2.1.

StO₂ Measurement Equipment

Two spectrometer designs were used during the course of this study. Both designs used a closed cell polyethylene foam light scattering calibrator, Plastazote® LD 45 (Zotefoams Inc, Walton, KY), to capture reference light intensity for sampled light at each wavelength prior to placing a reflectance probe on the tissue measurement sites. Reference intensities were collected before each experiment and stored for subsequent attenuation measurements.

A first full spectrum spectrometer, Biospectrometer-NB (Hutchinson Technology Inc, Hutchinson, MN), consisted of a stabilized 25 W tungsten halogen light source and a reflective grating charged-coupled device (CCD) array photodetector providing wavelength measurements of 550–1000 nm with an increment of approximately 0.5 nm. Six 400 μm glass-receiving fibers were coupled to the grating in a slit pattern to achieve a bandwidth resolution of 15 nm full width half maximum (FWHM). A Windows notebook computer was interfaced to custom software allowing data acquisition of StO ₂ and sample spectrum signals were block averaged and updated every 5 s or less. The software provided autoclock rate adjustments and dark measurements for optical signals in the 680–800 nm range. Fiber optic probes, 3 m in length, used numerous 200 or 400 μm glass optical send fibers to provide tissue illumination for probe spacings of 8 and 15 mm. Known transmission peaks of didymium glass calibrated the CCD array pixels to a corresponding wavelength prior to each experiment.

A second commercially available spectrometer, InSpectra™ Tissue Spectrometer Model 325 (Hutchinson Technology Inc, Hutchinson, MN), provided tissue attenuation measurements at four discrete wavelengths: 680, 720, 760, and 800 nm. Figure 2 details the optical components within the InSpectra tissue spectrometer. To exclude ambient light from the detected signal, four light emitting diodes (LEDs) were simultaneously pulsed on 760 Hz for 10 μs at a drive current near 2 A. The photomultiplier tube (PMT) detection signal passed through an analog filter circuit that isolated and integrated the 760 Hz LED pulse signal. At 48 Hz, the integrated analog signal was converted to a digital signal representing ambient compensated light intensity. To minimize instrument drift, such as time- and temperature-dependent changes in LED output, PMT sensitivity, and dark signal offset, the digital count signal was subdivided into three controlled time intervals that were sequentially and continuously repeated.

Figure 2

Optical schematic for the InSpectra tissue spectrometer. (a) Detection side optical components: L1 and L2 are aspheric collimating lenses, L3 is a plano convex condensing lens, SV is a stepper motor with a shaft mounted shutter vane, M1 is a 99% reflecting mirror, M2 is a combining mirror which transmits 99% of L1 light and reflects 1% of M1 light, M3, M4, M5, and M6 are, respectively, 800, 760, 720, and 680 nm dichroic mirrors, F1, F2, F3, and F4 are 10 nm FWHM interference filters for respective 800, 760, 720, and 680 nm center wavelengths, F5 is a 600 nm shortpass filter, PMT is photomultiplier tube, and PD is photodiode. (b) Removable probe optical components: SF is a sample light ferrule encompassing the receive optical fiber, RF is a 3-m-long 400-μm-diameter glass optical receive fiber, FF is a feedback light ferrule encompassing a 200-μm-diameter glass optical fiber, AT is a mixture of TiO ₂ and epoxy which attenuates and mixes the LED light, LED1, LED2, LED3, and LED4 are respective 680, 720, 760, and 800 nm light emitting diodes coupled to 400 μm glass optical fibers, MX is a 30-cm-long 1500 μm plastic optical fiber which couples and mixes light from the four fiber coupled LEDs, J1 is an operator removable optical/electrical connector, J2 is an optical fiber splice housing, and J3 is the probe distal tip which mounts to tissue with a polyethylene foam adhesive backed patient interface (not shown).

At an update rate of 3.5 s, the drift compensated sample light signal (I _DC ) was obtained as follows:

All four send fibers were coupled to a 1000 μm, 30 cm length plastic optical fiber [MX of Fig. 2(b)] to mix the discrete light wavelengths prior to being launched into the measurement sample. Probe spacings of 12, 15, 20, and 25 mm were used. Several InSpectra spectrometer design methods^{46 47 48 49 50} were used to meet accuracy and usage requirements. The dual beam signal stabilization design and method⁴⁶ was of primary importance since each LED and PMT can have unique time and temperature drift characteristics. Providing a continuous drift compensation signal measured with the same light source and detectors used for StO ₂ measurements limits StO ₂ drift, measured 24 h after an initial reference intensity measurement, to ±2 StO ₂ units. Mixing the discrete light sources into one optical beam with the plastic mixer fiber⁴⁷ prior to illuminating tissue removes any path length bias from having the discrete sources separated and located directly on tissue. This feature is important because when probe spacing is narrowed, say from 25 to 3 mm, a 0.4 mm center-to-center separation of optical fibers becomes a larger percentage of the probe spacing, which correlates directly to optical path length. With random orientation of the source fiber directly on the measurement sample, the blood phantom as detailed in Sec. 2.3 and Fig. 3(configuration A), observed probe-to-probe variability was near 10 StO ₂ units with a 25 mm probe and 30 StO ₂ units with a 3 mm probe. With homogenous mixing, probe-to-probe variability for all spacings of 3 mm or greater was found to be less than 3 StO ₂ units. Wavelengths of homogeneous emitted light are more evenly absorbed by local surface structures, such as moles or hairs on skin, resulting in less StO ₂ change as the probe is moved along the tissue. Other design aspects either relate to the StO ₂ algorithm method⁴⁸ (fully described in Sec. 2.2) or pertain to instrument use. The InSpectra automatically senses when a probe is correctly placed on the foam calibrator.⁴⁹ A second derivative method for sensing total amount of hemoglobin⁵⁰ is used to determine when a probe is placed on blood containing tissue. The total hemoglobin amount measurement and ambient light signal via a photodiode [PD of Fig. 2(a)] prevent false StO ₂ readings when the probe is not correctly placed on tissue.

Figure 3

Schematic for the isolated blood circuit used to calibrate and test StO ₂ measurement equipment. Configuration A represents the dual layer flow cell used to generate a calibration curve relating scaled 2D₇₂₀ to blood hemoglobin oxygen saturation, test StO ₂ reproducibility among multiple instruments, and test StO ₂ sensitivity to variable hemoglobin concentration and blood layer thickness. Configuration B represents a hollow flow cell used to contain a mixture of blood and Intralipid for testing StO ₂ sensitivity to probe spacing, scattering changes, and water concentration. In configuration C, isolated blood perfused organs with inhibited oxygen consumption were used to characterize StO ₂ performance in different tissues.

2.2.

StO₂ Algorithm

Tissue attenuation (A) measurements were calculated as log (reference intensity/sample intensity) for each measured wavelength. At a fixed wavelength gap interval, the second derivative of attenuation (2D) was obtained at each wavelength (λ) nm using an algebraic simplification of the difference between two first derivative attenuation (D) measurements calculated at a similar gap interval:

A wavelength gap of 40 nm was used to calculate the second derivative attenuation at two wavelengths, 720 and 760 nm. These two second derivative attenuation signals are related to the four measured attenuation wavelengths as follows:

For each tissue spectrum measurement a scaled 2D₇₂₀ value was used to predict tissue %StO ₂ from a predetermined empirical calibration relationship:

An in vitro model of human blood flowing over a fixed scattering layer was used to relate scaled 2D₇₂₀ spectrum measurements to co-oximeter %SO ₂ in order to create the empirical calibration curve (further detail provided in Sec. 2.3).

2.3.

%StO₂ Equipment Calibration and Reproducibility

Scaled 2D₇₂₀ from Eq. (8) empirically correlated to blood hemoglobin oxygen saturation (%SO ₂), measured with an IL 482 Co-Oximeter (Instrumentation Laboratory of Lexington, MA), using an isolated blood circuit coupled to a custom dual layer flow cell apparatus [Fig. 3(configuration A)].

The blood circuit (Fig. 3) consisted of a Minimax™ Hardshell Reservoir 1316 and Minimax™ Fiber Oxygenator 3381 (Medtronic, Minneapolis, MN) connected to a peristaltic pump and controlled temperature water bath to maintain a normothermic, 37.0±0.5 °C, blood flow of 300 mL/min. Blood pH was maintained to 7.40±0.05 with 8 wt % sodium bicarbonate as needed. A custom flow cell, to mimic blood-perfused tissue, interfaced the optical probes between the $\frac{1}{4}$ inch tubing which carried blood between the outlet port of the membrane oxygenator and the return port of the reservoir. A stopcock valve fitting placed between the flow cell outlet and reservoir return allowed syringe sampling of blood for co-oximeter %SO ₂ measurements.

The dual layer flow cell apparatus distributed the blood within the optical path of the measurement probes and provided a variable thickness of blood flowing over a block of LD45 Plastazote foam of sufficient volume to contain all emitted light paths. A top plastic film layer, 0.05 mm Mylar®D (DuPont Teijin Films, Hopewell, VA), separated flowing blood from contacting the measurement probes while a similar bottom plastic film separated the blood from the foam scattering material. Top and bottom plastic plates, housing the two plastic film layers and forming a sealed piston-like pressure vessel, were moved relative to each other in order to manipulate blood thickness and achieve signal attenuation close to blood perfused arm tissue measured with a 25 mm probe. In the absence of the flowing blood, the flow cell apparatus provides signal attenuation close to adipose tissue.

Human blood, 8.5 g/dL Hbt, flowing through the isolated blood circuit was equilibrated to variable concentrations of O ₂ entering the membrane oxygenator. At 23 separate blood equilibrium conditions, spanning near 0–99 %SO ₂ with increments near 5 SO ₂ units, scaled 2D₇₂₀ [Eq. (8)] measurements were paired with co-oximeter %SO ₂ in order to model a calibration curve to an average spectral relationship. The INTERP1 function of Matlab® (The Mathworks Inc., Natick, MA) was used to fit and interpolate the calibration data (not shown) to provide a lookup table matching scaled 2D₇₂₀ to %StO ₂ for all StO ₂ values ranging from 0 to 99% with an increment of 1%.

A single calibration curve lookup table was developed and was common to all probe spacings used, from 8 to 25 mm. To account for spectral bandwidth effects, separate calibration tables were developed for both the full spectrum (Biospectrometer-NB) and discrete wavelength (InSpectra) spectrometer designs. The calibration curves were developed using first run prototypes of the InSpectra and Biospectrometer designs. Spectrometers and probes providing the results of this study were not used to develop the calibration curves.

Four InSpectra spectrometers with 12, 15, 20, and 25 mm probes were tested using 10 U/mL heparinized human blood, 9.5 and 7.0 g/dL Hbt, within the previously described isolated blood circuit apparatus [Fig. 3(configuration A)]. Gas cylinders having 5% CO ₂, 0–21% O ₂, and the balance N ₂ were used to equilibrate blood hemoglobin oxygen saturation to 12 randomly selected target levels at both total hemoglobin concentrations. At each test condition, four paired readings of co-oximeter %SO ₂ and spectrometer %StO ₂ were obtained. This human blood study was conducted by the Mayo Clinic Center for Applied Vascular Biology and Intervention at Rochester, MN within the guidance of Good Laboratory Practice for NonClinical Laboratory Studies, 21 CFR Part 58.

2.4.

%StO₂ Performance in Simulated Tissue

Since hemoglobin and myoglobin have similar absorbance spectra, no attempt was made to distinguish between myoglobin and hemoglobin contributions to the in vitro and modeled data. The influence of total hemoglobin concentration, probe spacing with a variable optical path length, tissue optical scattering properties, and background water absorption were investigated using four simulated tissue models.

A pure hemoglobin theoretical model was used to test the inherent robustness of the algorithm method to total hemoglobin concentration and optical path length in a nonscattering absorbing medium. The model utilized the Lambert–Beer equation, which relates optical attenuation at a specific wavelength (A_λ) to path length distance (I), an absorbers concentration (C), and the absorber’s absorption coefficient (μ_aλ):

A second in vitro blood model was performed to determine the effect of variable total hemoglobin concentration and blood volume (thickness) on StO ₂ accuracy within a scattering environment. A 25 mm probe coupled to an InSpectra spectrometer was attached to the dual layer flow cell and an isolated blood circuit apparatus previously described [Fig. 3(configuration A)]. StO ₂ error relative to paired IL 482 Co-Oximeter %SO ₂ measurements were recorded for bovine blood layer thicknesses of 0.8, 1.4, and 2.0 mm at total hemoglobin concentrations of 12 and 6 g/dL. Additionally, 720 and 760 nm second derivative attenuation and attenuation at 680, 720, 760, and 800 nm were recorded for each thickness and Hbt level.

A third in vitro blood model was used to determine how changes in the background scattering attenuation and water concentration might affect StO ₂ accuracy for different probe spacings. The model consisted of isolated bovine blood, having intact red blood cells, diluted to 0.1 mM Hbt concentration with isotonic saline and 0, 0.2, 0.4, 0.6, and 0.8 wt % concentrations of Intralipid (Fresenius Kabi Clayton L.P., Clayton, NC). Four InSpectra spectrometers attached to 12, 15, 20, and 25 mm probes were used to measure changes in StO ₂ as the Intralipid solution was incrementally added to the blood. The hollow flow cell of the previously described isolated blood circuit [Fig. 3(configuration B)] was used. Additionally, experiments involved 0.1 mM bovine hemoglobin, 0.4 wt % Intralipid, isotonic saline, and variable amounts of water and deuterium oxide. Deuterium oxide (D ₂ O) has previously been used to adjust water concentration of tissue phantoms.⁵² Four spectrometers having 12, 15, 20, and 25 mm probes were used to measure StO ₂ accuracy at high, low, and middle range %SO ₂ targets for the Intralipid/blood solutions with 70% and near 100% background water concentration. Prior to Intralipid measurements, 800 mL of bovine blood equivalent to 5 g/dL Hbt was pumped through the isolated blood circuit apparatus and the inlet gas pO ₂ was adjusted until a target %SO ₂ equilibrium was obtained and measured with an IL 682 Co-Oximeter. With the inlet gas pO ₂ held constant, a portion of the blood was removed and replaced with calculated amounts of 20 wt % Intralipid, deionized water, deuterium oxide and sodium chloride. The molecular weight of hemoglobin was assumed to be 64 000 for all dilution calculations. The pH of the Intralipid/blood mixture was then adjusted to within 0.02 units of the starting bovine blood pH so the %SO ₂ of the Intralipid/blood mixture could be assumed equivalent to the previously measured whole blood %SO ₂ value.

A fourth theoretical model was used to determine how changes in background water concentration might influence StO ₂ accuracy for possible combinations of water and total hemoglobin concentration representative of different in vivo tissues. A single layer infinite slab diffusion theory equation⁵³ was used to create computer simulated tissue attenuation (A) spectra at variable inputs of a tissue absorption coefficient (μ_a), a scattering coefficient (μ_s ^′) and a probe spacing (ρ). This model was previously used to evaluate NIRS algorithm performance^{41 51} and has the form:

For this single layer tissue model, the tissue absorbance coefficient (μ_a) was estimated from absorbers thought to have the most significant spectral contribution for the %StO ₂ algorithm wavelength region, 680–800 nm. Within this region, water has a nonlinear spectral contribution that is amplified due to its high concentration in tissue, 70 wt %⁵² or 43 M considering lean tissue density of 1.1 kg/L.⁵⁴ Although fat has a lipid specific absorption peak near 930 nm, for this study it is assumed to have an effect similar to water, considering adipose tissue is 20% water.⁵²

The upper range of Hbt values modeled, 0.10–0.20 mM, is consistent with lean resting gastrocnemius muscle measured with a NIRS phase resolved spectrometer.⁵⁵ The optical scattering profile examined (Fig. 4) is for 1 wt % Intralipid.⁵⁶ For all modeled variables, μ_s ^′ was at least 20 times μ_a and in accordance with one validity constraint for the diffusion model, μ_s ^′≫μ_a. A second model assumption, probe spacing much greater than 1/μ_s ^′, was applied to at least one order of magnitude difference.

Figure 4

Model inputs for simulating StO ₂ repeatability in tissue using a single layer diffusion theory equation [Eq. (10)]. The absorption coefficient curves are for 0.1 mM Hbt with 70% water and 0–95% hemoglobin oxygen saturation. The scattering coefficient curve is estimated for 1 wt % Intralipid.⁵⁶

The absorption coefficients of HbO ₂, HHb,⁵¹ and water⁵⁷ were linearly combined to simulate the absorption coefficient of muscle. The water concentrations modeled were 0, 25%, 50%, 75%, and 100%. The contribution of HbO ₂ and HHb was weighted according to four assumed Hbt levels—0.05, 0.10, 0.15, and 0.20 mM—and 20 input %SO ₂ values—0–95% incremented at 5% (Fig. 4).

A 25 mm probe spacing, 0.15 mM Hbt, and 70% water calibration relationship was developed between output scaled 2D₇₂₀ [Eq. (8)] and input %SO ₂. When the modeled water concentration changed from 70% to 100%, the calibration data provided a best fit to the high StO ₂ 25 mm 0.1 mM Intralipid/blood empirical data obtained at near 100% water concentration. With the calibration held constant, the water and hemoglobin concentration was then varied to determine predicted %SO ₂ values relative to the known input %SO ₂ value. The difference between the predicted %SO ₂ output and %SO ₂ input represented the predicted %StO ₂ measurement error at concentration combinations of Hbt and water.

2.5.

Dyshemoglobin Effects

An InSpectra spectrometer coupled to a 25 mm probe was connected to the isolated blood circuit apparatus [Fig. 3(configuration A)]. Bovine blood, 12±1 g/dL Hbt and 10 U/mL heparin, was primed into the circuit. In multiple 60 mL syringes, 14 mL of blood was mixed with 46 mL of carbon monoxide for 30 min to produce near 100% carboxyhemoglobin (COHb). During the 30 min period, the syringe gas space was expelled and replaced two to three times. Capped syringes were horizontally positioned to provide maximum contact area between the blood and gas layers. Five minutes after adding a 14 mL COHb sample to the blood, five paired readings of spectrometer %StO ₂ and co-oximeter %SO ₂ and %COHb were recorded. Methemoglobin was less than 1% for all measurements. After a low %SO ₂ equilibrium experiment, the blood was purged and the circuit was primed with fresh blood in order to complete a similar high %SO ₂ experiment.

Likewise, 0.5 mL doses of isotonic sodium nitrite solution were added stepwise to fresh blood, allowing %StO ₂ error measurements at variable levels of methemoglobin (MetHb). COHb remained less than one percent for all methemoglobin equilibrium conditions. An IL 682 Co-Oximeter was used to obtain all %SO ₂, %COHb, and %MetHb measurements.

2.6.

Isolated Blood-Perfused Canine Hind Limb and Porcine Organs

Isolated canine hind limb experiments were completed (n=2) at Utah Biomedical Test Laboratory (Salt Lake City, UT) in accordance with Good Laboratory Practice for NonClinical Studies, 21 CFR Part 58. The Guide for the Care and Use of Laboratory Animals: DHHS, PHS, Revised 1985, was used for animal husbandry before and during the testing period.

The canine hind limb was surgically removed and kept viable by connection of femoral vein and artery cannulae to a normothermic (37±1 °C) isolated blood circuit with a blood pump and membrane oxygenator as described⁵⁸ [also reference Fig. 3(configuration C)]. Blood pH was maintained between 7.1 and 7.45 with additions of 8 wt % sodium bicarbonate.

Isolated porcine heart and kidney protocols, n=6 and n=2, respectively, were approved by the University of Minnesota Animal Use Committee, in accordance with established guidelines for the treatment of laboratory animals.

Isolated potassium chloride arrested porcine hearts had a blood inlet cannula inserted in the brachiocephalic artery, while venous blood was allowed to outflow from a drain tube inserted through the base of the right atrium. A triple lumen pulmonary artery catheter was placed in the coronary sinus to facilitate venous blood sampling. The heart was clamped over a hollow reservoir draining back to a Minimax™ Hardshell Reservoir 1316 and Fiber Oxygenator 3381 [Fig. 3(configuration C)]. A centrifugal pump (Medtronic BP-50 Biopump®, Minneapolis, MN) maintained an average perfusion pressure of 80 mm Hg. Potassium cyanide, 15 mg/mL, in 10 mL doses was infused as needed to inhibit tissue oxygen consumption in order to equilibrate the arterial and venous hemoglobin oxygen saturations to a difference of 5 %SO ₂ units or less. The oxygen ratio within the gas flow to the membrane oxygenator was varied to adjust arterial hemoglobin oxygen saturation within a range of 5%–99%. An IL 482 Co-Oximeter provided organ inflow (SaO ₂) and outflow (SvO ₂) blood hemoglobin oxygen saturation measurements that were paired with 8 mm probe %StO ₂ measurements at middle, upper, and lower left ventricle locations. If the paired SaO ₂ and SvO ₂ differences were greater than five %SO ₂ units, the readings were not used, and additional cyanide was added through the arterial tubing sample port.

In two isolated porcine heart experiments, the kidney was harvested and similarly perfused using a bifurcated arterial line connected to two separate cannulae. Venous outflow from the kidney was gravity fed to a blood reservoir. The same 8 mm probe and full spectrum spectrometer (Biospectrometer-NB) was used to obtain the heart and kidney %StO ₂ measurements.

2.7.

Human Volunteers with Induced Limb Ischemia

This study received approval from the United Hospital (St. Paul, MN) Institutional Review Board and was conducted and monitored in accordance with recognized Good Clinical Practices, 21 CFR parts 50, 54, 56, and 812.

Twenty-six volunteers, aged 27–56 years (mean=37.5 and SD=±7.3 years), consented to participate. Individuals having a body mass index⁵⁹ of ⩾30 were excluded from the study. Single unpaired StO ₂ measurements were sequentially obtained on the subject’s thenar eminence, a 12 mm probe, dorsal compartment of the forearm, a 20 mm probe, and the tibialis anterior muscle, 25 mm probe, at baseline, ischemia, and recovery. Baseline measurements were obtained with the subject in a supine position after 5 min of resting stabilization. Ischemia was induced via inflation of a pneumatic tourniquet 50 mm Hg above systolic blood pressure. Subjects having systolic pressures greater than 140 mm Hg were excluded from participation. Prior to ischemia, both the test leg and arm were elevated for 2 min to facilitate venous blood drainage. The test arm was also exercised using a graspable squeeze ball during the first 30 s of cuff ischemia. Upper extremity measurements began 2.5 min after induction of the cuff pressure. Lower extremity ischemia measurements began 5 min after cuff inflation. Inflation time did not exceed 15 min for either extremity. The mean StO ₂ value and its standard deviation were computed for each probe depth and tissue bed at each of the three blood flow state conditions.

2.8.

Statistical Methods

Scatter plots and linear regression models were used to describe correlations between %StO ₂ and reference %SO ₂ measurements. The squared Pearson correlation coefficient⁶⁰ was used to describe the degree of regression model fit. Squaring the correlation coefficient and multiplying it by 100 describes the percent variability in observed StO ₂, explained by changes in the reference %SO ₂ measurement. Statistical significance was determined at p value<0.05. StO ₂ error is expressed as the difference between measured or predicted %StO ₂ and the reference %SO ₂ method. Variability in human tissue measurements were expressed using mean %StO ₂ with one standard deviation of the measurement population.

For the human isolated blood study, a repeated measurement analysis of variance test method⁶¹ was used to determine significant mean effects. Repeating %StO ₂ measurements, four spectrometers and four replicate measurements per spectrometer, at each %SO ₂ target, provided better estimates of how total hemoglobin concentration and probe spacing affected StO ₂ measurements. Root mean square error for all paired spectrometer StO ₂ and co-oximeter SO ₂ measurements was calculated as follows:

3.1.

StO₂ Measurement Equipment

The Biospectrometer-NB spectrometer design acquired broadband attenuation spectra and allowed optimization and minimization of the wavelengths needed for a StO ₂ measurement (detailed in Sec. 3.2). With an algorithm method defined [Eq. (8)], the InSpectra spectrometer was developed specifically to improve ease of use and reduce operator induced measurement errors.

3.2.

%StO₂ Algorithm

A plot of published pure component HbO ₂ and HHb absorption spectra⁵¹ and second derivative absorption transformations using both narrow (1 nm) and wide (40 nm) wavelength gaps (Fig. 5) reveals the reasons for choosing an algorithm incorporating a 40 nm gap second derivative transformation. The absorbance profile of HbO ₂ [Fig. 5(a)] is nonlinear within the 680–760 nm wavelength region. With a wide 40 nm gap, the HbO ₂ second derivative at 720 nm is approaching maximal amplitude [Fig. 5(b)], whereas the corresponding 1 nm gap amplitude [Fig. 5(b)] is effectively zero. Therefore the larger 40 nm gap HbO ₂ spectral contribution allows a more precise estimate of %SO ₂. The 760 nm HHb specific 40 nm gap second derivative amplitude [Fig. 5(d)], being larger than the corresponding 1 nm gap amplitude [Fig. 5(c)], also provides noise tolerance. The 1 nm gap spectra-required curve fitting of the absorption spectra to provide visually presentable second derivative spectra, while the 40 nm gap second derivative spectra required no absorption smoothing. The problems associated with noise, especially when measuring HbO ₂ with 1 nm gap amplitudes, has been discussed elsewhere.⁶² Similar spectral results were found (Fig. 6) when comparing blood attenuation spectra near 0 and 100 %SO ₂ measured with the previously described isolated blood circuit apparatus [Fig. 3(configuration A)].

Figure 5

Transmitted pure HbO ₂ and HHb spectra from published source.⁵¹ (a) Absorption coefficient, (b) comparison of narrow and wide gap second derivative absorption coefficient for HbO ₂, (c) 1 nm gap second derivative absorption coefficient, (d) 40 nm gap second derivative absorption coefficient, (e) 1 nm gap second derivative absorption coefficient scaled to 760 nm, and (f) 40 nm gap second derivative absorption coefficient scaled to 760 nm.

Figure 6

Re-emitted HbO ₂ and HHb bovine blood spectra acquired at 15 nm FWHM resolution with a Biospectrometer-NB spectrometer using the dual layer flow cell apparatus [Fig. 3(configuration A)]. (a) Relative attenuation, (b) comparison of narrow and wide gap second derivative attenuation for HbO ₂, (c) 1 nm gap second derivative attenuation, (d) 40 nm gap second derivative attenuation, (e) 1 nm gap second derivative attenuation scaled to 760 nm, and (f) 40 nm gap second derivative attenuation scaled to 760 nm.

Second derivative blood attenuation spectra is sensitive not only to hemoglobin oxygen saturation but also total hemoglobin concentration (Hbt) and probe spacing (Fig. 7). Increasing the total hemoglobin [Figs. 7(a) and 7(b)] or the probe spacing [Figs. 7(a) and 7(c)] results in an increased signal amplitude. Therefore, a ratio of two second derivative attenuation measurements was used in order to provide a %SO ₂ specific measurement that would be inherently robust to optical path length and Hbt (Table I).

Figure 7

Re-emitted bovine blood second derivative attenuation spectra for multiple %SO ₂ ranging from approximately 5%–98% obtained with a Biospectrometer-NB spectrometer and the hollow flow cell apparatus [Fig. 3 (configuration B)]: (a) 7 g/dL Hbt with a 5 mm probe, (b) 13 g/dL Hbt with a 5 mm probe and (c) 5 g/dL with a 10 mm probe. The direction of high and low SO ₂ as shown in (b) applies to all panels.

Table 1

The 760 nm second derivative attenuation is ideally suited for this method because it exhibits maximal amplitude change with variable %SO ₂, at fixed Hbt and path length, and has no spectral contribution from HbO ₂. With a prerequisite 760 nm second derivative wavelength, the 720 nm second derivative wavelength was chosen as the second wavelength for its relatively close proximity to 760 nm and its sensitivity to both HbO ₂ and HHb. Since the 40 nm gap interval equaled the distance between the numerator and denominator second derivative wavelengths, a reduction from six to four measurement wavelengths occurred [see Eqs. (6) and (7)], simplifying spectrometer design. Although a similar reduction of wavelengths would result from using an 800 nm second derivative measurement relative to 760 nm, the 720 nm scaled point was chosen because of its larger and more varied second derivative amplitude with respect to %SO ₂ [Figs. 5(f) and 6(f)].

3.3.

%StO₂ Equipment Calibration and Reproducibility

Several approaches were possible for developing a calibration curve relating scaled 2D₇₂₀ measurements [Eq. (8)] to hemoglobin oxygen saturation in tissue. Currently no measurement standard exists for measuring tissue hemoglobin oxygen saturation. An approach using an in vivo calibration would involve assumptions regarding how the NIRS signal best represented a balance between invasively measured arterial and venous %SO ₂ near the measured tissue site. Therefore, an in vitro calibration method was adopted because it provided a controlled test environment for characterizing repeatability of StO ₂ measurements among the multiple spectrometer designs manufactured.

One in vitro method reviewed involved immersing a 5 mm probe in diluted whole blood, 5–12 g/dL Hbt, having sufficient volume to contain nearly all optical path lengths. This method was used to demonstrate the variability of second derivative amplitudes versus %SO ₂, Hbt and probe spacing (Fig. 7), but was not used to develop the depicted calibration curves [Fig. 8(a)] because the test environment provided too much signal attenuation with probes greater than 10 mm spacing. Although dilution of total hemoglobin significantly below 5 g/dL would reduce absorption and allow long path length measurements, the blood scattering properties would be significantly reduced as the red blood cell count decreased.

Figure 8

Spectrometer stored %StO ₂ calibration curves and resulting correlation to co-oximeter human blood %SO ₂ using the dual layer flow cell apparatus [Fig. 3(configuration A)]. (a) Empirically calibrated relationship between %StO ₂ and measured second derivative attenuation ratio for the two-spectrometer designs utilized in this study. The average spectral relationship was obtained at 8.5 g/dL Hbt with four spectrometers. Data used to generate the calibration curves are not shown. (b) Correlation of four previously calibrated InSpectra spectrometers to co-oximeter %SO ₂ using 12, 15, 20, and 25 mm probe spacings. The 384 measurements were equally divided among human blood at 9.5±0.5 and 7.0±0.5 g/dL Hbt. All measurement points were obtained using a common calibration curve. (c) Cross-instrument correlation of one previously calibrated Biospectrometer-NB spectrometer to co-oximeter %SO ₂ using one 20 mm probe to obtain 47 bovine blood measurements at 10 g/dL Hbt.

Intralipid emulsion solution has been used to provide a constant scattering environment to dilute Hbt into assumed tissue levels, ⩽1 g/dL.⁶³ However, blood co-oximeters, the current standard for blood %SO ₂, have limited accuracy below 5 g/dL Hbt, even before considering the possible interference from Intralipid. Although it is possible to derive a calculated saturation (%SO ₂ c) from blood pH, pO ₂, and temperature without Intralipid interference, some discrepancies between %SO ₂ measured directly with an IL482 Co-Oximeter and %SO ₂ c derived from an IL blood gas analyzer have been noted by the authors. Usually %SO ₂ c is generally 10 units higher across the full range with whole bovine blood. Because of this discrepancy, it was uncertain whether Hill type equations⁶⁴ for predicting %SO ₂ from pH, pO ₂ and temperature could be accurately extrapolated to all possible %SO ₂ values at hemoglobin concentrations below the normal physiologic range of blood.

Therefore a dual layer flow cell and isolated blood circuit was used [Fig. 3(configuration A) and detailed in Sec. 2.3] for controlling hemoglobin oxygen saturation of an approximate 1 mm thick layer of blood flowing above a constant scattering layer of polyethylene foam. This enabled optimal co-oximetry measurements, allowing blood full range %SO ₂ to be accurately defined and the possible confounding effects of carboxyhemoglobin and methemoglobin to be investigated. The two-layer model additionally allowed all probe spacings of this study to be correlated to co-oximeter %SO ₂.

An alternative calibration could simply use mathematical modeling such as a diffusion theory equation as described in Sec. 2.4. In such a case the absorption coefficient data should be empirically measured with the same optical equipment used for measuring %StO ₂ to account for how center wavelength and bandwidth resolution influences a calibration curve.

Results of the co-oximeter human blood correlation study indicated a strong linear correlation of %StO ₂ to %SO ₂ for both the four wavelength (InSpectra) and full spectrum (Biospectrometer-NB) spectrometers [Figs. 8(b) and 8(c)]. Total hemoglobin concentration, 7.0 and 9.5 g/dL, was not a significant source of measurement variability (p=0.993). Multiple measurements using four InSpectra systems revealed that the root mean square errors of individual measurements were 1.6 StO ₂ units within a 70%–99% range and 2.9 StO ₂ units for a 0–70% range. The calibration curves of Fig. 8(a) are nearly vertical at low StO ₂ and horizontal at high StO ₂. Thus at low StO ₂ a small change in scaled 2D₇₂₀ causes larger StO ₂ change when compared to high StO ₂. This spectral relationship implies that StO ₂ error at low StO ₂ might be drastically worse than at high StO ₂. However, Fig. 7 shows that second derivative amplitude approaches zero at high StO ₂ and is significantly large at low StO ₂. Therefore scaled 2D₇₂₀ is more precise at low StO ₂ since measurement errors, due to noise for example, are a smaller proportion of the 720 and 760 nm second derivative amplitudes. The more precisely measured 2D₇₂₀ spectral measurement at low StO ₂ offsets its higher spectral sensitivity to StO ₂ thus providing a 0–70% root-mean-square (rms) error only about two times higher than the 70%–100% error.

Significant error sources affecting StO ₂ reproducibility for the human blood correlation study include center wavelength and bandwidth variation since a common calibration curve was used for all devices, nonlinearity of light detection system relative to optical density, uncorrected drift in system light detection and illumination efficiency, as well as shot noise inherent in the low light level detection circuitry.

3.4.

StO₂ Performance in Simulated Tissue

Since no in vivo %StO ₂ standard exists, in vitro tests combined with theoretical mathematical models provided a means to examine the sensitivity of the StO ₂ measurement method to changes in total hemoglobin concentration, optical path length, optical scattering, and water concentration.

The isolated blood circuit in vitro model [Fig. 3(configuration A)] consisted of two primary absorbers, HbO ₂ and HHb, combined with a polyethylene foam background scattering environment. Despite changes in the relative attenuation [Fig. 9(a)] and second derivatives [Fig. 9(b)], the measured saturation remained independent [Fig. 9(c)] of blood path length or thickness and hemoglobin concentration. These results suggest that, in a scattering environment, the 2D₇₂₀ measurement [Eq. (8)] is inherently robust to changes in Hbt and path length as found in the nonscattering Beer–Lambert model results (Table I).

Figure 9

Effect of variable blood layer thickness and hemoglobin concentration on InSpectra %StO ₂ accuracy at 55%StO ₂ using the dual layer flow cell apparatus [Fig. 3(configuration A)]. (a) Relative attenuation measurements at each wavelength for variable blood thickness and Hbt, (b) 12 and 6 g/dL Hbt vs nonscaled second derivative attenuation values at multiple blood thicknesses, and (c) %StO ₂ error, mean with ±1 standard deviation limits, for variable blood layer thicknesses and Hbt. For panel (c) the 12 and 6 g/dL Hbt summary data were slightly shifted along the x axis to provide better visualization of results.

Intralipid results of Fig. 10 show that despite significant changes in Intralipid concentration, 0–0.8 wt %, a less than ten StO ₂ unit change was observed for a wide range of probe spacings, 12–25 mm. Light scattering (μ_s ^′) of Intralipid solution is linearly related to Intralipid concentration and is estimated to be 8 cm⁻¹ at 800 nm for the maximum 0.8 wt % Intralipid concentration tested.⁵⁶ Red blood cell concentration (hematocrit) also linearly relates to scattering and provides a μ_s ^′ of about 2 cm⁻¹ at 800 nm for 0.1 mM total hemoglobin, near 2% hematocrit.⁶⁵ At 0.8 wt % Intralipid and 0.1 mM Hbt the effective scattering of the Intralipid/blood solution (μ_s ^′) is estimated as 10 cm⁻¹ at 800 nm assuming that the overall scattering coefficient is a linear sum of its individual components. The range of optical scattering tested, μ_s ^′ of 10 cm⁻¹ or less, agrees with in vivo estimates for human tissues inclusive of muscle; about 8–10 cm⁻¹ at 800 nm measured with TRS on human legs and arms²⁵ and about 3–5 cm⁻¹ at 800 nm on human forearm measured with SRS combined with PMS.³⁸ For the 0.15 mM Hbt 0.4 wt % Intralipid test condition, it was observed that the 20 and 25 mm probe light intensity signals were approaching the threshold limit at which the monitor stops displaying StO ₂. Loss of signal prevented some 20 and 25 mm probe measurements [see Fig. 10(b)]. Errors in dark signal measurement result in a nonlinear light detection signal, which can create greater StO ₂ measurement error near the low signal threshold limit and could explain the wider range of StO ₂ shown in Fig. 10(b). The experimental results demonstrate that the StO ₂ second derivative algorithm method is insensitive to large changes in optical scattering. Insensitivity of StO ₂ to optical path length can also be concluded since the 12, 15, 20, and 25 mm probes, each having significantly different optical path lengths, gave StO ₂ readings within a range of about five StO ₂ units for the Intralipid concentrations tested.

Figure 10

Effect of variable scatter concentration on InSpectra %StO ₂ measurements using the previously described hollow flow cell apparatus [Fig. 3(configuration B)]. Four spectrometer systems with 12, 15, 20, and 25 mm probe spacings were used. Bovine blood was diluted to (a) 0.1 mM Hbt or (b) 0.15 mM Hbt within a mixture of 70 vol % H ₂ O, 30 vol % D ₂ O, 0.9 wt % NaCl. A fixed volume of 20 wt % Intralipid was stepwise added to the mixture. Red blood cells remained intact during all measurement conditions. For the 0.15 mM Hbt condition, insufficient detected light signal prevented 25 mm measurements at 0.6 wt % Intralipid as well as 25 and 20 mm probe measurements at 0.8 wt % Intralipid. For all measurement conditions actual blood hemoglobin oxygen saturation is unknown.

Additional Intralipid/blood experiments were performed in order to independently validate StO ₂ accuracy in a homogenous blood scattering media different then the dual layer flow cell used to develop the StO ₂ calibration curve. For these 0.1 mM total hemoglobin experiments [Fig. 11(b)], 0.4 wt % Intralipid was chosen to provide higher detected light intensity for a 25 mm probe spacing in order to better characterize StO ₂ error due to background water absorption. Overall μ_s ^′ is estimated to be 6 cm⁻¹ at 800 nm. Whereas the blood layer of the dual layer flow cell contains a high level of water (near 99%), the segregated polyethylene foam layer has no water. Due to the nature of the calibration system, the precise amount of water included in the spectral measurements used to develop the initial StO ₂ calibration relationship is unknown. The results of Fig. 11(b) indicate that at near 100% water concentration, StO ₂ error due to background water absorption is greater at higher levels of StO ₂ and approaches −20 StO ₂ units near 100% hemoglobin oxygen saturation. At normal levels of muscle water concentration (near 70%) the maximum StO ₂ error is estimated to be −10 StO ₂ at 100% hemoglobin oxygen saturation.

Figure 11

Predicted full range StO ₂ measurement error (solid lines) for 0, 25%, 50%, 75%, and 100% water concentration at (a) 0.05, (b) 0.10, (c) 0.15, and (d) 0.20 mM Hbt using theoretical light diffusion equation and published absorption and scatter coefficients as described in Sec. 2.4. At 0.1 mM Hbt, StO ₂ measurement results for near 100% H ₂ O (× symbols) and 70% H ₂ O /30% D ₂ O (○ symbols) were obtained with diluted bovine blood in a background of 0.4 wt % Intralipid and 0.9 wt % NaCl flowing in the hollow flow cell apparatus [Fig. 3(configuration B)]. Four InSpectra spectrometers with 12, 15, 20, and 25 mm probe spacings were used to obtain the four plotted measurements common to the level of %SO ₂ tested. At 5% SO ₂ and near 100% SO ₂ actual StO ₂ is estimated from a standard human blood dissociation curve for the measured solution pH, solution temperature and oxygenator outlet gas pO ₂. For all other intermediate levels of %SO ₂, actual %SO ₂ is estimated from co-oximeter measurements obtained with 5 g/dL blood flowing in the isolated blood circuit prior to diluting the blood to a 0.1 mM Hbt and adding D ₂ O and Intralipid. All dilution liquids were adjusted to 0.9 wt % NaCl. The oxygenator gas outlet pO ₂ remained constant before and after hemoglobin dilution and the Intralipid/blood mixture pH was adjusted to match the blood pH corresponding to the co-oximeter measurement sample. StO ₂ error represents measured StO ₂ minus estimated %SO ₂.

Figure 11 shows that the amount of StO ₂ measurement error due to variable water concentration is dependent upon the amount of water relative to the amount of total hemoglobin. Highest StO ₂ measurement errors occur at 0.05 mM total hemoglobin when water concentration exceeds 70%. Lean muscle tissue has a hemoglobin concentration range of 0.1–0.2 mM and a water concentration near 70%. Tissue comprising of adipose and muscle can have a hemoglobin concentration approaching 0.05 mM with water concentration limited to 50% or less.^{55 66 67 68} StO ₂ inaccuracy due to normal levels of tissue water appears to be less than ten absolute SO ₂ units.

At tissue hemoglobin concentrations of 0.1 mM or less, edema might cause a significant negative bias to StO ₂ measurements provided the added water occurs within the scattering media, in and adjacent to cells. A free space water layer accumulation, devoid of cellular bodies which cause scattering, is not represented in the measurement model and is expected to have less influence on StO ₂ accuracy since multiple light scattering events within the water layer are required to amplify the spectral contribution of water.

3.5.

Dyshemoglobin Effects

A second derivative transformation of attenuation spectra has commonly been used to limit interference from chromophores having constant background absorption. Wavelength regions are selected where the chromophore of interest has nonlinear absorption and possible confounders have near linear or constant absorption. Within 680–800 nm, the absorption spectrum of carboxyhemoglobin is more linear and less steep than methemoglobin,⁶⁹ which explains why the %StO ₂ algorithm is less sensitive to carboxyhemoglobin interference compared to methemoglobin (Fig. 12).

Figure 12

Effect of dyshemoglobin concentration on InSpectra %StO ₂ accuracy with 12±1 g/dL Hbt bovine blood entering the dual layer flow cell apparatus [Fig. 3(configuration A)]. (a) Carboxyhemoglobin and (b) methemoglobin induced measurement errors with best-fit regression curves.

The normal reference range for blood methemoglobin and carboxyhemoglobin is less than 2% total hemoglobin.⁷⁰ Within this reference range, carboxyhemoglobin and methemoglobin induced %StO ₂ error was limited to near +2 and +4 StO ₂ units, respectively. For smokers, the carboxyhemoglobin reference range is 2%–20%,⁷⁰ which results in a possible +6 StO ₂ unit bias.

Cyanosis is a clinical symptom associated with blood methemoglobin levels exceeding 10%.⁷⁰ At high StO ₂, elevated methemoglobin produced a limited bias near +10 StO ₂ units, whereas at low StO ₂, the bias linearly increased as methemoglobin concentration increased. Methemoglobinemia could confound the ability to diagnose low hemoglobin oxygen saturation in tissue. However, we conclude that under the vast majority of conditions, the levels of these hemoglobin derivatives will not contribute significant errors to the measured oxygen saturation.

3.6.

Isolated Blood-Perfused Canine Hind Limb and Porcine Organs

Isolated blood perfused organs provided a living tissue model in which to access the adequacy of the in vitro calibration method and the repeatability of %StO ₂ measurements among tissues having different optical absorption and scattering characteristics.

The reflected intensity signals used to derive the %StO ₂ measurements are thought to emerge from arterioles, capillaries, and venules⁴⁴ and the average traversed depth of the detected photons is approximately one half the probe spacing distance.¹³ The second derivative attenuation signal measured at 720 and 760 nm increases in amplitude with an increase in Hbt (Fig. 7). Therefore, it is reasonable to assume that the %StO ₂ signal is weighted towards the micro vascular structures producing the largest second derivative attenuation, and thus represent the largest overall contribution to blood volume within the measured tissue space. It is generally assumed that tissue hemoglobin oxygen saturation will be nearer that of the venous system, rather than the arterial system. However, since the distribution of blood volume within arterioles, capillaries and venules can change with limb position (venous pooling) and vascular resistance, it is difficult to represent %StO ₂ as a constant weighted measurement of SaO ₂ and SvO ₂ for the tissue being measured. Therefore potassium cyanide was used to inhibit oxygen consumption of the isolated organs in order to minimize the %SO ₂ gradient within the tissue micro vasculature, and hence reduce any artifacts introduced to the test by having to estimate accurately the exact compartment(s) measured by NIRS.

With the SaO ₂ and SvO ₂ gradient held to within a range of 0–5 %SO ₂ at a constant %SaO ₂, it was assumed that a weighted average of paired SaO ₂ and SvO ₂ measurements would provide an estimate of the micro vascular blood %SO ₂ to an uncertainty of ±2.5 %SO ₂ units. The correlation plots of %StO ₂ versus the estimated microvascular %SO ₂, defined as (0.33SaO ₂+0.67SvO ₂)/2, indicate a linear correlation generally within 10 %SO ₂ units of the identity line (Fig. 13). The observed correlation was similar among the canine hind limb, porcine heart, and kidney tissue. In all tissues >93% of StO ₂ variability was caused by changes in microvascular %SO ₂. Further, the y intercept was between +9 and +14 StO ₂ units and the mean difference between regression line and identity was +3.3, +10.1, and +6.7 StO ₂ units for respective hind limb, heart, and kidney.

Figure 13

Correlation results for isolated blood perfused organs titrated with cyanide to maintain %SvO ₂ to within five units of the manipulated %SaO ₂ [Fig. 3(configuration C)]. In all experiments paired %StO ₂, %SaO ₂, and %SvO ₂ readings were collected. (a) Correlation results, n=34, from two isolated canine limb experiments using four Biospectrometer-NB spectrometers with 15 mm probes. (b) Correlation results, n=113, from six isolated porcine hearts measured with an 8 mm probe connected to a Biospectrometer-NB spectrometer. (c) For two isolated kidneys, 25 paired readings were obtained with the same spectrometer system used for the isolated hearts.

The calibration curve representing the full spectrum spectrometers used for the isolated animal tissue measurements [Fig. 8(c)] suggests that some of the positive y intercept might be attributable to the calibration curve stored within the devices. This could arise from spectral bandwidth variations between the devices used to create the calibration relationship and the devices used for this study. The depicted calibration relationship [Fig. 8(c)] is for a 20 mm probe and diffusion theory analysis results (not shown) indicate an upward pivoting of the calibration slope for a shallow probe spacing. This is particularly noticeable in the 8 mm probe isolated kidney data [Fig. 13(c)]. In addition to the errors we predict from diffusion theory analysis, nonhomogeneous perfusion and distribution of cyanide might result in microvascular regions where %SO ₂ is outside the assumed uncertainty limit. Lack of homogeneous cyanide distribution might create a lower than expected StO ₂ due to localized high levels of oxygen consumption resulting in lower venule %SO ₂. Further, regions of restricted flow might create longer delays between downward %SaO ₂ adjustments and microvascular %SO ₂ equilibration resulting in higher readings than anticipated. Nevertheless, despite these small errors, it is clear that an in vitro calibration based on our two-layer model can be applied successfully to measure %SO ₂ in a range of in vivo situations.

Myoglobin (Mb) is another error source, as we calibrated %SO ₂ solely to hemoglobin oxygen saturation. However, linear regression results [Figs. 13(b) and 13(c)] demonstrate similar correlations of StO ₂ to microvascular %SO ₂ for both myoglobin and nonmyoglobin tissues (heart and kidney, respectively) using the same StO ₂ measurement device. These results indicate that myoglobin does not significantly bias %StO ₂ measurements. Myoglobin and hemoglobin have similar absorption profiles within the wavelength region measured, 680–800 nm. Considering the broad resolution of the StO ₂ spectrometers utilized, 10–15 nm FWHM, and wide second derivative gap, it is apparent that the described StO ₂ method and equipment cannot resolve the minor wavelength shifts that have been reported between HHb and Mb tissue spectra.⁷¹ However, even if the spectrometer cannot distinguish between Mb and HHb, the fact remains that the in vivo measures were tested only against co-oximeter hemoglobin saturation values. Therefore, myoglobin saturation changes could be detected by the spectrometer; but be silent to the co-oximeter.

Proton nuclear magnetic resonance (NMR) spectroscopy has been used to measure exercising gastrocnemius myoglobin and hemoglobin NMR signals simultaneously with an NIRS relative hemoglobin measurement.⁴⁴ There is conflicting evidence as to whether Mb does change significantly during muscle exercise and as to whether the kinetics of NIRS hemoglobin signals may be more representative of Mb changes during cuff ischemia.⁷² However, a recent modeling investigation suggests that myoglobin effects should not be discounted regarding the interpretation of NIRS derived hemoglobin signals in muscle.⁷³

3.7.

Human Volunteers with Induced Limb Ischemia

A study on healthy human limbs provided a means to characterize changes in StO ₂ with induced ischemia. Since adipose tissue thickness can influence how the measurement signal is weighted toward muscle, a variety of probe spacings were used in an effort to provide the best balance between measurement depth and measurement precision. Excessively wide probe spacing on lean muscle would cause a low detected light intensity and reduce measurement precision. Conversely, excessively narrow probe spacing could result in the StO ₂ measurement signal being derived solely from cutaneous and adipose tissues. As a guide for selecting probe spacing for a given population of test subjects, probe spacing design dimensions of at least two to three times the estimated sum of cutaneous and adipose thickness were chosen.

Table II shows that the InSpectra spectrometer reliably detected decreases in tissue hemoglobin oxygen saturation induced by ischemia. The absolute change between baseline and ischemia, and between ischemia and recovery was found to be statistically significant for males and females for every probe and muscle combination (p values<0.01). The most likely explanation for the statistically significant differences in dorsal mean StO ₂ measurements between males and females are differences in the thickness of the adipose tissue.

Table 2