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Optical methods for assessing skin flap survival
Optical methods for assessing skin flap survival
12
P. Va¨lisuo University of Vaasa, Vaasa, Finland
12.1
Introduction
The skin flap is a full thickness mass of skin, transplanted from a donor site to a recipient site. The blood circulation of the skin flap must be guaranteed to keep the skin flap alive. Some of the original blood vessels can be kept intact for local skin flaps, whereas a microsurgery is required to establish blood circulation for free skin flaps. Venous congestion and arterial occlusion are typical reasons for the failures of skin flaps (H€ olzle et al., 2005). Venous congestion leads to increase of blood fraction and haemoglobin concentration in tissue, whereas the declining blood flow and oxygen saturation are symptoms of arterial occlusion. If these vascular compromises are detected early enough, they can be operatively explored and fixed saving the skin flap (H€olzle et al., 2005). Therefore, skin flap vitality monitoring is very important during 24 h after the transplantation. The blood perfusion, blood flow and oxygen saturation are vital signals predicting the success of the operation. Smit et al. (2010) reviewed the most often used methods for skin flap assessment during 1999–2009. They reviewed implantable Doppler system, colour duplex ultrasonography, near-infrared spectroscopy (NIRS), microdialysis and laser Doppler flowmetry (LDF). The group concluded that Doppler systems and NIRS are the best methods available today. Payette et al. (2005) compared optical spectroscopy and laser Doppler assessment and concluded that the optical spectroscopy and spectroscopic imaging may have a lot of potential in evaluating skin flaps. Skin is a thin organ, easily accessible using non-invasive optical probes. Optical methods provide real-time information from a single point or over an area of skin. They can be used for measuring the blood perfusion, flow and oxygen saturation. This chapter reviews existing optical skin parameter assessment methods suitable for skin flap follow-up as well as some new methods that are not in regular clinical use yet.
12.2
Methods
There are many methods that can be used for measuring constituent skin parameters, including those vital signals predicting the success of the skin flap. Optical methods are popular because the skin is thin and easily accessible by non-invasive optical probes. The wavelengths approximately between red (600 nm) and near infrared (NIR) below Biophotonics for Medical Applications. http://dx.doi.org/10.1016/B978-0-85709-662-3.00012-9 Copyright © 2015 Elsevier Ltd. All rights reserved.
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950 nm, reside between the absorption peaks of haemoglobin and water and form the socalled diagnostic–therapeutic window (Tuchin, 2007), in which the light penetrates deeper into the skin than the other wavelengths. Therefore, many optical skin assessment methods operate in visual and NIR wavelength range. The optical methods are usually non-invasive, high-resolution, real time and relatively inexpensive. The existing methods are the results of research activities carried out for decades. Many useful clinical solutions have appeared, and there are dozens of interesting methods to analyse the fundamental parameters of the skin. Such as the pulse oximeter measures the oxygen saturation, LDF can measure the blood flow. However, the light interaction in skin is still too complex for a single method that would fit for all purposes. Some often used non-optical methods are, transcutaneous oxygen partial pressure (PO2) meter, Doppler ultrasonography, magnetic resonance imaging (MRI) and thermal imaging (Salman et al., 2005). The Doppler ultrasonography is only suitable for measuring the blood flow of large vessels. MRI is expensive and difficult to apply for most cases, but the transcutaneous PO2 meter is still in active use. While these nonoptical methods are good for many purposes, researchers are searching for optical methods that outperform them. Existing and emerging optical methods are often cost-effective, non-invasive, fast and reliable.
12.2.1 General properties behind optical measurements Optical probes consist of an emitter and a receiver or a matrix of receivers. The light from the emitter enters into the skin and scatters back to the receiver, carrying information about the skin, which can be analysed. The probe geometry determines the sampling depth and the coverage area of the measurement (see Figure 12.1). If the probe contains only one receiver, the measurement provides a single value at a time, representing an average of the probed area. If the receiver contains an array of sensors, it can produce an image over the area examined, where each value represents an average over a subset of the whole image. Some pointwise methods can be extended to imaging systems by mechanically scanning over a skin area. Medical imaging and pointwise measurements reveal information about the structure and the chemical composition of the tissue. The contrast in the measurement is achieved by one of the following five methods: (1) by structural changes in the tissue, (2) by changes in chemical concentrations in the tissue, (3) by the Doppler shift due to the moving particles in the blood circulation, (4) by observing changes in the laser speckle patterns due to moving particles in blood circulation and (5) by using an external contrast agent, such as indocyanine green (ICG). The methods, whose contrast is based on the movement of the particles, are called dynamic methods as opposed to chemical concentration and structure-based methods, which are called as static. When the target of the measurement is to follow the functions of the organ, the method is said to be functional. Optical functional measurements can be based on either dynamic or static methods. Usually functional measurements are made indirectly by observing changes in blood circulation or concentrations of nutrients and by-products of metabolic processes, such as cellular respiration. Active metabolism decreases the concentrations of nutrients and increases the concentrations of the by-products. Cellular
Optical methods for assessing skin flap survival
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d2
d1
Skin
(a)
(b)
(c)
Figure 12.1 The illumination geometries. In panels (a) and (b) the light is directed to the skin through an optical fibre, and the reflected light is detected with another fibre. The distance between the fibres affects the sampling depth of the probe. If the emitter and detector fibres are close to each other, the reflectance from the shallow parts of skin is increased. When the distance is larger in panel (a), the average sampling depth is deeper. Panel (c) shows a situation in which the skin is illuminated with planar light source, and the reflectance is collected from a point. This corresponds to the case in panel (b), where the emitter and detector are close to each other and the average sampling depth is shallow. The dashed horizontal line shows the diffusion limit, below which the measurement is challenging.
activity also typically affects local blood regulation by increasing blood perfusion. Functional measurements and functional imaging can be used in assessing the vitality of the skin flaps. As mentioned earlier, the blood perfusion and oxygen saturation are often important signals used to describe the skin flap vitality. Blood perfusion can be monitored by measuring the speed of red blood cells (RBCs) (dynamic method) or by measuring the blood concentration and oxygen saturation (static method). The speed of the RBCs and thus the speed of the blood flow is most often measured using Doppler shift of reflected light. The static methods are often based on the analysis of the reflectance spectrum when skin is illuminated with a wideband light source. The chromophores carried by blood change the reflectance spectrum and thus the blood circulation can be analysed. The pulsation of the arteries cyclically changes the blood volume in skin, which can be observed from optical transmission or reflectance signal. The pulsation itself is a good indicator of blood perfusion, and it can be also used for separating the absorption signal originating from arterial blood from other absorption sources. Furthermore, the metabolism of the cells can be indirectly observed by analysing the concentrations of nutrients and resulting chemicals using optical spectroscopy. However, the strong scattering of light in skin often makes the application of optical methods challenging. Most of the photons in the reflectance spectra are multiple scattered and have lost all of their direction information, and the length of the path the photons travelled in the skin is unknown. For these reasons, the change in chemical concentration changes the reflected light intensity nonlinearly, and the effect in reflectance spectrum of one chromophore is dependent on the concentrations of each other
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chemical in skin. Therefore, the spectrum analysis is often time-consuming and dependent on specific skin models. Some methods rely only on directly transmitted or single scattered ballistic photons, gating away the multiple scattered photons. These techniques provide sharp images and easy methods for chromophore concentration analysis. The disadvantages are that most of the light is lost in gating, and the proportion of ballistic photons is reduced exponentially when the depth is increased. Therefore, ballistic methods are best suited for superficial analysis of relatively translucent areas.
12.2.2 Diffusion limit When an orthogonally incident, collimated light beam, whose intensity is, I0, travels through a purely absorbing slab, whose depth is, d [cm], its intensity after transmission, I, is described according to the following equation: I ¼ I0 ema d ;
(10.1)
where ma [1/cm] is the absorption coefficient, which is the reciprocal of the average travelling distance of a photon between the absorption events. Similarly, if the material is purely scattering, the intensity of non-scattered light transmitted directly through the slab is given by the following equation: 0
I ¼ I0 ems d ;
(10.2)
where ms0 [1/cm] is the reduced scattering coefficient, which is a reciprocal of the mean free path between scattering events in the corresponding isotropic medium. The effect of both absorption and scattering can be accounted using the total interaction coefficient: mt ¼ ma + m0s . The reciprocal of the total interaction coefficient is the mean free path length between either scattering or absorption events: pmf ¼
1 1 ¼ : mt ma + m0s
(10.3)
When the thickness of the slab is increased, until it equals the mean free path length (d ¼ pmf), the intensity of the transmitted light I ¼ I0/e, and only 37% of the light passes directly through the slab. This thickness is said to be the diffusion limit. Superficial measurements and imaging above the diffusion limit are relatively easy, whereas more advanced techniques are needed to sample tissue below pmf. For human skin, the pmf is roughly about 1 mm (Farrell et al., 1992).
12.2.3 Dynamic methods In dynamic methods, the measurement signal is caused by Doppler shift of the photons scattered from the moving particles, as shown in Figure 12.2. Because the signal is relational to the speed of the blood flow in dynamic methods, they are especially applicable for blood perfusion measurements.
Optical methods for assessing skin flap survival
R–
R+
R
v Skin
(a)
(b)
335
Figure 12.2 The light reflectance from the skin with capillary loops. Panel (a) shows reflectance due to the scattering of moving RBCs. The RBCs moving upwards create a positive Doppler shift to the reflected signal R+, whereas the RBCs moving downwards cause corresponding negative Doppler shift to reflectance signal R. Panel (b) shows reflectance R due to scattering from tissue. The absorption of blood and tissue affects to the reflectance spectrum of R.
12.2.3.1 Laser Doppler flowmetry (LDF) The LDF was developed in the late 1970s for measuring the blood flow of tissue (Riva et al., 1972; Stern, 1975). Later, LDF established its position in blood perfusion monitoring. In LDF, the skin is illuminated by a coherent, collimated laser beam, using an optical fibre. The light reflected back from the tissue is collected by another fibre(s). Moving particles, such as RBCs, cause a Doppler shift that broadens the spectrum of the reflected light. This broadening of the spectrum can be used to calculate the speed and concentration of the RBCs. LDF is a proven, continuous, non-invasive and realtime method with many clinical instruments available. LDF is often used for skin flap assessment. The speed and the concentration of the RBCs in capillaries provide useful information about the function of the blood vessels after the surgery. Use of different probes makes the LDF method suitable for different situations. Typical probes are: surface probes, needle probes and single fibre probes, which can be inserted under the skin using a standard cannula. The disadvantages of LDF measurement are the small spatial region covered and the possible artefacts due to bending of the probe fibres and due to the pressure applied toward the skin by the probe. LDF measurement reaches only a few hundreds of micrometers below skin surface (Riva et al., 1972; Stern, 1975).
12.2.3.2 Laser Doppler imaging (LDI) The LDF measurements cover relatively small areas of skin. However, it is well known that the blood perfusion is spatially rather heterogeneous, and the LDF measurement sample covers only a small volume of tissue and depends, therefore, on the position of the sensor. Laser Doppler imaging (LDI) overcomes these issues by measuring the flow speed from many different locations over the imaged area. The scanning over a spatial area can be implemented mechanically by the use of rotating mirrors and stepper motors. The sampling depth is similar to that of LDF, about 200–240 mm (Wa˚rdell et al., 1993; Serov et al., 2002). LDI is a popular method for assessing skin flaps. It can provide information about the function of the blood circulation over a large area or the whole skin flap. LDI does
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not require the probe to be in contact with skin, which removes possible disinfecting problems. The disadvantages of LDI are rather moderate resolution, small penetration depth and slow imaging process.
12.2.3.3 Doppler optical coherence tomography (DOCT) Optical coherence tomographic (OCT) system uses a coherent laser beam to illuminate target skin area and a pointwise sensor to collect the reflectance. An interferometer is used to include only single scattered photons reflected from the specific depth. OCT systems can be used for scanning the tissue properties laterally over a skin area and axially over different depths, up to the diffusion limit. OCT can therefore provide a high-resolution profile of the scanned skin area. It is also possible to use a twodimensional charge coupled device (CCD)-sensor as a receiver to avoid lateral scanning altogether. This kind of full-field OCT has already been demonstrated, but it may not be available for clinical use in the near future. The most popular target for OCT is ophthalmology, studying the properties of retina since it is more translucent than many other tissues (Izatt et al., 1997; Chen et al., 1998). Typically, OCT uses only single wavelength within the diagnostic–therapeutic window to scan as deep as possible. Because only one wavelength is used, spectroscopic analysis cannot be made, but a picture of the skin structure can be achieved. OCT, as such, is not very useful for skin flap imaging, but Doppler OCT (DOCT) can measure the velocity of the scatterers using the Doppler shift of the reflected light together with structural scanning of the tissue. In that case, both the high-resolution structural image and the estimate of blood perfusion can be achieved. The disadvantages of OCT and DOCT are that they are rather slow to achieve data because the imaged area needs to be scanned both laterally and axially; therefore, high temporal resolution is not achieved. The image is a very high-resolution image, but only an average image over the area may be needed for skin flap assessment. DOCT can only obtain information above the diffusion limit.
12.2.3.4 Laser speckle imaging (LSI) and laser speckle correlation analysis Laser speckle imaging (LSI) was first used in the 1980s. It has since been used for imaging the cerebral blood flow (Briers and Fercher, 1982; Dunn et al., 2001). When tissue is illuminated with coherent laser light, the reflections from different parts of the uneven tissue surface reach the detector at different times, forming interference artefacts, called speckle patterns or speckle noise. When the light is reflected from moving particles, such as RBCs in the blood, the speckle pattern starts to fluctuate. Higher RBC velocity produces higher local variance to the speckle pattern. This phenomenon can be used for quantifying blood perfusion under the skin. The LSI has high lateral resolution (tens of micrometres), high temporal resolution (millisecond) and low cost. The measurement system can be constructed easily using standard components; only a
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standard laser and normal CCD camera are needed. The disadvantage is shallow measurement depth (Choi et al., 2004; Forrester et al., 2004). LSI may be a better method for assessing the validity of skin flaps than the LDI method (Stewart et al., 2005).
12.2.3.5 Diffuse correlation spectroscopy (DCS) Diffuse correlation spectroscopy (DCS) is an emerging technology for blood perfusion measurement of deep tissues. DCS was demonstrated in 1995 (Boas et al., 1995). It can directly observe the interference of moving scatterers, such as RBCs, and therefore the blood perfusion. DCS works similarly to LDF, but it can also handle multiple scattered light. The mathematical basis of the method is correlation transport theory, which is modified from the radiative transport theory to estimate correlation transfer through a turbid medium (Boas and Yodh, 1997). The measurement reaches more than 10 mm below the skin surface, and the acquisition time is approximately 100 ms with current technological solutions. The spatial resolution of the measurement is only about 1 mm because multiple scattered, diffuse photons are used (Yu et al., 2007).
12.2.4 Static methods The static methods (seen in Figure 12.1), measure the concentration of skin chromophores, such as oxygenated haemoglobin (HbO2) and deoxygenated haemoglobin (Hb). Often, the oxygen saturation level, SpO2, is also calculated using Equation (10.4). SpO2 ¼
½HbO2 : ½Hb + ½HbO2 0
(10.4)
When analysing the blood circulation, some of the static methods rely on the pulsation of the blood volume in the tissue due to the heartbeat and cyclic vasodilatation.
12.2.4.1 Reflectance pulse oximeter (RPO) The pulse oximeter has been part of standard medical care for decades. The operation principle of the pulse oximeter is as follows: The basic pulse oximeter consists of a clamp that has emitters on one side and receivers on the other side. Typical emitters are LEDs whose emission peak wavelengths are about Ired ¼ 660 nm (red) and INIR ¼ 940 nm (NIR). A standard photo diode can be used as a receiver, and may be attached to the other side of the clamp. The clamp is attached to some thin body part, such as a finger or an earlobe. The emitters illuminate the body part from the 0 other side, and the detector collects the light transmitted through the body part, Ired 0 0 0 and INIR. The intensities, Ired and INIR, depend on the concentrations of HbO2 and Hb. Hb absorbs red light relatively strongly, whereas HbO2 absorbs it weakly. The situation is the opposite in NIR range, where the absorption of HbO2 is stronger than the absorption of Hb. When these absorptions are known, the concentrations of HbO2,
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Hb and the oxygen saturation, SaO2, can be calculated when the intensities of trans0 0 and INIR , are measured. mitted red and NIR light, Ired The pulse oximeter also analyses the pulsation of the haemoglobin concentrations due to the cardiac cycle, and it causes the intensity of transmitted light to pulsate, too. It is assumed that the pulsation is caused by the changing volume of the arteries alone. The intensity of the transmitted light is at maximum during diastole and at minimum during systole. The difference between diastole and systole is then caused by the volume change of the arterial blood alone. By studying the systole and diastole differences in both red and NIR wavelengths, the pulse oximeter can estimate the oxygen saturation of arterial blood. In addition to the saturation, it usually measures the pulsation index (PI), too. PI is used for verifying that the sensor has good connection to the tissue, but it may also serve as an indicator of the strength of the blood perfusion (Alexander et al., 1989). The transmission pulse oximeter is not very useful for assessing skin flaps, but the oximeter can also work in reflectance configuration, where the emitter and the receiver are on the same side of the tissue. In reflection configuration, the emitter and the receiver are separated by a certain distance, d. The light from the emitter is guided to the skin, and part of it is scattered back from the tissue to the receiver. When the light beam from the emitter to the receiver travels in tissue, part of it will be absorbed due to the HbO2 and Hb chromophores in the same way as in transmission configuration (Mendelson et al., 1988). The pulse oximeter is one of the earliest applications of spectroscopy for clinical use. Many other static methods use similar techniques. However, the pulse oximeter itself may not be optimal for analysing the viability of skin flaps. First of all, the pulse oximeter operates only when the pulsation of the arterials is strong enough, which may not be the case in all skin flaps. Secondly, the SaO2 signal may not be the best indicator of problems in tissue blood circulation. The PI signal could be better, but its absolute value depends on the position of the sensor, and it is also sensitive to moving artefacts.
12.2.4.2 Photoplethysmography (PPG) The photoplethysmography (PPG) contains an emitter light source, typically LED, in the wavelength range of 600–900 nm and a receiver, such as a photodiode. The emitter illuminates the skin, and the receiver collects the light scattered out of the skin, the PPG signal, in either transmission or reflectance configuration in the same way as the pulse oximeter. The intensity of the received PPG signal is relational to the blood volume in the tissue, and it pulsates due to the cardiovascular cycle. The pulsation carries many kinds of information about the function of the cardiovascular system, the microcirculation in the tissue and the blood perfusion. Therefore, it might be useful in skin flap viability analysis (Challoner, 1979; Kamal et al., 1989; Allen, 2007). Some researchers have studied the use of PPG for skin flap vitality analysis (Bardach et al., 1979; Pickett et al., 1997; Chubb et al., 2012), and it has performed well in many studies. The PPG signal is relative, prone to moving artefacts and depends on the position of the sensor. These issues may be a hindrance for its widespread use in skin flap vitality assessment.
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Photoplethysmographic imaging (PPGI) is achieved using a camera instead of the photodiode as a receiver. The PPGI allows examination of the PPG signal from each pixel of the PPG image. The disadvantage of PPGI for skin flap assessment is that it works best in transmission configuration, which is difficult for skin flap measurements (Humphreys et al., 2005).
12.2.4.3 NIR spectroscopy (NIRS) NIR spectroscopy (NIRS) is a popular method for studying constitutive parameters of skin (Nystr€ om et al., 2003; Yu et al., 2007; Va¨lisuo et al., 2010). It has often been used for skin flap assessment, and compared with other methods (Payette et al., 2005; McKenna et al., 2009; Smit et al., 2010), NIRS is found to be a reliable, fast and cost-effective method for tissue viability analysis. The NIRS method is based on the analysis of the spectrum of light scattered out from the tissue, either in transmission or in reflectance configuration. A wideband light source is often used for NIRS, but multiple narrowband light sources are also used. Assume that the emission spectrum of the light from the light source is I0(l), where l is the wavelength. A biological tissue is illuminated with the light source I0(l). The light scattered out of the tissue has spectrum, I(l), which is guided to the spectrometer through a narrow slit. In the spectrometer, the light is spread into a spectrum using a graded mirror. The intensities of different wavelengths are then recorded by a sensor. The relative intensity of the received light, T(l) ¼ I(l)/I0(l), carries information about the concentrations of the chromophores in the skin. It can also be used as a PPG signal because it pulsates according to the cardiovascular cycle. The logarithm of the inverse of the transmission is said to be the absorption, A(l), of the tissue. See Equation (10.5). AðlÞ ¼ log e
l I ð lÞ : ¼ log e T ð lÞ I 0 ð lÞ
(10.5)
In case of no scattering, the relationships of the concentration of the chemical is directly relational to the absorption and the path length, l, of the photons in the tissue, according to the Beer–Lambert law (BLL), as shown in Equation (10.6) AðlÞ ¼ mðlÞl;
(10.6)
where m is the absorption coefficient of the tissue. In practice, the scattering in the tissue is strong, and the BLL does not hold, but the absorption, A(l), is nonlinearly dependent on the concentrations. Therefore, the analysis of the absolute values for HbO2, Hb and SO2 are not easy with NIRS. Often, the light propagation in skin is modelled using suitable skin models, such as the diffusion theory (Farrell et al., 1992), the Monte Carlo simulation (Prahl et al., 1989) or the Kubelka–Munk (Kubelka and Munk, 1931) model, and the model parameters are found by nonlinear optimisation methods. The measurement depth of the NIRS is dependent on the probe geometry (Va¨lisuo and Alander, 2008). The rule of thumb is that the larger is the spacing between the
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emitter and the detector, the deeper does the light travel before reaching the detector. When the spacing is increased, the intensity of the received light is reduced almost exponentially. Therefore, large emitter–detector spacing is often impractical. If the separation of the emitter and detector is approximately the same as the diameter of the emitter and detector fibres, the received signal is only slightly dependent on the scattering coefficient of the tissue. The optical probe can be further designed so that the BLL can be used by only adding a constant correction factor for the path length, l. This kind of probe was demonstrated and called differential path-length spectroscopy (Amelink and Sterenborg, 2004). NIRS seems to be a reliable and tested method for analysing tissue properties, such as HbO2, Hb concentrations and SpO2. NIRS is a flexible method; the sampling depth and many other parameters can be adjusted by designing suitable probes. The PPG signal is obtained as a side effect. It is also possible to apply NIRS invasively by implanting the emitter and receiver fibres, or a probe, inside the tissue. The NIRS method is potentially prone to the moving artefacts and bending of the measurement fibres.
12.2.4.4 Diffuse optical imaging (DOI) In diffuse optical imaging (DOI), an area of a tissue is illuminated with smooth and preferably diffuse light. The reflected light is recorded by using a CCD camera. The imaging can be said to be diffuse if the multiple scattered diffuse photons play an important role and the sampling depth exceeds the diffusion limit. The Fresnel reflection from the surface of the skin and single scattered photons in the point of illumination may considerably conceal the diffuse reflectance signal under the surface without providing any information about the parameters of the tissue. The Fresnel reflection and part of the single scattered photons can be filtered out using two polarising filters. One of the filters is assembled in front of the light source to make the incident light polarised. The other filter is situated orthogonally in front of the objective of the camera to filter out Fresnel reflection, which has kept its polarisation angle. The diffuse reflected signal below the surface of the skin is mostly multiple scattered, and the light loses its polarisation angle in several scattering events. Therefore, the Fresnel reflection and single scattered photons are mostly filtered out, and only the multiple scattered, diffuse reflected photons are allowed to the camera (Olivier et al., 2003; O’Doherty et al., 2007). This method is often called orthogonal polarisation imaging. DOI can use either one wavelength—standard RGB filters to acquire a three-band image—or include components that allow capturing of multi-spectral or hyperspectral image of skin. Already the standard RGB image contains a lot of spectral information, which can be analysed using methods similar to those used in NIRS. The disadvantages of using the RGB image are that the exact filter specifications of the camera are not usually available, only red components reside in the diagnostic– therapeutic window, blue and green do not penetrate deeply into the skin, and the number of channels is too small for many purposes. A multi-spectral image can be obtained, for example, using several narrowband LEDs with different wavelengths as illumination. The different LEDs with different peak wavelengths are used one by one, and one picture is taken in each wavelength.
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The captured images can be combined as one multi-spectral image. Hyper-spectral images can be taken using a wideband white light source and a special spectral camera. Several methods using DOI are: l
l
l
SIAscopy: Spectral Intracutaneous Analysis is a method wherein the skin is imaged using either RGB colours or RGB colours with some NIR channels. The images are analysed to produce a 2D map of HbO2, Hb, SpO2 and melanin concentrations. The method is based on the Kubelka–Munk skin model and solves the nonlinear skin model using precalculated look-up tables. The main purpose for SIAscopy is to diagnose skin cancer, but it may work for skin viability analysis as well (Claridge et al., 2003; Govindan et al., 2007). Spectrocutometry: The Spectrocutometry uses a linear Beer–Lamber model of the epidermis and a non-linear diffusion theoretical model of the dermis, which is linearised around an operating point to allow fast solution of the skin model at every pixel location. Spectrocutometry is used for calculating the HbO2, Hb, SpO2 and melanin fraction over the imaged area (Kaartinen et al., 2011; Va¨lisuo et al., 2011). TiVi: The purpose of the TiVi imaging is to estimate the concentration of RBCs over the imaged area. The method is based on linear equations with a fixing coefficient (McNamara et al., 2010).
DOI is used in many different forms. All of the methods are cost-effective, highresolution and moderately fast. The nonlinear relationship between the skin reflectance and the chromophore concentration makes it difficult to obtain absolute concentrations. However, SIAscopy and spectrocutometer aims to the absolute concentrations. SIAscopy has been on the market for several years, and there are many clinical articles about TiVi imaging. Spectrocutometer is emerging in future markets. All DOI instruments are potentially suitable for skin flap vitality assessments. The measurement depth and the lack of off-the-shelf clinical instruments may set some limitations to the application of these methods.
12.2.4.5 Diffuse optical tomography (DOT) Diffuse optical tomography (DOT) is an imaging mode that can produce a spectral image of an object residing several centimetres below the biological tissue. The illumination is handled using and array of light sources, and detection is handled using an array of detectors surrounding the imaged object. Each detector records the remitted light while one light source is illuminating the target. This procedure is repeated for each light source (Wang and Wu, 2007). Deep imaging depth can be achieved by utilising multiple scattered diffuse photons and NIR light within the diagnostic–therapeutic window. An optical tissue model, often based on diffusion theory, is used to infer localised tissue parameters. The DOT imaging system can work in three different modes: time-domain, frequency domain or DC-current mode. In time-domain mode, the light sources emit picosecond pulses, which are broadened by the tissue and observed by the detectors. In frequency mode, the light sources emit an amplitude modulated light signal that is observed by the detectors with smaller modulation depth, due to the multiple scattering in the tissue. In DC mode, the light sources emit only slowly time-invariant signals, which are again recorded by the detectors.
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Because the diffuse photons have lost their directional information, the axial and lateral resolution of DOT is poor—approximately 20% of the imaging depth. The acquisition time of a DOT image is short. If the light emitter array is tuneable or duplicated with several different wavelengths, a multi-spectral DOT image can be achieved that can be used in resolving HbO2, HbO concentrations and oxygen saturation SpO2. These parameters can be used to study the metabolism of deep skin structures, such as tumours (McBride et al., 1999). The spatial resolution provided by the DOT method is probably enough for skin flap assessment, and the fast acquisition rate is useful in analysing the blood circulation dynamics. The overall properties of the DOT method should be well suited for skin flap analysis.
12.2.5 Fluorescence methods In some cases, it is beneficial to assess blood perfusion using external contrast agents, such as ICG. ICG is a well-tested fluorescent dye that has been in use for decades. ICG can be excited with NIR light in the wavelength range between 600 and 800 nm, and it emits in the range of 800–950 nm. Both excitation and emission wavelengths are within the diagnostic–therapeutic window, which makes ICG suitable for imaging deep blood flow. ICG has often been used in neurosurgery to verify the blood flow in arteries after an operation. It has also been used in assessing the blood flow in muscles of the lower extremity as well as for many other purposes (Alander et al., 2012). ICG could be used for skin flap assessment by administering the ICG into the blood circulation through a vein. After a few seconds, the ICG enters all over the body, including the skin flap, if the blood circulation works properly. The liver removes the ICG from blood circulation within a few minutes. The ICG enters first in the arterials of the tissue, then in the capillaries and last in the veins. The dynamics of the ICG concentration in various parts of the skin flap and nearby normal skin can be measured and compared using fluorescent NIRS, absorption NIRS or fluorescent imaging. ICG assessment is an invasive method, and due to the weak fluorescent signal, very sensitive NIRS or DOI equipment with high-quality excitation and a fluorescence filter pair, are required. Still, equipment costs are quite moderate (Alander et al., 2012).
12.2.6 Summary Optical methods are usually relatively low cost, fast and offer high resolution. The optical measurements are available in many different modalities as reviewed in this chapter. The key properties of these modalities are shown in Table 12.1.
12.3
Future trends
The research on the optical methods for sampling skin parameters is actively continuing together with improvements in optical technologies. Many methods that are now under investigation will be standardised in the near future and will be available for
The summary of the key properties of optical methods available for skin flap assessment
Method
Dyn/static
Mode
Contrast
Depth (mm)
Resolution
Time
Availability
Cost
LDI LDF DOCT LSI DCS RPO PPG NIRS DOI DOT ICG
Dyn Dyn Dyn Dyn Dyn Static Static Static Static Static Static
Image Point 3D image Image Point Point Point Point Image 3D image Image
Doppler Doppler Doppler Doppler Doppler SpO2, PI Blood volume, PI HbO2, Hb, SpO2 HbO2, Hb, SpO2 HbO2, Hb, SpO2 CICG
0.1 0.1–5 0.1 0.1 >10 10 1 to >10 1 to >10 1 >10 1–10
Moderate – High High – – – – High Low Moderate
Minutes Real time Slow Fast
Available Available Available Available Emerging Available Available Available Available Emerging Available
Moderate Moderate High Low – Low Low Moderate Low Moderate Moderate
Real time Real time Real time Fast Slow Fast
Optical methods for assessing skin flap survival
Table 12.1
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clinical use. Because optical methods are relatively low cost, they will be integrated into other existing measurement systems. Also, combinational optical devices are likely to emerge. Currently, there are pulse oximeters with optoplethysmograpy function. LSI can easily be combined with DOI to provide information on the concentration of blood, oxygen saturation and blood perfusion. PPG imaging and DOI together could provide similar benefits for skin flap vitality assessment.
12.4
Sources of further information
The Biomedical Optics: Principles and Imaging book (Wang and Wu, 2007), describes the theories behind many optical measurement modalities in detail. This book is excellent for the technically oriented reader. The Tissue Optics book (Tuchin, 2007) explains the theories behind skin optics and contains many tables listing the optical properties of different biological tissue types. Introduction to the light interaction with skin (Baranoski and Krishnaswamy, 2008) is a good introduction to the theories behind various optical measurements. Handbook of Biomedical Optics (Boas et al., 2011) provides a recent view to the theories behind optical measurements of biological tissues. The coverage of various theories behind biomedical optics is excellent.
References Alander, J.T., Kaartinen, I., Laakso, A., Pa¨tila¨, T., Spillmann, T., Tuchin, V.V., Venermo, M., Va¨lisuo, P., 2012. A review of indocyanine green fluorescent imaging in surgery. Int. J. Biomed. Imaging 2012, 1–26. Alexander, C.M., Teller, L.E., Gross, J.B., 1989. Principles of pulse oximetry theoretical and practical considerations. Anesth. Analg. 68, 368–376. Allen, J., 2007. Photoplethysmography and its application in clinical physiological measurement. Physiol. Meas. 28, R1–R39. Amelink, A., Sterenborg, H.J.C.M., 2004. Measurement of the local optical properties of turbid media by differential path-length spectroscopy. Appl. Opt. 43, 3048–3054. Baranoski, G.V.G., Krishnaswamy, A., 2008. Light interaction with human skin: from believable images to predictable models. In: ACM SIGGRAPH ASIA 2008 Courses, SIGGRAPH Asia’08. ACM, New York, NY, pp. 10:1–10:80. Bardach, J., Voots, R.J., McCabe, B.F., Hsu, M.M., 1979. Photoplethysmography in the prediction of experimental flap survival. Ann. Otol. Rhinol. Laryngol. 88, 637–641. Boas, D.A., Yodh, A.G., 1997. Spatially varying dynamical properties of turbid media probed with diffusing temporal light correlation. J. Opt. Soc. Am. A 14, 192–215. Boas, D.A., Campbell, L.E., Yodh, A.G., 1995. Scattering and imaging with diffusing temporal field correlations. Phys. Rev. Lett. 75, 1855–1858. Boas, D.A., Pitris, C., Ramanujam, N., 2011. Handbook of Biomedical Optics. CRC Press, Boca Raton. Briers, J.D., Fercher, A.F., 1982. Retinal blood-flow visualization by means of laser speckle photography. Invest. Ophthalmol. Vis. Sci. 22, 255–259. Challoner, A., 1979. Photoelectric plethysmography for estimating cutaneous blood flow. Noninvasive Physiol. Meas. 1, 125–151.
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Chen, Z., Milner, T.E., Wang, X., Srinivas, S., Nelson, J.S., 1998. Optical Doppler tomography: imaging in vivo blood flow dynamics following pharmacological intervention and photodynamic therapy. Photochem. Photobiol. 67, 56–60. Choi, B., Kang, N.M., Nelson, J.S., 2004. Laser speckle imaging for monitoring blood flow dynamics in the in vivo rodent dorsal skin fold model. Microvasc. Res. 68, 143–146. Chubb, D., Whitaker, I.S., Rozen, W.M., Ashton, M.W., 2012. Continued observations in the postoperative monitoring of free flaps. Plast. Reconstr. Surg. 129, 222e–223e. Claridge, E., Cotton, S., Hall, P., Moncrieff, M., 2003. From colour to tissue histology: physics-based interpretation of images of pigmented skin lesions. Med. Image Anal. 7, 489–502. Dunn, A.K., Bolay, H., Moskowitz, M.A., Boas, D.A., 2001. Dynamic imaging of cerebral blood flow using laser speckle. J. Cereb. Blood Flow Metab. 21, 195–201. Farrell, T.J., Patterson, M.S., Wilson, B., 1992. A diffusion theory model of spatially resolved, steady-state diffuse reflectance for the noninvasive determination of tissue optical properties in vivo. Med. Phys. 19, 879. Forrester, K.R., Tulip, J., Leonard, C., Stewart, C., Bray, R.C., 2004. A laser speckle imaging technique for measuring tissue perfusion. IEEE Trans. Biomed. Eng. 51, 2074–2084. Govindan, K., Smith, J., Knowles, L., Harvey, A., Townsend, P., Kenealy, J., 2007. Assessment of nurse-led screening of pigmented lesions using SIAscope. J. Plast. Reconstr. Aesthet. Surg. 60, 639–645. H€ olzle, F., Rau, A., Swaid, S., Loeffelbein, D.J., Nolte, D., Wolff, K.-D., 2005. Simultaneous noninvasive monitoring for radial forearm and fibula flaps using laser Doppler flowmetry and tissue spectrophotometry. Mund Kiefer Gesichtschir. 9, 290–299. Humphreys, K., Markham, C., Ward, T., 2005. A CMOS camera-based system for clinical photoplethysmographic applications [WWW Document]. Proc. SPIE 5823, 88–95. http://spie.org/x576.xml. Izatt, J.A., Kulkarni, M.D., Yazdanfar, S., Barton, J.K., Welch, A.J., 1997. In vivo bidirectional color Doppler flow imaging of picoliter blood volumes using optical coherence tomography. Opt. Lett. 22, 1439–1441. Kaartinen, I.S., Va¨lisuo, P.O., Alander, J.T., Kuokkanen, H.O., 2011. Objective scar assessment—a new method using standardized digital imaging and spectral modelling. Burns 37, 74–81. Kamal, A.A.R., Harness, J.B., Irving, G., Mearns, A.J., 1989. Skin photoplethysmography—a review. Comput. Methods Prog. Biomed. 28, 257–269. Kubelka, P., Munk, F., 1931. An article on optics of paint layers. Z. Tech. Phys. 12, 593–601. McBride, T.O., Pogue, B.W., Gerety, E.D., Poplack, S.B., Osterberg, U.L., Paulsen, K.D., 1999. Spectroscopic diffuse optical tomography for the quantitative assessment of hemoglobin concentration and oxygen saturation in breast tissue. Appl. Opt. 38, 5480–5490. McKenna, J., Pabbies, A., Friesen, J.R., Sowa, M.G., Hayakawa, T., Kerr, P.D., 2009. Assessing flap perfusion: optical spectroscopy versus venous Doppler ultrasonography. J. Otolaryngol. Head Neck Surg. 38, 587–594. McNamara, P.M., O’Doherty, J., O’Connell, M.-L., Fitzgerald, B.W., Anderson, C.D., Nilsson, G.E., Toll, R., Leahy, M.J., 2010. Tissue viability (TiVi) imaging: temporal effects of local occlusion studies in the volar forearm. J. Biophotonics 3, 66–74. Mendelson, Y., Kent, J.C., Yocum, B.L., Birle, M.J., 1988. Design and evaluation of a new reflectance pulse oximeter sensor. Med. Instrum. 22, 167–173. Nystr€om, J., Lindholm-Sethson, B., Stenberg, L., Ollmar, S., Eriksson, J., Geladi, P., 2003. Combined near-infrared spectroscopy and multifrequency bio-impedance investigation of skin alterations in diabetes patients based on multivariate analyses. Med. Biol. Eng. Comput. 41, 324–329.
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O’Doherty, J., Henricson, J., Anderson, C., Leahy, M.J., Nilsson, G.E., Sj€ oberg, F., 2007. Subepidermal imaging using polarized light spectroscopy for assessment of skin microcirculation. Skin Res. Technol. 13, 472–484. Olivier, W.-A.M., Hazen, A., Levine, J.P., Soltanian, H., Chung, S., Gurtner, G.C., 2003. Reliable assessment of skin flap viability using orthogonal polarization imaging. Plast. Reconstr. Surg. 112, 547–555. Payette, J.R., Kohlenberg, E., Leonardi, L., Pabbies, A., Kerr, P., Liu, K.-Z., Sowa, M.G., 2005. Assessment of skin flaps using optically based methods for measuring blood flow and oxygenation. Plast. Reconstr. Surg. 115, 539–546. Pickett, J., Amoroso, P., Nield, D.V., Jones, D.P., 1997. Pulse oximetry and PPG measurements in plastic surgery. In: Proceedings of the 19th Annual International Conference of the IEEE Engineering in Medicine and Biology Society, vol. 5, pp. 2330–2332. Prahl, S.A., Keijzer, M., Jacques, S.L., 1989. A Monte Carlo model of light propagation in tissue. In: Muller, G.J., Sliney, D.H. (Eds.), SPIE Proceedings of Dosimetry of Laser Radiation in Medicine and Biology, pp. 102–111. Riva, C., Ross, B., Benedek, G.B., 1972. Laser Doppler measurements of blood flow in capillary tubes and retinal arteries. Invest. Ophthalmol. Vis. Sci. 11, 936–944. Salman, M., Glantzounis, G.K., Yang, W., Myint, F., Hamilton, G., Seifalian, A.M., 2005. Measurement of critical lower limb tissue hypoxia by coupling chemical and optical techniques. Clin. Sci. 108, 159–165. Serov, A., Steenbergen, W., de Mul, F., 2002. Laser Doppler perfusion imaging with a complimentary metal oxide semiconductor image sensor. Opt. Lett. 27, 300–302. Smit, J.M., Zeebregts, C.J., Acosta, R., Werker, P.M.N., 2010. Advancements in free flap monitoring in the last decade: a critical review. Plast. Reconstr. Surg. 125, 177–185. Stern, M.D., 1975. In vivo evaluation of microcirculation by coherent light scattering. Nature 254, 56–58. http://dx.doi.org/10.1038/254056a0. Stewart, C.J., Frank, R., Forrester, K.R., Tulip, J., Lindsay, R., Bray, R.C., 2005. A comparison of two laser-based methods for determination of burn scar perfusion: laser Doppler versus laser speckle imaging. Burns 31, 744–752. Tuchin, V.V., Society of Photo-optical Instrumentation Engineers, 2007. Tissue Optics Light Scattering Methods and Instruments for Medical Diagnosis. SPIE/International Society for Optical Engineering, Bellingham, WA. Va¨lisuo, P., Alander, J., 2008. The effect of the shape and location of the light source in diffuse reflectance measurements. In: IEEE Symposium on Computer-Based Medical Systems. IEEE Computer Society, Los Alamitos, CA, pp. 81–86. Va¨lisuo, P., Kaartinen, I., Kuokkanen, H., Alander, J., 2010. The colour of blood in skin: a comparison of Allen’s test and photonics simulations. Skin Res. Technol. 16, 390–396. Va¨lisuo, P., Kaartinen, I., Tuchin, V., Alander, J., 2011. New closed-form approximation for skin chromophore mapping. J. Biomed. Opt. 16, 046012. Wang, L.V., Wu, H., 2007. Biomedical Optics: Principles and Imaging. Wiley-Interscience, Hoboken, NJ. Wa˚rdell, K., Jakobsson, A., Nilsson, G.E., 1993. Laser Doppler perfusion imaging by dynamic light scattering. IEEE Trans. Biomed. Eng. 40, 309–316. Yu, G., Floyd, T.F., Durduran, T., Zhou, C., Wang, J., Detre, J.A., Yodh, A.G., 2007. Validation of diffuse correlation spectroscopy for muscle blood flow with concurrent arterial spin labeled perfusion MRI. Opt. Express 15, 1064–1075.
12
P. Va¨lisuo University of Vaasa, Vaasa, Finland
12.1
Introduction
The skin flap is a full thickness mass of skin, transplanted from a donor site to a recipient site. The blood circulation of the skin flap must be guaranteed to keep the skin flap alive. Some of the original blood vessels can be kept intact for local skin flaps, whereas a microsurgery is required to establish blood circulation for free skin flaps. Venous congestion and arterial occlusion are typical reasons for the failures of skin flaps (H€ olzle et al., 2005). Venous congestion leads to increase of blood fraction and haemoglobin concentration in tissue, whereas the declining blood flow and oxygen saturation are symptoms of arterial occlusion. If these vascular compromises are detected early enough, they can be operatively explored and fixed saving the skin flap (H€olzle et al., 2005). Therefore, skin flap vitality monitoring is very important during 24 h after the transplantation. The blood perfusion, blood flow and oxygen saturation are vital signals predicting the success of the operation. Smit et al. (2010) reviewed the most often used methods for skin flap assessment during 1999–2009. They reviewed implantable Doppler system, colour duplex ultrasonography, near-infrared spectroscopy (NIRS), microdialysis and laser Doppler flowmetry (LDF). The group concluded that Doppler systems and NIRS are the best methods available today. Payette et al. (2005) compared optical spectroscopy and laser Doppler assessment and concluded that the optical spectroscopy and spectroscopic imaging may have a lot of potential in evaluating skin flaps. Skin is a thin organ, easily accessible using non-invasive optical probes. Optical methods provide real-time information from a single point or over an area of skin. They can be used for measuring the blood perfusion, flow and oxygen saturation. This chapter reviews existing optical skin parameter assessment methods suitable for skin flap follow-up as well as some new methods that are not in regular clinical use yet.
12.2
Methods
There are many methods that can be used for measuring constituent skin parameters, including those vital signals predicting the success of the skin flap. Optical methods are popular because the skin is thin and easily accessible by non-invasive optical probes. The wavelengths approximately between red (600 nm) and near infrared (NIR) below Biophotonics for Medical Applications. http://dx.doi.org/10.1016/B978-0-85709-662-3.00012-9 Copyright © 2015 Elsevier Ltd. All rights reserved.
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950 nm, reside between the absorption peaks of haemoglobin and water and form the socalled diagnostic–therapeutic window (Tuchin, 2007), in which the light penetrates deeper into the skin than the other wavelengths. Therefore, many optical skin assessment methods operate in visual and NIR wavelength range. The optical methods are usually non-invasive, high-resolution, real time and relatively inexpensive. The existing methods are the results of research activities carried out for decades. Many useful clinical solutions have appeared, and there are dozens of interesting methods to analyse the fundamental parameters of the skin. Such as the pulse oximeter measures the oxygen saturation, LDF can measure the blood flow. However, the light interaction in skin is still too complex for a single method that would fit for all purposes. Some often used non-optical methods are, transcutaneous oxygen partial pressure (PO2) meter, Doppler ultrasonography, magnetic resonance imaging (MRI) and thermal imaging (Salman et al., 2005). The Doppler ultrasonography is only suitable for measuring the blood flow of large vessels. MRI is expensive and difficult to apply for most cases, but the transcutaneous PO2 meter is still in active use. While these nonoptical methods are good for many purposes, researchers are searching for optical methods that outperform them. Existing and emerging optical methods are often cost-effective, non-invasive, fast and reliable.
12.2.1 General properties behind optical measurements Optical probes consist of an emitter and a receiver or a matrix of receivers. The light from the emitter enters into the skin and scatters back to the receiver, carrying information about the skin, which can be analysed. The probe geometry determines the sampling depth and the coverage area of the measurement (see Figure 12.1). If the probe contains only one receiver, the measurement provides a single value at a time, representing an average of the probed area. If the receiver contains an array of sensors, it can produce an image over the area examined, where each value represents an average over a subset of the whole image. Some pointwise methods can be extended to imaging systems by mechanically scanning over a skin area. Medical imaging and pointwise measurements reveal information about the structure and the chemical composition of the tissue. The contrast in the measurement is achieved by one of the following five methods: (1) by structural changes in the tissue, (2) by changes in chemical concentrations in the tissue, (3) by the Doppler shift due to the moving particles in the blood circulation, (4) by observing changes in the laser speckle patterns due to moving particles in blood circulation and (5) by using an external contrast agent, such as indocyanine green (ICG). The methods, whose contrast is based on the movement of the particles, are called dynamic methods as opposed to chemical concentration and structure-based methods, which are called as static. When the target of the measurement is to follow the functions of the organ, the method is said to be functional. Optical functional measurements can be based on either dynamic or static methods. Usually functional measurements are made indirectly by observing changes in blood circulation or concentrations of nutrients and by-products of metabolic processes, such as cellular respiration. Active metabolism decreases the concentrations of nutrients and increases the concentrations of the by-products. Cellular
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d2
d1
Skin
(a)
(b)
(c)
Figure 12.1 The illumination geometries. In panels (a) and (b) the light is directed to the skin through an optical fibre, and the reflected light is detected with another fibre. The distance between the fibres affects the sampling depth of the probe. If the emitter and detector fibres are close to each other, the reflectance from the shallow parts of skin is increased. When the distance is larger in panel (a), the average sampling depth is deeper. Panel (c) shows a situation in which the skin is illuminated with planar light source, and the reflectance is collected from a point. This corresponds to the case in panel (b), where the emitter and detector are close to each other and the average sampling depth is shallow. The dashed horizontal line shows the diffusion limit, below which the measurement is challenging.
activity also typically affects local blood regulation by increasing blood perfusion. Functional measurements and functional imaging can be used in assessing the vitality of the skin flaps. As mentioned earlier, the blood perfusion and oxygen saturation are often important signals used to describe the skin flap vitality. Blood perfusion can be monitored by measuring the speed of red blood cells (RBCs) (dynamic method) or by measuring the blood concentration and oxygen saturation (static method). The speed of the RBCs and thus the speed of the blood flow is most often measured using Doppler shift of reflected light. The static methods are often based on the analysis of the reflectance spectrum when skin is illuminated with a wideband light source. The chromophores carried by blood change the reflectance spectrum and thus the blood circulation can be analysed. The pulsation of the arteries cyclically changes the blood volume in skin, which can be observed from optical transmission or reflectance signal. The pulsation itself is a good indicator of blood perfusion, and it can be also used for separating the absorption signal originating from arterial blood from other absorption sources. Furthermore, the metabolism of the cells can be indirectly observed by analysing the concentrations of nutrients and resulting chemicals using optical spectroscopy. However, the strong scattering of light in skin often makes the application of optical methods challenging. Most of the photons in the reflectance spectra are multiple scattered and have lost all of their direction information, and the length of the path the photons travelled in the skin is unknown. For these reasons, the change in chemical concentration changes the reflected light intensity nonlinearly, and the effect in reflectance spectrum of one chromophore is dependent on the concentrations of each other
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chemical in skin. Therefore, the spectrum analysis is often time-consuming and dependent on specific skin models. Some methods rely only on directly transmitted or single scattered ballistic photons, gating away the multiple scattered photons. These techniques provide sharp images and easy methods for chromophore concentration analysis. The disadvantages are that most of the light is lost in gating, and the proportion of ballistic photons is reduced exponentially when the depth is increased. Therefore, ballistic methods are best suited for superficial analysis of relatively translucent areas.
12.2.2 Diffusion limit When an orthogonally incident, collimated light beam, whose intensity is, I0, travels through a purely absorbing slab, whose depth is, d [cm], its intensity after transmission, I, is described according to the following equation: I ¼ I0 ema d ;
(10.1)
where ma [1/cm] is the absorption coefficient, which is the reciprocal of the average travelling distance of a photon between the absorption events. Similarly, if the material is purely scattering, the intensity of non-scattered light transmitted directly through the slab is given by the following equation: 0
I ¼ I0 ems d ;
(10.2)
where ms0 [1/cm] is the reduced scattering coefficient, which is a reciprocal of the mean free path between scattering events in the corresponding isotropic medium. The effect of both absorption and scattering can be accounted using the total interaction coefficient: mt ¼ ma + m0s . The reciprocal of the total interaction coefficient is the mean free path length between either scattering or absorption events: pmf ¼
1 1 ¼ : mt ma + m0s
(10.3)
When the thickness of the slab is increased, until it equals the mean free path length (d ¼ pmf), the intensity of the transmitted light I ¼ I0/e, and only 37% of the light passes directly through the slab. This thickness is said to be the diffusion limit. Superficial measurements and imaging above the diffusion limit are relatively easy, whereas more advanced techniques are needed to sample tissue below pmf. For human skin, the pmf is roughly about 1 mm (Farrell et al., 1992).
12.2.3 Dynamic methods In dynamic methods, the measurement signal is caused by Doppler shift of the photons scattered from the moving particles, as shown in Figure 12.2. Because the signal is relational to the speed of the blood flow in dynamic methods, they are especially applicable for blood perfusion measurements.
Optical methods for assessing skin flap survival
R–
R+
R
v Skin
(a)
(b)
335
Figure 12.2 The light reflectance from the skin with capillary loops. Panel (a) shows reflectance due to the scattering of moving RBCs. The RBCs moving upwards create a positive Doppler shift to the reflected signal R+, whereas the RBCs moving downwards cause corresponding negative Doppler shift to reflectance signal R. Panel (b) shows reflectance R due to scattering from tissue. The absorption of blood and tissue affects to the reflectance spectrum of R.
12.2.3.1 Laser Doppler flowmetry (LDF) The LDF was developed in the late 1970s for measuring the blood flow of tissue (Riva et al., 1972; Stern, 1975). Later, LDF established its position in blood perfusion monitoring. In LDF, the skin is illuminated by a coherent, collimated laser beam, using an optical fibre. The light reflected back from the tissue is collected by another fibre(s). Moving particles, such as RBCs, cause a Doppler shift that broadens the spectrum of the reflected light. This broadening of the spectrum can be used to calculate the speed and concentration of the RBCs. LDF is a proven, continuous, non-invasive and realtime method with many clinical instruments available. LDF is often used for skin flap assessment. The speed and the concentration of the RBCs in capillaries provide useful information about the function of the blood vessels after the surgery. Use of different probes makes the LDF method suitable for different situations. Typical probes are: surface probes, needle probes and single fibre probes, which can be inserted under the skin using a standard cannula. The disadvantages of LDF measurement are the small spatial region covered and the possible artefacts due to bending of the probe fibres and due to the pressure applied toward the skin by the probe. LDF measurement reaches only a few hundreds of micrometers below skin surface (Riva et al., 1972; Stern, 1975).
12.2.3.2 Laser Doppler imaging (LDI) The LDF measurements cover relatively small areas of skin. However, it is well known that the blood perfusion is spatially rather heterogeneous, and the LDF measurement sample covers only a small volume of tissue and depends, therefore, on the position of the sensor. Laser Doppler imaging (LDI) overcomes these issues by measuring the flow speed from many different locations over the imaged area. The scanning over a spatial area can be implemented mechanically by the use of rotating mirrors and stepper motors. The sampling depth is similar to that of LDF, about 200–240 mm (Wa˚rdell et al., 1993; Serov et al., 2002). LDI is a popular method for assessing skin flaps. It can provide information about the function of the blood circulation over a large area or the whole skin flap. LDI does
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not require the probe to be in contact with skin, which removes possible disinfecting problems. The disadvantages of LDI are rather moderate resolution, small penetration depth and slow imaging process.
12.2.3.3 Doppler optical coherence tomography (DOCT) Optical coherence tomographic (OCT) system uses a coherent laser beam to illuminate target skin area and a pointwise sensor to collect the reflectance. An interferometer is used to include only single scattered photons reflected from the specific depth. OCT systems can be used for scanning the tissue properties laterally over a skin area and axially over different depths, up to the diffusion limit. OCT can therefore provide a high-resolution profile of the scanned skin area. It is also possible to use a twodimensional charge coupled device (CCD)-sensor as a receiver to avoid lateral scanning altogether. This kind of full-field OCT has already been demonstrated, but it may not be available for clinical use in the near future. The most popular target for OCT is ophthalmology, studying the properties of retina since it is more translucent than many other tissues (Izatt et al., 1997; Chen et al., 1998). Typically, OCT uses only single wavelength within the diagnostic–therapeutic window to scan as deep as possible. Because only one wavelength is used, spectroscopic analysis cannot be made, but a picture of the skin structure can be achieved. OCT, as such, is not very useful for skin flap imaging, but Doppler OCT (DOCT) can measure the velocity of the scatterers using the Doppler shift of the reflected light together with structural scanning of the tissue. In that case, both the high-resolution structural image and the estimate of blood perfusion can be achieved. The disadvantages of OCT and DOCT are that they are rather slow to achieve data because the imaged area needs to be scanned both laterally and axially; therefore, high temporal resolution is not achieved. The image is a very high-resolution image, but only an average image over the area may be needed for skin flap assessment. DOCT can only obtain information above the diffusion limit.
12.2.3.4 Laser speckle imaging (LSI) and laser speckle correlation analysis Laser speckle imaging (LSI) was first used in the 1980s. It has since been used for imaging the cerebral blood flow (Briers and Fercher, 1982; Dunn et al., 2001). When tissue is illuminated with coherent laser light, the reflections from different parts of the uneven tissue surface reach the detector at different times, forming interference artefacts, called speckle patterns or speckle noise. When the light is reflected from moving particles, such as RBCs in the blood, the speckle pattern starts to fluctuate. Higher RBC velocity produces higher local variance to the speckle pattern. This phenomenon can be used for quantifying blood perfusion under the skin. The LSI has high lateral resolution (tens of micrometres), high temporal resolution (millisecond) and low cost. The measurement system can be constructed easily using standard components; only a
Optical methods for assessing skin flap survival
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standard laser and normal CCD camera are needed. The disadvantage is shallow measurement depth (Choi et al., 2004; Forrester et al., 2004). LSI may be a better method for assessing the validity of skin flaps than the LDI method (Stewart et al., 2005).
12.2.3.5 Diffuse correlation spectroscopy (DCS) Diffuse correlation spectroscopy (DCS) is an emerging technology for blood perfusion measurement of deep tissues. DCS was demonstrated in 1995 (Boas et al., 1995). It can directly observe the interference of moving scatterers, such as RBCs, and therefore the blood perfusion. DCS works similarly to LDF, but it can also handle multiple scattered light. The mathematical basis of the method is correlation transport theory, which is modified from the radiative transport theory to estimate correlation transfer through a turbid medium (Boas and Yodh, 1997). The measurement reaches more than 10 mm below the skin surface, and the acquisition time is approximately 100 ms with current technological solutions. The spatial resolution of the measurement is only about 1 mm because multiple scattered, diffuse photons are used (Yu et al., 2007).
12.2.4 Static methods The static methods (seen in Figure 12.1), measure the concentration of skin chromophores, such as oxygenated haemoglobin (HbO2) and deoxygenated haemoglobin (Hb). Often, the oxygen saturation level, SpO2, is also calculated using Equation (10.4). SpO2 ¼
½HbO2 : ½Hb + ½HbO2 0
(10.4)
When analysing the blood circulation, some of the static methods rely on the pulsation of the blood volume in the tissue due to the heartbeat and cyclic vasodilatation.
12.2.4.1 Reflectance pulse oximeter (RPO) The pulse oximeter has been part of standard medical care for decades. The operation principle of the pulse oximeter is as follows: The basic pulse oximeter consists of a clamp that has emitters on one side and receivers on the other side. Typical emitters are LEDs whose emission peak wavelengths are about Ired ¼ 660 nm (red) and INIR ¼ 940 nm (NIR). A standard photo diode can be used as a receiver, and may be attached to the other side of the clamp. The clamp is attached to some thin body part, such as a finger or an earlobe. The emitters illuminate the body part from the 0 other side, and the detector collects the light transmitted through the body part, Ired 0 0 0 and INIR. The intensities, Ired and INIR, depend on the concentrations of HbO2 and Hb. Hb absorbs red light relatively strongly, whereas HbO2 absorbs it weakly. The situation is the opposite in NIR range, where the absorption of HbO2 is stronger than the absorption of Hb. When these absorptions are known, the concentrations of HbO2,
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Hb and the oxygen saturation, SaO2, can be calculated when the intensities of trans0 0 and INIR , are measured. mitted red and NIR light, Ired The pulse oximeter also analyses the pulsation of the haemoglobin concentrations due to the cardiac cycle, and it causes the intensity of transmitted light to pulsate, too. It is assumed that the pulsation is caused by the changing volume of the arteries alone. The intensity of the transmitted light is at maximum during diastole and at minimum during systole. The difference between diastole and systole is then caused by the volume change of the arterial blood alone. By studying the systole and diastole differences in both red and NIR wavelengths, the pulse oximeter can estimate the oxygen saturation of arterial blood. In addition to the saturation, it usually measures the pulsation index (PI), too. PI is used for verifying that the sensor has good connection to the tissue, but it may also serve as an indicator of the strength of the blood perfusion (Alexander et al., 1989). The transmission pulse oximeter is not very useful for assessing skin flaps, but the oximeter can also work in reflectance configuration, where the emitter and the receiver are on the same side of the tissue. In reflection configuration, the emitter and the receiver are separated by a certain distance, d. The light from the emitter is guided to the skin, and part of it is scattered back from the tissue to the receiver. When the light beam from the emitter to the receiver travels in tissue, part of it will be absorbed due to the HbO2 and Hb chromophores in the same way as in transmission configuration (Mendelson et al., 1988). The pulse oximeter is one of the earliest applications of spectroscopy for clinical use. Many other static methods use similar techniques. However, the pulse oximeter itself may not be optimal for analysing the viability of skin flaps. First of all, the pulse oximeter operates only when the pulsation of the arterials is strong enough, which may not be the case in all skin flaps. Secondly, the SaO2 signal may not be the best indicator of problems in tissue blood circulation. The PI signal could be better, but its absolute value depends on the position of the sensor, and it is also sensitive to moving artefacts.
12.2.4.2 Photoplethysmography (PPG) The photoplethysmography (PPG) contains an emitter light source, typically LED, in the wavelength range of 600–900 nm and a receiver, such as a photodiode. The emitter illuminates the skin, and the receiver collects the light scattered out of the skin, the PPG signal, in either transmission or reflectance configuration in the same way as the pulse oximeter. The intensity of the received PPG signal is relational to the blood volume in the tissue, and it pulsates due to the cardiovascular cycle. The pulsation carries many kinds of information about the function of the cardiovascular system, the microcirculation in the tissue and the blood perfusion. Therefore, it might be useful in skin flap viability analysis (Challoner, 1979; Kamal et al., 1989; Allen, 2007). Some researchers have studied the use of PPG for skin flap vitality analysis (Bardach et al., 1979; Pickett et al., 1997; Chubb et al., 2012), and it has performed well in many studies. The PPG signal is relative, prone to moving artefacts and depends on the position of the sensor. These issues may be a hindrance for its widespread use in skin flap vitality assessment.
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Photoplethysmographic imaging (PPGI) is achieved using a camera instead of the photodiode as a receiver. The PPGI allows examination of the PPG signal from each pixel of the PPG image. The disadvantage of PPGI for skin flap assessment is that it works best in transmission configuration, which is difficult for skin flap measurements (Humphreys et al., 2005).
12.2.4.3 NIR spectroscopy (NIRS) NIR spectroscopy (NIRS) is a popular method for studying constitutive parameters of skin (Nystr€ om et al., 2003; Yu et al., 2007; Va¨lisuo et al., 2010). It has often been used for skin flap assessment, and compared with other methods (Payette et al., 2005; McKenna et al., 2009; Smit et al., 2010), NIRS is found to be a reliable, fast and cost-effective method for tissue viability analysis. The NIRS method is based on the analysis of the spectrum of light scattered out from the tissue, either in transmission or in reflectance configuration. A wideband light source is often used for NIRS, but multiple narrowband light sources are also used. Assume that the emission spectrum of the light from the light source is I0(l), where l is the wavelength. A biological tissue is illuminated with the light source I0(l). The light scattered out of the tissue has spectrum, I(l), which is guided to the spectrometer through a narrow slit. In the spectrometer, the light is spread into a spectrum using a graded mirror. The intensities of different wavelengths are then recorded by a sensor. The relative intensity of the received light, T(l) ¼ I(l)/I0(l), carries information about the concentrations of the chromophores in the skin. It can also be used as a PPG signal because it pulsates according to the cardiovascular cycle. The logarithm of the inverse of the transmission is said to be the absorption, A(l), of the tissue. See Equation (10.5). AðlÞ ¼ log e
l I ð lÞ : ¼ log e T ð lÞ I 0 ð lÞ
(10.5)
In case of no scattering, the relationships of the concentration of the chemical is directly relational to the absorption and the path length, l, of the photons in the tissue, according to the Beer–Lambert law (BLL), as shown in Equation (10.6) AðlÞ ¼ mðlÞl;
(10.6)
where m is the absorption coefficient of the tissue. In practice, the scattering in the tissue is strong, and the BLL does not hold, but the absorption, A(l), is nonlinearly dependent on the concentrations. Therefore, the analysis of the absolute values for HbO2, Hb and SO2 are not easy with NIRS. Often, the light propagation in skin is modelled using suitable skin models, such as the diffusion theory (Farrell et al., 1992), the Monte Carlo simulation (Prahl et al., 1989) or the Kubelka–Munk (Kubelka and Munk, 1931) model, and the model parameters are found by nonlinear optimisation methods. The measurement depth of the NIRS is dependent on the probe geometry (Va¨lisuo and Alander, 2008). The rule of thumb is that the larger is the spacing between the
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emitter and the detector, the deeper does the light travel before reaching the detector. When the spacing is increased, the intensity of the received light is reduced almost exponentially. Therefore, large emitter–detector spacing is often impractical. If the separation of the emitter and detector is approximately the same as the diameter of the emitter and detector fibres, the received signal is only slightly dependent on the scattering coefficient of the tissue. The optical probe can be further designed so that the BLL can be used by only adding a constant correction factor for the path length, l. This kind of probe was demonstrated and called differential path-length spectroscopy (Amelink and Sterenborg, 2004). NIRS seems to be a reliable and tested method for analysing tissue properties, such as HbO2, Hb concentrations and SpO2. NIRS is a flexible method; the sampling depth and many other parameters can be adjusted by designing suitable probes. The PPG signal is obtained as a side effect. It is also possible to apply NIRS invasively by implanting the emitter and receiver fibres, or a probe, inside the tissue. The NIRS method is potentially prone to the moving artefacts and bending of the measurement fibres.
12.2.4.4 Diffuse optical imaging (DOI) In diffuse optical imaging (DOI), an area of a tissue is illuminated with smooth and preferably diffuse light. The reflected light is recorded by using a CCD camera. The imaging can be said to be diffuse if the multiple scattered diffuse photons play an important role and the sampling depth exceeds the diffusion limit. The Fresnel reflection from the surface of the skin and single scattered photons in the point of illumination may considerably conceal the diffuse reflectance signal under the surface without providing any information about the parameters of the tissue. The Fresnel reflection and part of the single scattered photons can be filtered out using two polarising filters. One of the filters is assembled in front of the light source to make the incident light polarised. The other filter is situated orthogonally in front of the objective of the camera to filter out Fresnel reflection, which has kept its polarisation angle. The diffuse reflected signal below the surface of the skin is mostly multiple scattered, and the light loses its polarisation angle in several scattering events. Therefore, the Fresnel reflection and single scattered photons are mostly filtered out, and only the multiple scattered, diffuse reflected photons are allowed to the camera (Olivier et al., 2003; O’Doherty et al., 2007). This method is often called orthogonal polarisation imaging. DOI can use either one wavelength—standard RGB filters to acquire a three-band image—or include components that allow capturing of multi-spectral or hyperspectral image of skin. Already the standard RGB image contains a lot of spectral information, which can be analysed using methods similar to those used in NIRS. The disadvantages of using the RGB image are that the exact filter specifications of the camera are not usually available, only red components reside in the diagnostic– therapeutic window, blue and green do not penetrate deeply into the skin, and the number of channels is too small for many purposes. A multi-spectral image can be obtained, for example, using several narrowband LEDs with different wavelengths as illumination. The different LEDs with different peak wavelengths are used one by one, and one picture is taken in each wavelength.
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The captured images can be combined as one multi-spectral image. Hyper-spectral images can be taken using a wideband white light source and a special spectral camera. Several methods using DOI are: l
l
l
SIAscopy: Spectral Intracutaneous Analysis is a method wherein the skin is imaged using either RGB colours or RGB colours with some NIR channels. The images are analysed to produce a 2D map of HbO2, Hb, SpO2 and melanin concentrations. The method is based on the Kubelka–Munk skin model and solves the nonlinear skin model using precalculated look-up tables. The main purpose for SIAscopy is to diagnose skin cancer, but it may work for skin viability analysis as well (Claridge et al., 2003; Govindan et al., 2007). Spectrocutometry: The Spectrocutometry uses a linear Beer–Lamber model of the epidermis and a non-linear diffusion theoretical model of the dermis, which is linearised around an operating point to allow fast solution of the skin model at every pixel location. Spectrocutometry is used for calculating the HbO2, Hb, SpO2 and melanin fraction over the imaged area (Kaartinen et al., 2011; Va¨lisuo et al., 2011). TiVi: The purpose of the TiVi imaging is to estimate the concentration of RBCs over the imaged area. The method is based on linear equations with a fixing coefficient (McNamara et al., 2010).
DOI is used in many different forms. All of the methods are cost-effective, highresolution and moderately fast. The nonlinear relationship between the skin reflectance and the chromophore concentration makes it difficult to obtain absolute concentrations. However, SIAscopy and spectrocutometer aims to the absolute concentrations. SIAscopy has been on the market for several years, and there are many clinical articles about TiVi imaging. Spectrocutometer is emerging in future markets. All DOI instruments are potentially suitable for skin flap vitality assessments. The measurement depth and the lack of off-the-shelf clinical instruments may set some limitations to the application of these methods.
12.2.4.5 Diffuse optical tomography (DOT) Diffuse optical tomography (DOT) is an imaging mode that can produce a spectral image of an object residing several centimetres below the biological tissue. The illumination is handled using and array of light sources, and detection is handled using an array of detectors surrounding the imaged object. Each detector records the remitted light while one light source is illuminating the target. This procedure is repeated for each light source (Wang and Wu, 2007). Deep imaging depth can be achieved by utilising multiple scattered diffuse photons and NIR light within the diagnostic–therapeutic window. An optical tissue model, often based on diffusion theory, is used to infer localised tissue parameters. The DOT imaging system can work in three different modes: time-domain, frequency domain or DC-current mode. In time-domain mode, the light sources emit picosecond pulses, which are broadened by the tissue and observed by the detectors. In frequency mode, the light sources emit an amplitude modulated light signal that is observed by the detectors with smaller modulation depth, due to the multiple scattering in the tissue. In DC mode, the light sources emit only slowly time-invariant signals, which are again recorded by the detectors.
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Because the diffuse photons have lost their directional information, the axial and lateral resolution of DOT is poor—approximately 20% of the imaging depth. The acquisition time of a DOT image is short. If the light emitter array is tuneable or duplicated with several different wavelengths, a multi-spectral DOT image can be achieved that can be used in resolving HbO2, HbO concentrations and oxygen saturation SpO2. These parameters can be used to study the metabolism of deep skin structures, such as tumours (McBride et al., 1999). The spatial resolution provided by the DOT method is probably enough for skin flap assessment, and the fast acquisition rate is useful in analysing the blood circulation dynamics. The overall properties of the DOT method should be well suited for skin flap analysis.
12.2.5 Fluorescence methods In some cases, it is beneficial to assess blood perfusion using external contrast agents, such as ICG. ICG is a well-tested fluorescent dye that has been in use for decades. ICG can be excited with NIR light in the wavelength range between 600 and 800 nm, and it emits in the range of 800–950 nm. Both excitation and emission wavelengths are within the diagnostic–therapeutic window, which makes ICG suitable for imaging deep blood flow. ICG has often been used in neurosurgery to verify the blood flow in arteries after an operation. It has also been used in assessing the blood flow in muscles of the lower extremity as well as for many other purposes (Alander et al., 2012). ICG could be used for skin flap assessment by administering the ICG into the blood circulation through a vein. After a few seconds, the ICG enters all over the body, including the skin flap, if the blood circulation works properly. The liver removes the ICG from blood circulation within a few minutes. The ICG enters first in the arterials of the tissue, then in the capillaries and last in the veins. The dynamics of the ICG concentration in various parts of the skin flap and nearby normal skin can be measured and compared using fluorescent NIRS, absorption NIRS or fluorescent imaging. ICG assessment is an invasive method, and due to the weak fluorescent signal, very sensitive NIRS or DOI equipment with high-quality excitation and a fluorescence filter pair, are required. Still, equipment costs are quite moderate (Alander et al., 2012).
12.2.6 Summary Optical methods are usually relatively low cost, fast and offer high resolution. The optical measurements are available in many different modalities as reviewed in this chapter. The key properties of these modalities are shown in Table 12.1.
12.3
Future trends
The research on the optical methods for sampling skin parameters is actively continuing together with improvements in optical technologies. Many methods that are now under investigation will be standardised in the near future and will be available for
The summary of the key properties of optical methods available for skin flap assessment
Method
Dyn/static
Mode
Contrast
Depth (mm)
Resolution
Time
Availability
Cost
LDI LDF DOCT LSI DCS RPO PPG NIRS DOI DOT ICG
Dyn Dyn Dyn Dyn Dyn Static Static Static Static Static Static
Image Point 3D image Image Point Point Point Point Image 3D image Image
Doppler Doppler Doppler Doppler Doppler SpO2, PI Blood volume, PI HbO2, Hb, SpO2 HbO2, Hb, SpO2 HbO2, Hb, SpO2 CICG
0.1 0.1–5 0.1 0.1 >10 10 1 to >10 1 to >10 1 >10 1–10
Moderate – High High – – – – High Low Moderate
Minutes Real time Slow Fast
Available Available Available Available Emerging Available Available Available Available Emerging Available
Moderate Moderate High Low – Low Low Moderate Low Moderate Moderate
Real time Real time Real time Fast Slow Fast
Optical methods for assessing skin flap survival
Table 12.1
343
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clinical use. Because optical methods are relatively low cost, they will be integrated into other existing measurement systems. Also, combinational optical devices are likely to emerge. Currently, there are pulse oximeters with optoplethysmograpy function. LSI can easily be combined with DOI to provide information on the concentration of blood, oxygen saturation and blood perfusion. PPG imaging and DOI together could provide similar benefits for skin flap vitality assessment.
12.4
Sources of further information
The Biomedical Optics: Principles and Imaging book (Wang and Wu, 2007), describes the theories behind many optical measurement modalities in detail. This book is excellent for the technically oriented reader. The Tissue Optics book (Tuchin, 2007) explains the theories behind skin optics and contains many tables listing the optical properties of different biological tissue types. Introduction to the light interaction with skin (Baranoski and Krishnaswamy, 2008) is a good introduction to the theories behind various optical measurements. Handbook of Biomedical Optics (Boas et al., 2011) provides a recent view to the theories behind optical measurements of biological tissues. The coverage of various theories behind biomedical optics is excellent.
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