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BACKGROUND OF THE INVENTION
A. Field of the Invention
The present invention relates to a non-invasive method and apparatus for
measuring the concentration of D-glucose in the ocular aqueous humor. More
particularly, the present invention is a non-invasive technique for the in
vivo measurement of the glucose concentration in the ocular aqueous humor
employing the stimulated Raman effect.
B. Background of the Invention
Diabetes Mellitus is a major health problem in the world today because of
the physical complications which arise from living many years with
above-normal blood glucose levels. Currently, over 11 million people
suffer from diabetes in the United States alone. The two most common forms
of diabetes are Type I, juvenile-onset, and Type II, adult-onset. Type I
diabetes destroys the vast majority of the insulin-producing beta cells in
the pancreas, forcing its sufferers to take multiple daily insulin
injections. Type II diabetes is usually less severe than Type I as some
endogenous insulin production still occurs and, as a result, type II
diabetes can often be controlled by diet alone.
The body requires insulin for many metabolic processes; it is particularly
important to the metabolism of glucose. It is believed that many of the
physical complications associated with diabetes could be avoided if normal
blood glucose levels were maintained throughout each day. A diabetic's
blood glucose level can fluctuate widely around each meal. Maintaining
normal blood glucose levels and reducing these fluctuations requires using
some form of feedback to regulate the multiple daily insulin shots of Type
I diabetics or the diet of type II diabetics.
Currently, the blood glucose level can be determined by a chemical reaction
performed on a blood sample. Although the state of the art glucose
measurement devices are very accurate, the need for a blood sample for
each measurement limits their utility. The most dedicated diabetic patient
may take only 4 or 5 measurements per day, and many diabetics perform even
fewer. Because a diabetic's blood glucose level can fluctuate by a factor
of two or more in a period of an hour, this method cannot provide the
feedback necessary to maintain a normal blood glucose level throughout the
day.
A non-invasive blood glucose measurement technique would allow a large
number of daily measurements to be taken without the problems associated
with taking blood samples. Various schemes have been attempted to
non-invasively measure blood glucose level. Many promising techniques
attempt to measure the glucose level in the ocular aqueous humor because
it has been shown that the ocular glucose level directly correlates to the
blood glucose level and because the ocular aqueous humor provides a much
simpler spectroscopic environment than the blood.
D-glucose occurs normally and in abundance in both the blood and the ocular
aqueous humor. There are two anomers of D-glucose found in nature:
.alpha.-D-glucose and .beta.-D-glucose, which differ only in the
orientation of the groups attached to the C-1 carbon. Physically, these
two anomers of D-glucose can be distinguished by their optical activity;
i.e. based upon their ability to rotate the plane of polarization when
illuminated with plane polarized light. In general, the specific rotation,
[.alpha.], is defined as
##EQU1##
where .alpha. is the total optical rotation of the plane of polarization
measured in degrees, .iota. is the length of the sample in decimeters, and
d is the density in g/cm.sup.3. The specific rotations of
.alpha.-D-glucose and .beta.-D-glucose are 112 and 19 degrees,
respectively. In solution, one anomer is converted into the other as
necessary to achieve an equilibrium solution which has a specific rotation
of 52.7 degrees.
Since the specific rotation of D-glucose in solution is known, from
Equation (1) one can infer the concentration of D-glucose in a given
sample by measuring the total optical rotation. The accuracy and linearity
observed at very low D-glucose concentrations led March et al. to attempt
non-invasive measurements in the eyes of rabbits. See Rabinovitch, March
and Adams, Non-invasive Glucose Monitoring of the Aqueous Humor of the
Eye: Part I. Measurements of Very Small Optical Rotations. 5 Diabetes Care
1254 (May-June 1982); March, Rabinovitch and Adams, Non-invasive Glucose
Monitoring of the Aqueous Humor of the Eye: Part II, Animal Studies and
the Scleral Lens, 5 Diabetes Care 259 (May-June 1982). Unfortunately,
March and his colleagues experienced great difficulty in measuring the
concentration of D-glucose in the ocular aqueous humor. Many compounds in
the ocular aqueous humor other than D-glucose are optically active and
contribute to the rotation of the plane of polarization. In addition, the
cornea has birefringence, which causes a further rotation of the plane of
polarization of the incident light. See generally, Gough, The Composition
of and Optical Rotary Dispersion of Bovine Aqueous Humour, 5 Diabetes Care
266 (May-June 1982).
Both spontaneous and stimulated Raman spectroscopy are potentially useful
to measure the concentration of an Raman active molecule in a medium. With
spontaneous Raman spectroscopy a monochromatic laser beam is directed into
a Raman-active medium. Some of the incident beam is transmitted, some of
it is absorbed, and some of it is scattered. A small fraction of the
radiation scattered is shifted in frequency from the incident beam. The
amount of this relative frequency shift is related to the vibrational
states of the Raman active molecules in the medium. The problem with
spontaneous Raman scattering is that the Raman power is scattered in all
directions. This makes the detection of the scattered radiation difficult
for in vivo measurements.
Stimulated Raman spectroscopy (SRS) directs two monochromatic laser beams,
a pump laser beam and a probe laser beam, into a Raman active medium. If
the power of the pump laser is modulated, then the spontaneous Raman
scattered power will also be modulated, which will induce a signal on the
probe laser beam. Thus, rather than measuring the spontaneous Raman
scattered power directly, a measurement of an intensity fluctuation of the
probe laser beam can be made.
Stimulated Raman spectroscopy has been successfully used to measure very
low concentrations of certain selected organic liquids diluted by water
and other solvents. Owyoung and Jones performed a series of experiments
with benzene using stimulated Raman scattering techniques. Owyoung,
Sensitivity Limitations for CW Stimulated Raman Spectroscopy, 22 Optics
Communications 323 (September 1977); Owyoung and Jones, Stimulated Raman
Spectroscopy Using Low-Power CW Lasers, 1 Optics Letters 152 (November
1977). Their experimental set-up consisted of two lasers, a tunable pump
laser and a fixed frequency probe laser. The pump laser Power was
modulated while the probe laser power was held constant. The two laser
beams were combined and focused through a benzene cell. In the cell the
stimulated Raman effect caused a very small fraction of the Power at the
pump wavelength to be shifted to the probe wavelength. Thus, at the output
of the benzene cell the probe laser beam carried a small modulation signal
whose amplitude was directly proportional to the concentration of the
benzene in the cell. The probe wavelength was separated from the pump and
converted to an electrical signal by a photodiode. Both the probe signal
and the input pump modulation signal were fed into a synchronous detector
which greatly improved the signal-to-noise ratio. The pump laser is then
repeatedly tuned to new wavelengths to scan a range of wavelengths, thus,
obtaining a Raman spectra for the Raman-active liquid or gas. This is the
same type of spectrum obtainable by using a commercially available Raman
spectrometer.
Until the present invention, no one has developed a technique which would
allow for non-invasive in vivo measurement of the glucose concentration in
the ocular aqueous humor. March attempted a non-invasive technique
employing an energy wave transmitter, such as an infrared source located
on one side of the cornea and an associated detector on the opposite side
of the cornea. See U.S. Pat. No. 3,958,650. The wave source is aimed to
cause the radiation to pass through the cornea and the aqueous humor to
the detector. A transmitter is mounted adjacent to the detector and
coupled thereto for transmitting a signal that is a function of the
radiation level detected. This technique is seriously flawed. The
radiation detected will be a function of the concentration of all
substituents in the humor, not just glucose. The later optical rotation
technique of March, Rabinovitch and Adams suffers from a similar flaw.
Further, no one, until now, has determined whether stimulated Raman
spectroscopy may be successfully used to measure concentrations of glucose
in the ocular aqueous humor.
SUMMARY OF THE INVENTION
A non-invasive blood glucose measurement technique would allow more
frequent measurement of blood glucose concentrations without the problems
associated with taking blood samples. The present invention achieves this
goal by providing an apparatus and a method for non-invasively measuring
the in vivo concentration of an Raman active molecule in the ocular
aqueous humor by using stimulated Raman spectroscopy. The apparatus of the
present invention includes a means for emitting a probe laser beam and a
means for emitting a pump laser beam. Both means emit monochromatic laser
light and are separated in wavelength by a wavelength chosen to be within
a characteristic Raman shift spectrum for the Raman active molecule. By
setting the separation in wavelength between the pump and probe lasers,
one may select which one of a number of Raman active molecules will be
measured. In the preferred embodiment the selected Raman active molecule
is D-glucose and the separation between the probe wavelength and pump
wavelength is chosen to be 518 cm.sup.-1 in accordance with the
characteristic Raman shift spectrum for D-glucose.
The apparatus also includes a modulating means for modulating the output
Power of the pump laser beam. A power source is provided for the probe
laser for maintaining its power output substantially constant. The
modulated pump laser beam is then combined with the probe laser beam by a
means for directing the combined laser beams into the ocular aqueous
humor.
The introduction of light into the ocular aqueous humor stimulates Raman
radiation which shifts energy from the pump frequency to the probe
frequency, thereby inducing fluctuations in the probe laser beam directly
related to the concentration of the selected Raman active molecule in the
ocular aqueous humor. After the probe laser beam exits the ocular aqueous
humor, means are provided for detecting the probe laser beam and
converting it into a Raman electrical signal. The Raman electrical signal
is then compared to the modulation signal by a synchronous detector, a
dynamic signal analyzer, or a computer based synchronous detection system,
to produce a voltage representative of the concentration of the Raman
active molecule in the ocular aqueous humor.
The method of the present invention non-invasively measures the in vivo
concentration of an Raman active molecule in the ocular aqueous humor
using stimulated Raman spectroscopy. Two monochromatic laser beams are
provided, a probe laser beam and a pump laser beam. The probe laser beam
and the pump laser beam wavelengths are separated from each other by a
wavelength within a characteristic Raman spectrum for the Raman active
molecule being measured. In the preferred embodiment, the Raman active
molecule is D-glucose and the separation in frequency preferably chosen to
be 518 cm.sup.-1. The output power of the probe laser beam should be
maintained substantially constant, while the output power of the pump
laser beam is modulated by a modulation signal. The probe laser beam and
the modulated pump laser beam are combined and directed into the ocular
aqueous humor, thereby stimulating Raman scattered radiation. The probe
laser beam is detected after it exits the ocular aqueous humor and
converted into an electrical signal. The electrical signal is then
compared with the modulation signal to produce a voltage representative
concentration of the Raman active molecule.
It is, therefore, an object of the present invention to provide a system
for measuring an Raman active molecule in the ocular aqueous humor.
It is another object of the present invention to provide a system for
measuring very small D-glucose concentrations.
It is Yet another object of the present invention to make non-invasive in
vivo measurements of D-glucose concentration.
It is a further object of the present invention to allow multiple daily
measurements of D-glucose to be made non-invasively.
It is a still further object of the present invention to provide a
non-invasive glucose measurement system which is inexpensive to
manufacture, durable in construction, and efficient in operation.
These and other advantages will become apparent in the discussion below.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 is a graph of the spontaneous Raman spectrum for D-glucose.
FIG. 2 illustrates the stimulated Raman spectroscopy (SRS) wavelength
selection for the preferred embodiment.
FIG. 3 is a block diagram of one embodiment of the present invention.
FIG. 4 depicts the preferred amplitude modulation of the pump laser power
versus time.
FIG. 5 is a top view of an eye showing entering and exiting laser beams.
FIG. 6 illustrates the modulation of the detected probe laser power after
extraction versus time.
FIG. 7 illustrates the apparatus of the present invention for in vivo
measurement.
FIG. 8 is a block diagram an alternative embodiment of the present
invention.
FIG. 9 is a block diagram of another alternate embodiment of the present
invention using bulk optics.
DETAILED DESCRIPTION
When a monochromatic laser beam is incident on a Raman-active medium some
of the incident beam is transmitted, some of it is absorbed, and some of
it is scattered. A small fraction of the radiation scattered is shifted in
frequency from the incident beam. The amount of this relative frequency
shift is related to the vibrational states of the molecules in the medium.
The D-glucose molecule has several possible Raman active vibrational
states so that the Raman scattered power forms a spectrum which is
characteristic of D-glucose alone.
Turning now to the drawings in which like numerals denote corresponding
parts the preferred embodiment of the present invention is shown. FIG. 1
illustrates the characteristic spectrum for D-glucose dissolved in water
showing the relative intensity of spontaneous Raman power versus the
frequency shift. Each of the peaks in the spectrum corresponds to a
particular vibration of the D-glucose molecule, the largest peak occurring
at a frequency shift of 518 cm.sup.-1. The absolute intensities of the
peaks are directly related to the concentration of the Raman-active
molecule, in this case D-glucose.
Use of a single monochromatic laser beam to generate spontaneous Raman
scattering is difficult for in vivo measurements since the Raman power is
scattered in all directions. This problem can be solved by having two
monochromatic laser beams a pump laser and a probe laser) incident on the
chosen sample. Thus, in the preferred embodiment, the probe laser is at
the same frequency as the Raman scattered power from a large peak in the
D-glucose spectrum, as illustrated in FIG. 2. The pump laser is at a
frequency whose difference from the probe frequency is equal to the
frequency shift of the large peak selected for the probe laser. In order
to use stimulated Raman spectroscopy to measure the concentration of
D-glucose in the ocular aqueous humor, the frequency difference between
the pump laser beam and the probe laser beam must be chosen to coincide
with one of the peaks in the Raman spectrum for D-glucose. If the power
output of the pump laser is modulated, then the spontaneous Raman
scattered power will also be modulated which will induce a signal on the
probe laser beam. Rather than measuring the spontaneous Raman scattered
power directly, a measurement of an intensity fluctuation on the probe
laser beam is made. This method is called stimulated Raman spectroscopy.
As shown in FIG. 3, the present invention involves two lasers, a probe
laser 20 and a pump laser 22. Both lasers 20 and 22 emit monochromatic
laser beams. The relative wavelength difference between these two laser
beams is adjusted to be the same as the wavelength shift of one of the
largest spontaneous Raman peaks for D-glucose, 518 cm.sup.-1. Other peaks
unique to the D-glucose spectrum may be chosen, for example, 400
cm.sup.-1.
The probe laser 20 chosen operates at a wavelength of approximately 0.83
micrometers with an output power of 20 mW. The probe laser power output
should remain substantially constant over time in order to minimize errors
in measurement. Thus, the sensitivity of the system is directly related to
the ability to maintain the probe laser power output constant. The laser
diode chosen, an SDL-1401-H2, is slightly tunable with temperature.
The pump laser 22 chosen emits light with a wavelength of approximately 0.8
micrometers and an output power of 100 mW. The power output of the pump
laser should ordinarily be greater than that of the probe laser by about
5.times. depending upon the threshold conditions for the medium. The pump
laser wavelength is also slightly tunable with temperature to allow fine
adjustment for optimal signal level.
The actual wavelengths chosen for the pump and probe laser beams are not as
important as the separation between them. That separation should
correspond to a vibrational state of the Raman active molecule being
measured. For example, should it be desirable to concentrate on the
D-glucose peak at 400 cm.sup.-1, a different separation would be chosen
for the wavelengths output by the probe and pump laws. The particular
wavelengths of 0.8 and 0.83 were selected because of the availability of
commercial diode lasers of such wavelengths.
The pump laser beam is amplitude modulated by a biased square wave signal
from signal generator 26. The output of signal generator 26 is used to
modulate the current to the diode of the pump laser 22 thereby modulating
the amplitude of the output of the laser beam. The output of pump laser 22
is a biased square wave, with a maximum amplitude of approximately 100 mW
and a minimum of approximately 0 mW. An example of amplitude modulated
pump laser power over time is illustrated in FIG. 4.
It is possible to use other types of modulation to measure the
concentration of Raman active molecules using stimulated Raman
spectroscopy. Amplitude modulation was chosen over other types of
modulation, such as pulse width modulation, because amplitude modulation
is easier to generate. Further, the type of modulation chosen affects the
complexity of the detection scheme that must be used. The choice of
modulation technique also affects the power incident upon the eye. Any
technique which reduces this power reduces possible damage to the eye and
is to be preferred.
The probe laser beam and the amplitude modulated pump laser beam are fed to
a fiber optic coupler 28 via fiber optic pigtails 21. Only 50% of the
power in the probe and pump laser beams is coupled into the fiber optic
pigtails 21 due to their coupling efficiency. Another 50% of the power in
the probe and pump laser beams is lost when the fiber optic pigtails 21
are joined together by optic coupler 28. Thus, the maximum power which
could reach eye 40 is approximately 25 mW when starting with a laser power
output of 100 mW.
Fiber optic coupler 28 combines the probe and modulated pump laser beams
and directs them into an optional spatial filter 30. The spatial filter 30
converts the cross-section intensity to a Guassian Distribution which
allows the beam to be focused more precisely and insures the complete
combination of the two wavelengths before the laser beams are directed
into the eye 40.
The combined laser beams travel through fiber optic cables associated with
means for delivering the laser beams to the eye, preferably in the form of
a handset (not shown). The handset is held up against eye 40 and directs
the probe and modulated pump laser beams into the ocular aqueous humor.
The trajectory of incident laser beams 60 can be seen in FIG. 5. The laser
beams 60 are passed through the cornea 36 and the aqueous humor 38 in such
a manner as to bypass the lens 32 and the iris 34. Pump and probe beams 60
excite stimulated Raman radiation while inside the ocular aqueous humor
38. The scattered Raman radiation causes a small amount of the energy at
the pump frequency to be shifted to the probe frequency, thereby inducing
fluctuations in the previously constant power level of the probe laser
beam, as illustrated in FIG. 6. These fluctuations in the probe laser beam
power are directly related to the concentration of D-glucose in the ocular
aqueous humor 38. The now modulated probe laser beam and the pump laser
beam exit the eye and are coupled into an optical fiber in the
handset-cable assembly.
The coupled detected laser beams are passed through a set of cascaded
narrow band interference filters 42. These optional filters are centered
at the probe wavelength plus or minus about 5 nm, thereby filtering the
pump laser beam away from the Probe laser beam. Thus, if the probe
wavelength is set at 0.83 microns, the band width (BW) for the filters
would be BW=825 nm<.lambda.<835 nm. The filters are cascaded because the
desired reduction in the pump laser power cannot be achieved with a single
filter. Further reduction in the power of the pump wavelength may be
accomplished with a prism or grating. Thereafter, the modulated probe
laser beam is applied to a photo-voltaic diode in photodetector/amplifier
44, which outputs an electrical signal. The photo-voltaic diode provides a
current output in relation to the amount of light input. The
photodetector/amplifier 44 includes a low-noise amplifier which amplifies
the low current level output of the photo-voltaic diode to achieve a
voltage level compatible with the circuitry of the synchronous detector
46. A transresistance gain of greater than 10.sup.8 is achieved by the low
noise amplifier, which also filters off the large DC bias.
The operating band of the low noise amplifier is determined by the noise
spectrum of the probe laser. The noise spectrum of the probe laser is
relatively constant between 1-10 kHZ, and increases below this frequency
range. Restricting the passband of the low noise amplifier to this
frequency range helps eliminate undesired noise from the probe laser.
Those skilled in the art will understand that the passband chosen depends
upon the noise spectrum particular to the laser used as the probe laser.
Different lasers may be expected to require different low noise
amplifiers.
The output 45 of the photodetector/amplifier 44 is fed to synchronous
detector 46, such as a Princeton HR-8 PAR lock-in amplifier. The
synchronous detector 46 compares output 45 to the output 27 of signal
generator 26, and generates signal 49, which is representative of the
concentration of D-glucose in the ocular aqueous humor. The synchronous
detector output 49 may then be fed to any standard display element.
Use of the lock-in amplifier involves using the pump laser modulation
signal as an external reference signal and feeding the SRS signal into the
signal input of the amplifier. The bandpass filter for the external
reference may need to be tuned to the pump modulation frequency depending
on the model of the lock-in amp. A phase offset between the reference and
the SRS signal should be zeroed as indicated in the installation manual.
Then the appropriate gain setting for the SRS input signal should be set.
The time constant or integration time setting may vary depending upon the
noise present on the signal. The DC output signal may be read from the
meter in the unit or directed to an external display.
An alternative, and more costly method of generating a representative
electrical signal from the photodetector/amplifier 44 output is to feed
output 45 to a dynamic signal analyzer 48, such as an HP 3561 made by
Hewlett-Packard. Analyzer 48 measures the power contained in the detected
probe laser beam at the modulation frequency. This power is likewise
related to the D-glucose concentration. Thus, it is possible to use at
least two independent methods to calculate the D-glucose concentration.
The procedure for the dynamic signal analyzer involves connecting the SRS
signal to the input jack of the analyzer. The soft key programming is
generally discussed in the user's manual for the analyzer. The frequency
span should be set to a center frequency equal the pump modulation
frequency and a span of about 100 Hz which may vary depending upon signal
noise. Preferably, the unit should also be programmed to RMS average 50
samples. There is a setting for peak tracking which will display a
numerical value for the frequency peak in the local portion of the signal
spectrum which is currently displayed. The vertical scale units should be
set to "linear" and thus the value for the SRS signal corresponding to the
glucose concentration will be the value of the peak frequency component.
The numerical value is displayed on the screen of the analyzer.
Before using the apparatus of the present invention background measurements
should be made to establish a signal reference point and to be certain
that there are no spurious signals in the passband of the low noise
amplifier which forms part of the photodetector/amplifier 44. A noise
spectrum measured by the synchronous detector 46 with only the probe laser
turned on should be made. Another important background measurement is pump
laser power "leakthrough". Although the two narrow band interference
filters filter the pump wavelength, some small amount of pump power still
reaches the detector 46. Since this signal is present whenever the pump
laser is on, it will offset other measurements. In addition to the
presence of pump leakthrough, stimulated Raman scattering takes place in
the fiber optic cables 21 which carry the input power to the eye 40 and
this SRS signal offsets the SRS signal from D-glucose in the ocular
aqueous humor. These offset signals do not directly limit the sensitivity
of the D-glucose measurement since the SRS signal from the D-glucose adds
to the offset signals and thus the offset signals can be subtracted out in
the calculation of the D-glucose concentration. But the relative amplitude
of these offset signals does limit the detector gain which ultimately
limits the system sensitivity.
Occurring naturally in the ocular aqueous humor, water is also a
Raman-active molecule. There is no peak present in the Raman spectrum for
water at the frequency shift of 518 cm.sup.-1 ; even so, some broad
features of the water spectrum will produce an SRS signal at a shift of
518 cm.sup.-1. This SRS signal from water contributes to the offset signal
and should be subtracted out in the calculation of D-glucose
concentration.
A schematic of the present invention for in vitro measurement can be seen
in FIG. 7. The entire apparatus is mounted on a wheeled cart 68. The pump
laser (SDL-2412-H2) and its power supply 72 are mounted vertically on one
side of the cart, while the probe laser (SDL-2412-H2) and its power supply
70 are mounted vertically on the opposite side of the cart. The preferred
power supply for the probe laser is an LDX 3620 by ILX lightwave, Montana,
because of its low noise current to the laser diode and constant power
mode having an automatic compensation feature. Outputs from each of the
pump and probe lasers are fed by optical fiber pigtails 73 to a fiber
optic coupler 75. The pigtails are multi-mode fiber optic cables to match
those supplied with the lasers. The combined probe and pump laser beams
are then fed by fiber optic cables 76 from the fiber optic coupler 75 up
to the spatial filter 77 and from there into an eye, or alternatively,
into a glucose test cell 78, shown in place on a table behind the cart.
The spatial filter 77 is preferably a Model 900 from Newport Optics.
Glucose test cells 78 are used for in vitro measurement. The test cell 78
is machined plastic with special windows of a high quality, low impurity
glass with a special coating to reduce reflections of optical wavelengths
selected for the lasers. Preferred coatings include a magnesium fluoride
or a coating broadband near infrared coating like Newport #Ar.16.
A set of interference filters 79, previously described, receive the pump
and probe laser beams as they exit the glucose cell 78. Fiber optic cable
81 passes the detected laser beams from filters 79 to detector unit 74,
which houses both the photo-voltaic diode and the low noise amplifier. The
output from the detector unit 74 is then fed to the synchronous detection
unit. A switchable laser temperature readout 82 may be provided to monitor
the output of lasers 70, 72.
A block diagram of an alternative embodiment of the present invention is
shown in FIG. 8. The alternative apparatus incorporates both a probe laser
120 and a pump laser 122. These lasers are both monochromatic and both
operate at the same wavelengths discussed hereinabove with respect to the
preferred embodiment of FIG. 3. The pump laser 122 is modulated using an
AM modulation source 126, such as Hp 3314A.
The output of probe laser 120 is connected to an optical coupler 150 which
is reversed to split the probe laser output into two beams, one containing
5% and the other containing 95% of the probe laser power.
Another fiber optic coupler 128 combines ninety-five percent of the probe
laser power with the modulated pump laser beam and outputs the combined
beams to spatial filter 130. Using a handset (not shown) the combined
beams are then directed into the eye 140. Inside the ocular aqueous humor
of the eye 140 stimulated Raman radiation causes a portion of the power at
the pump frequency to be shifted to the probe frequency, thereby
modulating the probe laser beam. As the pump probe beams exit the eye they
are coupled into a fiber optic cable to transport the SRS optical signal
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