out of coupler 131. Alternatively, a monitoring detector
could be positioned in the reference arm fiber 113. The
photodetector 132 is used to measure the power of the source both
for eye safety purposes and to monitor the degradation of the
source.
[0048] Sample arm fiber 112 directs light into a delay line
implemented using corner cube 124. The corner cube is translatable
along an axis as indicated by arrows A to change the path length of
the sample arm. The path length of the sample arm is adjusted with
respect to the reference arm to select the depth in the tissue at
which the OCT image will be centered. The light exiting the corner
cube 124 is directed to a pair of scanning galvanometer mirrors 116
for scanning the beam in a plane perpendicular to the propagation
axis of the beam. The light is then passed through a lens doublet
defined by lens 150 and 152. A turning mirror 151 is interposed
between the lenses. It is preferable that the spacing between
lenses, and optical path length within lenses, along the OCT beam,
such as between lenses 150 and 152, be greater than the free space
OCT depth range so that reflections off the surfaces of the lenses
will not create interference effects that might be interpreted as
coming from structures in the eye.
[0049] The light beam is then turned towards the eye using a
dichroic beam splitter 160. The dichroic beamsplitter 160 functions
to directs light from the OCT path to the common optical path used
by various subsystems and redirects OCT return light back along the
OCT path while redirecting other light backscattered from the eye
along a different path. The light is directed into the eye with a
lens 162. Lens 162 is designed to correct for spherical aberrations
in the eye. In order to compensate for refractive error, we adjust
position of lens 162 with respect to the remainder of the optics.
Light entering the eye is reflected back from various structures in
the eye such as retinal layers. The reflected light travels back on
the same path to input port III of coupler 111. Coupler 111 directs
seventy percent of the light reflected from the source and returned
to port II to the combining coupler 131 while thirty percent of the
light returned from the source returns to the light source through
port I. Alternatively, the 70/30 coupler 111 may be an 80/20
coupler or a 90/10 coupler.
[0050] Coupler 131 functions to combine light returning from the
sample arm and arriving from the reference arm to create
interference effects. A majority of this combined light is directed
out of the coupler 131 at port III to a spectrometer 200. Further
information about interferometer designs having a transmissive
reference path can be found below in the section APPARATUS FOR
OPTICAL COHERENCE TOMOGRAPHY.
[0051] A polarization paddle is provided to optimize signal
strength. OCT depends on interference between sample and reference
beams, and the interference of these beams produces a modulation in
power to the extent that the polarizations of the beams match.
Specifically, if one uses the common Poincare sphere representation
of polarizations, the amplitude of the interference fringes is
proportional to the cosine of half the angle on the Poincare sphere
between the Poincare sphere representations of the sample and
reference polarizations.
[0052] Rotating birefringent elements are a common method of
controlling polarization. In fiber optics, bending the fiber is a
convenient means for creating birefringence, and rotating the
orientation of the bends rotates the axis of birefringence. Such an
assembly is often called a polarization paddle. (See, for example,
chapter 9 of "Polarized Light in Fiber Optics" Edward Collett, (c)
PolaWave Group Lincroft N.J. 2003).
[0053] Perfect polarization matching requires three polarization
paddles to compensate for arbitrary polarizations in the sample and
reference arm. In practice, perfect polarization compensation is
not required; only sufficient polarization alignment is necessary
to enable detection of interference. A single paddle is sufficient
to compensate for most polarization differences seen in practice so
that detection loss is no more that 1-2 dB. The single paddle
reduces equipment cost and simplifies the design and control and
improves exam throughput efficiency.
[0054] In the preferred embodiment, the single paddle is located in
the sample arm but it could be located in the reference arm. A
motor and hardware is included for a rotatable paddle to provide
polarization compensation for the fiber. The fiber can be mounted
with a U-shape bend onto the paddle as shown in FIG. 4, or in the
more traditional circular loop. The paddle can rotate out of the
plane of the paper with the U-shape remaining in one plane at all
times. The design parameters are the three radii (R.sub.1, R.sub.2
and R.sub.3) and the three angles (.alpha..sub.1, .alpha..sub.2,
.alpha..sub.3). These parameters are chosen to meet birefringence
requirements as discussed below.
[0055] If the paddle is located in a single-pass portion of the
interferometer, such as the reference path of a Mach-Zehnder
design, then one expects theoretically that a quarter-wave paddle
will be effective in matching polarizations. Suppose we are given
two arbitrary polarizations and an adjustable quarter-wave plate
affecting the first polarization. Imagine the locations A and B of
the polarization states on the Poincare sphere if we removed the
quarter wave plate. The goal is to move A as close as possible to B
by rotating A on the sphere by 90.degree. about an equatorial axis
of our choosing. Choose an axis x which puts both points in the
hemisphere x>0. Sighting along the z axis, we can see both
points on the same hemisphere, and want to move A by either
+90.degree. or -90.degree., whichever will bring A closer to B.
This choice corresponds to choosing to place either the fast or
slow axis of the wave plate in the direction corresponding to +x.
It is always possible to move the first point so that it is (1) in
the same x>0 hemisphere as the second point, and (2) within
.+-.90.degree. azimuthally about the z axis. The resulting distance
between A and B is always less than 90.degree.. The <90.degree.
result is best understood by visualization, but can also be proven
using the law of cosines on the three directions A, B and x. The
resulting fringe amplitude, then, is at least cosine 45.degree. or
71% of what it would be with optimally matched polarization.
[0056] If the paddle is located substantially at the end of a
bi-directional path, then two passes through a rotatable one-eighth
wave paddle will have the same benefits as derived above for the
single pass through a quarter-wave paddle. If the paddle is located
in a bi-directional path but located such that the light
experiences significant uncontrolled birefringence, such as by the
sample, experimentation and simulation have shown that three-eights
of a wave of birefringence more robustly restores the interference
fringe amplitude.
[0057] The spectrometer 200 is of the type disclosed in U.S. patent
Ser. No. 11/196,043, filed Aug. 3, 2005, (publication 2007/0030483)
incorporated herein by reference. Briefly, the spectrometer is in a
folded Littrow configuration. Light enters the spectrometer and is
directed to a grating 230. Grating 230 is preferably blazed for 840
nm, with approximately 1200 lines/mm to give adequate spectral
dispersion. Light reflected from grating 230 is directed to a pixel
camera 250. The dispersion of the grating and the imaging lenses
discussed below are chosen to spread wavelength from approximately
800 to 900 nm over the sensitive length of camera 250. A set of
three lenses 240, 210 and 220 is located between the grating and
the camera. The light passes through these three lenses both on the
path to the grating and on the return path to the camera. While
each of the lenses contributes to focusing and correction, lens 220
is the primary lens for focusing and collecting light from the
grating. The grating is tilted in a way to induce conical
diffraction which causes the returning light beam to be displaced
away from the fiber input, slightly out of the plane of the figure,
and towards the camera. The primary function of lens 210 is to
correct for effects of conical diffraction. Lens 240 functions
primarily as a field flattening lens.
[0058] The SD OCT system is capable of generating two- or
three-dimensional images of the retina in a manner known in the
prior art. The subject system has some additional capabilities that
will be described below. However, one added feature is the ability
to rapidly switch between imaging the retina to imaging of the
cornea. This capability is provided by including an extra lens 180
which can be moved into the path of the light in the sample arm to
permit focusing of the light onto the cornea. In conjunction with
the movement of the lens 180 into position, the position of corner
cube 124 is changed to allow the path length difference between the
sample arm and reference arm to correspond to the position of the
cornea. An advantage of this system is that the information about
the cornea can be easily obtained without having to reposition the
patient.
[0059] As an alternative to moving corner cube 124, mirrors can be
moved to switch in and out an extra fold in the optical path
length, as illustrated for example in FIG. 5 of U.S. patent
application Ser. No. 11/243,665, filed Oct. 5, 2005, (publication
2007/0076217) incorporated herein by reference. Other alternatives
to moving corner cube 124 are the rapid-scan optical delay (RSOD)
devices disclosed in International Patent Application No. WO
2005/033624 and in U.S. Pat. No. 6,654,127, either of which can be
configured to provide a change in group-delay, with relatively
small phase delay. The small phase delay is advantageous here
because changes in phase delay move the interference fringes across
the camera 250, which reduces the fringe intensity if the fringes
move during exposure of camera 250. The group-delay devices with
small phase delay cause relatively less motion of the fringes the
device is settling after a quick move.
[0060] While the optical delay does not need to be particularly
rapid for centering the imaging region, the ability to change the
optical delay rapidly is useful for adjusting the optical delay on
a scan-by-scan basis. For example, an RSOD can be used to flatten
the retina. Nominally, the retina will appear as a curved surface
in a 3-D image of the eye. Using a predetermined delay profile, an
RSOD can adjust the optical delay on each A-scan and flatten the
curved surface. When sufficiently fast computation elements are
available along with hardware feedback paths, on-the-fly optical
delay adjustments can be computed from A-scan to A-scan.
Alternatively, a sparse scan can obtain a select complement of
A-scans which can be used to identify the retina in each A-scan and
compute a fit to the retinal surface (say a spherical or parabolic
fit), which can be used to generate a delay profile for the RSOD.
Clearly, combinations of pre-determined profiles and on-the-fly
computations can also be used to direct the RSOD to modify the
optical delay on an A-scan by A-scan basis in a SD OCT system.
[0061] The second main sub-system is a fundus viewer. The preferred
fundus viewer technology is the line scanning ophthalmoscope LSO
40. The LSO 40 includes a relatively narrow band light source
which, in the preferred embodiment, is a super luminescent diode
410 emitting light at about 755 nm with bandwidth about 5 nm. The
light source is polarized. Light from source 410 is passed through
shaping optics 415 to create a line of light. The line of light is
directed to a beam splitter 420 which redirects the light to
scanning galvanometer mirror 430 for scanning the line in one axis
perpendicular to the plane of propagation of the light. Beam
splitter 420 comprises a reflective strip, so that illumination
from source 410, focused to a line along this strip, is directed to
the eye, while light returning from the eye largely passes around
this strip toward imaging lens 480. The illuminating light is
directed to a dichroic beam splitter 440 which is reflective of
light at 755 nm and transmissive at shorter wavelengths. The light
is passed through a lens doublet 450 to a dichroic beam splitter
460. Beam splitter 460 is reflective of light in the 700 nm
wavelength region and transmissive for light at 755 nm and 550 nm
wavelengths. Beam splitter 460 needs to have high transmissivity at
755 nm only for the polarization of light used in the LSO
subsystem. The design of the dichroic coatings on beam splitter 460
is easier if only one polarization state needs to be optimized.
(Analogous design optimization is available for beam splitters 460
and 160.) Light is then passed through beam splitter 160 into the
eye. Beam splitter 160 is reflective of light over 800 nm and
transmissive of light at shorter wavelengths. The beamsplitter 420
is nearly conjugate to the cornea, so that an image of the LSO
light source reflected from the patient's cornea is formed on the
reflective strip, thus blocking this corneal reflection from the
imaging optics.
[0062] Light from the LSO 40 is reflected by the eye and returns on
the same path to splitter 420. A portion of the reflected light is
transmitted through splitter 420 and it is imaged via a lens 480
onto a line scan camera 490. Commercially available line scan
cameras offering line rates around 10 kHz are appropriate for
camera 490. As the galvanometer 430 is scanned, different portions
of the retina are illuminated and imaged, so that a two-dimensional
image of the patient's retina is built up from successive exposures
of the camera. With 512 lines in a frame, a frame rate of 20 Hz is
achieved. The scan range of galvanometer 430 is easily variable to
adjust the field of view of the LSO.
[0063] The third main sub-system is an iris viewer 60. The iris
viewer is used primarily to align the patient's eye with the
optical axis of the device. The iris viewer includes an LED 610
positioned near lens 162 for illuminating the eye. Preferably, the
LED generates light having a wavelength of about 700 nm. The
reflected 700 nm light is captured by lens 162 and travels back
through splitter 160 to splitter 460 where it is reflected back
through a series of lenses to a CMOS camera 620. The LED can be
polarized, or its output filtered by a polarizer, so that the light
reflected from the iris is largely polarized, and beam splitter 460
optimized to reflect only one polarization state. Imaging the iris
in polarized light has the side effect of revealing birefringence
of the cornea. The camera 620 generates an output which is supplied
to a monitor that will display an image of the iris. As discussed
below, this image is used to position the patient.
[0064] The fourth main sub-system is a fixation system 50. Fixation
system 50 includes a display pad 510 for generating fiducial marks
that will be projected onto the patient's eye. The patient will be
asked to fixate her eye on these fiducial marks. Pad 510 generates
light at a visible wavelength preferably between 450 and 600 nm.
The light from pad 510 is conditioned by lens system 520 and
directed through dichroic beam splitters 440, 460 and 160 and
focused into the eye via lens 162. The preferred fixation target is
a variable sized, 2D fixation target. A 2D fixation target provides
both a center fixation target and the ability to rapidly change
visual stimuli for analysis of eye response. Preferably, the target
size is variable from a point target to an oversized target
embedded in a 120.times.120 pixel display covering a field of view
of 30 degrees.
[0065] As shown in the system block diagram of FIG. 2, a host
computer 90 is used to interface between the operator 35 and the
integrated system to control the subsystems, either directly or
through a controller such as a motor controller board 85, and to
send and receive subsystem data, either directly or indirectly
through and intermediary controller such as a frame grabber board
45. As shown in the electrical block diagram of FIG. 3, the host
computer 90 provides the system with input devices (such as
keyboard 91, mouse 92, or equivalents such as trackball or
joystick) and output devices (such as monitors 93 and printers 94)
as well as Input/Output devices such as digital storage devices
such as hard drives (not shown), CDs (not shown), DVDs 95, etc.,
connection ports such as serial ports (not shown), parallel ports
(not shown), USB ports 96, fire-wire ports (not shown), and the
like, and network connections 97 to local, peer-to-peer,
distributed, or even the world-wide web. The live iris camera 98 is
preferably a direct PC peripheral, though it could also be
integrated into the frame grabber, providing yet another image
input to that device. The architecture of the frame grabber board
can enable real time tracking by image processing an image on the
host PC and updating the galvo X and Y offsets on the frame grabber
in real time. In this embodiment, the galvo scan pattern is
corrected and the appropriate region is imaged, even in the
presence of eye motion.
[0066] In order to improve the functionality of the device, a
specific effort was made to insure that the various sub-systems
worked together in a cooperative manner. For example, the OCT, LSO
and iris viewer are all telecentric systems, so that adjusting
focus does not change the magnification of the image. The optical
systems that focus on the retina, the LSO, OCT, and fixation, are
parfocal so that they are simultaneously in focus on the retina
after compensation for refractive error. The systems that focus on
the retina use different wavelengths, so their focus adjustments
are calibrated to compensate for the different focal lengths of the
human eye at these various wavelengths. The systems that image the
retina, OCT and LSO, are confocal systems, meaning that small areas
of illumination are swept across the retina and images of these
areas directed to matched sensitive areas on the detectors.
Confocal imaging reduces glare from corneal reflections and
scattering from other ocular media such as a cataract.
[0067] This design takes care to minimize polarization dependence
in the optics along the OCT beam path. For example, differences in
optical delay between the polarization states, known as
polarization mode dispersion (PMD) cause OCT images with different
depths for each polarization state. Given that the polarization
state changes on transmission through the eye, polarization is
difficult to fully control, and PMD generally leads to broadening
of the axial (depth) resolution in the OCT image. The dichroic beam
splitters along the OCT path reflect, as opposed to transmit, the
OCT beam because the beam reflected from dichroic coatings
typically has less PMD than the transmitted beam. Smaller
polarization dependent effects, such as fractions of a wave of
birefringence, are also controlled. The beam splitters are placed
in locations where the OCT beam is telecentric, meaning the chief
rays of the OCT beams for various positions of the scanner 116 are
parallel, so that the angular-dependent polarization effects of the
beam splitters do not change as the OCT beam is scanned.
[0068] In use, the first step is to align the patient with the
device. In the preferred embodiment, the patient's head is put into
a motorized headrest. A suitable headrest is described in U.S.
patent application Ser. No. 10/843,767, filed May 12, 2004
(publication No. 2005/0254009) which is incorporated herein by
reference. The doctor will ask the patient to view the fiducial
marks generated by the pad 510. At the same time, the doctor will
observe the eye via a display (not shown) associated with the
camera 620 of the iris viewer. Initially, the distance between the
patient's eye and the lens 162 is adjusted for best focus of the
iris. Once the proper spacing has been achieved, the separation
between the lens 162 and eye is held constant while the position of
the eye with respect to the OCT system is varied to position the
center of the OCT image at the desired depth within the eye. Lens
162 is not carried by the motorized chin support 80 (FIG. 2).
Rather, a separate translation system is provided which is
operatively linked to the motion of the chin support during this
positioning step. Further information about the approach used to
position and align the patient's eye can be found below in the
section METHOD OF PATIENT ALIGNMENT FOR MULTIFUNCTIONAL FUNDUS
IMAGING.
[0069] The primary purpose of the iris viewer is to help the
operator center the patient's pupil so that the OCT and LSO beams
pass through the pupil to the iris. A continuous view of the iris
is helpful in keeping the patient's pupil centered during retinal
imaging. Note that the iris viewer can also be used to help
position the OCT beam around cataracts. Further, it can be used to
help collect OCT images through different portions of the pupil,
collecting light at different scattered angles from the retina.
[0070] Once the patient has been aligned, a wide variety of OCT
images can be generated. The Fourier transform of the signals from
the spectrometer provide A-scan information at each X and Y
position of the beam. (In some methods scans are repeated at the
same X and Y position to reveal time-dependent effects including
pulsatile flow, Doppler shifts, etc.) This data can be collected
and stored. Some A-scans can be acquired for purposes other than
imaging. For example, the scanning system 116 can direct the OCT
beam in a circle outside the aperture of lens 162, during which
time the camera records the reference signal only, with no signal
from the sample, thus collecting a background signal for use in
processing. The processor can then generate and display other image
information (such as B-scans, en face images, Doppler images, etc.)
familiar to the doctor. In addition to some of the more
conventional imaging modalities currently available on existing
systems, the subject apparatus has been configured to provide
additional functionality.
[0071] For example, the system is configured to generate fundus
type images based on OCT data. This approach is described in U.S.
patent application Ser. No. 11/219,992, filed Sep. 6, 2005
(publication 2006/0119858) and incorporated herein by reference. In
this type of analysis, the intensity information over the depth
range for the OCT data at any particular X/Y location is integrated
to generate a pixel in the fundus image. The integration of
intensity over a depth range to generate the fundus pixel may be
performed by either accumulating intensities prior to compression
(nominally logarithmic) for display or by compression of
intensities prior to accumulation. The fundus image can be
continuously displayed for the doctor to help interpret the OCT
images and position the device. This fundus image based on OCT data
is especially valuable for registration of the location of the
underlying OCT cross-sections, to an en-face view of the
retina.
[0072] The OCT may be used to generate maps with three-dimensional
rendering of elevation, topographical maps or color or grayscale
maps. U.S. patent application Ser. No. 11/717263, Mar. 13, 2007,
and incorporated by reference, discloses a variety of approaches
including collecting compound OCT scans for high definition scans
and a data cube to provide context for high definition scans. Also
disclosed are standardization techniques for orientation,
diagnostic metrics of texture and heterogeneity, retinal fluid
maps, etc.
[0073] The section METHOD FOR COMBINING B-SCANS ("THICK B-SCAN"),
below, discloses the concept of combining adjacent B-scans to
reduce noise and speckle and give an enhanced visual
impression.
[0074] The software can be set up to generate elevation maps of
tissue with respect to fitted reference surfaces. This approach is
described in U.S. patent application Ser. No. 11/223,549, filed
Sep. 9, 2005 (publication 2007/0103693) and incorporated herein by
reference.
[0075] The system may also be set up so that the chromatic
dispersion of the sample and reference paths are different from
each other to create a variation in the relative group delay as a
function of optical frequency between the sample and reference
paths. Thereafter, the measured interference spectrum can be
multiplied by a complex phase factor to compensate for the
mismatch. In this manner, the image contrast between reflections
from the sample and image artifacts can be increased so that the
doctor can better discern actual tissue images. Further information
on this approach is set forth in U.S. patent application Ser. No.
11/334,964, filed Jan. 19, 2006, (publication 2006/0171503)
incorporated herein by reference.
[0076] As noted above, one problem associated with prior art
systems relates to errors resulting from the movement of the
patient's eye during imaging. Errors of this type are reduced in
the subject system because the scanning speed is much faster. For
comparison, the time needed to scan the eye using our current
Stratus system is on the order of 2 seconds, while the subject
system can cover the same scan region in only 0.026 seconds.
[0077] The increase in speed is so great that new scanning
sequences can directly collect 3-D imaging data without the need
for intervening tomograms. In some cases, performing scanning
sequences collecting data in 2-D planar tomograms and then building
a 3-D volume from the 2-D slices is preferable because then
existing software can be used for visualization, reducing costs and
time-to-market. Nonetheless, direct collection of 3-D voxel data in
real-time using spectral domain optical coherence systems is now
possible and the 3-D volume can be rendered directly for
display.
[0078] In addition to increasing the scanning speed, other
approaches have been developed to still further reduce problems
associated with eye movement during measurement. For example, U.S.
patent application Ser. No. 11/331,567, filed Jan. 13, 2006
(publication No. 2006/0164653) and incorporated herein by reference
discloses the concept of taking a few partial, fast OCT scans and
using this information to provide registration information during
the slower, more complete OCT scans. In another approach, the LSO
system 40 can be used to generate guideposts that can then be
compared in the processor to the OCT images. The information can be
used to control the scanning of the galvanometer mirrors 116 in
real time to compensate for patient eye movement. Alternatively,
the LSO data can be used in post-processing to properly register
the data acquired from the OCT system. More information about this
approach can be found in U.S. patent application Ser. No.
11/389,351, filed Mar. 24, 2006 (publication 2006/0228011) and
incorporated herein by reference.
[0079] It is also desired that the device exhibit long term
repeatability and stability in the field. In the past, external
targets where used by the doctor to facilitate alignment and
calibration. The subject system has been provided with an internal
calibration system to simplify the process of making sure the OCT
and LSO systems are coaxially aligned. More specifically, a target
710 is provided which preferably includes fiducial marks such as
crosshairs or horizontal and vertical alignment bars (see FIG. 6).
An auxiliary mirror 720 is provided located just beyond the zone
where the OCT and LSO beams are scanned when measuring the patient.
During a calibration step, the galvanometer mirrors 116 and 430 are
positioned so that the light from the OCT and LSO systems strike
mirror 720 and are directed to target 710. The reflected light is
imaged by the two respective detection systems. The driving systems
for mirrors 116 and 430 are adjusted until the images overlap.
Information defining the position of the galvanometer mirrors is
stored and used to calibrate the device (see FIG. 7). The
crosshairs in target 710 can be grooves to facilitate imaging by
the OCT device. Target 710 preferably includes reflectors at
various depths for calibration of axial length measurements by OCT
and confirmation of the axial resolution of the OCT system.
[0080] Yet another use for internal calibration is a galvanometer
test. After scanning many times (up to billions of cycles), the
motors of optical scanning galvanometers mechanically wear. Before
catastrophic failure, their bearings and lubricants deteriorate and
affect performance. These motors are typically driven by servo
amplifiers that attempt to minimize the difference between the
actual galvanometer motor position and a commanded position. The
actual position is provided either as an analog or digital signal.
In ophthalmic scanning, image quality and repeatability is directly
dependent on galvanometer performance, so it is important to be
able to characterize the closed loop response and adapt performance
to achieve the desired response. The desired position may be
achieved by adapting the loop filter of the servo or the command
signal. In cases where there is no adequate internal alternative
available to achieve the required performance, the system can issue
a request service. The service request may be either through a
notice to the user on the system or a notice across a network to
either administrative personnel or directly to the service
organization.
[0081] This internal calibration may be performed on an internal
schedule, such as monthly or weekly or on every n-th boot (where
n-is a positive integer), or through remote service and
diagnostics.
[0082] The arrangement of the sub-systems leads to some novel
combination. For example, and as noted above, the OCT system, LSO,
iris viewer and fixation system are all parfocal. The iris viewer,
which preferably displays a continuous image of the iris, greatly
facilitates the alignment of the OCT system measurement system.
This approach can be compared to the prior art approach, often used
in fundus camera, of using the retinal imaging system, here an LSO,
to first image the iris at a distance spaced significantly from the
optimal positioning necessary to obtain an OCT image. In order to
then position the device to obtain a OCT and LSO images of the
retina, the doctor would have to carefully move the patient and
imaging system closer together, along a line without deviation so
that the imaging paths remain centered on the pupil. This
adjustment would be difficult because the doctor would no longer
have the image of the iris displayed.
[0083] FIG. 2 is a block diagram of the overall system. It shows
the operator 35 interacts with the computer 90. The computer 90
acquires Iris Viewer image data directly from the Iris Viewer 60
while fundus and OCT image information are received through the
frame grabber board 45. Alternatively, in some instances, one or
more high performance graphics boards can substitute for the frame
grabber board 45 however, because of the multiple sources of image
frames, specialized hardware was preferable for the preferred
design. The controller board 85 performs the real-time system
control for the OCT system 20, the fundus viewer 40, the fixation
target 50, the iris viewer 60, and the motorized chin rest 80.
[0084] FIG. 5 is a flow chart of one embodiment of the OCT scan
states during operation of the instrument. The OCT system remains
is a system idle state until the operator indicates the start of a
new acquisition 920. Since this chart is concerned with the scan
states of the OCT system, the flow in this chart assumes that the
iris image is already in focus. Then the first task is to align the
scan. First we perform a background scan 930 to determine current
scan position. The scan is then aligned 940 by setting the OCT
reference depth using the Z-motor and moving the combined ocular
lens and eye with synchronized x-, y-, and z-motion (using the
chinrest and Z-motor in combination) to optimize the focus of the
fundus image. The system process 950 intermittently returns the OCT
system to the background scan state 940 to ensure that the
alignment scan state 940 is performed using correct background
information. The system remains in the alignment scan state until
the operator starts the acquisition state 970, unless returned to a
background scan state or a system timeout 960 occurs. The system
timeout state 960 is entered after a fixed interval in which the
operator has not determined to acquire data. In this embodiment,
once the operator decides to acquire data, first a background scan
is performed 980 and the acquisition scan is performed 990. On
acquisition scan completion, the system enters a review data state
995 and the operator can review the acquired data through various
image display and analysis tools. On completion of the data review,
the system is returned to the system idle state 910.
Apparatus for Optical Coherence Tomography
[0085] The following embodiments describe interferometers for use
in the invention of record. Coherence-domain imaging techniques
such as OCT preferably use light sources with short axial coherence
length, but with spatial coherence in the transverse directions.
Superluminescent diodes, which are similar in structure to diode
lasers, have short temporal coherence and broad spatial coherence.
By design, they do not lase because there is insufficient optical
feedback. Superluminescent diodes are typically sensitive to
optical back-reflection of output light potentially causing output
power fluctuations and shortened lifetime.
[0086] The most common interferometer configuration for OCT is the
Michelson interferometer. Most Michelson interferometers return
some reference arm light to the source. The light returning to the
source can be diverted by the use of non-reciprocal optical
elements. To avoid the expense of non-reciprocal optical elements,
one can control the polarization state of the light and divert
light returning to the source based on its polarization state.
[0087] Some interferometer topologies allow the reference path to
be completely in fiber, allowing simple construction. Other
interferometers using essentially the same topology allow the
reference path to be nearly completely in fiber, only deviating
from continuous fiber to insert simple free-space optics, such as a
leakage optical attenuator. In OCT, the optical group delays must
be approximately matched between sample and reference paths. This
matching is typically accomplished by adjusting the reference
optical path length. If the reference path is completely in fiber
then the sample path length can be varied instead, as noted in U.S.
Pat. No. 5,321,501, c. 12,11. 16-21.
[0088] The OCT apparatus disclosed herein efficiently collects
light from the eye, uses a reflective sample path, returns no
reference light to the source, and does not require circulators or
other non-reciprocal elements.
[0089] Extreme split ratios in the fiber couplers can be avoided
and one configuration allows a safety monitor tap close to the
sample arm tap.
[0090] FIG. 8 illustrates the topology of an interferometer for OCT
that reduces reflections of the light back into the source. Low
coherence light source 101 is typically a superluminescent diode
(SLD) which typically tolerates back reflection of less that
10.sup.-3 of its output light. The SLD is coupled to source fiber
110 that routes light to directional coupler 111a. The optimal
directional strength of the coupling depends on system design
choices and may be 90/10 (as shown in FIG. 8) or 70/30 (as shown in
FIG. 1) or other as availability permits. Directional coupler 111a
splits the light into sample fiber 112a and reference fiber 113a.
The sample path includes delay apparatus 114 to adjust the length
of the sample path; shown in more detail in FIG. 9. The delay
apparatus couples the light from fiber 112a to a free-space OCT
beam 115. Transverse scanner 116 deflects the OCT beam and
preferably creates a focus in the beam near the region of interest
in sample 30a.
[0091] Some light scattered from sample 30a returns through the
scanner and delay apparatus to sample fiber 112a. Coupler 111a
routes this light through loop 117a to fiber coupler 131a, where it
is interfered with the reference light. The combining coupler 131a
provides two outputs. These outputs could be used for balanced
detection (U.S. Pat. No. 5,321,501 FIG. 10) in which both detector
200 and detector 142c are used to collect light for OCT.
Alternatively, the coupling ratio of coupler 131 a can be adjusted
to send most of the interfered light to a single OCT detector 200.
Each OCT detector can be a single photodetector for use in
time-domain OCT or swept-source OCT, or a spectrometer for use in
spectral domain OCT.
[0092] Optional tap 121 diverts a fraction of the reference light
to detector 122, which may be used to monitor the source power.
(Some reasons for monitoring include safety of the sample and
detection of degradation in the source 101.) The tap removes some
fraction of optical power from the reference fiber 113a, reducing
the power that reaches coupler 131a. Sensitivity in OCT can reach
the shot-noise limit if the reference power is large enough to
bring the interference signal above receiver noise, but not so
large as to bring intensity noise or beat noise above the level of
shot noise. The reference power is approximately determined by the
source power, and the coupling ratios in directional couplers 111a
and 131a, and adjusted by choice of tap 121.
[0093] The coupling ratios in directional couplers 111a, 131a and
121 are chosen to set a safe level of illumination to the sample,
and to set the appropriate reference power at the detector or
detectors. For example, in the case of ophthalmic OCT of the retina
using light with wavelengths near 850 nm, the safe exposure level
is approximately 0.5 mW, and the optimum reference level at the
detector is approximately 0.005 mW. Sources are available in this
wavelength range having output power of approximately 5mW. For
these conditions one would use a coupling ratio near 90%/10% in the
splitting coupler 111a so that 10% of the source power reaches the
sample. 90% of the scattered light will then be routed to loop
117a. In the case where there is a single OCT detector 200, the
combining coupler 131 a preferably routes most of the sample light
to that detector. The splitting coupler routes 90% of source light,
4.5 mW, to reference fiber 113a, while only 0.005 mW is required at
the detector. One could use a combining coupler 131a that couples
0.1% of the reference light into the single OCT detector 200, but
in manufacture it is difficult to control the 0.1% coupling factor.
A preferred solution is to use a 99%/1% split ratio in combining
coupler 131a, and take advantage of the additional degree of
freedom in tap 121 to adjust the reference power. Nominally,
tapping 89% of the power form reference fiber 113a will provide an
appropriate reference level of 0.005 mW at OCT detector 200, in
this example.
[0094] As an alternative to adjusting the tap ratio of optional tap
121, one can adjust the reference level by including attenuating
fiber (U.S. Pat. No. 5,633,974) in the reference path.
[0095] FIG. 9 illustrates one of several other possible
interferometer topologies. Low coherence light from source 101 is
divided by coupler 111b between fiber 112b and reference fiber
113b. Sample-routing coupler 141 further splits the light from
fiber 112b between monitor 142 and sample fiber 112c. Light in
sample fiber 112c is delayed by apparatus 114 and scanned by
scanner 116 across sample 30b, and some light scattered from the
sample is returned through these devices to sample fiber 112c. Some
of the returned light, preferably a large fraction, is routed by
sample-routing coupler 141 to combining coupler 131b, where it is
interfered with the reference light. Again, one or both of the two
outputs of combining coupler 131b can be used to detect the signal
for OCT.
[0096] Considering an example as for FIG. 9, with a 5 mW source,
the appropriate sample power can be achieved if splitting coupler
111b directs 90% of the source light to fiber 112b and 10% to the
reference path, and sample-routing coupler 141 couples 12% of the
light in fiber 112b to the sample fiber 112c. This split ratio in
coupler 141 routes 88% of the light returned from the sample from
fiber 112c to fiber 117b and toward the detector. The appropriate
reference level is obtained if the combining coupler 131b couples
1% of the power in the reference fiber to the detector, allowing
99% of the light in fiber 117b to reach detector 200.
[0097] Preferably, the path length to the sample is changed while
maintaining the OCT beam focus and without changing the range of
sample to be scanned. One practical solution in ophthalmic imaging
is placement of the path length adjustment between the output of
the interferometer and the scanner. However, adjusting the path
length can cause the OCT beam to move transversely, offsetting it
from the center of the entrance aperture to the scanner. In typical
scanners, this offset causes a phase shift in the OCT beam as the
beam is scanned. Such phase shifts cause signal loss or positioning
artifacts in frequency-domain techniques of OCT.
[0098] FIG. 10a illustrates the delay stage. Light from sample
fiber 112 is collimated by lens 120 to form OCT beam 128. Alignment
mirror 123 is one way to align the direction OCT beam 128 to be
parallel to the travel of the moveable corner cube 124; the
importance of this alignment is discussed later. Alternatively,
beam 128 can be aligned by mechanically moving lens 120 in
conjunction with the light emitting end of sample fiber 112. A
moveable corner cube is one way to vary the optical path length of
the OCT beam, in order to approximately match the optical group
delays between the sample path and reference path. Adjustment
mirror 125 directs the OCT beam to scanner 116. The scanner 116 can
be implemented using a pair of rotatable mirrors 126 and 127, which
in conjunction with scan lens 155 scan the OCT beam across sample
30.
[0099] If the OCT beam is not centered on the axis of rotation of
scan mirrors 126 and 127, then as these mirrors rotate the optical
path length to the sample is changed, as explained for example by
Podoleanu (Podoleanu, A. G., G. M. Dobre, et al. "En-face coherence
imaging using galvanometer scanner modulation." Optics Letters
23(3): 147-149 (1998)). The effect of the scanner on the sample
path length is doubled because the return path of light scattered
from the sample back to fiber 112 is also affected. This change in
optical path length causes a phase shift in the interferogram. A
continuous phase shift corresponds to a shift in optical frequency,
and such a frequency shift due to relative motion is generally
termed a Doppler shift. This Doppler shift has undesirable effects
on the data collection by frequency-domain OCT techniques, as
explained by Yun et al. (Yun, S. H., G. J. Tearney, et al. "Motion
artifacts in optical coherence tomography with frequency-domain
ranging." Optics Express 12(13): 2977-2998 (2004)).
[0100] Adjusting mirror 125 can be tipped and tilted to center the
OCT beam on the axis of rotation of mirrors 126 and 127. The
Doppler shift due to scanning can be easily be measured by the OCT
system, so as to provide a Doppler signal to be nulled by
adjustment of mirror 125. One way to measure this signal is to
provide a non-moving sample 30, repeatedly scan the OCT beam across
the sample, and record closely-spaced OCT interferograms. Pairs of
neighboring interferograms should be recorded from locations of
tissue that are close compared with the optical resolution of the
scanner, so the sampled regions significantly overlap. Pairs of
neighboring interferograms differ largely in the phase shift caused
by the optical path length change associated with transverse
scanning. The phase shift between neighboring interferograms is
thus a measure of the phase shift associated with the scanner, and
provides a signal which is zero when the OCT beam is properly
centered. Scanning the beam in alternate directions produces an
alternating phase shift associated with the scanner, allowing one
to distinguish this phase shift from other effects, such as the
Doppler shift due to unintentional motion of the sample. Scanning
each of mirror 127 and 126 separately produces a phase shift
proportional to the misalignment of the OCT beam off the respective
axes of rotations of these mirrors.
[0101] If the center of the beam is mis-positioned by as little as
0.5 mm, then the phase shift induced by rotation of the galvo is
significant. The galvo rotates 0.7 degrees mechanical per
millisecond during a 20-degree cube. The motion of the mirror at
the beam center is 6 mm/s, moving 240 nm during a 40 .mu.s
exposure. This motion is sufficient to cause significant fringe
washout. The misalignment tolerance follows from the acceptable
sensitivity loss due to fringe washout. The sensitivity loss due to
axial motion can be found from Yun et al [Optics Express 12(13):
2977-2998 (2004)] and in terms of decibels the loss in sensitivity
is 2.9 dB (q .DELTA.z).sup.2=18 dB (.DELTA.z/.lamda.).sup.2.
Requiring the axial sensitivity loss to be less than 0.5 dB yields
that the .DELTA.z due to mirror misalignment should be less than
0.167*.lamda.=0.14 .mu.m for .lamda.=f nm. During the exposure of
one A-scan from in a 128.times.128 cube covering 20.degree., we
move the beam 0.16.degree. in the patient's field of view. For a
typical optical setup, the pupil will be imaged on the scanning
mirrors, but with magnification typically 2.5, so that the mirrors
need rotate only 1/5 of the angular sweep of the beam at the
patient's pupil. Thus, the scan mirror rotates by 0.03.degree.
during the exposure of one A-scan. The resulting misalignment
tolerance is then 0.14 .mu.m/[2 tan(0.03.degree.)].apprxeq.135
.mu.m. Note that the tolerance to lateral misalignment, between the
OCT beam and the rotation axes of the scanners, scales with the
size of the image of the pupil on the scanner.
[0102] In some applications, there is a fast scan direction and a
slow scan direction. For example, mirror 127 may scan rows across
sample 320 and mirror 126 may move less often to move the OCT beam
between scan rows. In these situations one degree of freedom is
relatively more important in the adjustment of mirror 125. In
general there is one direction of scan that is relatively faster
than another, and in general there is one direction for which
stable alignment is relatively more difficult. The design will
preferably choose the more stable alignment direction to be the
direction associated with phase shifts due to the faster direction
of scan.
[0103] The measured phase shifts associated with scanning each
mirror 126 and 127 can provide feedback to drive adjusting mirror
125 to the position that gives a null phase shift. Such a feedback
system would allow the apparatus to self align during operation, if
the subjects are relatively still.
[0104] Alternatively to feedback using OCT, the correct position of
the OCT beam can be marked by other means. For example, a beam
splitter can direct a small fraction of light from OCT beam 128 to
a position-sensitive detector that is preferably close to the
location conjugate to the rotation axes of mirrors 126 and 127. The
proper position of the OCT beam is associated with the signal
values output from this position-sensitive detector during a
condition of correct adjustment. In operation, the system can
adjust mirror 125 to restore the signal from the position-sensitive
detectors that corresponds to correct adjustment of the OCT beam
position.
[0105] The correct adjustment of the OCT beam on the scanner can be
adversely affected by motion of the delay stage 114. If the OCT
beam 128 is not aimed to be parallel to the direction of motion of
corner cube 124, then the transverse position of the
retro-reflected beam will change upon translation of the corner
cube. FIG. 10a illustrates two positions of the corner cube 124 and
the respective positions of the retro-reflected OCT beam. In FIG.
10a, the drift in adjustment has been corrected using the feedback
mechanisms discussed previously so that the beam position out of
the corner cube after translation of the corner cube coincides with
the position of the beam out of the corner cube before translation.
FIG. 10b illustrates two positions of the corner cube 124 and the
respective positions of the retro-reflected OCT beam. Beam 133 is
the retro-reflected OCT beam directed to scanning mirrors 126 and
127 before translation of corner cube 124, while beam 135 is the
retro-reflected OCT beam directed to scanning mirrors 126 and 127
after translation of corner cube 124. It is illustrated here that
the beam out of the corner cube is not only delayed, but it is also
translated unless the entrance beam is properly aligned.
[0106] Feedback correction of the adjustment mirror 125 will be
easier, and possibly un-necessary, if the OCT beam is well aligned
to the direction of travel of delay stage 114. Such alignment can
be implemented by appropriate tip and tilt of alignment mirror 123.
One method for alignment of mirror 123 is an extension of the
method used to adjust mirror 125. The phase shift associated with
scanning (or the position of the OCT beam on a position-sensitive
detector) can be measured for two locations of corner cube 124.
Mirror 123 is aligned null any change in phase shift (or
position-sensor signal) with motion of the corner cube 124.
[0107] The delay rail that moves cube 124 is preferably mounted in
an effectively kinematic way, to avoid misalignments of the OCT
beam caused by strains in the optical mounts, such as those caused
by thermal expansion. For example, FIG. 11 shows such a rail 171
mounted to plate 172 via bolts 175 and 176. If the bolt holes in
plate 172 become improperly spaced due to thermal expansion of
plate 172, then rail 171 could become bent out of plane of the
figure; and the OCT beam could be shifted transversely from its
desired location. Cutting holes 191 and 192 into plate 172 allows
the remaining plate material in 193 to flex, so that the bolts 175
and 176 can maintain the proper spacing to match their holes in
rail 171, and so that rail 171 remains straight. The mounting is
effectively kinematic because it effectively relaxes the axial
(z-axis in the figure) constraint on the position of rail 171.
Without holes 191 and 192, bolts 175 and 176 imposed competing
constraints on the axial position of the rail; the flexibility of
material 193 relieves the redundant constraint.
[0108] Alternatives to expansion holes include: applying a heat
sink to the support plate 172 or manufacturing the system so that
excess heat does not accumulate at plate 172. Alternatively, plate
172 can be manufactured from materials with sufficient strength to
support the rail, but a low enough expansion coefficient to prevent
unacceptable flexing of the rail 171. Alternatively, combinations
of these mechanisms or others can be used to ensure proper
alignment of the corner cube 124 during system operation.
[0109] Corner cube 124 is often constructed from solid glass, using
internal reflections to guide the beam. The remaining surface of
the corner cube can produce weaker reflections. Such reflections
are undesirable in an OCT system because if they either return to
the fiber 112, or follow paths parallel to the main OCT beam, they
can produce additional interference signals corresponding to
different optical delays from that of the main beam. The additional
interference signals can result in ghost images. If a corner cube
is used in the longitudinal delay device, intentional misalignment
or anti-reflection coating can be used to reduce reflections.
[0110] The OCT interferometer of FIG. 8 has its reference path
entirely in fiber, while the sample path contains some air. (Air
space is required for example in delay line, scan optics and
working distance from optics to the sample). The different
chromatic dispersions of fiber and air cause the relative optical
delay between reference path and sample path to vary across optical
frequency (see, for example, co-pending U.S. patent Ser. No.
11/334,964, filed Jan. 19, 2006, publication 2006/0171503,
incorporated herein by reference). This variation in optical delay,
if not corrected, leads to an uncertainty in the optical path
length to each scattering center in the sample, worsening the axial
resolution of the resulting tomograms. At certain wavelengths, such
as near 1300 nm, the chromatic dispersion of fiber is near zero, so
an interferometer configuration in FIG. 1 requires no correction
for mismatched dispersion. For many applications of OCT different
wavelengths are preferred, such as retinal OCT in which absorption
by water in the eye would absorbs 95% of 1300 nm scattered from the
retina, before that light exits through the front of the eye.
[0111] In order to manage the mismatch in chromatic dispersion,
some elements in the sample path, which tends to have lower
dispersion than the all fiber reference path, can be constructed
using highly-dispersive glasses. For example, flint glass has
significantly greater chromatic dispersion than optical fiber, so
constructing corner cube 124 from flint glass significantly reduces
the mismatch in chromatic dispersion. Each 1 mm of flint glass
substituting for crown glass in the sample path approximately
balances the chromatic dispersion mismatch resulting form the
inclusion of 6mm air in the sample path. Substituting sufficient
flint glass for crown glass can also overcompensate, and reverse
the sign of dispersion mismatch, if desired.
[0112] Previous OCT devices required balanced chromatic dispersion
between sample and reference paths. If the reference path is
entirely, or nearly entirely, in fiber and some of the sample path
is in air, there is typically a mismatch in chromatic dispersion.
Such devices perform best when using wavelengths for which the
chromatic dispersion of the optical fiber is nearly zero. This
restriction limits the device applications to those where the
operation wavelength is chosen based on chromatic dispersion
properties and not based on subject penetration or image
optimization. Optical devices can be built to compensate for
dispersion but often at the cost of optical loss, so these devices
in the sample path would typically reduce sensitivity.
Alternatively, one can numerically compensate the chromatic
dispersion mismatch. Numerical compensation has benefits as
described in the above cited U.S. Patent Publication No.
2006/0171503. These benefits work best when the physical dispersion
mismatch is within bounds, so, so even with numerical compensation
some method of controlling the dispersion is desired.
[0113] Having the reference path entirely or primarily in fiber
does increase the opportunity for polarization mode dispersion
(PMD) in the fiber (Raja) which causes an undesirably variability
in optical path length with respect to the polarization state. When
building any fiber interferometer one often has to make splices,
which can fail, and to make the lengths of the fibers correct to
match the optical path length of the reference and sample arms.
Therefore one wants to be able to re-cut and re-splice the fiber.
This is typically facilitated by placing extra loops in each of the
sample and reference arms, with one loop from each arm removed each
time the fiber is re-cut and re-splice. Therefore, one wants to
have a considerable number of fiber loops. This leads to additional
length of fiber and the potential for considerable PMD. The desire
to fit the fibers in a small space increases the polarization mode
dispersion, because bending induced polarization mode dispersion
increases with smaller bend radius.
[0114] PMD can be reduced by careful routing of the fiber; for
example, the PMD caused by bends in a horizontal plane can be
compensated by following the horizontal bend with a vertical bend
that provides approximately the opposite PMD. Such local
compensation has the advantage that the net birefringence change is
zero when a loop is removed, as for re-splicing of the fiber.
Another advantage is that the compensating birefringence has the
same temperature-dependence as the birefringence to be compensated,
as they arise from the same physical cause. FIG. 12 illustrates one
way to achieve these compensating bends 144. Each pass of fiber
through the configuration drawn approximately compensates the
birefringence in a larger fiber loop (not shown) having a total
bend of 360-degrees with bend sections having 24 mm bend radius.
The 201-degree bends of radius 14 mm fiber-bend in an orthogonal
plane have approximately equal-magnitude birefringence of opposite
sign.
[0115] In summary, the interferometer configurations of FIGS. 8-12
and their equivalents, make efficient use of light returned from
the eye, compared to a Michelson interferometer. Such
configurations enable setting appropriate reference signal levels
at the detector. By using a variable delay in the sample arm, no
reference arm reflector is required, which reduces the number of
fiber-couplings and avoids delay-dependent variations in the
reference arm signal level. No fiber moves or bends with this
configuration. Although various embodiments that incorporate the
teachings of the present invention have been shown and described in
detail herein, those skilled in the art can readily devise many
other varied embodiments that still incorporate these
teachings.
Method of Patient Alignment for Multifunctional Fundus Imaging
[0116] The following embodiment included in one variation of the
present invention describes a method of patient alignment for
fundus imaging. This embodiment uses a suitable headrest, like the
one described in U.S. patent application Ser. No. 10/843,767, filed
May 12, 2004 (publication 2005/0254009) which is incorporated
herein by reference. In this method, whose optical paths are shown
in FIG. 13, one moves the patient head relative to an ocular lens
162 to set the human pupil at the entrance pupil of the instrument,
then moves the ocular lens and patient head together to correct for
refractive error. In FIG. 13, the vertical line 165 left of the eye
and ocular lens indicates the position of the retinal image formed
by the ocular lens; this line is the retinal conjugate. The figure
illustrates a system with a large number of optical components to
the left of a retinal conjugate, with the ocular lens and eye to
the right of the retinal conjugate. In the description below, the
terms headrest and chinrest are used interchangeably, referring to
a suitable headrest as described in U.S. Patent Publication
2005/0254009 capable of head support and functionality that moves
the patient's eye 30 at least along the optical axis, denoted as
the z axis.
[0117] The chief ray of the scanning OCT beam and the rays of light
used in the fundus imager both form ray pencils with a vertex at
the center of the entrance pupil of the instrument. The scanning
galvanometers in an OCT scanner or scanning ophthalmoscope
determine the location of the vertex of the set of chief rays in
the scanning beam. To get the beams into the eye the entrance pupil
of the instrument must overlap the pupil of the eye.
[0118] It is advantageous to simultaneously maintain a focused
image of the pupil of the eye for guidance in positioning of the
eye so that the OCT and fundus microscope optical paths pass
through the pupil of the eye.
[0119] The refractive error of the human eye varies over a range of
approximately .+-.20 diopters. Therefore, there is a need to focus
any OCT sample beam and the imaging optics of a fundus microscope
to compensate for the refractive error of the human eye.
[0120] While making these two adjustments, it is advantageous to
keep the working distance small (for better field of view without
excessive size of optics) but yet large enough for patient safety.
This leads to designs where the entrance pupil is at a fixed, safe,
distance from the closest lens to the patient.
[0121] The Visucam non-mydriatic fundus camera, uses a separate
off-axis iris view for alignment. The two adjustments are 1)
camera-to-patient distance to set the working distance, guided by
the iris camera and 2) refractive correction by moving a lens
within the camera.
[0122] Use of an ocular lens with a slit-lamp comprises moving the
biomicroscope portion of the slit lamp to focus on the retinal
image formed by the hand-held ocular lens.
[0123] Fundus cameras typically move an internal lens for
compensation of refractive error. Typically a retinal conjugate is
formed in the instrument, at a location depending on the patent's
refractive error; at this location the pupil of the eye is
typically imaged at infinity. A moveable lens within the instrument
is moved to focus on this retinal conjugate. The pupil of the eye
is typically imaged at the back focal plane of this imaging
lens.
[0124] U.S. Pat. No. 5,537,126 describes how to move the beam
scanning mechanism with the moveable lens, so as to keep the vertex
of bundle of chief rays of the scanning beam at the back focal
plane of the moving lens.
[0125] Rather than move the moveable lens and beam scanner, we move
the ocular lens and the patient together so that the retinal
conjugate is formed at a standard location with respect to the
remainder of the optics of the instrument.
[0126] An alternative solution would be to use a variable-power 1:1
relay system to re-form the retinal conjugate at a standard
location. Another alternative is to use moveable mirrors to fold
the optical path (in the shape of a trombone, for example) and
extend the optical distance using the moveable mirrors so as to
bring the retinal conjugate to a standard location.
[0127] The configurations described here have the advantage that
the angular magnification, from the human pupil to the scanning
mirrors, remains nearly constant in the face of compensation for
refractive error. This feature means allows the scan range of the
OCT beam, in terms of angle in the visual field, to be determined
based on the turning angles scanning mirrors, without need for
correction based on the motion of lenses for refractive error
compensation.
[0128] FIGS. 14a and 14b provide a table listing the alignment
steps for imaging, 14b providing optional steps. Experience has
shown that all these steps can be conducted pretty much in any
order, with the exception that the working distance has to be set
first. Otherwise, the adjustments can be performed in any
order.
[0129] All motor positions (chinrest x, y, z, ocular, polarization,
z-motor) can be recorded for every patient and restored upon repeat
visits.
[0130] The preferred optical coherence tomography device will
contain a feature called "pupil following". This is not pupil
tracking, but rather a mechanism that moves the head (and therefore
the pupil) when the fixation target is moved.
[0131] When we move the fixation target in the optical coherence
tomography device, the patient rotates their eye to follow the
fixation target. While they do so, their pupil shifts because the
center of rotation is behind the pupil. Therefore the chinrest has
to be moved sideways in order to compensate for this.
[0132] The current implementation does the following: [0133] 1. It
moves the fixation target continuously (rather than in one large
step where the fixation target suddenly is located somewhere else
and the patient has to search) [0134] 2. It compensates pupil shift
based on a simplified eye model and rotation. For a "nominal"
patient's eye the operator would never see the pupil move at all.
For a real patient there is some adjustment necessary, but it
helps.
[0135] Additionally, a corneal scan can be performed. One can
insert a flip-in diverging lens 180 (FIG. 1) after the
galvanometers, so as to form a virtual point source near the pupil
conjugate. This results in a beam waist near the pupil of the
subject. One can set the power of the lens so that the beam waist
is on the cornea of a typical eye, and move the z-motor by a
typical eye length simultaneously with addition of the lens, so as
to quickly switch between retinal and corneal OCT imaging.
[0136] When using this method, the iris view and LSO image are not
disturbed, and the patient continues to see the fixation
target.
Method for Combining B-Scans ("Thick B-Scan")
[0137] SD-OCT greatly enhances data acquisition speed by
simultaneously acquiring position and scattering intensities for
all scatterers along an A-Scan. The time savings may be used to
simply allow faster completion of the same exams or the time may be
used to acquire more data, such as acquisition of higher density
volumes. Acquisition of more data provides the opportunity to
combine data in order to improve some feature or parameter, such as
combining data by spatially compounding in order to reduce speckle.
OCT-tomograms generally suffer from degraded image clarity due to
image speckle and noise. Structures whose dimensions approach the
resolution limit of the imaging system display speckle
discontinuities. For example, the external limiting membrane in
retina cross sections shows speckle discontinuities when imaged
using medium resolution OCT. Speckle reduction is generally
achieved by compounding the image cell using various data acquired
by means that vary the speckle property, generally either frequency
compounding (by viewing the speckle generating cell by means of a
different optical frequency) or spatial compounding (by viewing the
speckle generating cell from a different spatial location, usually
a different angle.) (See U.S. Pat. No. 6,847,449 and Schmitt, J.
M., S. H. Xiang, et al. "Speckle in Optical Coherence Tomography."
Journal of Biomedical Optics 4(1):95-105 (1999).) This embodiment
describes a method of speckle reductions which combines elements of
adjacent A-scans to produce a speckle reduced B-scan with reduced
noise and enhanced visual impression. A B-scan resolution cell is a
cell within the B-scan that is resolvable in the displayed
image.
[0138] While this embodiment generally derives one or more A-scans
from a collection of A-scans, its implementation and speckle
reduction advantage is more easily described as a method of
determining a B-scan from one or more B-scans. In its simplest
instantiation, a single B-scan is created by over-sampling the
image region. (While over-sampling the original B-scan is not
necessary, it is easy to visualize compounding without resolution
reduction using over-sampled data.) A new B-scan is derived from
the image data of the acquired B-scan by laterally filtering the
acquired B-scan at each depth. The new B-scan may be decimated
either after or during the filtering step so that it is no longer
over-sampled. The resulting B-scan is speckle reduced along the
laterally filtered direction, but retains specular features
acquired in the transverse direction (orthogonal to the B-scan).
For transverse smoothing in the direction orthogonal to the B-scan
(and the creation of a "thick B-Scan"), one or more adjacent
B-scans can be used. Data at a fixed depth in an A-scan can be
combined with data from the same depth in other A-scans.
Algorithmically, these combinations are simpler when the B-scans
are acquired in a fixed grid of parallel planes (B-scans), as in
FIG. 15a. However, when combining data for speckle reduction, the
resulting image contains fewer, or at least different, artifacts
when the A-scans are not rigidly oriented on a regular grid. FIGS.
15a-d are used to illustrate compounding B-scans. In FIGS. 15a-d,
the dots represent A-scans and lines represent B-scans. FIG. 15a
shows the traditional scan pattern. In FIG. 15b, adjacent B-scans
are acquired with the A-scan acquisition shifted by 50% of the
spacing between A-scans in one of the B-scans. The shift of 50% of
the spacing between A-scans is particularly advantageous when
precisely two B-scans are used to acquire a new, interpolated
B-scan. FIG. 15c represents combining three (3) rows to form a
single B-scan while FIG. 15d shows three (3) columns being combined
to form a single B-scan.
[0139] In general, A-scans from M rows and N columns can be
combined to form a single computed A-scan. One method of combining
is bi-linear interpolation. An alternative combination is obtained
if a median filter is used. Alternatively, the value of any depth
point in the computed A-scan can be viewed as a weighted sum of the
neighboring A-scans. The weighted sum can include depth points. In
general, the weights should be set so that the majority of the
support for each computed pixel lies within one or two speckle
diameters along each axis. The smaller the scope of this support,
the greater the resolution (though this technique cannot improve
the resolution beyond that of the imaging system), while the larger
the scope of this support, the greater the speckle reduction. Two
A-scans are speckle diverse if they are separated by approximately
more than 1/2 the diameter of a speckle cell. Preferably speckle
diverse A-scans are separated by a speckle diameter, however,
smaller separations can achieve some speckle reduction. Similarly,
a collection of A-scans are speckle diverse at a point in a
direction if the collection contains A-scans which are separated by
approximately more than 1/2 the diameter of a speckle cell in that
direction. Again, preferably they are separated by a speckle
diameter.
[0140] The combination can be performed either during acquisition
or post acquisition. FIGS. 16a and 16c represent patterns expressly
designed for combination during acquisition. FIG. 16b is included
to depict the difference in the data collected using the prior art
acquisition sequence 16b and the embodiment of the invention
depicted in FIG. 16c. Here the A-scan locations wiggle (are
dithered) about a centerline. The span of the wiggle is greater
than a speckle diameter, as shown in FIG. 16c. As depicted, any 4
A-scans are speckle diverse both tangent to and orthogonal to any
plane orthogonal to the y-axis. Here the compounding can be
performed during acquisition with only a required memory of a few
A-scans. Combination can limit computation by using a simple FIR
filter, such as a boxcar or lowpass filter, or non-linear filter,
such as a median filter. Combinations may also use be complex,
using higher order statistics derived from neighboring A-scans. The
primary advantage of compounding during acquisition is that the
data is acquired at sufficiently high speed that motion during
acquisition can be ignored. For the target SD-OCT system described
herein, typical modulation parameters are nominally on the order of
10 .mu.m for the period of the wiggle and 10 .mu.m amplitude. A
range of 5-20 .mu.m is typical. Clearly, any differences in the
system design affecting the resultant speckle size would also
affect the nominal modulation parameters.
[0141] The deliberate decrease of discrimination in the orthogonal
direction to a B-scan may constitute a new OCT display modality of
tomographic data ("thick B-scan"). The thick B-scan modality is
useful for viewing layer-like structures with a thickness close to
the speckle limit. Further new displays combine the thick B-scan
and one or more standard B-scans. For example, a thick B-scan
derived from data from three (3) B-scan planes, say B1, B2, and B3,
can be displayed essentially also showing B.sub.1, B.sub.2, and
B.sub.3. We display the intensity of the thick B-scan with the hue
determined by B.sub.1, B.sub.2, and B.sub.3; where the hue is blue
if the intensity of B .sub.1 is closest to that of the thick
B-scan, the hue is yellow if the intensity of B.sub.2 is closest to
that of the thick B-scan, and the hue is red if the intensity of
B.sub.3 is closest to that of the thick B-scan. Any such display
provides a suitable presentation of the deviation of the
contributing B-scans from the combined one and can even offer some
spatial interpretation of the two dimensional data even in a
printed version.
[0142] This embodiment enables the selection of a B-scan with:
reduced speckle and noise; no significant loss in lateral
resolution in the direction of the B-scan; no increase in the
number of detectors or, in some cases, scan-time; optionally
increased density of A-scans in the presented B-scan; and
optionally adjustable lateral resolution orthogonal to the B-scan
direction. While B-scans are nominally thought of as planar
sections, they can be any curved surface. The most typical B-scans
are planar cross-sections and circle scans. Circle scans are scans
covering a cylindrical surface whose perpendicular cross sections
are nominally circular. Circle scans are particularly useful for
determining retinal nerve fiber layer health, where the macular
nerve head lies within the circle scan (nominally at the center of
the cylindrical surface of the B-scan.)
Adaptive Compensation of Galvo Response
[0143] The motors of optical scanning galvanometers ("galvos")
mechanically wear out after being scanned back and forth many times
(up to billions of cycles). Before they fail completely and
catastrophically, their bearings and lubricants may deteriorate
gradually over a long period of time. These motors are typically
driven by servo amplifiers that attempt to minimize the difference
between the actual galvo motor position and a commanded position,
provided as either an analog or a digital signal. In ophthalmic
scanning, image quality and repeatability is directly dependent on
galvo performance.
[0144] Two embodiments which ensure consistent scanning performance
(i.e. consistent position response for a given command sequence)
over the lifetime of an application are: [0145] 1) Characterize the
closed-loop response and adapt the loop filter of the servo in
order to keep the closed-loop response constant, and [0146] 2)
Characterize the closed-loop response of the galvo system and adapt
the command signal given in order to achieve the desired position
response.
[0147] In both of these embodiments, it is essential to
characterize the complete closed-loop response of the system. This
can be done, for example, by having some software on an instrument
that gives a white noise command to each galvo and digitizes the
position response. The Fourier transform of the position response
would provide the closed-loop frequency response.
[0148] In approach 1), a mechanism is needed to adjust the tuning
of the servo filter. If the servo is digital, this can be
accomplished using software running on an instrument. This can also
be accomplished if the servo is analog but has digitally settable
potentiometers in the servo circuit. Appropriate adaptive filter
algorithms can be achieved using known techniques.
[0149] In approach 2), the closed-loop response is allowed to
change over time. The desired command signal at any given time can
be determined by applying the inverse of the closed-loop response
to the desired position response.
[0150] In practice, it is best to utilize high-acceleration,
asymmetric galvo waveforms for the purpose of "clearing" the galvo
motors. These waveforms are designed to cause the balls in the
galvo bearings to skid to a new position and avoid pit formation
where the balls rock back and forth in the galvo raceways.
Reliability testing has shown that these clearing moves can prolong
the life and performance of the galvos.
Calibration Test Eye
[0151] In order to align, calibrate and test an ophthalmic
instrument, it is desirable to have an artificial test eye. Various
artificial eyes have been used throughout the ophthalmic industry,
some very simple with poor imaging quality, others are more
complicated, imitating the structure of the human eye (cornea and
lens) and achieving high optical quality at great cost.
[0152] FIG. 17 illustrates a test eye using one single refractive
surface and a stop. This arrangement achieves very high optical
quality over a 47-degree full field of view.
[0153] One embodiment is obtained using a single piece of optical
glass, with a stop placed in the center of curvature of the first
surface to avoid coma, astigmatism and lateral color. The imaging
surface is curved to match field curvature and any pattern on the
imaging surface is graduated in arc-mm to compensate for the
distortion.
[0154] An alternate embodiment is obtained using aspheric surfaces
to further reduce spherical aberration.
Co-Focus of Fundus Imager and Fixation Target
[0155] Ophthalmic instruments imaging the retina (fundus camera,
LSLO, CSLO, OCT) use an internal fixation target to align the eye.
It is desirable to co-focus the imaging optical path with the
internal test target optical path so that the fundus image seen by
the optician is in focus at the same time as the fixation target
seen by the patient. The following paragraphs describe a method and
apparatus which achieves this result.
[0156] Standard practice is to focus the imaging path and the
fixation path in the same plane. FIG. 18 illustrates the preferred
focus described herein. In this embodiment, the preferred focus of
the fundus imager is anterior to the photoreceptor layer of the
eye. The preferred focus for the fundus imager is at the blood
vessel layer, while the preferred focus for the fixation target is
at the photoreceptor layer. Therefore, we focus the imaging path
and the fixation path in different planes separated by a distance.
This enables the patient to see a sharper fixation target during
the eye examination when the fundus is being imaged by the doctor.
The sharper fundus image improves the patient's attention on the
fixation target, thereby decreasing eye motion and creating a
higher quality fundus image by reducing motion artifacts.
[0157] The standard design of camera lenses achieves near zero
longitudinal aberration across all wavelengths within its design
parameters. One embodiment achieves the desired result by
implementing a camera lens with a known positive longitudinal
chromatic aberration. That is, the lens longitudinal aberration is
the sum of human eye chromatic aberration and the desired focus
shift at the specified wavelength.
[0158] FIGS. 19a-c compare an ophthalmic instrument aligned with
the prior art approach to an ophtalmic instrument aligned using the
disclosed method. FIG. 19a shows the actual alignment of the Prior
Art system with a fundus imaging system and a fixation target which
is not chromatically aligned. The vertical arrow of FIGS. 19a-c
represents the refractive power of the cornea and crystalline lens,
combined as a single lens. While the system is ostensibly designed
to focus both the fundus imager and the fixation target systems in
the same plane, because the chromatic aberration of the eye optical
components is not accounted for, the actual focus has the infrared
fundus imager focus posterior to the visible fixation target. That
is, the focus of the fundus imaging system is too deep. FIG. 19b
shows the same system with the fundus imaging system properly
chromatically aligned, but without the preferred fundus imaging
focus alignment. FIG. 19c shows the same system with the preferred
fundus imaging focus alignment. The desired alignment can be
accomplished during manufacturing either by using different lenses
for aligning different optical systems or a single lens with the
desired optical properties of each system. The first alternative
has the advantage of utilizing readily available lenses and the
disadvantage of changing lenses during final system alignment. The
latter alternative has the advantage of not requiring manufacturing
to change lenses in the test fixture during alignment with the
disadvantage of requiring a special purpose lens which may require
re-design in case of system component changes. FIG. 19c provides
the design parameters for one such special purpose lens
[0159] Other imaging systems combined in one instrument are
commonly aligned in a common plane in the prior art and can also
benefit from further embodiments of this invention. For instance,
in color fundus cameras with imaging arrays, multiple sensors are
used for different wavelengths of illumination. In those cases, it
is possible to adjust the axial positions of the sensors relative
to each other, so that each sensor is optically conjugate to the
source of scattered light. If light from all wavelengths is
scattered from the same depths, then the optical system is
compensating for chromatic aberration to have all wavelengths at
the best focus simultaneously. Alternatively, one sensor configured
to receive visible light may be conjugate to layers anterior to the
retinal pigment epithelium, and a sensor configured to receive
near-infrared light may be conjugate to the choroidal blood vessels
posterior to the retinal pigment epithelium. Therefore, different
layers of the retina may be imaged on different sensors
simultaneously.
[0160] In the case of retinal OCT systems, a fundus camera and
confocal scanning optics provide simultaneous imaging of the
retina. For example, the Stratus OCT (Carl Zeiss Meditec, Inc.,
Dublin, Calif.) employs a fundus viewing system that is
color-corrected so that both the OCT beam and fundus viewer, for a
range of visible and near-infrared wavelengths, are both conjugate
to the same depth in the retina.
[0161] Scanning imagers, such OCT scanners, may not provide
sufficient speed to produce real-time images to allow technicians
to align the OCT scan area with the desired region to be imaged,
for instance the foveal region of the retina. To assist this
placement, a continuously displayed image of the retina is desired.
Illumination with near-infrared light, for instance in the range of
700-900 nm, provides an image of the fundus without causing patient
discomfort and/or constriction of the pupil. The scattering
efficiency and absorption in the retinal layers above the RPE is
relatively low in the near infra-red. Therefore, it is difficult to
produce images of retina in these layers with near infra-red light.
However, these layers are clinically very important, for instance
to characterize retinal pathologies such as macular holes, and
scanning systems, such as OCT, are often focused on these
layers.
[0162] One option is to provide an infra-red fundus viewer, such as
with an array sensor in a configuration similar to that in Stratus
OCT. The infra-red image in this case has the best contrast below
the RPE, where the choroidal vessels are imaged. Therefore, the
fundus viewer is adjusted so that it is conjugate to the choroidal
vessels when the OCT scanner is conjugate either to the RPE or to
layers anterior to the RPE such as the inner plexiform layer. This
can be achieved, for instance, by first adjusting the OCT and
fundus viewers to be conjugate to the same layer in the retina or
to a test fixture, then by shifting the axial location of the
fundus viewer sensor so that it is conjugate a specific distance
posterior to the RPE corresponding to tissue posterior to the RPE.
This distance would typically be in the range of 0.2 to 1.0 mm. For
example, in reference to U.S. Pat. No. 7,140,730 B, FIG. 3, the CCD
could be shifted along the optical axis as described above.
Alternatives to the alignment method, such as using a flip-in
spacer to make the adjustments, and alternatives to the design
adjustment, such as shifting lenses instead of the sensor, are
obvious to those skilled in the art. Providing best focus at
different depths simultaneously offers two advantages. First, the
best focus of the fundus image, occurring posterior to the RPE,
corresponds to the best focus of the OCT scanner at the desired
depth so that the user can use cues from the fundus image, such as
the sharpness of the image, to adjust the focus of the OCT. Second,
the fundus image provides the best possible features for use as
landmarks in placing OCT scans.
[0163] Another option is to provide two scanning imaging systems.
The first system, such as an OCT scanner, is slower than the second
system, such as a scanning laser ophthalmoscope (SLO), a line
scanning laser ophthalmoscope (LSLO), or a line scanning
ophthalmoscope (LSO). The second system provides video-rate images
of the area to be scanned, such as the foveal region of the retina.
When the second system is a confocal imager, even near-infrared
light can be used to provide good contrast images of the blood
vessels anterior to the RPE. To improve the contrast and sharpness
of those images further, the second scanner is adjusted so that it
is conjugate to a layer somewhat anterior to conjugate of the first
scanner. For example, the second system may be conjugate to blood
vessels anterior to the RPE when the first system is conjugate to
the RPE layer. This offset is typically in the range of 0.2-0.5 mm
depending on the expected state of pathology in the eye.
[0164] Another option is to provide an infra-red fundus viewer,
such as with an array sensor, with an OCT scanner in a
configuration similar to that in Stratus OCT, where both the fundus
viewer and OCT scanner are optically conjugate to the same layer.
The new instrument achieves the desired separation of focal planes
by first aligning the infra-red fundus viewer to its best focus on
the choroid (tissue posterior to the RPE). When you are ready to
capture OCT data, we then automatically shift the focus of both the
infra-red fundus viewer and the scanner to the desired scan depth.
This shift can be achieved, for instance, by motorizing at least
one of the two imagers and then using the motor to shift the sensor
or a lens axially before acquiring the scanned image.
Alternatively, this focal shift can be accomplished by flipping in
a lens or swapping out one lens for another to shift the focal
plane the desired distance.
[0165] FIG. 20 shows, in pictorial form, a conventional arrangement
of two imaging subsystems. A scanning imaging subsystem includes a
pencil beam light source 1001 and scanning mechanism 1002, scanning
lens 1111, chromatic beam splitter 1050, focusing lens 1150, lens
of the eye 1160 and a scanning area which sweeps out a surface
bounded above by 1205, bounded below by 1215 and including focal
point 1210. As the scanning mechanism sweeps the pencil beam across
the scanning area, the focal point 1210 moves about a region on the
retina. A second imaging subsystem is a fundus camera with area
illumination light source 1021 and beam splitter 1040 (in this case
depicted by a pin-hole mirror). The second imaging system includes,
common to the scanning imaging subsystem, optics path elements:
chromatic beam splitter 1050 focusing lens 1150, lens of the eye
1160 and a surface of points near the retina, including point 1221,
which will be brought to focus on detector 1030. When one says that
the two imaging systems are focused at the same plane, what is
meant is that the scanning area swept by the scanning imager is
essentially the same (or contains or is contained in) surface which
the second imaging systems brings into focus at detector 1030. Lens
system 1133 functions to focus the fundus camera image on the CCD
camera detector 1030.
[0166] In one embodiment of the current invention, the CCD camera
detector 1030 is moved to be conjugate to a point posterior to
1210.
[0167] FIG. 21 illustrates an arrangement of two imaging
subsystems. The residual portion of the first imaging system which
is not depicted resides in 1010. This can be a beam imaging system,
either scanning or fixed, or an area illuminating light source,
like a fundus illuminating source. The residual portion of the
second imaging system which is not depicted resides in 1020. This
can also be a beam imaging system, either scanning or fixed, or a
line or area illuminating light source, like an SLO or fundus
illuminating source, respectively. The first imaging subsystem is
conjugate to the focus at 1210 while the second imaging subsystem
is conjugate to the focus at 1220. The lens of the eye 1160, the
lens 1150, and one side of the chromatic beam splitter 1050 are
part of a common optical path for both subsystems. Lens 1110
completes the focus so that the first imaging subsystem is
conjugate to 1210. Lenses 1120 and 1130 complete the focus of the
second imaging subsystem so that detector 1030 is conjugate to
1220. Beam splitter 1040 (here depicted as a pin-hole mirror) acts
to redirect the source for the second imaging subsystem to align
with the common optical path, while allowing reflected light back
to detector 1030.
[0168] FIG. 22 illustrates a simplification of FIG. 21. The effect
of lenses 1150 and 1160 are combined into lens 1150'. The lens 1120
is replaced by lens 1120' so that the light is not collimated at
the pinhole mirror. Lens 1130 is replaced with lens 1130' so that
1220 is conjugate to detector 1030. Lens 1110 is replaced with lens
1110' so that 1210 is conjugate to the detector of the first
imaging subsystem.
[0169] FIG. 23 illustrates yet another configuration of the
invention. Here the second imaging subsystem focuses anterior to
the focus of the first imaging subsystem in order to illustrate
that either subsystem can image anterior to the other. The lens
1150' is replaced by lens 1150'', so that 1220 is essentially
conjugate to the pinhole mirror. Lenses 1120'' and 1130'' are
consistent with lens 1150' and are designed so that 1220 is
conjugate to detector 1030. Lens 1110'' is consistent with lens
1150' and is designed so that 1210 is conjugate to the detector of
the first imaging subsystem.
[0170] In FIGS. 21, 22, and 23, only the configuration is shown,
where the focal points 1210 and 1220 are not aligned. As previously
described, this can be accomplished in various ways, including the
changing out of lenses or movement of the detector.
[0171] It should be understood that the embodiments, examples and
descriptions have been chosen and described in order to illustrate
the principals of the invention and its practical applications and
not as a definition of the invention. Modifications and variations
of the invention will be apparent to those skilled in the art. The
scope of the invention is defined by the claims, which includes
known equivalents and unforeseeable equivalents at the time of
filing of this application.
* * * * *