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A novel laser system has been developed to study the effects of multiple laser pulses of differing wavelengths on cutaneous blood vessels in vivo , using the hamster dorsal skin flap preparation. The system permits sequenced irradiation with welldefined intrapulse spacing at 532 nm, using a long pulse frequency doubled Nd:YAG laser, and at 1064 nm, using a long pulse Nd:YAG laser. Using this system, we have identified a parameter space where two pulses of different wavelengths act in a synergistic manner to effect permanent vessel damage at radiant exposures where the two pulses individually have little or no effect. Using a two-color pump-probe technique in vitro , we have identified a phenomenon we call green-light-induced infrared absorption (GLIIRA), where a pulse of green light causes photochemical and photothermal modifications to the chemical constituents of blood and results in enhanced infrared absorption. We identify a new chemical species, methemoglobin, not normally present in healthy human blood but formed during laser photocoagulation which we believe is implicated in the enhanced IR absorption.
Keywords: Selective photothermolysis, telangiectases, hemoglobin, oxy-hemoglobin, met-hemoglobin, Nd:YAG laser.
The design of lasers for the contemporary vascular lesion marketplace must address a number of challenges in biomedical optics, the human-machine interface and basic economics. These challenges include:
 reducing the number of treatments necessary to clear a lesion, perhaps by improving method efficacy,
 reducing the cost per treatment, perhaps through reduction of the learning curve or technique sensitivity of the method, and
 reducing the severity and rate of incidence of adverse side effects, such as purpura, hyper- or hypopigmentation, scar formation and collateral damage in perivascular tissue.
Modern designs have brought treatment of facial vascular lesions to a highly advanced stage, where both adult and pediatric presentations may be treated with confidence both in terms of the prospective outcome and the intermediate healing time.
However, the effective treatment of larger cutaneous telangiectases and spider veins on the lower extremities (colloquially described as “leg veins”) using lasers is hampered by what appear to be mutually incompatible requirements. Proceeding per the methodology of Selective Photothermolysis proposed by Anderson and Parrish 1,2 one would examine the optical absorption spectrum of the target chromophores, in this case deoxy-hemoglobin and oxy-hemoglobin (henceforth deoxy-Hb and oxy-Hb), and be drawn to the intense blue-green-yellow absorption bands.
3,4 Factoring in the very strong absorption of epidermal melanin in the blue region of the spectrum, lasers operating in the range 520 – 590 nm would appear to offer the most promise, and indeed such wavelengths have been used with considerable success on small and medium calibre vessels.
5,6
# To whom correspondence should be addressed.
Phone (520) 621-4116, Fax (520) 621-8076. Email: barton@u.arizona.edu
.

One of the drawbacks of these yellow-green wavelengths is limited penetration depth, caused by scattering in the epidermis and dermis. This restricts the effectiveness of yellow-green lasers in the treatment of larger lower extremity vessels as these necessarily lie at greater depths compared for example to those on the face. The problem cannot simply be overcome by using increased radiant exposures, since side effects associated with increased heat generation by epidermal melanin absorption become more severe. Even with the advent of efficient epidermal cooling devices 6,7 this attenuating barrier imposes practical upper limits to the radiant exposure, particularly in higher order Fitzpatrick skin types.
Approaching this problem from the perspective of tissue optics, one might choose highly penetrating wavelengths in the near infrared spectrum, from approximately 700 – 1200 nm.
8,9 These wavelengths have smaller absolute scattering coefficients than yellow-green wavelengths, and also show more pronounced forward scattering anisotropy, further enhancing penetration into tissue. Lasers operating at these wavelengths have been shown to penetrate skin to 3 mm or more with useful fluence (e -1 depth) 10 , and have been successfully applied to laser epilation.
11 These near IR wavelengths also have their disadvantages however. Experiments show that the penetration of light at 1064 nm into blood held in glass capillaries is on the order of 2 mm for 50% absorption.
12 This low absorption probability predicates the use of extremely high radiant exposures, in some cases greater than 100 Jcm -2 , to achieve effective blood heating.
13 These energy levels can result in increased pain during treatment, and present an increased risk of collateral damage from laser energy not absorbed by the primary chromophore.
14,15
Compromise solutions using polychromatic devices at intermediate wavelengths have shown some efficacy, although there is evidence of an increased incidence of side effects at therapeutic fluences.
16,17
We describe here a new approach using a timed exposure sequence with one 532 nm green pulse and one 1064 nm IR pulse, taking advantage of the best features of both wavelength regimes. In vivo , we demonstrate the existence of a parameter space where the two pulses act synergistically to effect permanent vessel damage at comparatively modest radiant exposures. We also study the phenomenon in vitro , using cuvettes containing whole or diluted blood and laser-generated blood coagula. We show that a 532-nm induced photochemical / photothermal mechanism acts to enhance the absorption of the NIR 1064 nm light. We implicate a new chemical species, met-hemoglobin, in the green light induced infrared absorption (GLIIRA) effect based on the observation of changes in the absorption spectrum of blood in both dynamically evolving and asymptotic static samples.
The in vivo experiment is shown schematically in Figure 1 below.
Fiber Optic
Long-Pulse
532 nm Laser
Fiber
Optic
Computer /
Framegrabber
Long-Pulse
1064 nm
Laser
Fire Command
CCD
Camera
Delay
Generator
Trigger
Signal
Foot Switch Handpiece
Skin Flap
Window
Vasculature
Figure 1: Experimental Schematic for the in vivo experiments. The various components are described in the sections following.

1.
2.1 Animal Model.
Cutaneous blood vessels of the hamster dorsal skin flap model were used. This model has been used extensively to investigate laser-blood vessel interactions, and the surgical procedure has been described previously.
18 Briefly, the blood vessels were exposed by lifting a double thickness of dorsal skin, suturing it to an anodized aluminum fixture, and then removing a 1-cm diameter circle of skin from one thickness. The subdermal blood vessels, which were covered with approximately 50 µm of connective tissue, were protected with a glass window. During surgery, the animals were anesthetized with xylazine and ketamine, 3:4 ratio, 0.1 ml per 100g. Twelve hamsters received dorsal skin flap implants for this study. The animals were allowed to recover from surgery, then reanesthetized and placed in a clear plastic holder. During laser irradiation, the glass window was removed and the exposed subdermis irrigated with isotonic saline to prevent dehydration. Blood vessels were examined under a microscope 24 hours after irradiation and considered permanently damaged if there was no blood flow at this time.
2.
2.2 Lasers and Radiant Exposure Values – In Vivo.
A commercial long-pulse green laser (VersaPulse Cosmetic Laser (VPC) TM ; Coherent Medical Group, Santa Clara, CA) was used to provide the 532 nm laser pulses. Using an adjustable spot-size handpiece at a 3 mm diameter spot, the laser could operate at radiant exposures up to 16 Jcm -2 with a 10 ms pulsewidth. A commercial long-pulse 1064 nm laser (EDW-15;
Equilasers, Sunnyvale, CA) was used to provide the infrared pulses for the experiment. This laser could be configured to give output pulse durations from 0.1 – 10 ms, with pulse energies up to 10 Joules. The fiber-delivered output of this laser was imaged to a 3 mm spot spatially coincident with that of the 532 nm laser at the sample plane using a simple telescope and an adjustable holding fixture. The two lasers were synchronized so that an IR pulse followed a green pulse by a selectable time delay. A fiber-coupled photodiode (Det110; Thorlabs, Newton, NJ) mounted on the handpiece and configured to sense the arrival of the green pulse sent a trigger signal to a digital delay generator (DG535; Stanford Research Systems, Sunnyvale,
CA). This unit in turn sent a fire signal to the long-pulse 1064 nm laser. In this manner, the lasers could be fired with arbitrary time separation and with a pulse-to-pulse jitter of ± 500 microseconds. Pulse sequence timing was monitored using an oscilloscope (TDS 210; Tektronix Inc., Beaverton, OR). Laser pulse energies were monitored using a power meter
(PM30V1 volume-absorbing head, EM5200 readout; Molectron Detector Inc., Portland, OR).
A CCD camera (QN42H; Elmo, New Hyde Park, NY) with an f = 15 mm x0.3 macro lens (Elmo 9831-1) mounted in the holding fixture permitted observation of the interaction region from the window (subdermal) side of the preparation. The output of the camera controller (Elmo CC421E) was relayed to a VCR (JVC S3600U) and then to a video card installed in a laboratory microcomputer (Macintosh G3; Apple Computer Inc., Cupertino, CA). Images from the camera and VCR tape of the irradiation events were processed using NIH Image 19 to determine the blood vessel luminal diameter in the skin flap window. The scaling factor was determined by recording a millimeter ruler. Probit analysis was performed using the program
EZ Probit 20 to determine the probability of permanent damage versus radiant exposure. The probit analysis performed a sigmoid fit to a graph of frequency of occurrence of permanent damage as a function of radiant exposure. Student’s t-test was used to determine the statistical significance of differences in the 50% probability radiant exposure values (RE
50
) between categories.
2.3 In-vitro Model – Static System.
Blood samples were obtained from two apparently healthy Caucasian volunteers. The samples were kept refrigerated, and used within 72 hours after extraction. Samples were diluted where necessary using isotonic saline. Blood coagula were prepared by loading whole blood into custom-designed 20 mm wide x 100 and 200 micron pathlength demountable cuvettes
(Starna Cells, Atascadero, CA), and then “painting” the cuvette with 532 nm light from the VPC laser. Care was taken to ensure that the coagula completely filled the internal pathlength of the cuvettes, eliminating spurious reflections from interfaces in the sample. Total transmission and diffuse reflectance spectra of whole blood and blood coagula were recorded using a UV/VIS/NIR spectrophotometer (Cary 500; Varian Inc., Palo Alto, CA) equipped with an integrating sphere (DRA
400/500). Matched pairs of reflectance and transmission spectra were subjected to Inverse Adding Doubling (IAD) analysis to extract the absorption (
 a
) and reduced scattering coefficients (
 ’ s
).
22
2.4
2.4 In-vitro Model – Dynamic System.
For the two-color in-vitro dynamical experiments, appropriate blood dilutions were loaded into demountable cuvettes of various thicknesses from 0.1 – 1 mm (Starna 20/O-G Series), which were then placed into the central sample compartment of a standard double integrating sphere experiment.
23 The experimental arrangement is shown in Figure 2.

Figure 2: Experimental apparatus for the in vitro portion of this work.
Two integrating spheres (Labsphere RT-60-SF) were coupled using a specially designed sample holder angled to retro-reflect the front-surface Fresnel reflection from the cuvette back through the laser beam launch port. The 532 nm laser was transmitted to the experiment using a fiber-optic cable and the output recollimated using a simple imaging system. Incident pulse temporal waveforms and energies were monitored using a photodiode. Where necessary, the undeflected 532 nm energy in this experimental configuration was absorbed in a beam dump after the second sphere. Each blood sample was used once only to avoid complications resulting from localized microvaporization nucleated by residual coagulum particles. In practice it was also found that the diffuse reflectivity of a sample containing coagulum particles was substantially higher than that of unmodified blood. Two additional lasers were used to perform time-resolved pump-probe experiments in vitro . To probe the transient optical properties at 633 nm, a 15 mW HeNe laser (Melles-Griot 05-LHP-151; Irvine, CA) was launched through a port adjacent to the 532 nm launch port. The beam from this laser was overlapped with the spot from the 532 nm laser at the image / sample plane between the two spheres. To probe the transient optical properties at 1064 nm, a CW diodepumped 1064 nm laser (STAR - 150 mW, Coherent Laser Group, Santa Clara, CA) was used.

Scattered light in the spheres was sampled by fiber-optic patch cables leading to monitoring photodiodes (Thorlabs PDA-55.
The diodes were equipped with interference filters or low-pass cut-off filters to isolate the wavelength of interest and, where necessary, neutral-density filters. Where low signal levels were present, wide-area photodiodes with variable gain (Model
2031; New Focus, Santa Clara, CA) using primarily long-pass absorbing glass filters and only one interference filter close to the diode active area were employed. “Total transmission” measurements were made with a high-albedo scattering plug in the probe laser exit port of the rear sphere. “Undeflected transmission” measurements were made by removing the plug, placing the detector outside the rear sphere and measuring the collimated transmission plus forward scattered light contained within an approximately 2.8° half-angle cone. The “undeflected transmitted” beams at 633 and 1064 nm were monitored by a photodiode (Thorlabs PDA-55), placed behind an interference filter or low-pass cutoff filter for the wavelength of choice, and neutral density filters where necessary. While making these “undeflected” transmission experiments, we also monitored the signal in the rear integrating sphere. We denote these signals as “deflected transmission” measurements, as they constitute a measure of the energy deflected out of the normal path of the probe beam. Comparison of the shapes of the “undeflected” and
“deflected” signals gives a crude indication of the mechanism causing attenuation. If the signal intensity in both channels drops equally, we may hypothesize that absorption is the dominant contributor to the attenuation. However, if the
“undeflected” signal drops faster than the “deflected” signal, we could conclude that, in addition to absorption, energy is also being scattered out of the probe beam.
The light intensity in the reflectance sphere was calibrated using Fresnel reflections from both AR-coated and uncoated windows in the sample position, and angling the sample holder to create a reflection inside the sphere. The transmission of the AR-coated and uncoated windows was measured separately in a spectrophotometer (Cary 50). The transmission sphere was calibrated using neutral density filters in the sample position. The outputs of the sphere photodiodes and also the photodiodes monitoring the source laser power were sent to an oscilloscope (Tektronix TDS224). Waveforms were captured and processed using commercial software (Tektronix WaveStar™ and IGOR™; Wavemetrics, Lake Oswego, OR).
3.1 in vivo results
Initial studies were performed to determine promising 532 and 1064 nm laser parameters and interpulse spacing. The first series of irradiations was performed with the 532 nm laser set to 10 Jcm -2 , the 1064 nm laser set to 40 Jcm -2 and comparatively long delays between the 532 nm and 1064 nm pulses (70 – 150 ms). The green radiant exposure was chosen to give > 50% probability of permanent damage for these vessel sizes based on the earlier work of Barton et al.
24 using 532-nm only. Examination of the videotape post-irradiation showed that the initial 532 nm pulse caused complete stenosis of the vessel lumen of blood at these settings. The stenosis occurred within the minimum resolvable time period (one video frame, <
33 ms) of our technique. With no blood in the irradiation zone, the only visible indication of the 1064 nm pulse interacting with the tissue was a cylindrically symmetric contraction and relaxation of the connective tissue in the interaction region occurring over a period of several hundred milliseconds. Post-irradiation, at these settings, the ends of the vessels bordering on the interaction region were seen to be occluded by dark flow-stopping coagula. The effect of 1064 nm irradiation alone was ascertained in a separate series of experiments. Radiant exposures of 15 Jcm -2 were seen to have no effect on any arterioles or venules of 20 – 200 microns diameter, while radiant exposures of 42.5 Jcm -2 caused permanent damage of the larger arterioles and venules (150 - 200 µm in diameter).
Based on these initial results, and the in vitro results given in the next section, we concentrated this study on a parameter set with overlapping laser pulses. The 1064 nm laser was set to a pulse width of 4 ms, radiant exposure of 14.1 Jcm -2 , and a delay of 6 ms after the beginning of the 532 nm pulse. The 532 nm pulse width was set to 10 ms, and the radiant exposure varied from 4 to 11 Jcm -2 . Both lasers had spot sizes of 3 mm. This parameter space allowed direct comparison with the earlier 532 nm-only work.
24 Approximately 200 blood vessels were irradiated with this overlapping double pulse configuration in this study. Response to laser irradiation included no effect, embolized coagula, vessel constriction, enlargement, complete stenosis, and fixed coagula blocking flow. After 24 hours, some cases of vessel stenosis or occlusion from coagula resolved, and blood began flowing normally again. In rare cases, a blood vessel that appeared only partially constricted immediately after irradiation was completely stenosed at 24 hours. Blood vessels that had regions of complete stenosis or fixed coagula blocking flow were considered permanently damaged, whereas constricted vessels that still maintained some blood flow were not. The results of the probit analyses are given in Figure 3. Separate probit analysis was performed on arterioles and venules in each of three size categories. Also shown for comparison are the probit results from the earlier, green-only study.
24

12
10
8
16
14 arteriole (green and IR) venule (green and IR) arteriole (green only) venule (green only)
6
4
20-50 50-80 80-110
Blood vessel diameter/ µm
110-140
Figure 3: Probit results for the in vivo study. The legend “green only” refers to the original 532 nm results of reference 24. The legend “green and IR” refers to the present work.
3.2 in vitro static studies.
Figure 4 shows the IAD output for the absorption coefficient (
 a
) from 500 nm to 1700 nm for a 5x dilution of whole blood in isotonic saline measured in a 200 micron pathlength cuvette, and a laser-generated coagulum measured in a 100 micron pathlength cuvette. Figure 5 shows the IAD output for the reduced scattering coefficient (
 ’ s
) for a 5x dilution of whole blood
(200 micron cuvette) and a laser-generated coagulum (100 micron cuvette). Additional noise on the signal in the NIR region of the spectrum is an artifact of the integrating sphere electronics attached to the NIR detector.
100
Clot, hct = 49%
Blood, hct = 9.8%
10
1
0.1
600 800 1000 1200
Wavelength / nm
1400 1600
Figure 4: IAD output showing
 a
for a dilution of whole blood and the laser-generated coagulum. The discontinuity at 800 nm is caused by a detector and grating change in the spectrophotometer.
40
Clot, hct = 49%
Blood, hct = 9.8%
30
20
10
0
600 800 1000 1200
Wavelength / nm
1400 1600

Figure 5: IAD output showing
 ’ s
for a dilution of whole blood and the laser-generated coagulum. The discontinuity at 800 nm is caused by a detector and grating change in the spectrophotometer. The aberrant behavior below 600 nm is caused by the breakdown of the IAD approximation.

3.3 in vitro dynamic studies.
Figure 6 shows a representative plot of the time-resolved transient attenuation at 633 nm of a blood sample (2x dilution in isotonic saline, 200 micron pathlength) being irradiated at 10 Jcm -2 using the 532 nm laser. The traces are the diffuse reflected 532 nm light, the undeflected transmitted 633 nm light and the deflected transmitted 633 nm light. All traces have been low-pass filtered in the plotting program to remove 20 kHz ripple from the VPC flashlamp power supply. Figure 7 shows a representative plot of the time-resolved transient optical properties at 1064 nm of a blood sample (2x dilution in isotonic saline, 200 micron pathlength) being irradiated at 532 nm with 10 Jcm -2 . The traces are the diffuse reflected 532 nm light, the total transmitted 1064 nm light and the diffuse reflected 1064 nm light. All traces have again been low-pass filtered.
3.5
532 nm Remittance Signal
"Undeflected" 633 nm Transmission
3.0
"Deflected" 633 nm Transmission
Undeflected/Deflected Signal Ratio
2.5
2.0
1.5
1.0
0.5
0.0
-10 0 10
Time / s
20 30 40x10
-3
Figure 6: Time-domain transient absorption spectra at 633 nm for a 2x dilution of whole blood being illuminated by 532 nm light at a radiant exposure of 10 Jcm -2 .
0.25
0.20
0.15
0.10
532 nm Remittance Signal
1064 nm Total Transmission
1064 nm Diffuse Reflectance
95
90
85
80
75
5.5
5.0
4.5
0.05
0.00
-10 0 10 20
Time / s
30
70
65
40
4.0
3.5
50x10
-3
Figure 7: Time-domain transient optical properties at 1064 nm for a 2x dilution of whole blood being illuminated by 532 nm light at a radiant exposure of 10 Jcm -2 .

The most striking result from this initial study is the observation that two laser pulses, which would be substantially subtherapeutic in isolation, can be appropriately sequenced to produce highly effective laser photocoagulation and sclerosis of cutaneous vessels in vivo. For both arterioles and venules in every category, the RE
50
value for the two-wavelength case is substantially less than the 532 nm-only case (significance P < .005). We also note that the reduction in RE
50
values in the two-photon case is greater for arterioles than venules. As with the 532 nm-only case, for two-wavelength irradiation, the RE
50 values for arterioles are greater than venules of the same size category, and the RE
50
for the largest size category is greater than the smallest size category for both blood vessel types (P < .01 for both). Interestingly however, for the two-wavelength case, the smallest two size categories of arterioles do not have a significantly different RE50. One possible explanation has to do with the poor absorption of 1064 nm radiation. The penetration depth (1/(µ a
+ µ’ s
)) of 532 nm light is approximately 37
µm in blood, whereas for 1064 nm light it is approximately 610 µm for blood and 191 µm for coagulum. So although enough
1064 nm light was absorbed in the 20-50 µm blood vessels to significantly reduce the RE
50
as compared to the 532 nm-only case, it is possible that the increased light absorption and thermal relaxation time of the 50-80 µm blood vessel creates a more optimum match to the two wavelength condition.
A key to effective vessel destruction seems to be to tailor the initial green pulse to heat the blood in the target vessel to coagulation, but not to the point where the blood partially vaporizes and clears the lumen, or where mechanical constriction clears the lumen. If the lumen is clear when the second (IR) pulse strikes, this laser pulse has no chromophore with which to interact, and no additional damage to the vessel occurs. For the vessels studied here, the 532-nm-induced clearing of the lumen occurred in less than one video frame, or approximately 33 ms. By counting video frames, we have determined that the vessels stay clear for over 120 ms following 532 nm irradiation when the radiant exposure is high enough to clear the vessel.
A clue to the optimum timing sequence for the two pulses is given by the dynamic in vitro studies (Figures 6 and 7). The waveforms will be discussed at greater length elsewhere.
25-27 For the moment, we associate the sharp rise in reflectivity of the sample at 532 nm partway through the pulse to the onset of coagulation, following the interpretation of Verkryusse et al.
28
From Figure 7 it can be seen that approximately one third of the way through the green pulse, the attenuation of the 1064 nm light increases significantly. Data taken using a longer time-base on the oscilloscope indicate that the decrease in transmission persists for at least 50 ms after the end of the green pulse. The form of the time-resolved attenuation curve for
1064 nm is very similar in shape whether one monitors the total transmission, the “deflected transmitted”, or the “undeflected transmission”.
26 From this observation we may hypothesize that during the green pulse , absorption by the sample is the dominant contributor to the observed transient attenuation at 1064 nm. Losses due to scattering of laser energy out of the probe beam in the evolving coagulum appear to play a secondary role. This behavior should be contrasted with that at 633 nm
(Fig. 6), where the deflected and undeflected transmission signals display different behavior, indicating that scattering is more significant. In another experiment, we have determined that using 1064 nm light (from the EDW-15 laser) alone, a radiant exposure of 90 Jcm -2 is required to induce a coagulum in the cuvette. From this observation, we may reasonably conclude that the probe beams at mW power levels are not perturbing the sample.
For a fixed pathlength, the observed increase in absorption requires either an increase in absorber concentration, or an increase in the net absorption coefficient for the system. Since all absorbers in the system contribute linearly and additively to the observed total value, an increase in the net absorption coefficient for the system and a change in the shape of the absorption spectrum (Fig. 4) seem to imply the creation of one or more new absorbing chemical species. A clue to the possible nature of one of these species is given by visual inspection of the laser-generated coagula, which have a dark reddish-brown, chocolate-colored appearance. In healthy humans in-vivo , the iron atom complexed in the porphyrin ring is the Fe (II) (ferrous) oxidation state in deoxy-Hb. When a molecular oxygen ligand binds to the system, an electron transfer reaction occurs yielding the Fe (III) (ferric) oxidation state and the superoxide ion O
2
. The charge transfer is reversed when the hemoglobin liberates the oxygen molecule in tissue. However there is a third form of hemoglobin, called met-Hb, in which the iron exists in the Fe (III) oxidation state but where it is incapable of exchanging molecular oxygen in tissue. In a clinical setting, this species has been implicated in so-called blue baby syndrome and in nitrate-based fertilizer poisonings.
29
In vivo , met-hemoglobin is formed from the deoxy form of Hb, typically through direct chemical oxidation. We suggest that one of the species involved in the GLIIRA effect is met-hemoglobin, based on the following.
The IAD program output from processing the static (spectrophotometer) reflectance and transmission spectra of the laser generated coagulum shows distinct “shoulder” features between 600 and 800 nm. Figure 8 shows the IAD output for the

laser-generated coagulum and the blood dilution, and superimposed on this, the spectrum of met-hemoglobin from Zijlstra et al .
3 The blood dilution spectrum and the met-Hb spectrum have been scaled as a visual guide.
300
250
Coagulum, hct = 49%
met-Hemoglobin Absorption
Oxygenated Blood Absorption
200
150
100
50
0
500 550 600 650
Wavelength / nm
700 750 800
Figure 8: The absorption spectrum of the laser-generated coagulum in the 500 – 800 nm range, overlaid with the spectrum of blood and met-hemoglobin. The coagulum spectrum displays features seen in both oxy- and methemoglobin.
The shoulder feature at 632 nm present in the coagulum spectrum, but not in the native blood spectrum, is suggestive of met-
Hb. Also implicating met-Hb, at least as a transient intermediate, is the very large increase in attenuation at 633 nm and the significant increase in 1064 nm attenuation in the time-resolved spectra. The wavelength of maximum met-Hb absorption in the red spectral range is, fortuitously, 632 nm.
3 At this wavelength, the met-Hb absorbance is 3.8x higher than that of deoxy-
Hb and 30x higher than oxy-Hb, the major component of the starting material. If we use the 1000 nm results of Zijlstra et al. as an approximation to 1064 nm, met-Hb has an absorbance about 13x higher than that of deoxy-Hb and about 3x higher than that of oxy-Hb. In a separate experiment 25,27 , we have measured the time-domain remittance of the sample at 633 nm. This measurement indicates that the back-scattered photon flux from the sample actually decreases significantly at the onset of coagulation during the green pulse. This contrasts directly to the same experiment at 1064 nm (Fig. 7) where the backscatter increases. The most satisfactory explanation for a decrease in both transmission and remittance is that the absorption coefficient of the sample at 633 nm is increasing quite significantly, consistent with the met- hypothesis.
Making some assumptions, we can calculate the % conversion from oxy- to met- in the pump laser pulse in vitro . The assumptions are:
(a) we generate the met- form,
(b) we have the same ratio of absorbances for oxy-/met- at 1064 nm as found by Zijlstra et al . at 1000 nm, and
(c) absorption is the dominant factor in the 1064 nm attenuation during the green pulse, as suggested by the deflected and undeflected transmission measurements.
25-27
Based on the above, for the approximately 10% drop in transmission during the green pulse suggested by the data of figure 7, we calculate a conversion of between 70 – 80 % in the probe volume from the oxy- to the met- form. Also, assuming that we have the met- form, we note that this GLIIRA process has the potential to change 1064 nm from a relatively poor wavelength for vascular treatments to an exceptionally good one. A 3x increase in the absorption coefficient of the target will reduce the penetration depth of light (e -1 ) in whole oxygenated blood at 1064 nm from 1.25 mm to approximately 400 microns.
26 Finally on this theme, we recognize that even if scattering does play a role in the increased attenuation in the GLIIRA effect, we can consider it to be an advantage. In this case increased scattering constitutes an increase in the effective path length through the target, in turn increasing the probability for capture of an IR photon in the two-color treatment. Figure 7 shows that the

remittance at 1064 nm from the sample does increase at the coagulation point during the green pulse, indicating that the reduced scattering coefficient must be increasing. The detailed properties of the coagulation process are in fact much more complicated 25-27 , but to a first order approximation, we believe our discussion above represents the situation well.
The observations made here are corroborated to some extent by the results of Halldorsson 12 who investigated the optical properties of thermally denatured blood. He observed that, at 633 nm, the optical absorption of thermally denatured blood was higher than that of oxygenated blood. At the time, Halldorsson interpreted this as a sign of thermal deoxygenation, a conclusion consistent with the data at this wavelength. At 1064 nm however, he obtained the result, deemed “confusing” at the time, that the absorption of thermally denatured blood was also higher than that of oxygenated blood. If the mechanism were thermal deoxygenation, one would have expected the absorption at 1064 to decrease compared to oxygenated blood.
3
With the benefit of our data, his result can be explained by postulating the generation of met-Hb under conditions of thermal denaturation, which would lead to an increase in the absorption of the samples at both 633 nm and 1064 nm, consistent with his observations.
We have set out a more detailed discussion of the possible mechanisms for met-Hb generation elsewhere.
26 Although photochemical generation of the met- form has been observed for blood under UV irradiation 30,31 , this is the first example that we are aware of where visible light has generated this species. It is possible that the hemoglobin protein structure must be denatured in order for met-Hb to form. It is well known for example that protein conformational changes are involved in the oxy-deoxy transition.
32,33 This prerequisite might explain the onset of met-Hb production at the point where the sample starts to coagulate (the temperature at which protein denaturation takes place). If met-Hb is formed preferentially from oxy-Hb, it might also explain the pronounced effect of the sequenced pulses on the arterioles. These would tend to have higher oxygenation ratios than the corresponding venules, so the concentration of met-Hb formed during the green laser pulse may be higher. This would in turn make the arterioles more opaque to the 1064 nm light, and lead to higher vessel temperatures during the second laser pulse.
This observation of a cooperative effect between two laser pulses has some important implications for laser treatment of vascular abnormalities in humans if the phenomenon is also found in a clinical setting for this species.
34 In practice, based on the % conversion to met- calculated above, we may expect vessels with diameters around 2 – 3 mm to become completely opaque (transmission < e -5 ) to the second IR photon in our two-photon technique. This implies that we can treat with considerably reduced IR radiant exposures compared to the single wavelength case 13 , reducing the pain associated with the treatment and the risk of post-operative complications associated with collateral damage in the perivascular tissue. We may also expect a lower incidence of complications associated with melanin absorption in the basal layer of the epidermis (DE junction separation, post-operative pigmentation changes, etc.) in the case of the yellow-green exposure. The combination system may also allow the treatment of higher-order Fitzpatrick skin types, restricted previously because of pigmentationrelated side effects.
6 Alternatively, for commonly accepted treatment parameters in long-pulse 532-nm therapy, a higher probability for complete vessel damage may be expected in the presence of a second IR pulse. This could contribute to a more standardized treatment regimen.
The GLIIRA effect also offers some intriguing possibilities for site-selective photocoagulation of vascular abnormalities. We can speculate about a possible two-stage technique, in which a yellow-green laser is used to modify a volume of blood so that it contains a substantial fraction of met-Hb. Normal circulation then moves the modified sample to a point where we would like to stop the flow. The progress of the modified blood could be tracked using remittance at a wavelength strongly absorbed by met-Hb, 633 nm for example.
25,27 An infrared laser pulse could then be applied to the modified blood sample sealing off the vessel by taking advantage of the GLIIRA phenomenon.
Finally, it is interesting to speculate about the possibilities of other resonant two-photon cooperative phenomena in biological systems. In a series of seminal papers in the early 1990’s, Crim and co-workers 35,36 explored the possibilities of what they termed vibrationally-mediated chemistry . In this technique, absorption of an infrared photon by a chromophore modified the spectroscopic behavior of the chromophore, and made it more susceptible to uptake of a second photon at a wavelength at which the original chromophore would be transparent. Crim and coworkers demonstrated unique control of unimolecular photodissociation and bimolecular reactions using this technique. While it may not be possible to achieve the same level of state-specificity in larger systems, the absorption spectra of vibrationally excited biomolecules and transient intermediates in photochemistry are largely unknown. It may be that using resonant two-photon excitation, new therapeutic uses for existing laser wavelengths could be found.

We have shown that two sequential laser pulses, one of 532 nm (green) light and one of 1064 nm (near infrared) light, when applied to cutaneous vessels with an appropriate time delay, can produce a photocoagulation effect greater than that which would be expected from the individual isolated pulses. In vitro experiments have shown that photochemically-induced changes in the blood cause an enhancement of the absorption of red and near-infra-red wavelengths. We have implicated the production of met-hemoglobin as contributing to the enhancement of IR absorption. These observations offer intriguing possibilities for treatment of vascular disorders in humans, specifically a reduction in the total radiant exposure required for treatment, and a concomitant reduction in potential adverse side effects and pain associated with the procedure. Future work will include using a high-speed thermal imaging camera to determine the temperature at which the met-Hb reaction starts, which may in turn allow extraction of the activation energy for the reaction. Clinical trials in humans are underway.
Acknowledgments:
Dr. J. K. Barton has received an equipment loan from Coherent, but otherwise has no financial involvement with the company. We would like to acknowledge stimulating discussions with Drs. Christine Dierickx, Rox Anderson, Eric Bernstein and Jim Holtz (Coherent Star). We would like to thank Dr. Scott Prahl at the Oregon Medical Laser Center for enlightening discussions relating to the IAD process, and Dr. A. Roggan at the Free University, Berlin, Germany, for providing an electronic copy of blood optical properties. Finally, we would also like to acknowledge Sherry Tolman and Aaron Brown for their mechanical design and engineering skills.
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