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1 Department of Medical Physics
and Bioengineering, Two near-infrared spectroscopy (NIRS) methods
are available for measuring changes (
cerebral blood volume; near-infrared spectroscopy
OVER THE LAST DECADE, near-infrared spectroscopy (NIRS)
has found widespread application for the monitoring of tissue
hemodynamics and oxygenation (7, 11, 24). Because the technique allows the continuous and noninvasive measurement of changes in concentration of oxyhemoglobin ( An important step in the development of NIRS occurred when methods for
the measurement of absolute hemodynamic parameters were devised. In
1988, Edwards et al. (6) described the measurement of absolute blood
flow by using a small but rapid change in concentration of
HbO2 as an intravascular
near-infrared dye. The technique was first applied to the measurement
of neonatal cerebral blood flow and, in a subsequent
validation, was shown to correlate with the xenon-clearance method
(20).
By 1990, a method for the measurement of absolute cerebral hemoglobin
concentration (CHC) had been described (25). This method employs a
small, but this time slow, change in
HbO2 concentration and, for this
reason, is usually referred to as the
O2 method for estimating cerebral
blood volume (CBV). To date, independent validation of this
method has not been performed, although the technique has been
widely used after its initial application in neonatal medicine (15, 22,
26).
It is, therefore, possible to attempt an internal validation by
comparing changes in CBV measured from the continuous
The aim of the study reported here was to confirm these observations in
a more controlled animal model, and to investigate, from a theoretical
basis, the possible reasons for any differences. This paper will
describe measurements in which a change in fraction of inspired
O2
(FIO2) was used to measure
CBV (by using the O2 method) and a
change in fraction of inspired CO2
(FICO2) was used to change CBV.
In a nonscattering medium, the attenuation
A of light traveling a distance
d is simply given by the equation
ABSTRACT
Top
Abstract
Introduction
Discussion
References
) in total cerebral hemoglobin
concentration (CHC): 1) a continuous measurement of the
changes in total hemoglobin concentration
([Hb]tot)
and 2) the difference between two absolute measurements of
CHC, each derived from a small, controlled change in inspired
O2 fraction. This paper
investigates the internal consistency of these two methods by using an
experimental and theoretical comparison. NIRS was used to measure
[Hb]tot in five newborn piglets before and after a change in arterial
PCO2. [Hb]tot
demonstrated a low coefficient of variation of 2.8 ± 2.8 (SD) % which allowed changes in
CO2-cerebral blood volume reactivity to be clearly discriminated. However, a high coefficient of
variation of 22.8 ± 3.5% on the CHC measurements
obscured any CO2 reactivity
changes. A theoretical analysis demonstrates the effects of optical
pathlength, background absorption, scatter, and blood vessel diameter
on both methods. For more accurate monitoring of CHC, individual
measurements of optical pathlength and more accurate pulse oximetry are required.
INTRODUCTION
Top
Abstract
Introduction
Discussion
References
HbO2),
deoxyhemoglobin (Hb), and changes in the redox state of cytochrome
oxidase, its simplest application is as a trend monitor of global
tissue oxygenation status. If the
HbO2 and Hb signals are
summed, this provides a derived measurement of changes in total
hemoglobin concentration
([Hb]tot), which
can be used to reflect changes in blood volume.
[Hb]tot measurement
against the change calculated from two separate absolute concentration
measurements (
CHC = CHC2 
CHC1). One study (2) attempted such a comparison by using
CO2-induced CBV changes in a small
number of newborn infants and reported a significant difference in the
values obtained by each method: 0.89 ml · 100 g
1 · kPa1
for the absolute
O2 method vs. 0.22 ml · 100 g1 · kPa1
for the continuous-change measurement
([Hb]tot).
NIRS MEASUREMENT THEORY
where
each absorber i has a concentration
Ci and a specific absorption
coefficient µia.
(1)
In a scattering medium, such as biological tissue, the scattering acts to increase the path traveled, and, as shown by the diffusion Eq. 1, attenuation is no longer a linear function of the µa. For small changes in absorption, however, changes in attenuation can be approximated by
|
(2) |
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(3) |
is the factor by which the optical pathlength has been increased
due to scattering [the so-called differential pathlength factor
(DPF) (4)]. Using Eq. 3, with
the known specific µa of
hemoglobin and the other chromophores in tissue (water, cytochrome), it
is possible to derive changes in the concentration of
HbO2 and Hb from the attenuation changes.
The O2 method of measuring total
CHC relies on using HbO2 as a NIR
"dye" which can be measured in both the peripheral and cerebral
circulation. A small, slow change in
FIO2 is induced, the effects
of which can be measured peripherally, by using a pulse oximeter, as
a change in arterial O2 saturation (SaO2) and in the brain by using NIRS
as a change in cerebral HbO2
concentration. If cerebral blood flow, CBV, and
O2 consumption remain constant
during the maneuver, the overall
µa of the blood is affected
without altering the
[Hb]tot.
HbO2 is, therefore, equivalent
to the product of the total CHC and
SaO2. The total CHC can then be
easily computed as
![CHC = <FR><NU>&Dgr;[HbO<SUB>2</SUB>] − &Dgr;[Hb]</NU><DE>2 ⋅ &Dgr;fSa<SUB>O<SUB>2</SUB></SUB></DE></FR>](/content/vol85/issue5/fulltext/1915/img004.gif)
|
(4) |
fSaO2 is the fractional change in
SaO2. The term
[Hb]diff is often used
to describe the difference between the oxy- and deoxyhemoglobin
concentration (HbO2 [Hb]). CHC can therefore be computed from
the gradient of the plot of
[Hb]diff and
SaO2.
It is important to remember that NIRS measures changes in chromophore concentration in micromolar units. Estimates of blood volume are obtained from these measurements by converting the concentration data into the more conventional clinical units of milliliters/100 grams. This conversion requires knowledge of the concentration of red blood cells (the hematocrit). The hematocrit varies with vessel size and is lower in smaller vessels and capillaries. Lammertsma et al. (13) measured the cerebral-to-large vessel hematocrit ratio as 0.69, and Sakai et al. (19) found a value of 0.76. The value measured depends on the distribution of vessel sizes considered. Sakai et al. also found that the hematocrit may change with blood volume. For a 15% increase in blood volume (induced by 5% CO2 inhalation), the hematocrit dropped by 5%. For these reasons, the present study considers only the measurement of hemoglobin concentration, rather than blood volume.
Experimental Procedure
Five newborn piglets (age <24 h) were studied. Anesthesia was induced by using 5% isofluorane, mechanical ventilation was established, and the isofluorane level was reduced to 1.5%. Ventilation was maintained at FIO2 of 40% (balance N2). The optodes from a Hamamatsu NIRO 500 spectrometer were placed on either side of the head at spacings between 3.8 and 4.5 cm (Table 1). By using pulsed laser diodes at four wavelengths (779, 821, 855 and 908 nm), the spectrometer was used to measure changes in the concentration of cerebral HbO2 and Hb. For newborn piglets, a differential pathlength factor (4.57) was used to convert the data to micromolar units (R. Springett, personal communication). SaO2 and heart rate were measured by using a pulse oximeter (Novametrix 500) via a probe positioned on a skin flap on the right leg. Mean arterial blood pressure was recorded via the umbilical artery, and analog outputs from the oximeter and blood pressure transducer were linked directly to the spectrometer for display and storage alongside the NIRS data. All data were sampled every 2 s.
|
A computer-controlled gas blender (8) was used to supply a controlled, accurate mixture of O2, N2, and CO2 to the time-cycled ventilator (Vickers model 77) with preset pressure. Initially, FICO2 was set to 0%, and at least six CHC measurements were made by slowly varying FIO2 over several minutes. FICO2 was then increased to either 2.5 or 5%. Once a new stable baseline had been reached, the CHC measurements were repeated. Arterial blood samples were used to measure arterial PCO2 throughout the study.
Experimental Results
An example of the NIRS and SaO2 data collected during a single CHC measurement in one piglet is shown in Fig. 1. In this example, SaO2 was reduced to ~50%. However, only the initial portion of this change (in which SaO2 is >90%) is used for the calculation of CHC. As previously described, CHC is calculated from the gradient of the[Hb]diff and
SaO2 plot shown in Fig.
2.
[Hb]tot is included
on this plot to demonstrate that, for small changes in
SaO2, the cerebral circulation is not
disturbed. At least three CHC measurements were made at each of the two
levels of PaCO2 for each piglet. The
summarized results of these measurements are shown in Fig.
3 and detailed in Table 1. The normocapnic PaCO2 levels for each piglet ranged
between 3.8 and 6.8 kPa.
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|
|
The coefficient of variation for the CHC method was 22.8 ± 3.5%
(mean ± SD) compared with 2.8 ± 2.8% for the
[Hb]tot method. A
statistical calculation (assuming a power of 90% at the 95% significance level) was performed on the CHC data to determine the
minimum change in CHC which could be detected
(min[CHC]d) under
these experimental conditions. The values for this
min[CHC]d are given
in Table 1. In four of the five piglets,
min[CHC]d was
higher than the change in
[Hb]tot and CHC
because of alteration of PaCO2.
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THEORETICAL MODEL |
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To investigate this problem theoretically, we used diffusion theory (1, 18) to calculate the expected attenuation of light from known values of absorption and scattering of tissue and blood and also the source-detector separation. The tissue was assumed to be a uniform slab of 40-mm thickness, with an optode spacing of 40 mm (similar to that used in the experimental study). Data were calculated by using absorption and scattering properties at four wavelengths, chosen to be similar to those used by the NIRO 500. The values used for the µa of hemoglobin and water are shown in Table 2. Data are taken from Matcher et al. (17) for blood and from Matcher et al. (16) for water and then converted to natural log base (ln). The µa of blood was calculated by assuming a hemoglobin concentration in the blood of 1.68 mM (10.8 g/100 ml) and a mean saturation of 80%. To calculate blood volume, we used a saturation decrease of 3%.
|
The transport scattering coefficient,
µ's, is defined as
µs (1 g,
where g is the anisotropy factor of blood). This was assumed
to vary linearly with wavelength 
and was calculated from the
following expression [approximated from data measured on human
neonatal brain (23)], with in nanometers
|
(5) |
|
(6) |
Recent papers (10, 14) have shown that, when the blood is confined to vessels, embedded in a high-scattering, low-absorbing background, the effective absorption of the tissue will vary inversely with the vessel diameter. To investigate the effect of blood vessel size, we used the expression derived by Liu et al. (14) for the effective µa
![⟨&mgr;<SUB>a</SUB>⟩ = &mgr;<SUB>a</SUB> + f<SUB>v</SUB>(&mgr;<SUP>b</SUP><SUB>a</SUB> − &mgr;<SUB>a</SUB>) exp [−<IT>r</IT>(&mgr;<SUP>b</SUP><SUB>a</SUB> − &mgr;<SUB>a</SUB>)]](/content/vol85/issue5/fulltext/1915/img007.gif)
|
(7) |
It has been suggested that, in the neonate, because skull bones are not fused, the head could expand slightly with a blood volume increase, leading to a change in the optode spacing. To investigate the effect of optode movement, the source-detector separation used in the diffusion equation was varied by ±1 mm.
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THEORETICAL RESULTS |
|---|
Optical Pathlength and Background Absorption
Both methods for measuring concentration change rely on knowledge of the differential pathlength factor, because the hemoglobin concentrations calculated are inversely proportional to . Thus, if
changes during the measurement, for whatever reason, the concentration measurement will be affected.
Changing the µa of the medium
will alter by a small amount. On the basis of diffusion theory, it
is possible to calculate that, for typical tissue optical properties of
µa = 0.02/mm and µs = 1.0/mm, a 15% increase in
absorption will lead to a 4% decrease in the pathlength .
The continuous
[Hb]tot is
calculated directly from changes in attenuation. Thus, this variation
in pathlength will give rise to a 4% error in the measured change in
concentration (i.e., a 15% change in concentration would be measured
as a 14.4% change).
For the CHC method, the difference between two absolute
concentrations is calculated, and the pathlength change will lead to an
error in the absolute concentration (i.e., the second concentration will be 4% too small). If the difference is calculated for an initial concentration of C and an increase of 15%
|
(8) |
The change in pathlength with increased blood volume is dependent on the absorption properties of the bloodless tissue. If the background absorption of blood tissue is zero, then a 15% increase in blood volume results in a 15% increase in absorption. However, if the background absorption of the tissue is nonzero, then a change in the blood volume will have a lesser effect on the total absorption.
Figure 4 shows the theoretical estimate of
the measured change in hemoglobin concentration for the two techniques
as calculated for different background
µa (for differing values of
k in Eq. 6). As expected from the above discussion, for low
background absorption, the CHC method underestimates the real value.
As the background absorption increases, the continuous
[Hb]tot
method gives smaller values.
|
Effect of µs
As an example of the effect of a change in µs, Fig. 4 also shows the effect on the blood volume measurements of a 1% drop in µs simultaneous with an increase in blood volume.Blood Vessel Diameter
Figure 5 shows the effect on the [Hb]tot measurements of restricting the change in volume to blood vessels of different diameters. As blood becomes concentrated in fewer, larger vessels, the observed concentration change, as measured by both theCHC and
[Hb]tot methods,
decreases.
|
Movement of Optodes
Figure 6 shows the effect of a ±1-mm optode movement on the measurement of[Hb]tot and CHC
for a fixed change in blood volume of 15%. This model predicts that
optode movement will have a greater effect on the
[Hb]tot than on the
CHC signal; i.e., with zero optode movement,
[Hb]tot = 15% and
CHC = 10%; however, in the presence of optode movement of 1 mm, the
observed [Hb]tot = 29%, whereas the CHC signal will still show a change of 10%. These data have been calculated by assuming a movement of the optodes occurring between measurement of volumes but also assuming that the
optode positioning is constant during any given CHC
measurement.
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DISCUSSION |
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Top
Abstract
Introduction
Discussion
References
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Experimental Results
It can been seen clearly that the variability of the CHC values is far greater than the variability of the[Hb]tot data. By
using the CHC values obtained at the lower
CO2 levels, the sensitivity of the
measurement was calculated and was found to be inadequate to accurately
resolve the changes in CHC caused by the increase in
FICO2.
For the data presented, although [Hb]tot always
shows an increase with
FICO2,
CHC does not show a significant change.
The calculation of CHC in Eq. 4
involves finding the regression between
([HbO2] [Hb]) and fSaO2. To avoid
changes in blood volume and flow, the changes of
SaO2 are usually restricted to
<10%. These changes are measured by using pulse oximeters with typical accuracy of ±2% (21). Therefore, it is likely that the error on the SaO2 reading may be
responsible for a large degree of the noise in the CHC measurement. A
further problem can be the limited analog-output step size of pulse
oximeters, because some output data only in steps of 1%.
The noise in the measurement of
[Hb]diff is
approximately the same as that for measurement of
[Hb]tot. However, in
the conversion to an absolute concentration, the former is divided by a
small factor (fSaO2) which acts to
multiply the noise. This is further compounded by errors in the
measurement of SaO2.
Unfortunately, although greater SaO2 changes would potentially lead to more accurate estimations of blood volume, a large change in SaO2 might also result in physiological effects, such as changes in blood volume or flow, which would invalidate the basis of the measurement. Also, for clinical use of the technique, there are obvious limits on the change in FIO2 which can be used without compromising the patient's health.
Intersubject variability in DPF has been shown to be on the order of
12-17% (5). However, in these studies, DPF acts purely as an
equal scaling factor on both the CHC and
[Hb]tot data to
convert the units from micromoles/centimeter to micromoles and, as
such, will not contribute to the discrepancy shown between the two measurements.
Theoretical Results
The effect of background tissue absorption can be understood by considering the fact that, when blood volume increases, the tissue volume sampled by the light decreases. Hence, any attenuation caused by tissue absorption will decrease. This acts to lessen the overall increase in attenuation with blood volume and will thus tend to lead to an underestimate of the hemoglobin concentration change. Increasing the number of wavelengths used would make the fitting to the Hb spectral shape more accurate and thus decrease this error (17).The theoretical modeling of changes in µs demonstrates a significant effect on the measurement of CHC. However, it is difficult to think of physiological circumstances under which such a change in µs might be observed. It would seem unlikely that the µs of tissue would change by more than a fraction of a percent under normal physiological circumstances. The overall µs might vary as the blood volume changes, due to a change in the µs of blood. Kienle et al. (12) measured the anisotropy factor of blood g at 820 nm as 0.993 and its µs as 5.5/mm for a hematocrit of 0.01 (which agrees well with Mie theory calculations for red blood cells). This corresponds to a µs of 134/mm for a hematocrit of 0.41 and a µ's of 0.94/mm. Because the transport µs of neonatal brain tissue is ~1.0/mm (23), the change in µs due to a small increase in blood volume will be negligible. Changes in scattering caused by the depolarization of neurons both in the normal brain during cortical activation and under pathological conditions of stroke are presently the subject of much investigation (3, 9). Typically the magnitude of any changes seen under these conditions is very small, particularly compared with changes resulting from a change in µa; hence, such changes would also have a negligible effect on the described model.
There is some uncertainty in the value of g for blood, and since it is so close to 1.0, uncertainty in g will lead to large uncertainty in µ's. However, because a blood volume increase of 15% is only 0.3% of the total (~2%) volume, the maximal decrease in scattering would be 0.3% if blood were nonscattering. A 1% increase in scattering could be caused only if blood had a µs 5 times that of the surrounding tissue, which is clearly not the case.
As expected, the modeling of the effect of vessel diameter demonstrates that the measured hemoglobin concentration change decreases as the blood becomes concentrated in fewer, larger vessels. However, because the majority of vessels inside the brain are <0.2-mm diameter, this is unlikely to have a major overall effect. It is possible, however, that a significant change in blood volume confined to the larger pial arteries and veins on the brain surface could lead to an underestimate of the hemoglobin concentration change.
The effect of optode movement is clearly more significant in the
measurement of
[Hb]tot than
CHC. In practice, it is probable that the optodes will be subject to
continuous small random movements, which will give rise to greater
noise and uncertainty in the measurement. This effect is likely to be
larger for the CHC measurement, for the reasons previously stated,
because CHC is derived from a difference of two measurements.
In conclusion, the measurement of absolute hemoglobin concentration has been shown to be inherently more liable to noise than measurements of changes in hemoglobin concentration.
The absolute value of CHC measured with this technique is dependent on
the µa of the tissue surrounding
the blood vessels. The
[Hb]tot measurement
is less sensitive to the background
µa, but it is affected more by
any changes that occur in the µs
of the tissue. Measures which might lead to an improved understanding of the CHC data would be 1) to
monitor the optical pathlength during the experiment,
2) to improve the accuracy of the
pulse oximetry, and 3) to use more
wavelengths to improve fitting to the hemoglobin absorption spectra.
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ACKNOWLEDGEMENTS |
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We thank the Engineering and Physical Sciences Research Council, UK (Grant GR/K07386), the Medical Research Council, and Hamamatsu Photonics KK for financial support.
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FOOTNOTES |
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Address for reprint requests: C. Elwell, Dept. of Medical Physics and Bioengineering, Univ. College London, 1st Floor, Shropshire House, 11-20 Capper St., London WC1E 6JA, UK (E-mail: celwell{at}medphys.ucl.ac.uk).
Received 16 June 1997; accepted in final form 1 July 1998.
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 F. Vernieri, F. Tibuzzi, P. Pasqualetti, N. Rosato, F. Passarelli, P. M. Rossini, and M. Silvestrini Transcranial Doppler and Near-Infrared Spectroscopy Can Evaluate the Hemodynamic Effect of Carotid Artery Occlusion Stroke, January 1, 2004; 35(1): 64 - 70. [Abstract] [Full Text] [PDF]
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