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TECHNICAL FIELD
The present invention pertains to the field of laser photocoagulation and,
in particular, to apparatus for predetermining laser coagulation dosages
and for monitoring laser photocoagulation by detection of secondary
radiation.
BACKGROUND ART
Laser photocoagulators have become an important tool in treating eye
disease. As stated in an article entitled "Lasers In Ophthalmology: The
Path From Theory To Application" by M. L. Wolbarsht and M. B. Landers III
in Applied Optics, Vol. 18, No. 10, May 15, 1979, pp. 1516-1526,
specifically at p. 1516:
"The argon laser photocoagulator is routinely used by ophthalmologists all
over the world and is now standard practice for the treatment of many
retinal diseases. Indeed for some problems such as diabetic retinopathy,
not to use it borders on malpractice."
Although retinal photocoagulation has been used since before the advent of
lasers and is now a standard treatment of choice for some common retinal
disorders, the occurrence of complications is high, compared to other
kinds of laser therapy. Because of variations in vascularization and
pigmentation from place to place on the fundus, obtaining an optimum
application of laser energy requires a delicate adjustment of laser
energy, laser pulse duration, and lower beam spot size. An exposure which
is therapeutic in one location may be ineffective in another, and may
produce a hemorrhage in yet another.
Current clinical practice relies on visual assessment of pigmentation and
vascularization and upon visual estimates of coloration changes that occur
after a treatment pulse. However, such visual assessments are recognized
by the art to be inadequate for reliable determination of the optimal
laser application. Furthermore, there is recognition in the art of the
need for a means of monitoring laser photocoagulation during treatment. As
stated in an article entitled "Fundus Reflectometry: A Step Towards
Optimization Of The Retina Photocoagulation" by R. Birngruber, V. -P.
Gabel and F. Hillenkamp in Mod. Probl. Ophthal., vol. 18, 1977, pp.
383-390 at p. 383:
"Improvements in clinical retinal photocoagulation can be achieved by both
the optimal adaptation of the instrument to the problem, and variation of
the physical irradiation parameters . . . . It is obvious from the theory
of heat conduction, that in the range of exposure times from 10.sup.-3 to
10 sec of interest here, the energy necessary for a given reaction in the
irradiated area decreases markedly with shorter times. . . . It is
understood though, that the possibility of manual control of the effect
through visual observation of the coagulation site ceases for exposure
times below about 1 sec.
Monitoring the time development of the retinal blanching during and after
coagulation with suitable photodetectors should result in a more direct
measure of the influence of the important parameters such as energy and
exposure time on the retinal reaction. Such a technique could moreover
eventually lead to a method for an automatic control of exposure times,
even for very short times."
In searching for a means of monitoring the progress of laser
photocoagulation during treatment, it has been recognized in the art that
there is a connection between reflectivity of the irradiated tissues and
the effects of the photocoagulation, see for example an article entitled
"Time And Location Analysis Of Lesion Formation In Photocoagulation" by
Oleg Pomerantzeff, Guang-Ji Wang, Michail Pankratov, and Julianne
Schneider in Arch. Ophthalmol., Vol. 101, June, 1983, pp. 954-957. This
has suggested the use of reflectometry, i.e. measurement of light
backscattered from an illuminated spot on the retina, to monitor
photocoagulation. The reflected light could come from the photocoagulation
laser itself or from a secondary pilot laser.
In addition to attempts to monitor laser photocoagulation during treatment,
there have been attempts in the art to pre-determine the appropriate laser
dosages to apply for treatment of specific diseases. These attempts have
used reflectometry to determine the laser parameters. Such a use of
reflectometry is illustrated in an article entitled "A Method To
Predetermine The Correct Photocoagulation Dosage" by Oleg Pomerantzeff,
Guang-Ji Wang, Michail Pankratov, and Julianne Schneider in Arch.
Ophthalmol., Vol. 101, June 1983, pp. 949-953, at p. 949:
"The most common goal of photocoagulation in the macular area is closing
leakage from very small vessels and destroying new-formed vessels in the
sub-retinal space. . . . Yellow and green light are recommended for
treatment of the macula since the yellow pigment in the inner layers of
this area absorbs very little of these colors. . . . The reaction of
retinal tissue to the irradiation with a given power density varies
according to the local concentrations of blood and melanin. Therefore, to
avoid overtreatment and the risk of hemorrhage, it is desirable to know
this relative concentration in the target tissues before treatment is
applied, especially when red light is used. In this study, we suggest a
possible method to measure this relative concentration.
Absorbance cannot be measured directly in a living eye but it can be
measured indirectly by measuring reflectance. To do this we assume that
the light that is neither absorbed by nor scattered back from the retinal
or chloroidal layers reaches the sclera, which transmits only a negligible
percentage, and is reflected from it. This reflected light is partly
absorbed on its way back, and finally emerges from the retina into the
vitreous. Therefore, if we measure the power applied to the retina and the
power emerging back from the retina, the difference between the two is a
measure of the absorbance in the retina and choroid. In photocoagulation
it is also important to determine, if possible, the level within which
most of the melanin is concentrated."
at p. 950:
"Since the reflection by the retinal structures is most diffuse, we are
obviously not collecting all the light reflected from the retina. However,
we may assume that the ratio of the collected to the reflected light
remains the same, at least in the same eye."
and at p. 952:
"Not all the light emerging from the cornea is diffusely reflected. There
are also some discrete specular reflections that may eventually fail into
the entrance pupil of the measuring system, making the measurements
unreliable. . . . The reflectance, and consequently the absorbance,
depends not only on the retinal area and the selected wavelength, but also
on the angle at which the particular structure is irradiated. Therefore
the absorbance should be measured using the coagulating beam in its
coagulating position."
In sum, fundus reflectometry, i.e. measurement of light that is
backscattered from an illuminated spot on the retina, is used in the art
to pre-determine laser photocoagulation dosages as well as to monitor
laser photocoagulation during treatment. In theory, if the intensity of
the incident radiation is known, the absorbance of the tissue can be
calculated from the reflectance, and if the absorbance is known, the
amount of energy absorbed from a treatment can be predicted. However,
application of this theory involves a number of complications, such as:
(1) wavelength dependence of the scattering, (2) angular dependence of the
scattering, (3) the relation between the scattered light which leaves the
pupil of the eye and is therefore accessible for measurement and the
scattered light which is re-absorbed inside the eye and therefore cannot
be directly measured. Furthermore, in prior art apparatus constructed to
apply fundus reflectometry, reflection of the incident light from filters,
lenses and other optical transmission components is quite strong, being at
least a few percent of the incident radiation. This means that simple
reflectance from a target cannot be easily measured when the same optical
system is used to deliver the light to the target and to capture the
reflected light. However, when two separate optical systems are used for
simple reflectance measurement, it is difficult to insure that they are
both aimed at precisely the same target spot.
DISCLOSURE OF INVENTION
The above-described problems occurring in the use of reflectometry for
pre-determining appropriate laser dosages for laser photocoagulation and
for monitoring laser photocoagulation during treatment are advantageously
solved by embodiments of the present invention.
Embodiments of the present invention comprise means for applying radiation
to a target in an eye and means for detecting radiation produced by the
target having wavelengths different from the wavelength of the applied
radiation. Such radiation produced by the target is denoted hereinafter as
non-Rayleigh radiation (NRR). In various embodiments, the NRR is
fluorescence radiation, Raman radiation, two photon excitation radiation
and so forth. Furthermore, in a first embodiment, the means for applying
radiation to the target spot is the laser coagulation source itself; in
other embodiments the application means comprises one or more laser
sources other than the laser coagulation source; and in still further
embodiments the application means comprises means, cooperating with the
laser coagulation source, to provide laser radiation.
Embodiments of the present invention can be used to pre-determine the
effect of the laser coagulation by measuring absorbance prior to treatment
and also to monitor the laser coagulation during treatment.
BRIEF DESCRIPTION OF THE DRAWING
A complete understanding of the present invention may be gained by
considering the following detailed description in conjunction with the
accompanying drawing in which:
FIG. 1 shows, in pictorial form, an apparatus constructed in accordance
with the present invention.
PREFERRED EMBODIMENT FOR CARRYING OUT THE INVENTION
When optical radiation strikes a target, such as the fundus of the eye,
some of the radiation is transmitted, some is scattered and some is
absorbed. The fundus is a complex optical system consisting of several
different kinds of pigments arranged in a layered structure. It comprises
a highly non-uniform distribution of particles and surfaces which absorb,
scatter, and reflect light. Nevertheless, it is possible to measure the
amount of light absorbed by the fundus by measuring the amount of light
that is backscattered or reflected. Since light is scattered at all
angles, only a fraction passes back out through the pupil where it can be
measured. But, by making some reasonable assumptions about the angular
distribution of scattering, estimates of fundus absorption can be made
from measurements of light that is scattered back out through the pupil of
the eye.
The physical mechanisms responsible for scattering of light from molecules
provide that when light of a first wavelength strikes a molecule, most of
the scattered radiation is at exactly the same wavelength. Such scattered
radiation (e.g. reflected radiation) is called Rayleigh radiation.
However, some scattered radiation occurs at wavelengths different from the
first wavelength, denoted hereinafter as NRR. In fluorescence, for
instance, radiation is absorbed and re-radiated at many different
wavelengths. In Raman scattering, radiation is produced which has a
relatively small change in wavelength from the incident radiation, such
Raman radiation being characteristic of the scattering molecule. Raman
radiation may have wavelengths longer than the first wavelength, i.e.
"Stokes" Raman radiation, or it may have wavelengths shorter than the
first wavelength, i.e. "anti-Stokes" Raman radiation. Both Stokes and
anti-Stokes Raman radiation are typically much smaller in intensity than
Rayleigh radiation. Furthermore, anti-Stokes Raman radiation is typically
much less intense than Stokes Raman radiation. There are, however, some
circumstances in which both types of Raman radiation may become very much
stronger than usual. In particular, this occurs when the wavelength of the
indicent radiation coincides with an absorption band of the molecule. In
this instance, resonance Raman scattering occurs with an intensity many
orders of magnitude greater than normal Raman scattering.
I have discovered that apparatus for pre-determining laser dosage and for
monitoring laser coagulation can advantageously be provided by monitoring
NRR, and Raman radiation in particular. An important reason for this
advantage is the fact that this radiation is produced mainly in the target
(Glasses used in lenses and optical fibers may produce fluorescence and
Raman radiation. Moreover, certain tissues in the eye, such as the lens,
may produce NRR. However, the levels of NRR produced by such sources would
ordinarily be very small in comparison with the NRR from the target and
may be regarded as a negligible "background" signal). Consequently, by
detecting NRR instead of Rayleigh radiation, background scattering from
optical transmission components and areas of the eye other than the target
is significantly reduced.
Embodiments of one aspect of the present invention comprise a laser
coagulation source, a beam delivery system for transmitting laser
radiation from the source to the laser coagulation target spot on the
retina, and an NRR monitor which receives the NRR from the laser
coagulation target spot on the retina after it is transmitted back through
the beam delivery system. A major advantage of such an embodiment,
obtained by using the same optical system for delivering the light to the
target spot and for monitoring the backscatter therefrom, is that the same
target spot is "seen" by the delivery and the monitoring systems. Such
embodiments are useful for monitoring during treatment because they
provide NRR detection at the same time that laser coagulation radiation is
being applied for treatment. The NRR is detected by an apparatus, such as
a monochromator, which is tuned to select suitable NRR and to reject the
strong Rayleigh radiation. This provides a scattering response with a
sufficiently high signal-to-noise ratio as to be useful in estimating
target absorbance.
FIG. 1 shows an embodiment of the present invention where laser source 1 is
a 5 watt, CW argon ion, Spectra Physics Argon laser, model no. 164/168.
Laser radiation produced by laser source 1 impinges upon filter 2. Filter
2 selects the 514.5 nm line and passes the filtered radiation as beam 100,
which filtered radiation is used for photocoagulation. The filtered
radiation in beam 100 passes through shutter 3 in order to form a pulse
for coagulation. The pulse of radiation then passes through beam splitter
4 and is focused by lens 5, for example, an 18 mm focal length lens, into
optical fiber 6, for example, a Fiberguide Industries, Inc. fiberoptic
cable 80/125 (denoting core/cladding diameter in microns).
The filtered radiation, output from cable 6, impinges upon zoom lens 7 (for
example, Chugai Boyeki Company, Ltd. Model No. M10Z118 zoom lens), which
zoom lens is held by slit lamp 8, for example, a Zeiss slit lamp. Zoom
lens 7 is used in this embodiment to allow a physician to vary the spot
size of the laser radiation on the target in the eye and the slit lamp is
provided as a light source for the physician to examine the eye.
The radiation emerging from zoom lens 7 impinges on mirror 9 where it is
diverted through shutter 10, for example, a black anodized aluminum
shutter, and from there into eye 11 of a patient. In an instrument
designed solely for clinical photocoagulation, shutter 10 would not
normally be present. Photocoagulation pulses would be controlled only with
shutter 3, as mentioned previously hereinabove.
Carotenoid retinal pigments in the eye generate Raman radiation having a
frequency shift of approximately 1500 cm.sup.-1 from the 514.5 nm laser
excitation. This corresponds to a Stokes radiation wavelength of 558 nm
when the laser excitation is at 514.5 nm. Thus, NRR, including Raman
radiation in the 1500 cm.sup.-1 Stokes band at 558 nm, is generated in eye
11 and directed back through the above-described system to impinge upon
beam splitter 4. Beam splitter 4 directs the backscattered radiation,
containing mostly radiation at the wavelength of laser source 1 along with
relatively weak NRR, through band rejection filter 200 and into
monochromator 12. In this embodiment, monochromator 12 should be tunable
so as to cover wavelengths in the range of roughly 2000 cm.sup.-1 on
either side of the 514.5 nm argon laser line. Furthermore, it must have a
very high rejection ratio for the 514.5 nm Rayleigh radiation background.
The radiation impinging upon monochromator 12 is stripped of 514.5 nm
Rayleigh reflected light by 514.5 nm band rejection filter 200. As shown
in FIG. 1, band rejection filter 200 comprises Omega, Inc. optical filter
21, one millimeter slit 22, Oriel, Inc. Model No. 5748 variable
interference filter 23, and Edmund Scientific Corporation "Edscorp
.times.15" microscope eyepiece focusing lens 24. Lens 24 focuses the
radiation emerging from band rejection filter 200 onto 0.006 inch wide
entrance slit 25 in single grating monchromator 12. RCA 931B
photomultiplier tube photodetector 13 receives the output from
monchromator 12 and, in response thereto, provides an electrical output
signal to electrical readout 14. In various embodiments electrical readout
14 may comprise strip-chart recorders or microprocessors or other types of
computers and the like for evaluating the output of photodetector 13.
During therapeutic photocoagulation, retinal tissue undergoes changes in
optical characteristics that include changes in the level of Raman
backscatter. This change is monitored by electrical readout 14. The
resulting electrical signal provides an objective measurement of the
degree of tissue photocoagulation at each instant, and may be used to
modulate the laser intensity to achieve an optimum treatment. The
evaluation of the changes may involve comparing previous readings, then
making a decision based thereon and thereafter resetting the output of
laser source 1 during treatment or establishing appropriate settings for
the output of laser source 1 for subsequent treatment. Furthermore, the
laser pulses may be controlled by varying the response of shutter 3 to
increase or decrease the laser pulse duration or intensity, and so forth.
The control of laser source 1 or shutter 3 discussed hereinabove may be
performed by applying signals 15 from electrical readout 14 directly to
laser source 1 and/or shutter 3. It should be clear to those skilled in
the art that various other means for controlling the intensity, duration
and placement of the laser coagulation radiation in response to output
from electrical readout 14 may be constructed without departing from the
spirit and scope of the present invention.
Embodiments of the present invention may also be used to measure the
temperature of laser irradiated tissue. This is done by tuning the NRR
detector to detect anti-Stokes Raman radiation. For the specific example
of argon laser 514.5 nm input, the anti-Stokes radiation corresponds to a
wavelength of 478 nm. The anti-Stokes radiation is weaker than the Stokes
radiation, however its intensity is proportional to the population of the
first vibrational level of the Raman-active molecules. Since the
population of this vibrational level increases with temperature according
to the Boltzman law, a measurement of the intensity of the anti-Stokes
radiation may be used to calculate temperature. Further, the duration,
intensity and position of the laser coagulation pulse may be controlled in
response to the temperature in the manner described hereinabove.
Clearly, many other varied embodiments may be constructed by those skilled
in the art without departing from the spirit and scope of the present
invention.
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