Near Infrared Spectroscopy
`
`BIOMEDICAL OPTICS RESEARCH LABORATORY
`UCL DEPARTMENT OF MEDICAL PHYSICS AND BIOMEDICAL ENGINEERING
`
`
`
`Search Medical
`Physics
`
`Near Infrared Spectroscopy
`By Clare Elwell and Jem Hebden
`
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`Introduction
`All of us are exposed to optical (i.e. visible and near-infrared) radiation from the
`sun and other sources throughout our lives. Assuming our eyes are shielded from
`excessive intensity, and our skin is protected from the ultraviolet content of
`sunlight, we accept this exposure in the knowledge that it is perfectly safe. Unlike
`x-rays, optical photons are insufficiently energetic to produce ionisation, and unless
`light is concentrated to such a high degree that it causes burning to the skin,
`optical radiation offers no significant hazard. The diagnostic potential of optical
`methods has been widely known since Jöbsis [1] first demonstrated that
`transmittance measurements of near-infrared (NIR) radiation could be used to
`monitor the degree of oxygenation of certain metabolites. This led to the
`development and increasingly widespread use of clinical near-infrared
`spectroscopy (NIRS), which offers a safe, non-invasive means of monitoring
`cerebral function at the bedside without the use of radioisotopes or other contrast
`agents [2].
`
`Human tissues contain a variety of substances whose absorption spectra at NIR
`wavelengths are well defined, and which are present in sufficient quantities to
`contribute significant attenuation to measurements of transmitted light. The
`concentration of some absorbers, such as water, melanin, and bilirubin, remain
`virtually constant with time. However, some absorbing compounds, such as
`oxygenated haemoglobin (HbO2), deoxyhaemoglobin (Hb), and oxidised
`cytochrome oxidase (CtOx), have concentrations in tissue which are strongly
`linked to tissue oxygenation and metabolism. Increasingly dominant absorption by
`water at longer wavelengths limits spectroscopic studies to less than about 1000
`nm. The lower limit on wavelength is dictated by the overwhelming absorption of
`Hb below about 650 nm. However, within the 650-1000 nm window, it is possible
`with sensitive instrumentation to detect light which has traversed up to 8 cm of
`tissue.
`
`Absorption
`The absorption of light intensity in a non-scattering medium is described by the
`Beer-Lambert Law. This law states that for an absorbing compound dissolved in a
`non-absorbing medium, the attenuation (A) is proportional to the concentration of
`the compound in the solution (c) and the optical pathlength (d):
`
`A = log10 [Io/I] = a.c.d ,
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`where A is the attenuation measured in optical densities, Io is the light intensity
`incident on the medium, I is the light intensity transmitted through the medium, a is
`the specific extinction coefficient of the absorbing compound measured in
`micromolar per cm, c is the concentration of the absorbing compound in the
`solution measured in micromolar, and d is the distance between the points where
`the light enters and leaves the medium. The product ac is known as the absorption
`coefficient of the medium µa. In a medium containing several different absorbing
`compounds (except at very high concentrations not usually met in biological
`media) the overall extinction coefficient is simply the linear sum of the contributions
`of each compound:
`
`A = log10 [Io/I] = [ a1.c1 + a2.c2+ a3.c3 + ... + an.cn ] d .
`
`A compound which absorbs light in the spectral region of interest is known as a
`chromophore. Each chromophore has its own particular absorption spectrum which
`describes the level of absorption at each wavelength. The principle chromophores
`in tissue are as follows:
`
`i) Water
`
`Figure 1: The absorption spectra of pure water. As shown above in figure 1, the
`absorption of light by water is relatively low between 200 - 900 nm. Beyond 900
`nm absorption starts to rise with increasing wavelength, a spectral peak being
`visible at 970 nm. The high concentration of water in living tissue, typically 80% in
`adult brain tissue [3], (equivalent to 56 molar) determines the wavelength region in
`which spectroscopic interrogation of tissue is possible by strongly limiting the tissue
`thickness through which light can penetrate. For this reason, the water spectrum is
`said to demonstrate a "window" of transparency between 200 and 900 nm within
`which spectroscopic measurements can be made. For the purposes of most
`clinical measurements the water concentration in tissue can be thought of as
`constant, and as such water acts as a fixed constant absorber.
`
`ii) Lipids
`
`Although the distribution of lipid in tissue is dependent upon tissue type, it can also
`be thought of as a constant absorber with changes in its concentration throughout
`the course of a clinical measurement being unlikely. The absorption spectrum of
`lipid is approximately the same as that of water and it can comprise 10 - 40 % (i.e.
`several molar) of tissue.
`
`iii) Melanin
`
`Melanin, found in the epidermis layer of skin, is a highly effective absorber of light,
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`especially in the ultraviolet region of the spectrum. Although this absorption can be
`considered to be constant and oxygen independent, the concentration of melanin
`in tissue will directly effect the reflectance of light from the skin and therefore the
`transmission of light into the tissue below.
`
`iv) Haemoglobin
`
`Figure 2: The absorption spectra of HbO2 and Hb. Figure 2 above shows the
`specific extinction coefficients of oxygenated haemoglobin (HbO2) and
`deoxyhaemoglobin (Hb) in the wavelength range 450 - 1000 nm [4]. The difference
`in the absorption spectra explains the well recognised phenomena of arterial blood
`(containing approximately 98% HbO2) having a bright red appearance while
`venous or deoxygenated blood appears more blue. In the NIR region of the
`spectrum the absorption of both chromophores decreases significantly compared
`to that observed in the visible region. However the absorption spectra of Hb and
`HbO2 remain significantly different in this region allowing spectroscopic separation
`of the compounds to be possible using only a few sample wavelengths. An
`isobestic point where the specific extinction coefficients of the two compounds are
`equal can be seen at around 800 nm, which can be used to calculate haemoglobin
`concentration independent of oxygen saturation. The typical value for haemoglobin
`concentration in, for example, adult brain tissue is 84 micromolar.
`
`There are other haemoglobin compounds which have a characteristic absorption in
`the near infrared, although their concentrations in tissue are low and in many
`cases almost non existent in normal blood. These compounds include
`carboxyhaemoglobin, (HbCO), which may be present in significant quantities in the
`tissue in some subjects, but has a low specific extinction coefficient in the NIR
`rendering its effect on most in-vivo measurements negligible. Haemiglobin (Hi) is
`present in very low concentrations and sulfhaemoglobin (SHb) is not present at all
`in normal blood. The combined error in ignoring these compounds in the
`measurement of the total haemoglobin signal is probably less than 1% in normal
`blood and in the majority of clinical conditions encountered. However it is worth
`remembering that some of these forms of haemoglobin, especially Hi, may become
`significantly raised in some diseases of the liver or in malaria.
`
`v) Cytochrome c oxidase
`
`Cytochrome oxidase (CtOx) is the terminal enzyme in the cellular respiratory chain,
`and is located in the mitochondrial membrane. The enzyme contains four redox
`active groups, two haem iron (a and a3) and two copper (CuA and CuB) centres.
`These four metal centres change their redox state (i.e. accept or donate electrons)
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`during electron turnover of the enzyme. The oxygen binding site of the enzyme is
`the binuclear unit which is formed of the CuB and haem a3. It is the donation of
`electrons from this unit to oxygen which accounts for the great majority of oxygen
`consumption in biological tissue. The CuA and haem a centres donate electrons to
`this binuclear unit and are therefore not directly involved in reduction of oxygen.
`However absorption of NIR radiation by cytochrome oxidase occurs primarily at the
`CuA centre, the oxidised spectrum having a characteristic shape, with a broad
`peak centred around 830 nm which is missing in the reduced enzyme. In the short
`term the total tissue CtOx concentration does not vary and NIRS measurements of
`changes in CtOx thus measure alterations in the redox state concentration of CuA
`within cytochrome oxidase.
`
`Figure 3: The difference absorption spectrum between the oxidised and reduced
`forms of CtOx.
`
`Since the total CtOx concentration does not alter, NIRS measurements need only
`be made of the change in redox state, so it is only necessary to know the
`difference spectrum between the oxidised and reduced forms of the enzyme. This
`difference spectrum is shown in figure 3. It can be seen that the magnitude of the
`specific extinction coefficients are similar to those of haemoglobin, but since the
`concentration of cytochrome oxidase in living tissue is usually at least an order of
`magnitude below that of haemoglobin [5], the measurement of cytochrome oxidase
`with optical techniques is by no means as easy as that of haemoglobin. When
`oxygen limits the rate of oxygen consumption by cytochrome oxidase, the CuA
`centre becomes more reduced. Therefore the absorbance of NIR light by
`cytochrome oxidase may be used as an indicator of oxygen availability at a cellular
`level and ultimately of cell metabolism.
`
`Scattering
`Scatter of light in tissue is due to the chaotic variation in refractive index at a
`microscopic and macroscopic scale. This occurs at membrane boundaries of the
`cells themselves as well as at boundaries between various organelles inside the
`cell. Index mismatching will occur between intra and extracelluar fluid, or
`intracellular fluid and fluid inside the nucleus of the cell or other enclosed particles
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`such as mitochondria, ribosomes, fat globules, glycogen and secretory globules.
`As with absorption, the volume of a particular scatterer within the tissue is as
`important as its scattering ability. Evidence suggests that cell membranes are the
`most important source of scattering in brain tissue since they account for a large
`proportion of the solid content of the tissue.
`
`Scatter is by far the most dominant tissue-photon interaction at NIR wavelengths.
`The effect of scattering is to substantially increase the pathlength travelled by
`photons within tissue, and therefore significantly increase the probability of
`absorption occurring. When NIR radiation is scattered in tissue virtually all the
`collisions are elastic, and the direction in which the scattered photon travels is
`dependent upon the size of the scattering particle, the wavelength of the light, and
`the refractive indices of the scattering media through which it is travelling.
`
`The attenuation (A) due to single scattering is proportional to the number density
`of the scattering particles (N), the scattering cross section of the particles (s) and
`the optical pathlength (d):
`
`A = log10 [Io/I] = N.s.d .
`
`The product Ns is known as the scattering coefficient of the medium (µs), and is
`the probability per unit length of a photon being scattered. The reciprocal is the
`mean free path between scattering events. The scattering coefficients of human
`tissues are generally within the range 10 - 100 mm-1, roughly one hundred times
`greater than those for absorption [6]. The most highly scattering tissues include
`bone, cerebral white matter, and skin dermis.
`
`For multiply scattering media such as tissue, the simple formula given above no
`longer applies. In order to fully describe scatter of light in tissue, it is necessary to
`consider the probability of a photon being scattered in a given direction at each
`interaction. The probability of a photon, incident along a unit vector p being
`scattered into a direction q is described by the phase function f(p,q). For a random
`medium it can be assumed that this probability is independent of p and only
`depends on the angle between the incident and scattered directions, e. Thus the
`phase function can be conveniently expressed as a function of the scalar product
`of the unit vectors in the initial and final directions, which is equal to the cosine of
`the scattering angle cos(e). The anisotropy in the probability distribution is
`commonly characterised in terms of the mean cosine of the scattering angle g.
`
`In biological tissues, scatter occurs principally in a forward direction, corresponding
`to an anisotropy in the range 0.69 >g >0.99 [6]. Despite the forward scatter, typical
`values of scattering coefficient ensure that light travelling through more than a few
`millimetres of tissue loses all of its original directionality, and can be treated as
`effectively isotropically distributed. Thus it is convenient to express the
`characteristic scatter of tissues in terms of a transport scatter coefficient:
`
` µs´ = µs (1 - g) ,
`
`which represents the effective number of isotropic scatters per unit length, and is a
`fundamental parameter in diffusion theory.
`
`The Modified Beer-Lambert Law
`When a highly scattering medium is considered, the Beer-Lambert relationship
`must be modified to include (i) an additive term, G, due to scattering losses and
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`(ii) a multiplier, to account for the increased optical pathlength due to scattering.
`The true optical distance is known as the differential pathlength (DP) and the
`scaling factor as the differential pathlength factor (DPF):
`
` DP = DPF . d ,
`
`where d is the geometrical distance. The modified Beer-Lambert law which
`incorporates these two additions is then expressed as:
`
`A = log10 [Io/I] = a.c.d. DPF + G .
`
`Unfortunately G is unknown and is dependent upon the measurement geometry
`and the scattering coefficient of the tissue interrogated. Therefore this equation
`cannot be solved to provide a measure of the absolute concentration of the
`chromophore in the medium from a measure of absolute attenuation. However if G
`does not change during the measurement period, it is possible to determine a
`change in concentration (c2-c1) of the chromophore from a measured change in
`attenuation (A2-A1):
`
`(A2-A1) = (c2-c1).a.d. DPF .
`
`Note that the differential attenuation is actually measured, giving rise to the
`terminology differential pathlength and differential pathlength factor. The
`quantification of the change in concentration still depends upon the measurement
`of the geometrical distance d and the differential pathlength factor, i.e. the true
`optical pathlength which the scattered light has travelled. Although d is simple
`enough to measure, as it is purely the geometrical distance between the points
`where the light enters and leaves the medium, determination of DPF is more
`difficult. There are a number of different techniques which can be used to measure
`DPF in tissue, as briefly described in the following section.
`
`Measurements of DPF
`a) Time of Flight Method
`
`The development of picosecond pulse lasers and ultrafast detectors during the past
`twenty years has made possible the direct measurement of the time of flight of
`light through tissue [7]. The system currently used at University College London is
`a Ti:sapphire laser pumped by a diode-pumped CW laser. The system can
`produce a single pulse with a duration of approximately 2 ps, and with suitable
`mirrors the laser can be tuned between 740 nm and 920 nm. The laser beam is
`split and part of the laser output is taken directly to the streak camera as a time
`reference. The other part of the beam is directed through the tissue sample. The
`temporal reference and the signal which has traversed the tissue sample are
`recorded simultaneously on the same streak image. The geometrical distance d
`between the centre of the transmitting fibre and the centre of the detecting fibre
`bundle is accurately measured. As with conventional spectroscopy measurements,
`it is important to minimise movement of the tissue between the fibres and various
`stereotactic devices have been used to stabilise the tissue under interrogation.
`
`The time difference <t> between the light entering the tissue and the mean time of
`that which has traversed the tissue is measured from the streak image and is then
`used in the calculation of the differential pathlength factor in a simple time of flight
`approximation:
`
` DPF = DP / d = cv.< t> / d.n ,
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`where cv is the speed of light in a vacuum, and n is the refractive index of the
`tissue (usually taken as 1.40 [8]). Time of flight systems of the type described
`above are large, expensive, and confined to dedicated optical laboratories, which
`precludes routine clinical use on neonates. To date therefore, this method has
`generally been restricted to measurements on post mortem infants and on adult
`volunteers [9-11].
`
`b) Intensity Modulated Optical Spectrometer
`
`Figure 4: Measurement in the time and frequency domains.
`
`When NIRS is applied to measurements of tissue oxygenation in the fetal head
`during labour and childbirth or in muscles during exercise, large changes in the
`geometrical distance d may occur during a study. The ultimate goal of a reliable
`accurate bedside spectrometer can therefore only realistically be achieved when
`real time measurement of the total light path can be incorporated. By making
`spectroscopy measurements in the frequency rather than the time domain, it has
`been possible to develop a new method of continuously monitoring the total path
`which the NIR light has travelled in the tissue of interest. Figure 4 demonstrates
`schematically the principles of the time and frequency domain measurements of
`DPF. A continuous laser source can be easily modulated at all frequencies from
`DC to a few hundred MHz and the phase shift between the light entering and
`exiting the tissue can be recorded. It has been shown [12] that if P is the phase
`shift measured in radians, then for modulation frequencies less than 200 MHz, the
`total distance travelled through the tissue, DP, is given by:
`
` DP = P. cv / 2.pi. f. n ,
`
`where f is the modulation frequency. Optical pathlengths at one or two wavelengths
`have been reported from a measurement of phase shift of light modulated at a
`single frequency of 220 MHz [13].
`
`An intensity modulated optical spectrometer capable of working at several
`wavelengths and over a wide range of modulation frequencies has been developed
`at UCL for use in a number of clinical studies [14-16]. In this spectrometer four
`intensity modulated laser diode sources produce light in the NIR region enabling
`the user to measure the optical pathlength at each wavelength and simultaneously
`perform the conventional spectroscopy measurements for determining the change
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`in concentration of Hb, HbO2 and CtOx. In this way the change in measured
`attenuation can be corrected for pathlength variations in real time. Since the
`distance measured is the total optical pathlength, DP, this instrument negates the
`need for manual measurement of the geometric distance d, and therefore
`represents a dramatic improvement in the accuracy of pathlength and
`spectroscopy measurements. Research with this type of instrument will prove
`useful not only in studies where the optical fibres are likely to move (e.g. during
`fetal measurements or exercise studies) but also in improving the accuracy of
`oxidised cytochrome oxidase measurements which, due to the chromophore's
`relatively low concentration in tissue are particularly vulnerable to errors in optical
`pathlength estimation. In addition, since this type of system is capable of
`measuring DP at four wavelengths, a more accurate correction for the pathlength
`variations known to occur with absorption coefficient, µa, (and hence wavelength)
`can also be made.
`
`Although this type of system allows a continuous measurement of DP combined
`with the normal spectroscopic measurements, the first clinical application of the
`instrument has been to collect data on the absolute values of DP in different
`groups of subjects. If the geometrical distance d is measured at the time of the
`study, the DPF can easily be calculated from the DP measurement, and compared
`with previously recorded values. The portable nature of the instrument has also
`allowed the measurement of cranial DPF to be made on a large group of live
`neonates at the cotside [15]. In addition a study has recently been completed to
`investigate the age dependence of cranial DPF in humans. The intensity
`modulated spectrometer described above was used to measure cranial DPF in a
`total of 283 subjects whose age ranged from 1 day to 50 years [16]. The results
`suggest a slowly varying age dependence of DPF following the relation:
`
` DPF780 = 5.13 + 0.07Y 0.81 ,
`
`where DPF780 is the differential pathlength factor at 780 nm and Y is the age of
`the subject in years.
`
`Factors influencing the total optical pathlength
`It has been shown that DP is dependent upon the following factors:
`
`a) Tissue Type
`
`Measurements of DPF have been made on neonatal head, and adult head,
`forearm and calf. A marked difference is seen between these four tissue types.
`This difference is to be expected since the DPF is directly dependent upon the
`proportion of, for example, soft tissue, muscle and bone in the illuminated tissue.
`
`b) Absorption coefficient (and wavelength)
`
`The time of flight system has been used to demonstrate the relationship between
`DPF and wavelength (and hence absorption coefficient) in the adult head, forearm
`and calf (in vivo) and the infant head [17]. The DPF was estimated from mean time
`of flight measured between 740 nm and 840 nm. In all tissues the DPF generally
`decreased with increasing wavelength, although also exhibiting the absorption
`characteristics of haemoglobin, with a variation of typically 12% over the
`wavelength range. For this reason it is always important when quoting DPF values
`to also quote the wavelength at which the DPF measurements have been made.
`The spectral dependence of DPF must also be taken into account in the algorithm
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`used to convert from attenuation (OD) data at a range of wavelengths to changes
`in chromophore concentration.
`
`c) Geometry of Optodes
`
`It has been demonstrated [18] with the aid of computer simulation, that on a
`spherical object DPF is dependent upon angular position. DPF may vary
`significantly between an emitter-detector angle of 180 to 60 degrees, and even
`more rapidly for lesser angles. In contrast, an experimental study [10] showed that
`in all tissues DPF initially falls with increasing geometrical distance, d, the value
`becoming almost constant for source-detector spacings above 2.5 cm. This
`discrepancy between the theoretical and experimental results can be explained in
`part by the fact that the theoretical model did not take into account the
`inhomogeneity of the tissue illuminated. This has been confirmed by modelling of
`multilayered tissues where the DPF has been shown to vary with angle in the same
`way as that observed experimentally [19]. Much work is currently being done to
`further refine the methods used for prediction of DPF, particularly in realistic tissue
`models [20].
`
`Spectroscopic Measurements of the Brain
`
`Figure 5: NIRS measurement across the head.
`
`Figure 5 shows a schematic of the experimental set up for the spectroscopic
`measurement across a head. The optical fibres which carry the NIR light to and
`from the head are terminated with small prisms which direct the light normally on to
`the surface of the tissue. The geometrical distance d, known as the interoptode
`spacing (IOS), is usually measured with a pair of callipers directly over the
`measurement site. Note this distance is the chord (straight line) distance rather
`than the length of the arc between the two points. This assumption is based upon
`the fact that light inside the brain becomes essentially diffuse within a few
`millimetres of entering the tissue, at which point it becomes an isotropic source
`[18], even if the angle between the source and detector is less than 180 degrees.
`
`The differential pathlength factor has been measured in the adult head using both
`the UCL time of flight system and the UCL intensity modulated optical
`spectrometer, and a value of approximately 6 was obtained. Therefore, for an IOS
`of 4 cm, the mean distance which the light actually travels in the head is
`approximately 24 cm.
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`The chromophores of interest within the tissue, whose concentration vary with
`oxygenation are HbO2, Hb and CtOx. The specific extinction coefficients (a) for
`these chromophores are expressed in units of per micromolar of chromophore per
`litre of tissue per cm. Once d, a and DP are known the change in chromophore
`concentration can easily be computed from the measured change in attenuation.
`However for the simultaneous computation of the changes in concentration of a
`number of chromophores from the changes in attenuation at a number of
`wavelengths, a matrix operation must be performed incorporating the relevant
`extinction coefficients for each wavelength and chromophore. For each wavelength
`it is assumed that the linear changes in attenuation for each chromophore can be
`linearly summed. The result of these computations is the value of the absolute
`change in concentration of each chromophore in the non arbitrary units of
`micromolar of chromophore per litre of tissue.
`
`Since the absolute concentration of chromophore is unknown (and cannot be
`determined due to the effects of light scattering within the tissue), all
`measurements are expressed as absolute concentration changes from an arbitrary
`zero at the start of the measurement period. Thus using this technique the
`quantified changes in tissue oxygenation can be non invasively monitored.
`Furthermore the quantified changes in the concentration of Hb and HbO2 in the
`units micromolar can be used to measure absolute haemodynamic parameters
`such as cerebral blood flow [21] and cerebral blood volume [22].
`
`References
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`Cope, M, and Delpy, DT (1996): Pediatr. Res. 39, 889-894.
`17. Essenpreis, M, Elwell, CE, Cope, M, van der Zee, P, Arridge, SR, and
`Delpy, DT (1993): Appl. Opt. 32, 418-425.
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`
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