INTRODUCTION

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THE BLOOD SUPPLY to the inner retina is well regulated to maintain an adequate O₂ supply to this tissue. Despite this circulation's key role in maintaining viability of the inner retina, little is known about retinal O₂ consumption and how it changes in response to disease or other challenges. Noninvasive optical methods for determining the percent blood O₂ saturation (SO₂) in large retinal vessels in humans and animals have been reported. These methods depend on quantitating differences between oxyhemoglobin (HbO₂) and deoxyhemoglobin (Hb) light absorption spectra (16) at different wavelengths. Hickam and co-workers (7, 8) and Laing et al. (12) investigated the use of photographic densitometry of large retinal vessels in vivo to monitor arterial and venous retinal blood SO₂. Both groups concluded that accurate measurements could be made using the ratio of vessel optical densities (ODs) at two wavelengths, but their methods depended on external calibration by means of independent arterial SO₂ measurement. This type of measurement is laborious and inherently limited by the nonlinearity and variable reproducibility of photographic film. These investigators applied the photographic method to study relationships between retinal circulation and arteriovenous SO₂ differences (7) and assessed its applicability for animal measurements of retinal SO₂ (12).

Delori (4) developed a retinal oximetry technique based on photoelectronic measurements of retinal vessel OD at three wavelengths. His method used a mechanical scanner in a modified fundus camera equipped with a photomultiplier detector to monitor sequentially the light reflected inside and outside vessels at three wavelengths. By using the theoretical relationship between OD and light absorption of whole blood described by Anderson and Sekelj (1) and others (11, 15), he was able to avoid the need for external calibration. This method was applied in a study of retinal vein SO₂ in optic atrophy (6).

We describe a new technique for the determination of retinal vessel SO₂ that uses digitally recorded retinal images obtained simultaneously at two wavelengths: one sensitive to changes in the percentage of HbO₂ present in the blood and the other an isosbestic wavelength for HbO₂ and Hb. Figure 1 shows the effect of differing HbO₂ content on reflected light from retinal vessels near the optic disk. Reflected light intensities from arteries and veins at the isosbestic wavelength (569 nm) are comparable, whereas at 600 nm the high percentage of HbO₂ increases the reflectance of the artery. Our method utilizes the approximately linear relationship that has been found between Hb SO₂ and the OD of hemoglobin in solution (16) and blood (8, 12) as a basis for determining retinal vessel SO₂. We use a vessel tracking algorithm to calculate apparent ODs of blood contained in retinal vessels from reflectance measured on the vessel and in the surrounding retinal tissue. Using assumptions about the optical path through the blood and the hemoglobin concentration to determine actual light absorption in the blood, Delori's method employed similar measurements from vessels to estimate SO₂ (4). Our approach differs from those used previously by the use of digital image analysis of reflectance at two wavelengths to determine empirical relationships between ratios of vessel density and external measurements of the systemic SO₂. The method is simpler in theory and practice and requires less complex instrumentation than those previously described. It can be adapted to most currently available fundus cameras. We also show that it is possible to compensate for the variability in fundus pigmentation among individuals and races. This technique should be useful for studying diseases that affect retinal O₂ utilization and for evaluating potential therapeutic interventions.

View larger version (174K): [in this window] [in a new window]   Fig. 1.   Dual-wavelength image of retinal vessels near optic disk obtained with oximeter. In 600-nm image, arteries (A) appear lighter and veins (V) darker; at 569 nm, both vessel types are dark. Nerve fiber layer can be seen in both images, with greater contrast in 569-nm image.

	METHODS

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Study individuals. Ten healthy nonsmoking men aged 18-40 yr without history of dyshemoglobinemia consented to participate in the study. One subject was unable to complete the oximetry portion of the study. Two subjects were excluded from the oximetry portion of the study, because images obtained from one subject revealed sclerosis of the arterial wall, and in the other subject the artery demonstrated extreme instability of caliber during measurements. Oximetry data from five Caucasians of northern (n = 3) and southern (n = 2) European descent and from two African-Americans were included in the study. Guidelines of the Association for Research in Vision and Ophthalmology for human investigation were followed, and our institution's Human Investigation Committee approved the study protocol.

Breathing gas-monitoring protocol. We placed an Ohmeda 3700 (version J) ear oximeter probe on the massaged earlobe. A Mapleson C breathing circuit and face mask were modified by inserting an inspiratory valve just proximal to the mask, allowing all exhaled gas to be vented through the expiratory valve placed just distal to the mask. This arrangement of components kept rebreathing of alveolar gas and dead space to a minimum. A polarographic inspired O₂ analyzer with a Clark electrode (Oxychek model 2000, Critikon) was calibrated to room air and inserted just proximal to the inspiratory valve (Fig. 2). Subjects were allowed to breathe room air during the first series of fundus image recordings. We then used flows of 10 l/min to deliver various O₂-N₂ mixtures (size H gas cylinders). After the initial image recordings during room air breathing, subjects breathed a gas mixture containing 14% O₂, which corresponds to an arterial Hb SO₂ of ~94%, or a mixture of 10% O₂, which corresponds to an Hb SO₂ of ~84%. In some individuals the 14% O₂ mixture was given first, then the 10% O₂ mixture. A mixture containing 8% O₂ was then given, which corresponds to an Hb SO₂ of ~80%. Finally, subjects breathed 100% O₂, which produces an arterial Hb SO₂ of 100%. When a steady state was achieved at each step (defined by an unchanging pulse oximeter reading for 2 min), we noted the oximeter saturation reading and obtained retinal images with the fundus camera.

View larger version (11K): [in this window] [in a new window]   Fig. 2.   Breathing circuit for administering gas mixtures. Velcro straps were used to fix face mask tightly over subject's mouth and nose. Reservoir bag was inflated with gas mixture before breathing started. FGF, fresh gas flow; IV, inspiratory valve; EV, expiratory valve.

Simultaneous dual-wavelength imaging system. The retina was illuminated using the internal xenon flash of the fundus camera (model RC/W, Kowa, Tokyo, Japan). Flash energy was supplied by an external strobe power supply (model 404, Norman) at a setting of 200 J, giving a retinal exposure per flash of ~100 mJ/cm² (40° field). Flashes were synchronized with the recording system, and wavelengths above and below the recording wavelengths were eliminated with a broad-band filter (580 nm center, 60 nm half-width). The fundus camera was modified to form an intermediate image (II2) just before the exit aperture by using a 6 diopter lens in one of the filter wheel apertures (Fig. 3). A slit placed at this plane contained an ~1 mm wide × 2 mm high image of the retina. To avoid central light artifacts present in the fundus camera, we positioned the slit off the optical axis. Light forming the intermediate image in the slit was separated into two identical-length paths with image-splitting optics, ultimately forming two laterally displaced images at the solid-state camera detector (Fig. 4). The beam was split with a dichroic mirror having a transition at 595 nm (Omega Optical, Brattleboro, NH). Narrow (4.5 nm half-width) band-pass filters with 569- and 600-nm center wavelengths (Omega Optical) were placed in each optical path. We achieved lateral displacement of the images with a series of front surface mirrors that could be moved on rails by means of a microcaliper mechanism. An opaque light baffle blocked stray light between the two images, and a two-element achromatic converging lens (Edmund Scientific, Barrington, NJ) placed before the beam splitter refocused the images at the image sensor. We recorded images with an 18-bit air-cooled digital camera (model OMA, EG & G) equipped with a 1,024 × 1,024-pixel charge-coupled device detector. The camera was controlled using manufacturer-supplied software (HIDRIS). Each image element was formed by 3 × 3 pixels, and the temperature of the charge-coupled device was maintained at 60°C.

View larger version (14K): [in this window] [in a new window]   Fig. 3.   Optical train of Kowa fundus camera as used with image splitter. A negative lens (DL) contained in 1 filter wheel (W) position produces an intermediate image (II2) at a slit mask (S) that projects past back plate of fundus camera. A second positive lens after slit (CL) refocuses image at camera after passing through optics of image splitter. Xe, xenon flash; F, illumination filter; M, mirrors; A, mirror aperture; OL, ophthalmological lens; VFOL, variable-focus objective lens; IS, image sensor; L, lens.

View larger version (32K): [in this window] [in a new window]   Fig. 4.   Optical train of image splitter. Beam passing through slit is refocused with lens (L). Beam is split into 2 paths with a dichroic beam splitter (DBS) and redirected using front surface mirrors (M1-M7) to produce laterally displaced images at sensor. Optical path lengths are made equal by adjusting distances of M1 and M2 (left) and M5 and M6 (right) from main optical axis along tracks (T). Images are positioned side by side by on sensor by small adjustments in position of M2 and M6 relative to M1 and M5. A baffle (B) blocks stray light at image sensor from 2 beams. F1 and F2, filters with 569- and 600-nm center wavelengths, respectively; CCD, charge-coupled device.

Measurement wavelengths. O₂-sensitive images were obtained at 600 nm, where light absorption of HbO₂ and Hb differ by approximately fourfold (16). A second image was obtained at 569 nm, which is an isosbestic wavelength for Hb and HbO₂. Filter half-amplitude bandwidths were 4.5 nm. We employed the isosbestic wavelength, rather than a second O₂-sensitive wavelength, to assess stability of vessel measurements.

View larger version (22K): [in this window] [in a new window]   Fig. 5.   Points on absorption spectra of oxyhemoglobin (HbO2, ) and deoxyhemoglobin (Hb, ) near measurement wavelengths [data from Van Assendelft (16)]. Solid lines in HbO2 spectrum show segments over which curve fits were performed to obtain product of spectrum and filter function. A single curve was fit to deoxyhemoglobin spectrum. Transmission spectra of filters were used to select O2-sensitive (600-nm) and O2-insensitive (569-nm) images (dashed lines). Solid lines, polynomial curve fits. OD, optical density.

Fundus recording. Eyes were dilated with 1% topical tropicamide and 2.5% phenylephrine. After the pupil diameter stabilized, the room was darkened and the subject's fundus was illuminated using the tungsten aiming light of the fundus camera. Positioning of eye gaze to enable imaging the retinal area of interest was established using a movable illuminated fixation target against a dark background. Optimal focus was also achieved using the aiming light. The actual data recording at each breathing gas mixture was made using the xenon flash. Three to five measurements were recorded at each gas mixture and stored to disk memory.

Image analysis. A computer program scanned the digital image to obtain average reflected light intensity from identical vessel segments in both laterally displaced images. Figure 6 shows the path obtained by tracking the minimum reflection inside vessels and the path of the extravascular reflection at a fixed distance from the minimum reflection in both images. To begin a scan, the initial x-y coordinate of the vessel segment was identified in the 569-nm image, where vessel contrast was greater than in the 600-nm image, and x- and y-offsets to the same anatomic point on the 600-nm image were determined. If light reflex artifacts were present in the center of the vessel, the starting point was chosen to lie to the side of the reflex. Between 30 and 120 rows or columns of elements in the 569-nm image were scanned, depending on the orientation of the vessel, to establish the path of the minimum reflected light intensity along the vessel. Edges of the blood column were identified using gradient detection filters (Sobel operators). The edge was defined at pixels for which the sum of filter responses peaked; this procedure did not mistake smaller light gradients of the central vessel reflection for the edge of the blood column. Minimum intensity inside the vessel was then found by scanning between the edges. In arteries, only the minimum-valued element on each line was used; in veins, larger diameter and smaller central reflexes allowed averaging minimum image intensities within a neighborhood of points. The final measurement of light reflection was obtained from the average of minima along the vessel segment. Extravascular light reflection was determined from pixels at a fixed distance outside the vessels. The algorithm could be preprogrammed to average from both sides or only one side of the vessel if strong reflections from nerve fibers were present on one side. We then made similar measurements from the 600-nm image by adding the x- and y-offsets to the vessel path determined from the 569-nm image. Dark pixel values from an area of the detector outside the fundus image were subtracted from measured intensities.

View larger version (166K): [in this window] [in a new window]   Fig. 6.   Computer-generated scans along retinal vessels in 600- and 569-nm images. Scans inside vessels (thin lines) follow minimum value of reflectance and usually remain to 1 side of central vessel reflections. Scans outside vessels (wide lines) occur at a fixed distance from inner scan. Top and bottom vessels are veins. Middle vessel is an artery. Dark pixel readings are obtained from space between images.

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	RESULTS

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Stability of recording. To test the stability of simultaneous reflectance recordings at each wavelength over time, we obtained sets of images while holding systemic arterial SO₂ at different values. Figure 7 shows reflected intensities from inside and outside vessels at both wavelengths, which were obtained from images recorded at 3-s intervals in one subject who was told to maintain a fixed eye gaze. At each more hypoxic level, there was greater variation in the reflected light obtained from successive images. This pattern was typical of our recordings and is consistent with our experience that optimal alignment of the eye and eyelid is more difficult to maintain under hypoxic conditions. Because the 569- and 600-nm images were recorded simultaneously, changes in light intensity tended to move in parallel. Vessel ODs averaged from these images for each level of SO₂ and corresponding values of ODR obtained using method I (see below) are shown in Table 2. Theoretically, OD₆₀₀ should show an inverse monotonic relationship with SO₂, whereas OD₅₆₉ should remain constant. In practice, these relationships were usually not present in the data for reasons that may be attributable to small changes in focus. However, analyses given below show that when O₂-sensitive ODs are normalized by ODs obtained simultaneously at the O₂-insensitive wavelength, OD₆₀₀/OD₅₆₉ does follow an inverse monotonic relationship with the externally recorded systemic SO₂. For all subjects in the study, this relationship is fit well by a linear model.

View larger version (21K): [in this window] [in a new window]   Fig. 7.   Reflected light intensities from images obtained during different levels of systemic O2 saturation (SO2). Symbols (left to right) represent light intensity values for a fixed systemic SO2 from each successive image. Circles and squares, intensities outside and inside vessels, respectively. Open symbols, 600-nm reflectance; filled symbols, 569-nm reflectance.

Retinal oximetry in arteries. To characterize the response of our oximeter, we first determined its response in retinal arteries at fixed known values of arterial Hb SO₂. We expected that subject-to-subject variations in the amount of pigment contained in extravascular structures would have an effect on our determination of ODs and, therefore, measurements of SO₂. Thus we compared our results using three different methods for analyzing the data, each of which treats variations in background pigmentation differently.

Method I: direct reflectance analysis. In method I, OD₆₀₀/OD₅₆₉, determined by direct calculations of vessel OD from each of the subjects, was plotted against arterial systemic SO₂ and linear regression analysis was performed between these variables. Slopes of regression lines, which are estimates of O₂ sensitivity, were 0.00427-0.01242 [0.00787 ± 0.00098 (SE)] and y-intercepts (ODR at SO₂ 0%) were 0.514-1.50 (1.02 ± 0.115). The goodness-of-fit R was 0.9581 to 0.9999, with an average of 0.98318 (Table 3). These results show that a linear model gives excellent agreement with the change in ODR vs. arterial SO₂ between 77 and 100%. Figure 8A shows ODR points obtained from all seven subjects with associated line fits, along with the O₂ sensitivity estimated from the average of the slope and y-intercepts. By this method, maximum and minimum slopes varied by a factor of 2.91 and the coefficient of variation in regression slope (SE/mean) was 0.125. We found that standard errors of ODRs were typically <3% of the mean when results from four or more images were averaged.

View larger version (17K): [in this window] [in a new window]   Fig. 8.   OD ratios (ODRs) from retinal arteries vs. systemic SO2 (7 subjects). A: ODR obtained from direct calculation of vessel OD by using method I. B: ODR corrected for retinal pigmentation (ODRcor) by using method II. Dashed lines, line fits through points from each subject; solid line, line obtained by averaging slope and intercept from all subjects: 0.00787 ± 0.00098 (SE) for method I and 0.00504 ± 0.00029 (SE) for method II. C: lines obtained from linear regression between ODR from intravascular reflectance model (ODRirm, method III) and systemic SO2. Maximum and minimum values of slopes, representing O2 sensitivity, are 0.00616 and 0.00437 per unit of SO2 (%), which is approximately the same as range obtained using method II. Vessel diameter indicated for each line fit is expressed in µm (12.33 µm/pixel).

Method II: correction for extravascular pigmentation. The analysis using method I may underestimate vessel ODs, particularly at 569 nm, because pigmentation in tissue surrounding vessels causes strong light absorption. To assess the effects of pigmentation on light reflectance at our measurement wavelengths, we determined the ratio of EVR at 569 and 600 nm near vessels of the temporal circulation in nine subjects. EVR averaged from five Caucasian subjects was 0.51 ± 0.25. The two lowest values were seen in subjects with blue irises. In one Indian subject, EVR was 0.6, and in three African-American subjects EVR averaged 0.81 ± 0.06. Our values of EVR are in agreement with previous reflectance spectra (5), which showed a greater reduction in reflectance at longer than at shorter wavelengths with increasing pigmentation. When regression slopes obtained by method I are plotted against EVR (Fig. 9, solid symbols), these variables vary inversely (P = 0.057). Reflectance from extravascular retinal tissue was greater at 600 than at 569 nm in all subjects.

View larger version (17K): [in this window] [in a new window]   Fig. 9.   Relationship between ratio of extravascular reflection at 569 and 600 nm (EVR) and slopes of line fits between ODR and SO2. , Slopes obtained by direct determination of ODR by method I; , slopes obtained by method II with correction for choroidal pigmentation. AA, African-American; C, Caucasian; BK-BR, black hair and brown eyes (refers to 3 neighboring points); BR-BL, brown hair and blue eyes; R-BL, red hair and blue eyes.

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View larger version (18K): [in this window] [in a new window]   Fig. 10.   Relationship between measurement parameters and vessel diameter. Top: ODRcor at 100% O2 (method II) vs. diameter in pixels. Linear regression slope = 0.0129/pixel (R = 0.7739, P = 0.045). Bottom: slope of regression between ODRirm and systemic SO2 (method III) vs. diameter. Linear regression slope = 18.4/pixel (R = 0.8224, P = 0.023).

Method III: intravascular reflectance model. An estimate of SO₂ that is independent of effects of extravascular pigmentation can be made by comparing reflectance from only within blood vessels at the two wavelengths and replacing measured reflectance outside vessels with an estimate of reflectance from unpigmented fundus. If arterial density at 100% SO₂ is set primarily by light absorption by HbO₂, an appropriate estimate of unpigmented reflectance (UR) would set vessel ODs at each wavelength to values predicted from the HbO₂ absorption spectrum. We propose a model that yields this value of UR from Hb extinction coefficients (broad-band coefficients in Table 1) and measured reflectance from inside vessels at each wavelength when blood is 100% saturated with O₂ (intravascular reflectance model)

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We used Eq. 5 with measurements of I_in obtained from images recorded during 100% O₂ breathing to calculate UR by an iterative method. We then applied this value of UR to measurements during normoxic and hypoxic recordings of each subject to calculate ODR derived by the intravascular reflectance model (ODR_irm) at each breathing set point. Extravascular intensity scanned near the vessel at 600 nm was used to compensate for differences in illumination at each point. Figure 8C shows regression lines that were fit to these sets of points for each subject, along with diameters of each artery. Initial assumptions cause the lines to meet near 100% SO₂, and thus this plot shows more clearly the range in slopes from our subjects. It was not possible to predict slopes from subject coloration or from the EVR measured in each subject. In this model, regression intercepts and slopes are codependent; correlation between the slope and vessel diameter was significant to the 0.05 level (Fig. 10, bottom). These results again suggest that there is a small dependence of the O₂ sensitivity on size of vessels in the range 80-200 µm.

Venous SO₂. Figure 11 shows ODRs of an artery-and-vein pair recorded from the temporal retinal circulation in one subject. Although a priori there is no reason to expect that venous SO₂ should vary linearly over the range of arterial saturations studied, we found that venous ODRs obtained during room air breathing and hypoxia fell on straight lines, whereas the ODR at 100% O₂ breathing was below the line, a finding similar to that reported by Hickam and Frayser (7). Presumably, the high concentration of dissolved O₂ at 100% significantly reduces dissociation of HbO₂ during capillary passage, lowering the ODR. This effect was seen even though vasoconstriction during hyperoxia reduces blood flow, allowing more extraction of O₂, which, by itself, would have increased the ODR.

View larger version (12K): [in this window] [in a new window]   Fig. 11.   Values of ODRcor obtained for an artery-vein pair plotted against systemic SO2. Vein slope = 0.00315/%SO2 (R = 0.994, excluding 100% O2 point). Artery slope = 0.00808/%SO2 (R = 0.992). Error bars, SE.

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	DISCUSSION

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We have described a new imaging method for measuring retinal vessel Hb SO₂. The use of an image splitter and high-resolution camera to image vessels simultaneously at two wavelengths is an extension of the earlier work of Hickam et al. (7, 8) using photographic film and offers significant advantages. Almost perfect linearity and reproducibility of the solid-state camera enable detection of small differences in reflected light associated with changes in vessel SO₂. Equally important, we found that a ratiometric approach effectively cancels effects of subject motion, which otherwise interferes with the recording process. We used the image recorded at 569 nm as a reference image with which to cancel changes unrelated to blood oxygenation in the 600-nm image. With these recording techniques it was possible to obtain estimates of blood OD by relatively simple vessel-tracking methods.

We found in each subject that the ODRs, when plotted against 77-100% systemic SO₂, fell along straight lines, with R close to 1. These results agree with previous ratio measurements obtained from retinal vessels by reflectance (8, 12) and are predicted over our range of SO₂ by assuming the quasi-linear D-E characteristic described by Delori (4). Using reflectance pulse oximetry, de Kock et al. (3) showed a nonlinear relationship at <85% saturation in vitro; however, the method of de Kock et al. relied on pulsatile changes in reflectance from vessels and surrounding retinal tissue to detect arterial blood, rather than directly imaging vessels to determine apparent ODs.

Differences in choroidal pigmentation among subjects initially influenced the sensitivity of our oximeter to changes in retinal vessel SO₂, producing variability in slopes of line fits against known systemic SO₂ values. We were able to reduce the effect of light absorption by extravascular pigment by substituting reflectance at 600 nm, where pigment light absorption is substantially weaker than at 569 nm (method II). This correction significantly reduced variability in regression coefficients. Elimination of effects of extravascular pigment by using only measures of intravascular reflectance (method III) did not result in significant further reduction in the variation of O₂ sensitivity among subjects. Neither method II nor III addressed effects of pigment behind the vessel that may still exert an influence on the apparent vessel OD. At longer wavelengths, including 600 nm, retinal reflectance has been shown to be more variable than at shorter wavelengths among subjects of different coloration (5). However, we used our longer wavelength to correct for the influence of choroidal pigmentation, because the degree of light absorption at 569 nm by these pigments is significantly greater than at 600 nm. The stronger light absorption at 569 nm produces a larger numerical effect on the calculation of vessel OD than do variations at 600 nm.

After correction for pigmentation, remaining scatter in ODR parameters was correlated with vessel diameter. Hickam et al. (8) and Delori (4) found that their measures of SO₂ were dependent on vessel size: Hickam et al. concluded that similarly sized vessels near the optic disk could be accurately measured with a single calibration, and Delori incorporated vessel diameters into his theoretical analysis. We adjusted measured ODR values for vessel size. Further improvement in accuracy of our technique by using method II possibly could be obtained if vessel size dependence of O₂ sensitivity (i.e., effect on slope) were also included. However, the degree of correlation between the regression slope for ODR_cor (method II) and diameter in our data was not statistically significant; hence, we have used the mean value of slopes to define O₂ sensitivity. We found with method II that the vessel ODR_cor obtained at 100% O₂ is dependent (P < 0.05) on vessel diameter, and it is easy to measure. If this parameter is obtained, retinal arterial SO₂ can be readily determined by our method. A statistically significant correlation did exist between slopes obtained by our intravascular reflectance model (method III) and systemic SO₂; however, this second method required more computation. Because methods II and III reduced effects of pigmentation on O₂ sensitivity by comparable amounts and because the simpler method II yielded reasonable estimates of venous SO₂ in veins of different diameter, we believe that method II is most suitable for oximetry. The results with both methods indicate that vessel diameter has a slight influence on measurement of vessel blood SO₂ and that compensation for the effect of pigmentation is necessary before effects of vessel size can be addressed.

We have based our determinations of vessel SO₂ on the O₂ sensitivity found from means of regression slopes in arteries, with correction for vessel size using the diameter dependence of arterial ODR at 100% O₂. Determination of the actual O₂ responses in arteries during graded hypoxia is not practical in a clinical setting or each time a new measurement is needed. There has also been no practical way to independently validate retinal venous saturation measurements in human subjects, making it difficult to interpret O₂ extraction data. Confidence in the accuracy of venous SO₂ measurements relies on agreement reached using different instrumental and analytic methods. The mean venous saturation determined by our method is in general agreement with values obtained by the three-wavelength method of Delori et al. (4, 6) and in close agreement with results from the earlier work of Hickam et al. (7, 8) using dual-wavelength ratio analysis from photographic images. Hickam et al. and we assumed that the arterial calibration between ODR and externally measured systemic saturation could be applied to venous ratios by extrapolation toward values of lower saturation. This assumption requires that the relationship be linear over the range of saturation found in arteries and in veins. Shortly after Hickam et al. reported their measurements of venous SO₂ in human subjects, Laing et al. (12) published their "falling saturation" experiment in the rabbit, confirming that retinal arterial ODRs vary linearly with blood SO₂ between 20 and 97%, which covers the normal saturation range of both vessel types. The practical problem with use of arterial calibration data for venous measurement is that any error in the best-fitting regression line leads to inaccurate estimates outside the range of the calibration. Our determination of venous saturation is subject to this kind of error. We note, however, that we were able to obtain goodness-of-fit values close to 1 (0.94 to 0.99) in regression line fits for all our subjects, which increases confidence that a dual-wavelength method yields the expected linear relationship between ODR and SO₂ and that the slope of the best-fit line is a good predictor of other points outside the range of calibration.

Verification of our oximeter response depends on the accuracy of external SO₂ measurements. Clinical studies have demonstrated that the accuracy of pulse oximeters under steady-state conditions is within ±2 saturation points within 100 to 70% saturation (2, 9, 13). We chose to use an ear probe, rather than a finger probe, because of its ability to respond accurately and quickly to a changing systemic arterial saturation. Kagle et al. (9) demonstrated that the Ohmeda 3700 version J ear probe is highly accurate during hypoxic conditions in normal volunteers, giving a goodness-of-fit R of 0.98 for regression of ear probe readings vs. arterial saturation. The slope and intercept for the African-American subjects did not differ significantly from that for the Caucasian subjects, a finding also described by Cecil et al. (2) using the Ohmeda 3700 finger probe. One of the most significant limitations of the pulse oximeter is signal failure due to poor perfusion or hypothermia. The Ohmeda 3700 features a graphic display of waveform and signal strength to aid in minimizing inaccurate readings. We monitored the quality of ear probe signals during our 2-min stabilization period with steady-state oxygenation conditions before recording oximeter readings and obtaining retinal images and used only readings with a strong signal.

Sources of possible error by our technique include effects of changes in red cell aggregation and orientation on light transmission, which was observed at different blood shear rates by Klose et al. (10). This effect of flow must be considered in evaluating oximetry measurements by reflectance. We note that for light transmission of the blood at our measuring wavelengths to change by amounts necessary to affect our results by using a ratio analysis, shear stress would need to change by over two orders of magnitude. This degree is greater than that associated with autoregulatory changes in the larger vessels. Errors could also arise if either light absorption in extravascular pigment or the efficiency of light scattering from blood and surrounding tissue were to change between measurements. We assume that these effects remain constant at fixed sites and thus will not influence measurements of SO₂ when identical vessel segments are imaged before and after changes. The small amount of contrast of arteries against surrounding tissue at 600 nm, which is consistent with the low light absorption by HbO₂, suggests that there is not a strong difference in amounts of light scattered by blood and surrounding tissue backward to the fundus camera, which could also contribute to the vessel OD. Although light absorption appears to govern the changes in vessel OD in our images, determinations of the actual blood OD are complicated by light scattering within the blood and by multiple angles of incident light (4), and thus our calculations by Eq. 1 give only an apparent OD for the vessel. Our analytic approach avoids assumptions about the optical path through the vessel, relying instead on empirical relationships between image parameters and vessel SO₂ obtained from subjects of different coloration. This approach does give reasonable estimates of Hb SO₂ in the larger retinal arteries and veins and should be applicable to 50- to 200-µm vessels.

Perhaps more interesting for understanding effects of metabolism or blood flow changes on retinal function is the ability to determine the changes in SO₂ before and after interventions (14). The change is more straightforward to obtain, since many variables potentially affecting absolute measures of SO₂ will cancel. However, error could occur if vessel diameter changed significantly between measurements. The vessel image should be checked in each recording to ensure that the vessel remains well focused and that diameter changes do not occur. If they do, then the change in diameter should be included in the analysis. The relatively strong vasoconstriction response during breathing pure O₂ causes the larger retinal vessels to constrict by <15% (7), an amount that does not seriously effect the O₂ sensitivity of our method. We conclude that Hb SO₂ in retinal vessels and arteriovenous differences in SO₂ can be determined from simultaneous dual-wavelength measurements. The technique may find application in studying effects of experimental interventions or treatments on retinal O₂ utilization.