Biophotonic methods in microcirculation imaging

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Medical Laser Application 22 (2007) 105–126 www.elsevier.de/mla

Biophotonic methods in microcirculation imaging Martin J. Leahya,, Joey G. Enfielda, Neil T. Clancya, Jim O’Dohertya, Paul McNamaraa, Gert E. Nilssonb a

Biophotonics Laboratory, OFSRC, Department of Physics, University of Limerick, Limerick, Ireland Institutionen fo¨r medicinsk teknik, Linko¨pings universitet, Linko¨ping, Sweden

b

Received 16 May 2007; accepted 29 June 2007

Abstract Visible and near-infrared light, particularly in the wavelength region of 600–1100 nm, offer a window into human and animal tissues due to reduced scattering and absorption. We review the main biophotonic methods applied to visualisation and assessment of the microcirculation and document the progress made over the past 10 years in particular. Applications, particularly in human skin, are of special topical importance due to an improved knowledge of its role and its value as a surrogate for other organs in drug testing at a time when drug development is under severe pressure. r 2007 Published by Elsevier GmbH. Keywords: Biophotonics; Microcirculation imaging; Laser speckle; Laser Doppler; Perfusion monitoring; Capillaroscopy; Tissue viability imaging (TiVi); Photo-acoustic tomography (PAT)

Introduction Direct (optical) observation of skin and other tissues is of course the oldest method of biomedical imaging. We know that the ancient Greeks already understood that the pallor of the skin was a significant discriminator between health and disease. However, the decomposition of white light [1], the Doppler effect [2] and the identification of haemoglobin as the substance which changes its colour and that of the skin when carrying oxygen [3] provided a scientific basis for reliable measurements of the microcirculation developed over the past 150 years. Over the past 30 years in particular, tremendous progress has been made in single point biophotonic measurements of microvascular state/activity and more recently imaging of same. Corresponding author.

E-mail address: [email protected] (M.J. Leahy). 1615-1615/$ - see front matter r 2007 Published by Elsevier GmbH. doi:10.1016/j.mla.2007.06.003

The vital role of blood supply, and the oxygen it carries, in the health of the individual has ensured that many different techniques for assessing it have been investigated. The principle upon which the methods are based vary widely, as does the suitability, cost of the materials and technology necessary. A small number of the methods are continuous and even fewer are truly non-invasive. Non-invasive in vivo techniques have obvious advantages for the user in providing information without disturbing the normal environment. In the same way, it poses considerable difficulties for developers through the need to make accurate measurements in complex environments subject to enormous (biological) variability. The imaging technologies discussed here can be separated by physical resolution and sampling depth (Fig. 1) with optical coherence tomography (OCT) and ultrasound included for completeness. Pixel resolution is improved by two orders of magnitude in commercially

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available units by going from laser Doppler perfusion imaging (LDPI) to laser speckle perfusion imaging (LSPI) and again from LSPI to tissue viability imaging (TiVi). Many techniques have been proposed for imaging the microcirculation, from X-rays to thermography, and still new techniques are emerging e.g., OCT, PAT, TiVi, etc. Fig. 1 shows the resolution and sampling depth of the various techniques. The depth of interdisciplinary knowledge required to specify the appropriate technique often means that the wrong technique is used, or the correct one is used inappropriately. This often leads to unfair criticism of techniques used in situations for which they were never designed. The techniques used in the clinical setting tend to have the common advantages of being mobile, inexpensive or otherwise available as well as the required technical specification. The rush (fashion) for evermore-sophisticated technologies, which suggests that expense is somehow an advantage, is not borne out by clinical use. There is a clinical need for simple to use, inexpensive technologies for bed-side monitoring and to get to (research) patient groups who cannot attend the small number of locations which house this elite equipment. To understand why optical imaging of the skin and microcirculation (Fig. 2) is not universally applied, it is instructive to consider how light interacts with tissue and which situations allow for direct visualisation of the skin. Light microscopy and direct visual inspection are hampered by scattering, regular reflection from the surface and absorption by superficial chromophores. Bloodless melanin-free skin provides little absorption but significant scattering across all colours of visible light. Since returning photons are scattered many times within the skin they are diffuse therefore the skin appears white. Regular, mirror-like reflection accounts only for a few percent of the incident photons. Ultrasound has many advantages due to its low scattering in this tissue matrix and can provide excellent images for larger deeper vessels. However, it requires higher frequencies for smaller vessels causing difficulties

Fig. 1. Microcirculation imaging domains.

Fig. 2. Microcirculation in the skin.

close to the surface. It is possible to view capillary vessels directly in nailfold plexus and to view the retinal vessels directly due to the lack of significant absorption/ scattering outside of these vessels. However, absorption is the main contrast mechanism and back-scattering is required to return photons to the detector/eye.

Nailfold capillaroscopy One of the most useful methods of assessing blood flow in a small number of blood vessels in the skin is by direct observation using capillary microscopy [4]. Capillaries are the smallest of the blood vessels and their purpose is to link the arterial and venous systems together, allowing the exchange of carbon and oxygen between the tissue and blood cells. Only in certain places in the body do the capillaries come close enough to the surface of the skin to be naturally visible i.e., without the use of specialised optical equipment or optical clearing agents. This is one of the reasons that the study of the skin overlapping the base of the fingernail and toenail is so important. In the nailfold, the capillaries come within 200 mm of the surface of the skin, meaning the methods of examining them are simple. Another reason is that the fingernail is easily fixed in position, free from any movement, due to arterial pulsations or respirations and the capillaries run parallel to the skin surface. Using a microscope with a magnification of between 200  and 600  , it is possible to clearly see the structure of the blood vessels and because each vessel is sufficiently transparent, the red blood cell (RBC) motion, in a single capillary, can be measured. If a video camera is attached to the microscope [5,6] the motion can be examined in a frame-by-frame procedure yielding accurate velocity information for RBC flow in the capillary. The flow may be altered by the fixing procedure and/or the

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Fig. 3. Nailfold capillary pattern of a healthy adult [9]. The finger nail is marked as N. A typical capillary density for a healthy adult is approximately 30–40 mm2 [10]. Reproduced with permission from Wiley-Blackwell Publishing Ltd.

heating effect of the light source. Examining the capillary condition and density can aid the diagnosis of certain diseases. Capillary density and diameter are dependent on age, with younger children having a lower density and capillary thickness than adults [7]. The presence of abnormal vessels, in addition to these factors, has been attributed to various diseases. Some conditions can be detected very early as capillary deformations can be observed before other symptoms occur [8]. Fig. 3 is an image of a nailfold capillary arrangement of a healthy adult [9]. The capillaries are hairpin-like loops, arranged into regular rows. Each loop consists of two parallel limbs; a thinner arterial limb and a thicker venous limb. The walls of each blood vessel itself are transparent and the RBCs are seen passing through the capillaries [11].

Measurements A disease-specific ‘pattern’ can be identified by analysing the geometry and density of the capillaries and the presence of abnormal or very large blood vessels [10]. Other relevant factors include avascular areas (absence of capillaries) [9] and the RBC(s) red blood cell velocity (RBV) through the capillaries [12]. A typical value of RBV would be approximately 1 mm s1 [12]. Examples of a vascular areas, and giant and dilated capillaries can be seen in Fig. 4. The geometry of each capillary is defined by taking measurements at specific points. An example of this is shown in Fig. 5. The most difficult calculation in nailfold capillaroscopy is that of the RBV. There are many different methods of determining this. One such method is laser Doppler flowmetry (LDF). There are two main advantages of nailfold capillaroscopy over this method. The first is that LDF is not limited to a specific area,

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Fig. 4. Nailfold capillaries in a patient with dermatomyositis. Examples of avascular areas, A; giant capillaries, G and dilated loops, D are shown [9]. Reproduced with permission from Wiley-Blackwell Publishing Ltd.

Fig. 5. Measurements of dimensions of a capillary [10]. The arterial capillary diameter (a). The venous capillary diameter (b). The loop diameter (c). The width of the capillary (d). The distance between the limbs (e).

meaning the measurements are not taken on a specific RBC. As a result, the measurements are made less accurate by cells with various velocities [13]. The second drawback is that the signal produced is in proportion to microvascular perfusion, which is the product of both the velocity and concentration of the RBCs. It is impossible to measure the absolute units of blood flow [14]. In nailfold capillaroscopy, because individual RBCs can be seen, direct measurement of the RBV is allowed. There exist several methods of assessing RBC perfusion. Measurements can be performed in a frame-by-frame analysis of specific patterns of RBC flow (RBCs or plasma gaps) [15]. By tracking the pattern over two consecutive frames, the velocity can be approximated. A drawback of this method is that these patterns are difficult to detect in larger vessels. Other methods include cross-correlation methods, which can be further sub-divided into two methods. The first works by measuring the average intensity at two independent regions of the capillary. Using a temporal cross-correlation and the distance between the regions, the RBV can be calculated [16]. The second

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method works by creating a series of spatial intensity distributions on a particular section. Using a temporal spatial correlation and the time separation between consecutive frames, the velocity can be calculated [17]. Sourice et al. made use of the spatiotemporal autocorrelation function to develop software to perform calculations of the RBV [18]. Experimental considerations A typical experimental set-up would consist of a microscope, monitor and video recording system [8,11,12,18]. Fibre-optic illumination is often used [7,10]. The finger can be immobilised easily by use of a clamp [12]. To increase the transparency of the skin, a clearing agent such as immersion oil is usually applied to the nailfold [7,8,11]. A fluorescent solution such as fluorescein can be used to distinguish the capillaries from the surrounding skin [18]. The software used for analysing the images and the preparation of the subjects is dependent on what is being measured. Ohtsuka et al. [8] showed evidence that for patients with primary Raynaud’s phenomenon, the loop diameter, capillary width and capillary length (inner length+outer length/2) were greater in the patients who went on to develop undifferentiated connective tissue disease, than those who did not. Bukhari et al. [10] studied the dimensions and density of capillaries and showed that there was a significant increase in all dimensions except the distance between the limbs in the patients with Raynaud’s phenomenon and systemic sclerosis (SSc) compared with a control group. Also, capillary density is reduced in patients with SSc. Bhushan et al. [19] found that there was a decrease in capillary density in the nailfold of patients with psoriasis compared with a control group. Wong et al. [20] showed that a characteristic pattern of capillary abnormalities occurred in patients with scleroderma. Aguiar et al. [12] studied the RBV, functional capillary density and capillary morphology differences in patients with primary Sjo¨gren’s syndrome and a healthy control group. It was found that the affected patients had a higher functional capillary density and lower RBV than the control group. Limitations and improved analysis The measurements required for analysing the capillary dimensions and patterns take much more time than the examination itself. Computerised systems have been developed [21–23] for analysing images of capillary networks and improving their quality by excluding disturbances caused by hair, liquid, reflections, etc. Combining videocapillaroscopy and various mathematical methods, the statistical properties of the capillaries can be analysed automatically by extracting capillary count, position, density, size, etc. from the images. Sainthillier et al. [21] and Zhong et al. [22] used

triangulation methods to allow statistical analysis of distances between nearest neighbouring capillaries, which is useful for looking at areas that are susceptible to development of necrosis. Hamar et al. [23] are currently developing a Markov chain-based detection algorithm for updating capillaroscopy images to reduce them to a greyscale image containing only the capillaries. These technologies are essential to the development of nailfold microscopy into a more sophisticated method of disease diagnosis. Conclusion While nailfold capillaroscopy is important in the diagnosis of certain connective-tissue diseases, some of the methods described above cannot be applied to other areas of the microcirculation for the simple reason that the capillaries are not clearly visible. This necessitates other methods of measuring the condition and density of the capillaries, and the velocity of the RBCs.

Laser Doppler Blood Flowmetry (LDBF) LDBF refers to the general class of techniques using the Doppler effect to measure changes in blood perfusion in the microcirculation non-invasively. Perfusion has previously been defined as the product of local speed and concentration of blood cells [24]. The principles behind the technique were first developed by Riva et al. [25] who developed a technique to measure RBCs speed in a glass flow tube model; however, the development of Laser Doppler perfusion monitoring (LDPM) for assessing tissue microcirculation was first demonstrated in 1975 by Stern [26]. The technique operates by using a coherent laser light source to irradiate the tissue surface. A fraction of this light propagates through the tissue in a random fashion and interacts with the different components within the tissue. The tissue consists of both static and moving components. The light scattered from the moving components will undergo a frequency shift which can be explained by the Doppler effect while the light scattered from the static components will not undergo any shift. It is almost impossible to resolve the frequency shifts directly. To overcome this, the backscattered light from the tissue is allowed to fall on the surface of a photo-detector where a beat frequency is produced. The typical frequency of this beat detected using a wavelength of 780 nm ranges from 0 to 20 kHz. However, for a 633 nm source the range is smaller, typically 0–12 kHz [24]. The apparent contradiction with Eq. (3) below is caused by the deeper vessels probed by the longer wavelengths, which contain faster flowing blood. From this beat frequency, it is possible to determine the speed and concentration of the blood cells.

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Theory This frequency shift produced by Doppler scattering is determined by several factors (Fig. 6). These factors are the scattering angle (y), the wavelength of light in the tissue lt and the velocity of the moving scattering object ~ vs . If we assume that a photon injected with propagation ! ! vector ki gets scattered in the direction ks , the resulting scattering vector is given by   y ~ ~ ~ q ¼ ki  ks ¼ 2k sin . (1) 2 The Doppler shift in frequency df between the incident frequency fi and the scattered frequency fs for a single scattering event is given by     4p y f s  f i ¼ df ¼ ~ vs  ~ q¼ cos f. (2) vs sin lt 2 However, from Eq. (2) the maximum frequency shift occurs when the incident light is normal to the scatterer (f ¼ 901) and all the light is scattered in the forward direction (y ¼ 901). In this case, we have   4p df ¼ (3) vs . lt Eq. (3) shows that the maximum resulting frequency shift that occurs due to a scattering event is directly proportional to the velocity of the scattering particles. In a system of moving particles such as blood cells in tissue, the photon can undergo more than one Dopplershift, in which case the detected frequency difference is the sum of the individual scattering events. For n scattering events, it can be shown that the Doppler shift is given by Eq. (4) [24]: !  i n X 4p i y i df ¼ (4) i vs sin 2 cos f . lt i¼0 Eq. (4) shows how the frequency of the light that is scattered from moving blood cells is frequency shifted. It must be noted that if the backscattered light is examined it contains photons that have been frequency-shifted differently due to the physiological and statistical distributions of ~ vs , y and f resulting in a distribution of frequencies rather than a single Doppler-shifted frequency. This results in a Doppler shift power spectrum, P(o). This power spectrum can be recorded by heterodyne light mixing spectroscopy [27], a technique that makes use of the interference between the

Fig. 6. Process of a single Doppler shifting event.

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shifted and unshifted light to generate a time-varying photocurrent whose spectrum contains only the Doppler shift beat frequencies. The frequency content of the fluctuating portion of this photocurrent is related to the blood cell average speed, while the magnitude is determined primarily by the number of moving blood cells within the scattering volume. The determination of the perfusion from the power spectrum as obtained from Bonner and Nossal’s work [28] can be implemented using the following equation [24]: " R o1 # Ro ð o2 oPðoÞÞ do  o21 NðoÞ do , (5) LDP ¼ C I2 where o ¼ 2pf ,

(6)

where C is a calibration factor, P(o) the power spectral density of the Doppler shift frequencies, and N(o) the power spectral density of the shot noise at intensity I. Light sources The incident laser light beam has a depth of penetration of approximately 1 mm depending on the wavelength and configuration of the equipment used. This estimate originates from mathematical modelling of photon diffusion through ‘imaginary tissues’ using Monte Carlo techniques. Therefore, all elements of the dermis may be included, from the superficial nutritional to deeper thermoregulatory vessels. For a given wavelength of light, the absorption spectrum of components of the tissue determines the interaction that occurs [29]. The penetration depth of light is predetermined by the path the light takes through the skin and is limited by the absorption and scattering effects of the tissue. Another limiting factor is the level of pigmentation of the skin [30]. The absorption of haemoglobin and water are lower towards the red end of the visible spectrum. Because of commercial availability, the HeNe lasers (632.8 nm) dominated in earlier LDPM; however, longer wavelengths produced by laser diodes (780 and 810 nm) are now preferred due to increased penetration depth and lower absorption by melanin in the near infrared which shows less dependence on skin colour [30]. Also, because the absorption of oxidised and deoxygenated haemoglobin is almost equal at this wavelength, any dependence on oxygen saturation is eliminated. However, systems exist that use dual wavelengths to obtain additional information about the tissue structure [31,32]. Techniques Laser Doppler perfusion monitoring (LDPM). This is an established technique for the real-time single-point measurements of perfusion changes in tissue. It works

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by illuminating the skin with a laser and from the backscattered light, extracting information about the velocity and concentration of moving blood cells in the tissue. The technique correlates well with other measurements of skin blood flow, such as venous occlusion plethysmography [33] and xenon washout [34]. LDPM offers substantial advantages over other methods in the measurement of microvascular blood perfusion. Studies have shown that it is both highly sensitive and capable of resolving changes1 of less than 0.1 s to local blood perfusion, making it useful for continuous perfusion monitoring. As the probe is not required to touch the surface of the tissue, the technique is potentially non-invasive so it does not affect the normal physiological state of the microcirculation. Furthermore, the small dimensions of the probes have enabled it to be employed in experimental environments not readily accessible using other techniques. The technique has its limitations. It is designed to record tissue perfusion at a single spot over time; however, cutaneous circulation is known to be heterogeneous and examining a small area of perfusion does not necessarily give representative data for the surrounding perfusion [35–37]. Another issue with the technique is that measurements obtained are intrinsically of a relative nature, the measurements are proportional to blood flow. However, the factor of proportionality will be different for different tissues and tissue sites. This means the system does not output exact measurements of blood flow in the system, which makes it only useful in examining changes in perfusion due to different stimuli in a single location. The technique has been adapted and improved to remove many of the preliminary problems and has been widely used both in research and as a clinical tool over the past 20 years in the measurements of cutaneous microcirculatory flow. Guidelines have been published on how to perform measurements using the technique [36]. There are many different types of probes available, each suited for different applications such as fibre probe systems [39] and integrated probe systems [40]. Laser Doppler perfusion imaging (LDPI). The need to study perfusion over larger tissue areas led to the construction LDPI [41,42]. LDPI is a technique that scans a series of perfusion measurements over a section of tissue. This is converted into a colour-coded image representing the distribution of perfusion over an area of interest. This technique offers some advantages over the single-probe technique; the blood flow is measured over an area rather than at a single site, which removes some of the difficulties that arise in the LDPM technique such 1 Discovery Technology International (DTI) OxiFlow System Brochure.

as movement artefacts and site-to-site perfusion variations due to the heterogeneous nature of tissue. The technique requires no contact with the tissue being examined making it of particular advantage when assessing open, and often infected, wounds. The technique has got limitations; one issue is the length of time taken to complete a full scan. This can introduce problems in trying to examine a rapid change in blood flow. However, more recent techniques such as the line-scanning and full-field laser Doppler offer to greatly improve this. LDPI provides arbitrary perfusion measurements [35] meaning that the tool is valid for comparative data only and that all measurements must be performed under the same conditions. It is also necessary to consider temporal variations in blood flow over short periods [38] (seconds to minutes) and over extended periods due to seasonal changes [43] while attempting to monitor long term changes in perfusion to a region. Another issue with the technique is that due the scanner not being held in contact with the tissue, the distance between the imager and the tissue affects the measured level of perfusion. If this distance changes between subsequent measurements it will affect the results. There are conflicting opinions on whether the measured perfusion increases [44] or decreases [45] as the distance from the tissue increases. It may be that the algorithms used for signal processing by different machines determine this [46]. In order to make comparable measurements between different systems, a standard calibration and system design is required [24]. Between 1997 and 2002, a standardisation project was undertaken by several international research institutes, in which a perfusion simulator for calibration and standardisation of LDPM and use of nomenclature was agreed. In 2002, a report on a project named HIRELADO (high resolution laser Doppler) from the standardisation group of the European society of contact dermatitis published guidelines on the measuring of skin bold flow [47]. This project addressed mainly the technical aspects of how experiments should be performed and what precautions should be taken to ensure valid results. There are several different techniques that have been developed for producing LDPI data, each with different advantages and disadvantages. Point raster technique. A raster scan laser Doppler image operates by recording a series of point perfusion measurements. This type was the first laser Doppler image system to be developed. The basic schematic is shown in Fig. 7. The imaging is achieved by using a moveable mirror to scan the laser beam over the area to be measured in a raster fashion. At each site an individual perfusion measurement is taken. When the scan is complete the system generates a two-dimensional (2D) map image from the single point perfusion measurements.

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Fig. 7. Schematic of raster scan laser Doppler imaging system. (Moor Instruments Ltd. MoorLDI Brochure.)

The problem with the technique is the time taken to make measurements. Each individual measurement can be achieved in 4–50 ms, however lowering the acquisition time results in increased noise in the signal [47]. For time considerations, it is important to minimise the number of measured points. This is best achieved if the step length between measurements equals the diameter of the laser beam. For smaller step lengths, the same physiological information is partly collected by neighbouring measurements, which increase the overall time without retrieving more information. The moorLDI system reports capability of capturing a 64  64 perfusion image in 20 s and has a spatial resolution of 0.2 mm at a distance of 20 cm. The system is capable of scanning a region up to 50 cm  50 cm and as small as a single point measurement. Line scanning technique. Laser Doppler line scanning imaging is a new development in laser Doppler image (LDI) technology by Moor instruments. This technique does not utilise the standard point raster approach but works by scanning a divergent laser line over the skin or other tissue surfaces while photodetection is performed in parallel by a photodiode array. This technique allows for a series of perfusion measurements to be performed in parallel. This greatly reduces acquisition time and is bringing laser Doppler Imaging one step closer to real time imaging. A schematic of the system is shown in Fig. 8. The moorLDLS utilises 64 diodes to obtain the perfusion measurements in parallel. The system is capable of measuring each line in times of 100–200 ms. Regions up to 20 cm  15 cm can be imaged and a 64  64 pixel perfusion measurements can be produced in 6 s. Full field scanning technique. A high-speed LDPI using an integrating CMOS sensor for full field scanning has been recently developed by Serov et al. [48]. The basic

Fig. 8. Schematic of line scanning laser Doppler system. (Moor Instruments Ltd. MoorLDLS Brochure.)

schematic of the system is shown in Fig. 9. A laser source is diverged to illuminate the area of the sample under investigation and the tissue surface is imaged through the objective lens onto a CMOS camera sensor. This new CMOS image sensor has several specific advantages: firstly, the imaging time is 3–4 times faster than the current commercial raster scan LDI systems. Secondly, the refresh rate of the perfusion images is approximately 90 s for a 256  256 pixel perfusion image. This time includes acquisition, signal processing, and data transfer to the display. For comparison, the specified scan speed of a commercial laser Doppler imaging system, moorLDITM (Moor Instruments Ltd., UK), is approximately 5 min for a 256  256 pixel resolution image. However, the scanning imager can only measure areas of up to approximately 50 cm  50 cm in size while the CMOS system at present does not allow imaging of the areas larger than approximately 50 mm  50 mm [49,50]. Applications The laser Doppler technique has many medical applications; however, it still has not been fully integrated into clinical settings and is mainly used in research. LDPI systems have recently been reported in the assessment of burns [51–54]. The use of LDPI is reported to outperform existing methods of assessment of burn wound depth and provides an objective, realtime method of evaluation. Accuracy of assessment of burn depth is reported to be up to 97% using LDPI, compared with 60–80% for established clinical methods [51]. Fig. 10 shows the typical results obtained-high perfusion corresponds to superficial dermal burns, which heal with dressings and conservative management;

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burns with low perfusion require surgical management i.e., skin grafts. Correct assessment is of particular importance here as 1 in 50 grafts fail; if this failure is caught within 24 h, 50% of these cases can be saved [55]. The use of LDPI has also been investigated in different aspects of cancer research and treatment; it has been found that there are higher levels of perfusion in skin tumours, which can be detected by the LDPI system [56]. The technique has also been applied to examine how a tumour is responding to different treatments such as photodynamic therapy (PDT) [57]. Wang et al. applied the technique to the assessment of post-operative malignant skin tumours [58]. It has also been applied to many other fields such as examining allergic reactions and inflammatory responses to different irritants [59–61]. Bjarnason et al. applied the technique in the assessment of patch tests [62,63]. Ferrell et al. has used the technique to study arthritis and inflammation of joints [64]. It has also been applied in investigating the healing response of ulcers [65] and examining the blood flow pattern in patients at risk from pressure sores [66].

Fig. 9. Schematic of full field laser Doppler system [49].

Laser speckle perfusion imaging Introduction As the angles in LDP are more or less random and a wide range of velocities is present in the microcirculation, a continuous Doppler frequency spectrum is produced. It is interesting to note that a similar result can be achieved by considering the scattering particle to be a reflector of light moving with velocity v. The reflected light will interfere with the reference beam and the resultant intensity will vary with the difference in optical path, OPD, between the two beams. The number of interference fringes to pass the detector in time, t, will be given by OPD 2d ¼ , lt lt

(7)

where d is the actual distance moved by the particle. The number of fringes per second is then f ¼

2d 2v ¼ . lt l

(8)

Thus, the relationship of the resultant frequency to velocity is the same in each case. Indeed, several groups [67–69] have re-interpreted the technique in terms of time-varying laser speckle. The speckle pattern is a granular variation of light intensity obtained from any surface, such as the skin, which is rough on a scale comparable with the wavelength of light. The scattering inhomogeneities can be considered as point sources and the small differences in the distances travelled to the detector (eye/photodiode) cause bright or dark speckles to be seen at all points in space. Far-field speckle is that observed as the scattered light falls on a screen some distance from the surface in question while image speckle (of a greater importance from a biomedical imaging point-of-view) refers to the pattern observed on the surface itself [67]. The formation of these two distinct types of speckle is shown in Fig. 11.

Fig. 10. Use of LDPI in assessment of burn damage [51].

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So far, the source of the speckle has been described in terms of surface ‘roughness’ and undulations in height. However, light may also be reflected by living tissue where cells are responsible for backscattering. If the scatterers are in motion (such as RBCs), the pattern is seen to fluctuate, resulting in time-varying laser speckle. This type can be quantified in terms of the speckle contrast (K) [67]: K¼

s p1, hIi

(9)

where s is the standard deviation of the intensity variations and I the intensity. In practical situations, the standard deviation is less than the mean intensity. Briers [68] proposed a simple analogy to explain the physical cause of the fluctuations: a RBC in motion may be considered as the moving mirror in a Michelson interferometer, while a detector is placed at the centre of the resulting interference pattern. Each time the mirror (cell) moves through a distance of l/2, the detector ‘sees’ an interference fringe pass. That is, a fluctuation in light intensity is recorded. Variations in speckle contrast will then be dependent on the velocity distribution of the scatterers. Therefore, this velocity can be determined using a measurement of the statistical behaviour of the speckles over time. This is similar to the reading given by a laser Doppler measurement, and in fact it has been demonstrated that Doppler and time-varying speckle are two methods of arriving at the same result [68]. Re-writing Eq. (8) in terms of the frequency of the fluctuations (Df) measured by a detector in the image plane gives Df ¼ f

2v , c

(10)

where v is the velocity of a scatterer, f the frequency of the laser light and c the speed of light in a vacuum.

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When recorded over a finite integration time, the moving scatterers result in a speckle pattern that appears ‘blurred’ or decorrelated. The amount of decorrelation depends on the speed and volume of the RBCs in the tissue [69]. An early implementation of this effect in the measurement of blood flow involved the use of a linear CCD array and a helium–neon laser (expanded into a line using a cylindrical lens). The time-varying component of the speckle was quantified by comparing the intensity recorded at each pixel in successive scans. The average difference between the two measurements was dependent on flow – large in high flow regions and small in low flow areas. This is analogous to the relationship between correlation time and RBC velocity. A 2D map was created by scanning the line across a region of interest (ROI). Experiments showed a decrease in the blood flow parameter during occlusion of the finger, while an increase was observed across a scar on the back of the hand [70]. A more successful approach was developed using laser speckle contrast analysis (LASCA) under the assumption that, with a Lorentzian flow profile, the speckle contrast could be related to correlation time, tC (and hence, RBC velocity) [67]: sffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi    ffi tC 2T K¼ 1  exp , (11) tC 2T where T is the camera integration time.

Full-field measurements The set-up for full-field laser speckle measurements is shown in Fig. 12. The tissue is illuminated by an expanded laser beam and the resulting image speckle is recorded by a CCD camera.

Fig. 11. (a) Far-field speckle. The pattern is formed on a screen some distance away from the object. (b) Image speckle. The pattern at the object itself is observed directly. Reproduced with permission from Institute of Physics Publishing and Professor David Briers [67].

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Experimental considerations The source used is typically a low power expanded helium–neon (632.8 nm) laser in the 3–30 mW range [71,72] although the use of 660 nm [73], 514 nm [74], 780 nm [75,76] and 808 nm [77] wavelengths have also been reported. The detector area is of particular importance, as averaging of multiple speckles will result in reduced fluctuation amplitude. If enough speckles fall within the detector area (pixel area of CCD), a uniform constant intensity will be recorded [67]. However, if the average speckle size is much larger than the pixel size, an insufficient number of pixels will be available for a valid calculation of K. The speckle size is determined by the aperture size of the imaging device alone [67], so for imaging applications, the speckle size is usually chosen to be approximately the same as one pixel (15 mm). In practice, this is done by controlling the F-stop of the camera [78]: s  1:2ð1 þ MÞlF ,

(12)

where s is the speckle size, M the magnification of the lens, l the wavelength of the illuminating light and F the camera F-stop number. A similar condition applies to the integration time of the detector, which should be small in comparison with the correlation time of the intensity fluctuations to avoid ‘averaging-out’ of the signal [67]. Typically for studies of capillary perfusion, this is set between 5 and 20 ms [69,71,72,75,79]. Yuan et al. [80] found that the optimal integration time for rodent cerebral blood flow studies was 5 ms. This technique originally developed for retinal blood flow imaging [81] uses a short integration time as mentioned previously, to photograph the illuminated area and produces a high contrast speckle pattern. The contrast K, as described in Eq. (9), is calculated based on small areas of the pattern, typically a sliding 7  7 or

5  5 pixel window [67,71,72]. The higher the velocity of the RBCs, the higher the frequency of the fluctuations (Eq. (10)). This means that for a given integration time, more averaging will occur and there will be a drop in contrast. Conversely, an area where blood flow is low will appear as a high contrast reading. These contrast measurements can therefore be used to build up a 2D map of relative velocities as shown in Fig. 13 below. Limitations and improved analysis The effects of specular reflection can be eliminated with the addition of polarising filters [69,73,83–85]. Multiple scattering can have the effect of blurring the image by influencing the apparent size of the object [85]. Also, the effects of scattering by static tissue must also be considered [73]. As the laser speckle signal is dependent on the fluctuating part of the pattern, static tissue has the effect of reducing the signal-to-noise ratio. Zakharov et al. have developed a model taking this into account that includes a correction to the equation relating speckle contrast to correlation time [85]. Attempts to limit the effect of static scatterers have also been made in single point measurements using fibre optic probes by adjusting the source–detector separation [86]. This limiting effect may be one of the most significant reasons why the speckle technique is suited to more superficial measurements. Indeed, many of the applications reported, as described in the Microcirculation section below, involve measurements on surgically exposed tissue. The main difficulty with using imaging systems based on Eq. (11) and the LASCA system is that the relationship between speckle contrast and RBC velocity is a non-linear one. Forrester et al. have developed an improved version of Briers’ LASCA system called LSPI [69]. A different approach to producing the contrast image is taken by first calculating a reference image by averaging the speckles spatially and temporally. The intensity fluctuations due to blood flow can then be examined by calculating the intensity difference (ISD) between the captured speckle images and the corresponding pixels in the reference image (Eq. (13)). The result is then normalised by the reference in order to negate the effect of non-uniform illumination, or changes in laser output or tissue optical properties. For N successive speckle images and an ‘averaging square’ of i  j pixels, the speckle difference intensity, ISD at pixel (x, y): I SD ðx; yÞ ¼

Fig. 12. Full-field laser speckle contrast imaging. The tissue is illuminated by a laser that has been diverged by a lens. The speckle pattern is then recorded by the CCD camera and the speckle contrast, K, computed.

1 I REF ðx; yÞ " !# yþj NX xþi X MAX X I SP;N ðx; yÞ  I REF ðx; yÞ ,  N¼1

xi

yj

ð13Þ where IREF is the reference intensity.

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Fig. 13. Laser speckle contrast imaging. Images obtained from a commercial laser speckle imager, moorFLPI show false-colour maps of speckle contrast related to the velocity and concentration of red blood cells. Images of rodent microcirculation in the spinotrapezius muscle (a) and femoral artery (b) are shown. Courtesy of Dr. Egginton [82].

ISD is inversely proportional to the speed of the scatterers: in low flow regions, a sharp speckle contrast is observed and a large difference between the reference and speckle intensities will be recorded. In high-flow regions, a large amount of decorrelation occurs and the difference approaches zero. This inverse relationship can be well represented by a linear equation and the creation of a ‘blood flow index’, IBF(x,y):   1 I BF ðx; yÞ ¼ A0 þ A1 , (14) I SD ðx; yÞ where A0 and A1 are empirically determined constants. It was found that for capillary blood flow ranges and a camera integration time of 16 ms, the relationship between IBF and ISD was successfully modelled by Eq. (14). An in vitro blood flow model was devised to test the response of the LSPI system [83] consisting of a 0.95 mm bore glass tube fixed in a gelatine-set tissue phantom 1 mm below the surface. Red cell concentrations of 0.1–5% were imaged at flow rates in the range 0–800 mL min1 (corresponding to RBC velocities of 0–18.8 mm s1). These parameters cover the typical physiological range [87]. In comparison to a commercially available laser Doppler imager (moorLDI), it was found that at a constant flow rate, LSPI had a similar response to LDPI to changes in RBC concentration (correlation r2 ¼ 0.93). At a constant concentration (1%), LSPI also had a similar response to LDPI to changing RBC velocity, with its perfusion index increasing linearly over the range of flow rates (correlation r2 ¼ 0.99). Speckle imaging systems using the ‘averaging window’ suffer from a loss of resolution (for example, a 7  7 pixel window will produce a speckle image 1/49th the

resolution of the original). However, the temporal averaging technique used by the LSPI algorithm means that the full resolution of the CCD could be retained to produce high-detail images of the vascular structure. For higher speed acquisition however, a spatial averaging mode may be employed (i ¼ j ¼ 3) [69]. The main advantage over LDPI is the speed at which the images may be acquired. Acquisition speeds at video rates (25 frames s1) can be attained, making it possible to measure the blood flow response to occlusion and hyperaemia (typical LDPI scans are of the order of minutes). Also, the much higher resolution (limited only by the CCD, maximum LDI resolution is 256  256 (moorLDI)) enables more detailed imaging of tissue structures [69]. An example of this high spatial and temporal resolution is shown in Fig. 14. Microcirculation The early use of laser speckle imaging of the microvasculature has been reported [67]. However, more recent work incorporating the improvements outlined above, and in particular the advent of video rate acquisition has led to some interesting applications. Human retinal blood flow has been examined using the technique [77] in a similar method to Briers et al. and an in vitro calibration model was developed to relate speckle decorrelation to actual readings of velocity. In vivo experiments conducted in parallel with electrocardiogram measurements showed that the speckle contrast varied with the heart rate and RBC velocity reached a peak corresponding to that of the cardiac cycle. Yaoeda et al. [88] adapted the technique developed by Fujii et al. [70] to monitor blood flow to the optic nerve head showing differences between the right and left eyes.

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Fig. 14. High-resolution LSPI images: (a) monochrome photo of laser-illuminated tissue. The images (b)–(d) show blood flow in the finger tips in normal conditions (b), occlusion (c) and hyperaemia (d). Scan times for conventional laser Doppler scanners are too long to capture the hyperaemic reaction in this way. Reproduced with permission from the IFMBE [69].

The LSPI system developed by Forrester et al. [69] was used to measure changes in flow in the knee joint capsule of a rabbit before and after femoral artery occlusion, recording a 56.3% decrease after the manoeuvre. The system was also incorporated into a handheld endoscope for the simultaneous measurement of the same event. The response was validated as before with an in vitro model, showing the linear response over a large flow range (0–800 mL min1). Equipped with a motion-detection algorithm to limit motion artefact, the endoscopic set-up recorded a similar decrease (58.7%) [73]. Clinical data from human subjects has also been acquired using this endoscopic LSPI system, recording a decrease in blood flow in the medial compartment of the knee after application of a tourniquet. A dose-dependent response to the vasoconstrictor epinephrine was also recorded [89]. Choi et al. have demonstrated the use of laser speckle imaging in a rodent skin fold model [72] as shown in Fig. 15. It was possible to provide quantitative measurements of relative blood flow in areas of tissue that were otherwise obscured by attached muscle and fat. Selected areas of the tissue were irradiated with a 585 nm laser pulse to stop blood flow locally and this was seen in the LSI images (Fig. 15(c) and (d)). Similar animal models were investigated by Cheng [71] while testing the effect of varying doses of phentolamine, a vasodilator on intestinal tissue. In both cases, image processing was conducted offline. Extensive studies of cerebral blood flow in rodents have been undertaken using various laser speckle

imaging techniques to monitor the vascular response to stimuli and chemically induced changes. Cerebral blood flow was monitored in rats to demonstrate its heterogenic response to mild hypotension (induced by withdrawal of blood from the femoral artery). Regional variations in perfusion were recorded [90]. Blood perfusion measurements have also been made on the surgically exposed cerebrum of rodents that show a drop in signal over time after a photochemically induced infarction [76] while Royl et al. [91] recorded an increase in flow in the somatosensory cortex during electrical stimulation of the forepaw. Laser speckle measurements of cerebral blood flow through the intact rat skull (minus the skin) have been presented by Li et al. [75]. A temporal averaging algorithm was employed that used data from 40 consecutive speckle patterns and computed the difference between the intensity at a pixel of interest and the average intensity calculated for that pixel over the 40 frames. The resulting so-called laser speckle temporal contrast analysis image was found to be more immune to the effects of static scatterers (such as the skull) than spatial techniques such as LASCA. The effect of PDT and pulsed dye laser irradiation on the microvasculature in a rodent dorsal skin-fold model was examined using LASCA [92]. Images taken 18 h after treatment show vascular flow reduction in the exposed area (Fig. 16). The versatility of the technique has been illustrated by conducting laser speckle analysis in parallel with PDT using the 514 nm argon ion PDT laser as the source for

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Fig. 15. LSI images of rodent skin fold model. The microvasculature is clearly visible in the speckle image (b) through the attached fat and muscle seen in (a). 585 nm laser pulses were incident on the circled areas in (b) and (c) leading to an arrest in blood flow locally (d). Reproduced with permission from Elsevier [72].

Fig. 16. Laser speckle contrast maps of blood flow (a) prior to PDT treatment, (b) immediately after and (c) 18 h after. Reproduced with permission from Wiley-Liss, Inc., a subsidiary of John Wiley & Sons, Inc. [92].

the speckle measurements [74]. The processing algorithm was the speckle reduction technique developed by Forrester et al. [69]. Other animal models studied include temperature-induced changes in blood flow in the rodent [79] and wound-healing in porcine subjects [84]. Agreement with laser Doppler While accurate absolute measurements of RBC velocity cannot be calculated, all of the above LSI systems can successfully track relative changes in perfusion. Recent developments and in vitro experiments have shown that laser speckle images can respond in the same way as laser Doppler to changes in RBC concentration and velocity [83]. In vivo experiments have also shown a strong correlation between the techniques [91]. In fact, recent studies on burn scars show a correlation with an r2 value of 0.86 between the

techniques [93]. It should also be noted that the perfusion measurements correlated strongly with the clinical grading of the scar. As each imager used the same wavelength (632.8 nm), Stewart et al. stated that the penetration depth was in theory, the same for each [93]. However, Moor Instruments Ltd. report that the penetration depth of the ‘FLPI’ system is less than that of the corresponding LDI system. This is due to two factors: firstly, unlike the laser Doppler signal, the laser speckle signal is not frequency weighted, meaning that photons arriving from the faster-moving RBCs in deeper blood vessels cannot be isolated. Secondly, as the source in laser speckle systems is an expanded beam, the power density is much lower than the corresponding Doppler system; hence the depth at which noise dominates is more superficial. Laser speckle, which is a confounding phenomenon in some applications such as laser Doppler [95], is proving

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to be a very useful tool in blood perfusion measurements. Full-field images can be recorded in a single shot containing information on RBC concentration and velocity. While traditional laser Doppler images can take several minutes to acquire and are of limited resolution (256  256 pixels), LSI can show rapid changes in blood flow in real time. The first commercial speckle imager, moorFLPI, can acquire and process images at video rates (25 frames s1) at high resolution (576  768 pixels) [94]. The most popular processing algorithm reported is that based on speckle contrast maps (LASCA) [71,72,91] and it is this method that is used in the commercial device [94]. However, a newer algorithm has improved on this, responding linearly to increasing RBC velocity, limiting the effect of static scatterers and specular reflection, and compensating for variations in tissue optical properties. This ‘reduced speckle’ technique is promising as it follows laser Doppler measurements very closely but at a much higher resolution and speed [69,73,74,83,93]. The low-cost nature of the system, acquisition speed, resolution and compatibility with existing laser Doppler measurements make LSPI a very attractive technology for studies of blood perfusion in the microvasculature.

Polarisation spectroscopy Introduction The technique of polarisation spectroscopy allows the gating of photons returning from different compartments of skin tissue under examination [96]. Fig. 17 details the operation of basic polarisation imaging, showing that by use of simple polarisation filters, light from the superficial layers of the skin can be differentiated from light backscattered from the dermal tissue matrix. When monochromatic or white incoherent light is linearly polarised by a filter and is incident on the surface of the skin, a small fraction of the light (approximately 5%) is specularly reflected as surface glare (Fresnel reflection) from the skin surface due to refractive index mismatching between the two media. A further 2% of the original light is reflected from the superficial subsurface layers of the stratum corneum. These two fractions of light retain the original polarisation state, determined by the orientation of the first filter. The remaining portion of light (approximately 93%) penetrates through the epidermis to be absorbed or backscattered by the epidermal or dermal matrix. Approximately, 46% of this remaining light is absorbed by the tissue and not re-emitted, while 46% is diffusely backscattered in the dermal tissue. This backscattered portion is exponentially depolarised due to scattering events by chromospheres present in the tissue [97], and also by tissue birefringence due to collagen fibres [98].

The depth of penetration of polarised light is heavily dependent on the optical properties of the medium at each wavelength present [99]. Upon re-emerging from the tissue structure as diffusely reflected light, the remaining fraction of light is almost completely depolarised, and consists of two 22% fractions of parallel and perpendicular polarisations, with respect to the direction of the original filter. This light contains information about the main chromophores in the epidermis (melanin) and dermis (haemoglobin), while the surface reflections contain information about the surface topography, such as texture and wrinkles. On detection by a CCD or similar light collecting device, one can differentiate between detecting the surface reflections or a fraction of the diffusely backscattered light by placing another polarisation filter over the detector parallel or perpendicular with respect to the direction of the filter over the light emitter. With both filters oriented parallel or perpendicular with respect to each other, co-polarised (CO) or cross-polarised (CR) data, respectively, is obtained. This allows the gating of photons, and the technique is based on the assumption that weakly scattered light (the surface reflections) retains its polarisation state, whereas strongly scattered light will successively depolarise with each scattering event. It has been suggested that more than 10 scattering events are required to sufficiently depolarise light [100,101]. Current technology Polarisation gating has been employed to many different technologies, in order to investigate surface details of the skin structure by accepting light scattered

Fig. 17. Fundamental operation of polarisation spectroscopy. Remaining percentages of light intensity are illustrated by I, and 100% intensity is observed after the polarisation filter. SR represents specular reflection, a combined effect of Fresnel reflection and light returning after few scattering events from the upper layers of the epidermis. LP and DP represent linear and depolarised light, respectively, while ICO and ICR represents the remaining intensity, which contributes to CO or CR data, depending on the two possible states of polarisation filter PF2. abs. represents the percentage of light absorbed in the tissue. PF2 can be arranged with pass direction parallel (CO) or perpendicular (CR) to PF1.

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from the superficial layers of the tissue. Examples of technologies, which have applied polarisation filtering include polarisation sensitive optical coherence tomography (PS-OCT), LDPI, Raman spectroscopy, and simple microscopy. The use of polarisation filters has been shown to reduce tissue motion artefact in LDPI, while also reducing the overestimation in LDPI readings due to the amplification of specular reflection [102]. Detail of tissue birefringent axes from various layers via phase sensitive light recording can be extracted by PS-OCT, and the technique can be used to generate three-dimensional (3D) images of the polarisation state of backscattered light from skin tissue in vivo [103,104]. Raman spectroscopy of layered media will detect chemical signatures from multiple layers, and polarisation gating has been successfully employed to generate Raman spectra of only the superficial skin layers [105]. The combination of polarisation gating with multiphoton excitation microscopy allows the characterisation of fibrous structures in the skin layers, thus verifying the mechanics of transdermal drug delivery and the significance of dermal structural dermal changes [106]. Gating has also been used with laser speckle imaging in the reduction of specular reflection for investigation of the blood vessels in the microcirculation [69]. Polarisation cameras and probes employing the gating technique have been developed in order to investigate both the physiology and microcirculation of skin tissue [107,108]. The orthogonal polarisation spectroscopy (OPS) technology is a single wavelength (548 nm – an isobestic point of haemoglobin close to peak absorption) video microscopy technique operating at 30 frames s1 (fps) which increases contrast and detail by accepting only depolarised backscattered light from the tissue into the probe. Single blood vessels are imaged in vivo at a typical depth of 0.2 mm and magnification of 10  between target and image, and information about vessel diameter and RBC velocity are easily obtained. OPS in human studies is limited only to easily accessible surfaces, but can produce similar values for RBC velocity and vessel diameter as conventional capillary microscopy at the human nailfold plexus [109]. It has also been applied to study superficial and deep burns to the skin [110,111], adult brain microcirculation, and preterm infant microcirculation [112]. Polarisation gating images by sequential acquisition of superficial photons (CO) and photons undergoing multiple scattering (i.e., photons interacting with the reticular dermis – CR) have shown cancer, lesion and burn scar boundaries. A fluorescence method of imaging skin tumours stained with fluorescent dyes rejects excitation light by way of sequential CO and CR image acquisition with a tuneable monochromatic source, and details tumour boundaries similar to Mohs surgen boundaries [113,114]. A similar cost-effective method combining polarisation imaging with spectroscopy

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technique has also been developed to visualise water content in the superficial layers of the skin, and has shown good comparison to a well-recognised capacitance measurement technique [115]. Another method of polarised light imaging to examine surface skin pathologies in greyscale at 435  548 pixels is capable of distinguishing between freckles, nevii and carcinomas [116,117]. An extension of this preliminary technology produced a handheld cost-effective polarisation imaging camera, which operates at 7 fps at 400  400 pixels (15 fps at 250  250 pixels) [117]. Parallel and perpendicular images are acquired simultaneously from 2 CCDs, and an image processing algorithm is applied which examines only the histological aspects of the skin surface, with no investigation of the underlying microcirculation. The prototype technology has been successfully examined on melanoma margins. Image and data processing Image (or data) processing algorithms have generally employed an equation, which requires acquisition of both CO and CR images in order to investigate the superficial layers of the tissue [106,107,114–116]. The equation represents a normalised difference between the CO and CR images: P¼

CO  CR , CO þ CR

(15)

where P represents a composite image observing surface skin histology. The denominator in Eq. (15) represents the normalisation of the image, which cancels any nonuniform illumination. P images have been shown to be comparable to histopathology examination, the invasive gold standard of skin examination. Recent technology [118] analyses only the RBC concentration in skin microcirculation by way of single image acquisition or video acquisition by low-cost cameras utilising orthogonally placed polarisation filters. For single image acquisition, the flash of a consumer-end RGB digital camera is used as a broadband light source, and the camera CCD acquires an instantaneous image in three 8-bit primary colour planes. Colour filtering is performed on-camera by three 100 nm bandwidth colour filters, blue (E400 –500 nm), green (E500–600 nm) and red (E600–700 nm). This represents the first low-cost technique designed to image in real-time tissue haematocrit. Using polarisation imaging theory and Kubelka– Munk theory a spectroscopic algorithm was developed which is not dependent on incident light intensity, taking advantage of the physiological fact that green light is absorbed more by RBCs than red light. For each arriving image, the equation applied is M out ¼

M red  M green pððM red M green Þ=M red Þ e , M red

(16)

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where Mred and Mgreen represent the red and green colour planes of the image, and p represents an empirical factor to produce the best linear fit between output variable (called tissue viability index – TiVi index) and RBC concentration. Mout represents a matrix of maximum 3648  2736 TiVi index values for single image acquisition. Fig. 18 shows an example of a TiVi image colour coded in a similar method to laser Doppler images where high and low RBC concentration is represented by blue and red colour, respectively. The technique has high spatial and temporal aspects, with lateral resolution estimated at 50 mm [119]. The sampling depth in Caucasian skin tissue is estimated at approximately from Monte Carlo simulations of light transport in tissue [118]. Topical application of vasoconstricting and vasodilating agents demonstrate the capacity to document increases (erythema) and decreases (blanching) in RBC concentration [120]. This research group is currently developing a realtime system, whereby RBC concentration can be displayed online. Current technology can capture CR frames at 25 fps at a frame size of 400  400 pixels. Video studies will lead to increased temporal resolution, with 0.04 s between each image. Thus theoretically, the human pulse can be resolved from the time average (duplex) mode of the video imager, and also spatial data over time can be observed. Fig. 19 shows three processed video frames of the human fingers undergoing occlusion of the cephalic vein for the upper arm, with one frame pre-occlusion, under occlusion of 80 mmHg and one image post-occlusion. It can be seen that 30 s after the occlusion has been released; the pre-occlusion RBC concentration level has not been reached. Higher RBC concentration occurs under occlusion due to the RBCs being trapped by the occlusion of the cephalic vein, as the brachial artery pumps RBCs into the area. An example of the duplex mode is shown in Fig. 20, detailing the average TiVi index of a 50  50 ROI on the volar forearm for an

acquisition of 25 fps for cephalic vein occlusion. Clear vasomotion of the vessels can be observed at a rate of approximately 5 cycles min1 during and post-occlusion and release of the area. Future work will investigate pulsations of the vascular bed, and pulse models have already showed promising results. Limitations of this new technique include the lack of SI units in measurements, and inaccurate measurements on persons with dark skin due to high absorption of melanin. Light levels are also reduced due to the polarisation filters, with only 22.5% of total light intensity reaching the detector. The effect of melanin in the epidermis causes an increase in TiVi index. Image subtraction algorithms have been developed so as to display only the reaction on the skin site, while subtracting out the static components in the CR images. This is also helpful in removing the unwanted effect of deep veins in the volar forearm, a site, which is used commonly in routine skin tests. Owing to the availability of measuring tissue haematocrit only, low cost, ease-ofuse and portability, the TiVi imaging systems are an attractive alternative to expensive and cumbersome equipment such as LDPI and LSPI for investigation of tissue RBC concentration.

Photo-acoustic tomography In recent years, photo-acoustic tomography (PAT) has gained considerable interest. This is because the technique seeks to combine the best features of biophotonics with those of ultrasound; it offers high optical contrast and high resolution. The technique operates by irradiating a sample with a laser pulse. When the laser light is absorbed it will produce a pressure wave due to an increase in the temperature and volume. These pressure waves will emanate from the source in all directions. A high frequency ultrasound receiver can then be used to detect these waves and build a 3D picture of the structure. This is made possible by the

Fig. 18. Example of a CR and TiVi image. The CR image was taken of a UVB burn to the volar forearm of a young healthy volunteer. The image was taken 30 h after a 32 s exposure to an 8 mm wide UVB source. The TiVi image shows spatial heterogeneity in the blood distribution over the affected area. Notice that hairs have no effect on algorithm performance.

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Fig. 19. 3 video frames displaying RBC concentration in the fingers by occlusion of the cephalic vein by a sphygmomanometer placed on the upper arm. Increased RBC concentration occurs by the brachial artery pumping RBC in, but no return is possible due to occlusion the return veins.

Fig. 21. Schematic of photo acoustic tomography system [122]. Fig. 20. Time trace of a 50  50 pixel regions of the volar forearm of a subject. A pressure cuff on the upper forearm causes an occlusion of the cephalic vein. This demonstration shows clear vasomotion during occlusion and release, and further investigation will focus on pulsatile activity.

much lower speed of sound and lack of scattering or diffusion of ultrasound en route to the receiver. Diffuse, randomised scattering of the light before reaching the microvessels is not a disadvantage and is likely to be useful in providing relatively uniform illumination even to vessels which could not receive ballistic photons due to other vessels in the direct light path. In PAT light in the visible to NIR spectrum is used. This is because haemoglobin in the blood dominates the optical absorption in most soft tissue. If a tissue sample is irradiated by light in this range a map of the vessel network within the tissue can be produced. The pulse width is chosen such that it is much shorter than the thermal diffusion time of the system this way heat conduction can be neglected [121]. A schematic of the system is show in Fig. 21. The data from the transducer is processed by a computer and can be output in an image format. Image reconstruction is achieved by various reconstruction algorithms [123,124].

The technique is capable of diagnosing different physiology conditions. This is because different physiological conditions can change the light absorption properties of the tissue; for example reports of absorption contrast between breast tumours and normal breast tissue have been as high as 500% in the infrared range [124]. The technique is also capable of detecting haemoglobin levels in tissue with very high contrast. One group [126] has used the technique to acquire 3D data of the haemoglobin level in rat brain. This was achieved using a 532 nm laser source, as the absorption coefficient at this wavelength of whole blood is 100 cm1 is much larger than the averaged absorption coefficient of the grey and white matter of the brain 0.56 cm1. The main advantage of combining ultrasound and light is that you can get much improved resolution (diffraction limited) and can image vessels with 60 mm resolution [122]. The Wang lab at University of Washington in St. Louis are building a system which should get down to 5 mm resolution and have already been able to produce exquisite 3D images of microvascular architecture in the rat. One issue with the technique is the time it takes to produce large 3D scans of a region. Times of up to 16 h have been reported to image a region of radius 2.8 cm in

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the XY plane and a depth 4.1 cm in the Z plane [125]. However, the use of an ultrasonic transducer array would greatly improve imaging time.

Resu´men

Discussion and conclusion

La luz visible e infrarroja cercana, particularmente en el rango de longitudes de onda entre 600 y 1100 nm, ofrece una ventana espectral en los tejidos humanos y animales debido a su baja dispersio´n y absorcio´n. En este trabajo hemos repasado los principales me´todos biofoto´nicos aplicados a la visualizacio´n y al estudio de la microcirculacio´n, y documentamos el progreso obtenido en los u´ltimos 10 an˜os. La aplicacio´n de estos procedimientos, particularmente en el caso de piel humana, es de gran importancia debido al mejor conocimiento de su rol y de su valor como substituto de otros o´rganos en el estudio de drogas, en un momento en el que el desarrollo de nuevas medicinas se encuentran bajo una gran presio´n.

We have discussed the modus operandi and applications of capillaroscopy, LDPI, LSPI, PAT and TiVi. Each has had its own peak time and all retain niche applications. Proper application of each is dependent on user knowledge of light interaction with tissue and the basic workings of the device used. Choice of technology for in vivo imaging of the microcirculation is dependent on the sampling depth (nutritional or thermoregulatory vessels), resolution (physical and pixel), field-of-view, and whether structure or function is to be investigated.

Acknowledgements This research is supported in part by IRCSET (project RS/2003/135), the Irish Research Institute for Science, Engineering and Technology, and by VINNOVA (projects P26825-1 and P27840-1), the Swedish agency for Innovation Systems.

Zusammenfassung Methoden der Biophotonik zur Darstellung der Mikrozirkulation Der sichtbare und NIR-Bereich des elektromagnetischen Spektrums, besonders der Wellenla¨ngenbereich zwischen 600 und 1100 nm, bietet sich fu¨r optische Untersuchungen von humanem und tierischem Gewebe aufgrund einer reduzierten Streuung und Absorption der Proben in diesem spektralen Fenster an. In diesem Beitrag fassen wir die Methoden, welche im Bereich der Biophotonik zur Darstellung der Mikrozirkulation hauptsa¨chlich eingesetzt werden, zusammen. Hierbei soll besonders auf die Entwicklungen der letzten 10 Jahre eingegangen werden. Die Anwendung solcher bildgebenden Verfahren, besonders fu¨r menschliche Haut, ist von besonderem Wert – vor allem wegen des Erkenntnisgewinns u¨ber den Einsatz und die Bedeutung von Hautgewebe als Ersatz fu¨r anderes Organgewebe in Medikamententests. Dies ist in einer Zeit, in der die Entwicklung und Tests von Medikamenten unter einem großen o¨ffentlichen Druck stehen, von besonderer Relevanz. Schlu¨sselwo¨rter: Biophotonik; Bildgebung; Mikrozirkulation; Laser Speckle; Laser Doppler; Perfusionsmonitoring; Kapillarmikroskopie; Vitalita¨tsmessung (TiVi); Photoakustische Tomografie (PAT)

Me´todos biofoto´nicos en el monitoreo de la microcirculacio´n

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