PHOTOACOUSTIC DISCRIMINATION OF VIABLE AND THERMALLY COAGULATED BLOOD FOR BURN INJURY IMAGING

Transcription

1 PHOTOACOUSTIC DISCRIMINATION OF VIABLE AND THERMALLY COAGULATED BLOOD FOR BURN INJURY IMAGING A Thesis presented to the Faculty of the Graduate School at the University of Missouri-Columbia In Partial Fulfillment Of the Requirements for the Degree Master of Science by ROBERT JOHN TALBERT Dr. John A. Viator, Thesis Supervisor AUGUST 2007

2 The undersigned, appointed by the Dean of the Graduate School, have examined the thesis entitled: PHOTOACOUSTIC DISCRIMINATION OF VIABLE AND THERMALLY COAGULATED BLOOD FOR BURN INJURY IMAGING presented by Robert John Talbert, a candidate for the degree of Master of Science, and hereby certify that, in their opinion, it is worthy of acceptance. Dr. John A. Viator, Biological Engineering Dr. Gang Yao, Biological Engineering Dr. Scott Holan, Statistics

3 This Work Is Dedicated To My Mother Patricia Furst April 17, 1950 Present

4 ACKNOWLEDGEMENTS First and foremost, I would like to thank my advisor, Dr. John Viator, for giving me the opportunity to conduct this research. I would also like to thank him for the exceptional guidance and support he has provided me. I came into this Master s program with very little research experience, but I am now leaving with a wealth of knowledge and experience, thanks to him. Dr. Viator has given me the tools that are necessary to conduct and present both beneficial and successful research. I cannot thank him enough for all that he has done for me. I must thank Dr. Gang Yao for playing such a large role in my understanding of biomedical optics. Without the material I learned in his excellently crafted and taught biomedical optics courses, this work would not be possible. I would also like to thank Dr. Yao for serving as a member of my thesis committee. I would like to thank Dr. Scott Holan for the outstanding statistical contribution he has made to this work. His expertise was much appreciated. I would also like to thank Dr. Holan for serving as a member of my thesis committee. I would like to acknowledge Emily Spradling, an undergraduate student in our lab, for her contribution to this work. Even in the wake of her apartment being burned down, she still managed to produce excellent research. I cannot thank her enough. I would also like to thank Dr. Trista Strauch and August Rieke, of the Department of Animal Science, for providing the pig blood necessary to perform these experiments. In addition, I would like to thank Dr. Carol Lorenzen and Rick Sutcliffe, of the Department of Meat Science, for their help in providing pig skin. I would like to thank the members of the Department of Dermatology for providing additional support that was necessary for this work. I acknowledge my former lab members, Melissa Lyons and Kevin MacDonald, for their assistance with this work. ii

5 I would like to thank Dr. Sherman Fan for allowing us to share equipment with his lab. I would also like to thank the members of his lab for their support and friendship. I also want to acknowledge my fellow graduate students and compadres, Melvin Sims, Ryan Weight, and Lisa Huhman. They were very helpful and a blast to be around. I quite enjoyed working with the other undergraduate and high school students and would like to acknowledge them, too. I acknowledge the support of the Biological Engineering Department and the Christopher S. Bond Life Sciences Center for financial support and facilities during this study. Finally, I extend my greatest thanks to my family and my girlfriend, Dawn Stocker, who have been of the utmost support to me, in this, and in all of my endeavors. iii

6 TABLE OF CONTENTS ACKNOWLEDGEMENTS... LIST OF FIGURES... ABSTRACT... ii vii xi Chapter 1 Burns The Burn Problem The Burn Wound Anatomy and Physiology of Normal Skin Pathophysiology of Thermal Burn Injury Zone Classification of Burn Wound Burn Depth Classification and Treament Surgical Excision and Grafting Burn Depth Estimation Competing Technologies & Techniques Photoacoustic Estimation of Depth Photoacoustics Introduction History Biomedical Photoacoustics Photoacoustics in Biological Tissue Medical Imaging Other Applications Motivation for Burn Imaging Photoacoustic Theory iv

7 2.5 Photoacoustic System Laser System Acoustic Detection Photoacoustic imaging of viable and thermally coagulated blood Introduction Statistical Method for Classifying Coagulated and Non-coagulated Blood Measure of Future Performance Materials and Methods Spectral Analysis of Coagulated and Non-coagulated Blood Experimental Setup for Photoacoustic Measurements Classification Analysis of Coagulated and Non-coagulated Blood Results Spectral Analysis of Coagulated and Non-coagulated Blood Photacoustic Discrimination of Planar Blood Samples Photoacoustic Imaging Discussion Biophysical Changes in Hemoglobin Photoacoustic 543/633 nm Ratios Photoacoustic Classification and Imaging Conclusions Photoacoustic determination of optical absorption coefficient of thermally coagulated blood in the wavelength range from 580 to 700 nm Introduction Materials and Methods Extraction of µ eff from Photoacoustic Waveforms Approximation of µ a Photo- and Thermal Stability of Chlorazol Black Dye Experimental Set-up for Photoacoustic Measurements v

8 4.2.5 Phantom Construction Photoacoustic Spectroscopy of Intralipid Photoacoustic Spectroscopy of Thermally Coagulated Blood Results Photo- and Thermal Stability of Chlorazol Black Dye Photoacoustic Spectroscopy of Intralipid Photoacoustic Spectroscopy of Thermally Coagulated Blood Discussion Thermal Stability of Chlorazol Black Dye Photoacoustic Spectroscopy of Intralipid Photoacoustic Spectroscopy of Thermally Coagulated Blood Conclusion General Discussion and Conclusions Photoacoustic Imaging of Viable and Thermally Coagulated Blood Photoacoustic Determination of the Optical Absorption Coefficient of Thermally Coagulated Blood in the Wavelength Range from 580 to 700 nm Conclusion BIBLIOGRAPHY VITA vi

9 LIST OF FIGURES Figure Page 1.1 The anatomy of normal skin is shown here. The epidermis lies above the dermis, which lies above the subcutaneous fat (fat cells) The different zones of the burn wound are shown here. The zone of hyperemia lies below and to the periphery of the zone of stasis, which lies below and to the periphery of the zone of coagulation. The zone of stasis is a dynamic layer that can remain viable or become necrotic over a period of time The various classifications of burn depth are shown here A deep partial thickness burn is shown here. The burn wound extends deep within the dermis and causes extensive damage to the blood vessels and nerves. Treatment of this type of wound is complex and healing can take weeks Surgical excision, shown here, involves the removal of necrotic tissue from the burn wound. A dermatome is used to remove thin layers of necrotic tissue from the wound until only viable tissue is left Shown here, is the healthy skin from a donor site that has been meshed prior to being grafted to the recipient site vii

10 1.7 A punch biopsy, shown here, is used to remove a portion of skin from the burn wound. The specimen of skin is then histopathologically analyzed to determine burn depth A laser pulse of duration, τ p < δ/c s, is incident upon a planar absorber. Stress confinement is achieved, a thermoelastic expansion occurs, and a pressure wave is formed An initially transmitted pressure distribution, p 0 (z) trans, heads downward towards the acoustic transducer. The initially reflected pressure distribution, p 0 (z) refl, heads away from the transducer before being reflected back towards the transducer at the free boundary that exists between air and the absorbing medium Time-resolved, transmission mode pressure signal from planar absorber. The T portion of the waveform represents the initially transmitted pressure distribution, p 0 (z) trans. The R portion of the waveform represents the initially reflected pressure distribution, p 0 (z) refl Distance-resolved, transmission mode pressure signal from planar absorber. The surface of the absorber is located approximately 1.81 cm above the transducer Initial pressure distribution within the absorbing medium as a function of depth, z, into the medium The laser system is shown here. An Nd:YAG laser produces 1064 nm pulsed laser light, which is reduced to a frequency of 355 nm, via a second and third harmonic generator, before reaching the OPO. The OPO can convert the 355 nm laser light to a user-specified wavelength over a range from 410 to 2400 nm. (Photograph courtesy of R.M. Weight) viii

11 2.7 The activity of PVDF film is shown here. When PVDF film is subjected to a mechanical stress, such as that from a pressure wave, a net polarization occurs that is proportional to the stress. The net polarization creates a voltage, which can be measured using an oscilloscope The transmission set up was performed by placing blood samples in polyacrylamide cylinders placed on an acoustic transducer The experimental set-up used in the photoacoustic analysis of coagulated and non-coagulated blood. A burn phantom containing coagulated and non-coagulated blood was irradiated with 543 and 633 nm light from a optical parametric oscillator. A piezoelectric probe detected the photoacoustic pressure wave as it propagated away from the upper surface of the phantom (a) Schematic depicting the burn phantom and photoacoustic measurement setup. (b)photograph of the experimental set-up. The coagulated and non-coagulated blood vessels have been highlighted to indicate their presence just under the surface of the Intralipid/Direct Red solution Average optical absorption spectra for coagulated and non-coagulated blood. Arrows indicate the location of 543 and 633 nm Pressure waveforms resulting from 543 and 633 nm laser excitation for (a) non-coagulated blood and (b) coagulated blood. Peak-peak measurement points are indicated Two-dimensional images produced to highlight different blood types types. (a) Viable blood was highlighted in the image using P expected 543/633 = (b) Thermally coagulated blood was emphasized using P expected 543/633 = ix

12 4.1 (a) Acoustic amplitude as seen at the oscilloscope. (b) Acoustic amplitude as a function of depth into the sample The experimental set-up used to perform the photoacoustic measurements. A phantom containing the sample was irradiated by light from an optical parametric oscillator within the wavelength range, 500 to 700 nm. The induced acoustic wave was detected by a piezoelectric transducer on the side of the sample opposite to the laser irradiation Photoacoustic measurements were obtained by coupling the sample to a polyacrylamide gel cylinder, which was coupled to an acoustic transducer The change in the optical absorption spectra, µ a, of five samples of thermally heated Chlorazol Black dye is shown. The mean absorption coefficients are shown in bold black lines, while the 95% confidence bands are shown in gray Photoacoustically measured µ eff in (a) a relatively strong absorbing intralipid solution (gray dots) and (b) a relatively weak absorbing Intralipid solution (gray dots). The predicted µ eff values based on the known amount of added absorber, µ a, and the values of µ s obtained by Flock et al and van Staveren et al are also shown (a) An average value of µ eff (white diamonds), ˆµ eff (black diamonds), and ˆµ a (black dots) was found from the photoacoustic measurements of the coagulated blood of three separate pigs. (b) The average absorption coefficient, µ a, of coagulated blood from three separate pigs, which was derived from the approximation outlined in Section (a) Linear regression through the absorption coefficient values, µ a, within the 580 to 700 nm wavelength range. (b) Extrapolation of the regression to 500 nm x

13 PHOTOACOUSTIC DISCRIMINATION OF VIABLE AND THERMALLY COAGULATED BLOOD FOR BURN INJURY IMAGING Robert John Talbert Dr. John A. Viator, Thesis Supervisor ABSTRACT Early and accurate determination of burn depth is crucial to monitoring and treatment of the burn wound. One such treatment, surgical excision and grafting, involves removal of necrotic tissue from the wound and replacing it with healthy skin donated from another area of the body. We propose that a photoacoustically obtained depth profile of the burn wound, which delineates the boundary between necrotic tissue and viable tissue, would prove useful for this intervention. A simplified model of a dermal burn wound can be described as a layer of necrotic tissue, containing thermally coagulated blood, atop a layer of inflamed tissue that is characterized by the presence of viable (non-coagulated) blood. Using optical spectroscopy and photoacoustic spectroscopy, we show that it is possible to discriminate between coagulated and non-coagulated blood using a dual-wavelength photoacoustic method and, therefore, discriminate between the two layer types. A blood vessel phantom study confirmed the feasibility of this dual-wavelength photoacoustic technique. Finally, since little is known about the optical properties of thermally coagulated blood, we sought out to elucidate them. A novel photoacoustic method was used to derive the optical absorption coefficient, µ a, of thermally coagulated blood over the wavelength range from 580 to 700 nm. Additionally, we performed a linear regression on the 580 to 700 nm absorption spectrum and extrapolated it out to 500 nm, creating a theoretical 500 to 700 nm absorption spectrum for thermally coagulated blood. xi

14 Chapter 1 Burns 1.1 The Burn Problem Each year in the United States, it is estimated that 500,000 people seek medical attention for burns, 25,000 of whom will be admitted to a special burn care center due to the extent of their injuries. Of those hospitalized for burn injuries, approximately 1,000 will die [1]. Burns can result from heat, chemical agents, electricity, or radiation [2]. The focus of this work, though, is based on burn wounds that are the result of heat. Burn injuries present a host of complications and require victims to receive complex and specialized care in order to manage them [3]. Management of these complications includes compensation for fluid shifts and electrolyte imbalances, providing respiratory support, wound care, infection prevention, and occasionally treating sepsis and multiple organ failure [2, 3]. Over the past few decades, recent advancements in the management of burns, have been attributed to the marked improvement in the mortality rate and outcome of patients with severe burn trauma [2]. Essential to the diagnosis and treatment of these burn injuries, is the early assessment of burn extent and burn depth [4]. Assessment of burn extent, or the percentage of total body surface area (% TBSA) covered by burns, is important in determining nutritional and fluid requirements, estimating outcome, and deciding if admission to a specialty burn care center is necessary [3]. A method referred to as the Rule of Nines provides clinicians with a rapid and relatively accurate estimate of the extent of burn trauma. Although the Rule of Nines does have an element of inaccuracy it is not enough to be clinically irrelevant, as it is only necessary to get an estimate [3, 5]. Accuracy, however, is very important in the determination of burn depth and is the focus of this work. Many attempts have been made to develop an accurate means to do so and are described in Section 1.3. Before proceeding, though, 1

15 it is important to understand the relevance of burn depth. Section 1.2 addresses this by describing the physiology and anatomy of normal skin, the changes that take place when skin is subjected to thermal trauma, and how the various depths of injury are classified and treated. 1.2 The Burn Wound Anatomy and Physiology of Normal Skin Skin is the largest and one of the most important organs of the human body [3]. It serves a wide range of vital functions, which include [3, 6]: Mechanical protection - guards the body against the elements. Immunological protection - assists in presentation of antigens to immune system. Fluid, protein, and electrolyte homeostasis - controls the levels of these substances. Thermoregulation - prevents the excessive loss or gain of heat. Neurosensory - nerve endings enable the nervous system to process, receive, and interpret information from surroundings. Social-interactive - aids in social, interpersonal reactions. Metabolism - vitamin D production. Skin covers a surface area of 0.2 to 0.3 m 2 in newborns and 1.5 to 2.0 m 2 in adults [6]. It is composed of two distinct layers, the epidermis and dermis. The epidermis can range from 0.05 mm to over 1 mm thick, while the dermis is usually at least 10 times the thickness of the dermis [6]. In total, the thickness of the skin can range from 0.5 mm to 5 mm [7]. Thick skin is usually found on areas of the back and the soles of feet. Thinner skin can be found around areas such as the eyelids. 2

16 Men tend to have thicker skin than women. Children under the age of 5 and the elderly possess thinner skin, too [6]. As shown in Figure 1.1, the epidermis sits atop the layer of the dermis. The epidermis is a layer composed primarily of epithelial cells, more specifically keratinocytes [6]. These keratinocytes are constantly being sloughed off from the upper surface of the epidermis and are replaced by keratinocytes that arrive from the bottom, basal layer of the epidermis. At the basal layer of the epidermis, the keratinocytes begin as young and immature cells. As they migrate towards the surface, they mature by increasing their production of keratin. It is this increased amount of keratin production in the cells of the epidermal surface that imparts the epidermis with the important quality of being impervious to water and other environmental elements [3, 6]. Melanocytes can also be found within the epidermis. These cells produce melanin, which gives the skin its characteristic color and protects the body from ultraviolet irradiation [6]. Hair Epidermis Dermis Dermal Vasculature Nerve Ending Subcutaneous Layer (Fat Cells) Arteriole Venule Figure 1.1: The anatomy of normal skin is shown here. The epidermis lies above the dermis, which lies above the subcutaneous fat (fat cells). 3

17 Just below the epidermis and connected tightly by hemidesmosomes via the basement membrane and other associated fibrils, is the dermis [6, 7]. This area at which the two layers meet is referred to as the dermo-epidermal (D-E) junction. The dermis is a relatively thick layer composed of fibrous tissue, primarily fibroblasts, which excrete the proteins collagen and elastin into the intracellular matrix. Collagen imparts the dermis with stretchability and tensile strength, while elastin gives the dermis a certain amount resting tension [6]. Between these elements exists a ground substance, composed of glycosaminoglycans, proteoglycans and other similar macromolecules, which provide a semifluid matrix that lubricates the components of the dermis and provides a pathway for the diffusion of transient cells and nutrients through it [6]. The dermis can also be subdivided into two different layers, the papillary dermal layer and the reticular dermal layer. The papillary dermis is a thin layer of dermis located just below the D-E junction. The reticular dermis resides below this layer and constitutes a considerably larger portion of the dermis. In these layer lies two separate plexuses of vessels. In the papillary dermis, lies the superficial vascular plexus. Below that, in the reticular dermis, lies the larger blood vessels of the deep vascular plexus, which are connected to the superficial vascular plexus [6 8]. These vessels are supplied by a plexus of even larger vessels, the subdermal vascular plexus, which lies in the adipose tissue (fat cells) directly beneath the dermis [6, 7] Pathophysiology of Thermal Burn Injury Physical Injury Phase A direct, physical burn injury occurs when a portion of the skin is subject to enough thermal energy, for a long enough period of time, to cause irreversible damage to cells within that area. An early study by Moritz and Henriques [9] sought out to determine the relationship between the extent of injury and the temperature and time needed to create such injury [6]. Their findings indicated that when porcine skin was subjected to temperatures as low as 44 degrees Celsius, for an hour or longer, irreversible cellular damage was produced because of the denaturation of proteins and subsequent loss of repair mechanisms. Exposure to temperatures greater than 65 4

18 degrees Celsius, for a time as short as 2 seconds, was also found to produce irreversible damage to cells resulting in necrosis [6]. Even after removal of the thermal source causing direct injury, residual heat can continue to produce an indirect (secondary) thermal injury. This secondary trauma usually persists for 6-12 hours [10]. Inflammatory Reaction Phase Burn injuries are dynamic and progressive. In addition to the damage caused by heat, damage is also created as a result of the biochemical inflammatory reaction [10]. This reaction begins within in 1 hour of initial physical insult and continues for up to 72 hours. During the first stages of this reaction, the capillaries of the viable tissue, adjacent to the necrotic tissue, experience increased permeability. The result of this is the loss of fluid from the capillaries into the interstitial space, which is referred to as edema formation. It is a complex and not entirely understood process that causes the surrounding tissue to become ischemic and compromises oxygen delivery to that tissue [6, 10]. In addition to edema formation, the inflammatory reaction phase also causes progressive thrombosis of the microcirculation in the viable tissue adjacent to the necrotic tissue. The healthy tissue affected by this will, in turn, become necrotic as a result of compromised (hypoxia) or nonexistent (anoxia) oxygen delivery. This phase of the injury can last up to 72 hours [10]. Rejection of Necrotic Tissues At 72 hours, the physical injury and inflammatory reaction phases are followed by a rejection of necrotic tissue phase. This phase is the response of viable tissue cells to the disintegration of necrotic tissue and cells at the interface. The result of this rejection reaction can be: (1) disintegration of necrotic cells in the injury interface; (2) regeneration of viable cells in the injury interface; (3) microbial infection in the injury interface. The disintegration products of the necrotic cells is a liquefied product that contributes to the aggravation and accumulation of damage to the injured tissue. The liquefied products cause the viable cells to begin instinctive regeneration, which also increase the inflammation reaction. Although these pathological processes increase damage to the tissue and produce harmful systemic responses, 5

19 their presence is important, as they inhibit the growth of microbial flora within the burn wound. This rejection reaction process is the last process to occur [10] Zone Classification of Burn Wound To fully describe the morphology of burn injury pathology, Jackson subdivided the overall burn wound into three concentric zones [6, 11]: Zone of Coagulation The zone of coagulation, shown in Figure 1.2, is characterized by denatured proteins resulting from supraphysiological temperatures. All levels of protein architecture within this zone become altered [6]. This zone, as with the two other zones, is deepest in the center and becomes shallower towards the periphery of the injury location. Protein denaturation within this region leaves the cells irreversibly damaged [10]. The lumina of capillary vessels is obliterated and the blood they contained becomes thermally coagulated [6, 11]. This zone comprises the layer of burn eschar initially seen in a burn injury [6]. Zone of Stasis The zone of stasis lies below and to the periphery of the zone of coagulation (Fig. 1.2). It is a dynamic and important layer of the burn wound. The blood vessels within this region are initially patent and the cells are viable [11]. However, progressive microcirculatory thrombosis occuring as a result of heat injury to erythrocytes leads to circulatory stasis. The decreased blood flow within the region has the ability to cause cell death and tissue degeneration, essentially converting the zone of stasis into a necrotic eschar. Importantly, though, conversion to a necrotic layer is not always this zone s fate. Under optimal conditions circulation may be recovered and cell death ceased. However, those cells originally lost as a result of the circulatory stasis are not recoverable [6]. 6

20 Zone of Hyperemia Lying directly below and to the periphery of the zone of stasis is the zone of hyperemia (Fig. 1.2). This layer is a result of the inflammatory reaction that accompanies a thermal injury [6, 10]. This reaction causes blood vessels within the zone to become vasodilated, increasing the blood volume within the zone and giving it a reddish appearance. In the absence of infection or trauma, cells within this zone are fully recoverable from the pathomorphological conditions that exist, such as hyperemia, edema, anoxia, and exudation [6, 10]. Skin Surface Zone of Coagulation Zone of Stasis Zone of Hyperemia Figure 1.2: The different zones of the burn wound are shown here. The zone of hyperemia lies below and to the periphery of the zone of stasis, which lies below and to the periphery of the zone of coagulation. The zone of stasis is a dynamic layer that can remain viable or become necrotic over a period of time Burn Depth Classification and Treament As far back as 1676 attempts were made to classify the severity of burns by their depth [11]. Previous to that, assessment of burn injury was based upon the external appearance of the burn wound with no reference to depth. As a result, such diagnoses were unreliable and inaccurately characterized the burn injury. To properly treat burns, they must be classified by their depth. As shown in Figure 1.3, the depth of a burn depth and its required treatments can be classified into one of four categories: 7

21 Superficial Superficial burns involve only the epidermis (Fig. 1.3) [12]. They are rarely clinically significant other than being painful [6]. This type of burn is usually caused by UV radiation (sunburn) [12]. They have a red appearance and can blister [6, 12]. There is no direct blood supply to the dermis, so the erythema is due to vasodilation of the capillaries within superficial region of the dermis. This vasodilation is caused by irritation of the vessels resulting from thermal insult to the epidermis. Treatment required for such a thermal injury is minimal. It often consists of pain control, hydration, and moisturizing burned areas [12]. Superficial burns will heal within 3 to 7 days with no scarring [6, 12]. From Hettiaratchy, S. et al BMJ 2004; 329: , courtesy of BMJ Publishing Group. Figure 1.3: The various classifications of burn depth are shown here. Superficial Dermal (Superficial Partial Thickness) In a superficial dermal burn the epidermis is destroyed along with parts of the dermis (Fig. 1.3) [6, 12]. Their appearance is bright red with blister formation and a moist surface. They are often very painful because of the survival of nerve endings within the superficial dermis. Edema is also present because of the inflammatory 8

22 effects on the vasculature within the dermis [12]. As with a superficial dermal burn, the treatment required for a superficial dermal burn is minimal. Treatment above and beyond that for a superficial burn might include supportive dressing changes [6]. Bulge region stem cells located within the deeper areas of the dermis, around hair follicles and sebaceous glands, facilitate the re-epitheliazation process [6, 10, 13]. Scarring in such a burn injury is minimal due to the relatively quick wound closure from rapid re-epitheliazation [6, 10, 12]. However, pigmentation changes may result [12]. A significantly sized zone of stasis usually accompanies burns of this type and therefore the potential for conversion of deeper dermis to necrotic tissue does exist. Unless such a conversion occurs, healing should take place within 14 to 21 days [6, 12]. Deep Dermal (Deep Partial Thickness) Like a superficial dermal burn, a deep dermal burn extends through the entire epidermis (Fig. 1.3, 1.4) [6, 10, 12]. It is is different, though, in that it extends deeper into the dermis (reticular dermis) so that most of the dermis is destroyed [12]. The appearance of these wounds can be either moist or dry and exhibit a pale red or waxy white color. Although edema is present, little blistering occurs because the thick layer of eschar above the edematous layer resists lifting [6, 12]. Extensive damage to nerve endings results in loss of pain sensation [6]. However, pressure sensors which reside deeper within the dermis are subject to less damage and have the potential to respond to pressure. Treatment of deep dermal burns is crucial because of the many complications which result from such a drastic insult. Surgical excision and grafting, discussed in greater detail in Section 1.2.5, is the primary form of treatment [12]. Because of the depth from which migrating epithelial cells must travel, re-epithelization is greatly hindered in this type of wound [6]. Because of the greater length of time that the wound is exposed, great care must be taken to prevent infection. Prolonged wound exposure also draws out the inflammatory reaction phase, which causes collagen deposition to proliferate, thereby producing extensive scarring. After healing, the regenerated epidermal layer is often thin, fragile and nonfunctional. Healing of this 9

23 Hair Epidermis Dermis Dermal Vasculature Deep Partial Thicknes Burn Nerve Ending Subcutaneous Layer (Fat Cells) Arteriole Venule Figure 1.4: A deep partial thickness burn is shown here. The burn wound extends deep within the dermis and causes extensive damage to the blood vessels and nerves. Treatment of this type of wound is complex and healing can take weeks. type of wound can take many weeks and is dependent on the timing and level of intervention [6, 12]. Full Thickness Full thickness burns extend through the epidermis and dermis and into the subcutaneous fat [6, 10, 12]. Depending on the injury, the wound can appear charred, mottled, pale, waxy, yellow, brown or nonblanching red [12]. They are firm, dry, leathery and exhibit edema formation. All dermal appendages, including the nerve and pressure receptors, are damaged. Like a deep dermal burn, the treatment of these wounds is crucial. Surgical excision and grafting is important. Without it wound closure (healing) will be difficult because migration of epithelial cells can only occur from the edges of the wound [6]. The severeness of the injury also suppresses the local immune system and leaves the wound susceptible to infection [12]. It also has the potential to affect other organs within the body. Early fluid resuscitation and nutritional support are required to fuel the increased metabolic activity of the wound healing system. Without aggres- 10

24 sive treatment, healing is unlikely to occur. Even with treatment, healing will take a prolonged amount and result in proliferative scarring [6] Surgical Excision and Grafting Surgical excision and grafting of the burn wound, as mentioned in Section 1.2.4, is the popular method of treatment for deep dermal and full thickness burn wounds. It is used because no other intervention is available to treat burns of such magnitude [10, 12, 14]. In general, it is performed on burns that are not likely to heal on their own within 3 weeks [15]. Predictions as to whether or not a burn will heal within this time period are based upon initial measurement and continuous monitoring of burn depth, which is discussed in greater detail in Section 1.3 [4]. The first step in this procedure, as the name suggests, is to surgically excise the necrotic tissue (eschar) from the burn wound. It is important to remove this necrotic tissue because its presence increases the likelihood of infection. Excision of this tissue is accomplished using a dermatome. Submillimeter layers of necrotic tissue are shaved off until only viable tissue remains (Fig. 1.5). Removal of the necrotic tissue and presence of viable tissue is signalled by the bleeding that occurs from damage to the patent vessels within that region. After excision of the necrotic tissue, a graft of healthy skin from other regions of the patient (autograft) must be placed on top of the viable tissue that remains (recipient site) [16]. Skin substitutes, allografts, and xenografts may also be grafted in place of healthy skin from the patient. However, these substitutes only provide a temporary solution and are used when healthy skin is not able to harvested from the patient [17]. To perform an autograft, the patient s skin is removed from the donor site using a dermatome. The thickness of the skin removed from this site can vary depending on the requirements at the recipient site and the type of dermatome being used (electric, air-powered, or handheld). Generally, thicknesses range from to inches. Ideally, donor skin is removed from areas where the skin is unburned and the color will match the skin in the region of the recipient site [14]. Before grafting the donor skin to the recipient site, it is usually meshed (Fig. 1.6) to maximize the amount of re-epitheliazation that can occur [14, 16]. However, 11

25 From Papini, R et al BMJ 2004; 329: , courtesy of BMJ Publishing Group. Figure 1.5: Surgical excision, shown here, involves the removal of necrotic tissue from the burn wound. A dermatome is used to remove thin layers of necrotic tissue from the wound until only viable tissue is left. unmeshed skin is grafted to areas such as the face and hands to facilitate the most cosmetically appealing outcome as hypertrophic scarring will be reduced and a mesh pattern non-existent [16]. The timing of this surgical intervention is important [12, 14]. Ideally, the decision to submit a patient to this treatment should be made by the 10th day of injury. The best time to perform a surgical excision and grafting, though, is within the first five days of treatment. Doing so helps to minimize blood loss and prevent local and systemic infection that could result from the prolonged presence of necrotic tissue. Studies have also shown that early surgical excision and grafting can decrease hospitalization stays, hypertrophic scarring, and morbidity [3, 12]. 12

26 From Papini, R et al BMJ 2004; 329: , courtesy of BMJ Publishing Group. Figure 1.6: Shown here, is the healthy skin from a donor site that has been meshed prior to being grafted to the recipient site. 1.3 Burn Depth Estimation As discussed in Section 1.2.4, the severity of a burn is classified by the depth of damage. The decision to treat the burn wound and what intervention will be performed, primarily surgical excision and grafting, are reliant upon the early and accurate determination of burn depth [4]. Ongoing monitoring of burn depth is also necessary so that the treatment plan can be adjusted if the need arises. A method that could objectively determine the extent of burn damage would provide clinicians with a valuable tool in the monitoring and diagnosis of burn wounds. Furthermore, the accurate determination of burn depth has the potential to aid in the more efficient excision of necrotic tissue from the burn wound. Many technologies and techniques have been implemented to better determine burn depth. As with anything, they have their advantages and disadvantages. Some are better at determining one type of burn depth over another type of burn depth. 13

27 Some can only provide a qualitative measure of burn depth. Others can provide a quantitative measure of burn depth, but their resolution is poor or they are depthlimited. The purpose of this work, the photoacoustic determination of burn depth, occupies a niche where its advantages are matched by no other technology or technique. More specifically, photoacoustic methods have the ability to quantitatively determine the depth of superficial and deep dermal burns with high resolution. A brief case for this claim will be made in the Sect after a thorough review of the advantages and disadvantages of competing technologies is presented in the next section (Sect ) Competing Technologies & Techniques Clinical Observation To this day, clinical observation remains the standard for determination of burn depth [4]. Inexperienced observers can easily determine if a burn is of superficial or full thickness just by looking at the burn wound. However, burns of intermediate depth (superficial dermal and deep dermal burns) require a more experienced observer and additional physical diagnostic techniques, such as the pinprick and capillary refill test [3, 4, 11, 18]. As far back as the early 1800 s, differences in pain sensitivity have been used to determine the severity of burn wounds [11]. From this observation evolved the pinprick test of the mid-twentieth century, a technique that has changed little to this day. The pinprick test is carried out by applying pressure to the skin with the point of a hypodermic needle. The more pressure that is required to create the sensation of pain, if at all possible, signals the extent of nerve damage, which can be correlated with burn depth. Burns of greater thickness will exhibit less sensitivity to pain than those of shallower depth because they have a greater number of damaged nerve endings. This technique is quite capable of determining the depth of full thickness burns because they exhibit total analgesia. However, the technique is unable to accurately resolve the depth of superficial dermal and deep dermal burns because of the relative nature of pain. Another common test, the capillary refill test, classifies burn depth according 14

28 to the vascular perfusion of the wound [4]. It is performed by pressing a finger against the wound to temporarily squeeze blood out of the capillaries (blanch). The finger is then released and the time taken for the capillary to refill is observed. A full-thickness burn will exhibit no blanching. A deep dermal burn will blanch, but capillary refill, if it occurs at all, will be slow. Superficial dermal burns, though, will exhibit a much quicker capillary refill time. Although clinical observation can determine if a burn is of superficial, intermediate dermal, or full thickness, it lacks the ability to objectively quantify the depth of intermediate dermal burns. Because of this, experienced surgeons are only 50 % reliable in determining whether an apparent deep dermal burn will heal within 3 weeks [4]. Burn Wound Biopsy Histologic analysis of the burn wound is the most accurate method for determining the depth of a burn wound, for all depths [3, 4]. However, there are many problems associated with its use. A histological analysis is performed by first removing a circular, cross-sectional piece of the burn wound. This technique is referred to as a punch biopsy and is depicted in Figure 1.7. Once removed, the sample is fixed with paraffin, sliced into thin sections, stained with hematoxylin and eosin, then analyzed under a microscope. The depth of the burn is marked by of destruction of the vasculature, the presence of microthrombi, and denatured collagen [4, 19]. A number of disadvantages accompany its use and make it inapplicable to the clinical setting. The procedure is an expensive one that leaves permanent scars and requires a trained pathologist to differentiate denatured collagen from live collagen [3, 4]. Also, because of the destructive nature of the procedure, it cannot be done at many areas of the burn wound to obtain a good sampling of burn depth. The removal of tissue at a measurement point prevents the ongoing monitoring of the burn injury, too [4, 19]. It is important that monitoring is able to be done because the depth of a burn, especially those of intermediate depth, can change dynamically within the first 48 hours. Although histological analysis is an excellent method to diagnose burns of all depths, it has too many drawbacks to make it a clinically feasible technique [4]. 15

29 Punch Burn Area Excised Portion of Skin Epidermis Dermis Figure 1.7: A punch biopsy, shown here, is used to remove a portion of skin from the burn wound. The specimen of skin is then histopathologically analyzed to determine burn depth. Flourescein Flourometry Flourescein fluorometry is a technique used to estimate burn depth by evaluating perfusion within the burn wound [3, 4]. Flourescent dyes are intravenously administered to the patient and allowed to circulate through the patient s bloodstream for a period of time. Once the dye has been adequately distributed throughout the bloodstream, excitation light is directed at the burn wound. Emitted light from the excited fluorophores is collected and quanitified. A wound of shallower depth will have more intact vasculature than a wound deeper into the dermis. Its vasculature will also lie closer to the surface of the burn wound. Because of these two factors, it will produce a more intense fluorescent signal. This measurement, alone, is not enough. It must also be performed on an uninjured portion of the patient s body with similar skin type to make a comparison. A qualitative estimate of burn depth can then be made. The use of sodium flourescein for fluorometric measurements was used to suc- 16

30 cesfully differentiate intermediate depth burns from full thickness burns in a study of 63 burn sites on six patients [20]. A major drawback to the use of sodium fluorescein as the fluorophore, though, is that the blue light required to excite it is strongly absorbed by the skin and, therefore, inhibits the amount of light that can be fluoresced from deeper regions of the burn injury [3]. For this reason, perfusion fluorometry techniques have become centered around the use of indocyanine green dye (ICG) as the fluorophore, because it can be excited by infrared radiation (825 nm), a wavelength that is much less susceptible to attenuation by the skin. A study conducted using IR-excited ICG, correctly classified 64 porcine skin burn wounds as either deep dermal or full thickness burns [21]. The study concluded that the technique was 100 % accurate at determining if a burn wound would heal within 21 days. Green et al used the dual-wavelength (UV and IR) excitation of ICG to also delineate full-thickness burns from deep dermal burns [22]. Although Jerath et al reports great success in determining if a burn wound will heal within 21 days, others do not feel that the technique is clinically reliable [3, 4]. Fluorescein fluorometry accomplishes nothing more than could be accomplished by clinical observation [4]. Like fluorometric techniques, clinical observation is able to distinguish an intermediate dermal burn from a full thickness burn [3, 4]. However, neither have the ability to objectively quantify the depth of an intermediate burn, which is a necessity for operative planning. Laser Doppler Flowmetry Laser Doppler flowmetry is another technique that estimates burn depth by evaluating perfusion within the burn wound [4]. Single wavelength laser light is delivered to the skin where it interacts with moving structures (patent vasculature) and stationary structures. Some of the interacting light is backscattered out of the skin and collected. Light that has interacted with moving structures (dilating and contracting blood vessels) exhibits a Doppler shift in frequency whereas backscattered light from the stationary structures will exhibit no change in frequency. The mixing of these light waves is converted to an electrical signal, which is analyzed to provide and estimation of blood flow in comparison to an area of normal skin. 17

31 Green et al used laser Doppler flowmetry to evaluate the need for surgical excision and grafting on 13 potentially deep, dermal wounds on ten patients [23]. Although their method was unable to generate a quantitative estimate of burn depth, it predicted the need for surgical excision and grafting of the wounds 92% of the time after monitoring changes within vasculature over the first 72 hours. Another study found that Laser Doppler flowmetry could have a 94% accuracy in predicting whether or not a burn wound would heal within 3 weeks as compared to the 70% of a clinical diagnosis made by an experienced physician [24]. Laser Doppler has proven itself a great tool for monitoring the circulatory changes within the burn wound and predicting the need for surgical excision and grafting [4]. The fact that it requires no contact with the wound also confers it a great advantage. However, drawbacks to laser Doppler flowmetry exist and keep it from being fully implemented in the burn unit. Most notably, as with previously mentioned methods, it is unable to provide a quantitative estimate of burn depth, especially for burns of intermediate depth. Other disadvantages include the techniques sensitivity to temperature, patient anxiety, infection, curvature of the wound, and elevation of extremity [3, 4]. Ultrasound Unlike the previously described techniques, which determine the level of burn depth by tissue perfusion, ultrasound provides an estimate of burn depth based on changes in the acoustic impedance of thermally damaged dermis [25]. More specifically, it is the thermal denaturation of collagen that causes a change in the acoustic impedance of the damaged dermis. Because of this change in the acoustic impedance of the damaged dermis, an acoustically mismatched interface will exist between it and the healthy dermal tissue, which ultrasound has the capability to locate. Utilizing a crude, industrial ultrasound device, Moserova et al, showed that it was possible to detect differences in the acoustic impedance of the normal porcine skin and porcine skin scalded 5 and 15 seconds [4, 26]. Cantrell et al used a 0.1mm resolution ultrasound device to detect the burn interface within the dermis of a 18

32 number of experimentally induced burns on pigs [25]. The interface detected by this technique was the thermally denatured collagen-viable collagen boundary, which is not quite an accurate measure of depth, though. Since collagen denatures at 65 degrees Celsius and irreversible damage to cell membranes at 42 degrees Celsius, the extent of burn depth could be underestimated if its measurement was based solely on the denaturation of collagen. Another major disadvantage to these and other early ultrasound techniques is that they require contact with the skin [3]. A noncontact ultrasonographic was recently employed to overcome this limitation [27]. The technique proved to be 96% accurate in determining which burns would heal within 3 weeks and those that would not. Although ultrasonic methods are becoming more patient-friendly and some studies have shown that it has an ability to provide a quantitative measure of burn depth, other studies have not been able to draw the same conclusions [4]. It is felt that more precise instrumentation is needed to increase the resolution along with more clinical testing to determine it effectiveness [3, 4]. Reflectance Spectroscopy Reflectance spectroscopy is a technique used to determine burn thickness based on changes in the optical properties of thermally damaged skin [3]. The burn area is illuminated by a strong optical source and the backscattered light is measured. The optical properties affecting the backscattering are most influenced by the creation of a layer of eschar and variations in the volume and oxygenation of blood within the dermis [3, 28]. Backscattering effects can be evaluated over multiple wavelengths to analyze a number of optically contrasting elements within the burn wound. Afromowitz et al developed a real-time imaging system, the Imaging Burn Depth Indicator (IBDI), to estimate if burns would heal within 21 days on their own [29]. The IBDI collected diffuse reflectance data from the burn wound in the red, green, and near infrared wavelength bands. On a study of over 100 burn wounds, the IBDI proved to be more successful at predicting burn healing than the attending physicians. Eisenbeiss et al improved the technique by evaluating the backscattered light over four wavelengths, which allowed the the burn wound to be classified into 19

33 one of four burn depth categories [30]. In addition to increasing the reliability, he made the technique very easy, enough so that a non-experienced physician could even employ it. Reflectance spectroscopy, like many of the other techniques, lacks the ability to objectively quantify the depth of burns. Instead, burn depth is classified by the ability of the wound to heal within 3 weeks or it is lumped into a depth category. In addition to this, the validity and practicality of the technique remains to be seen [3, 4]. Polarization-Sensitive Optical Coherence Tomography (PS-OCT) Optical Coherence Tomography (OCT) is a non-invasive, interferometric technique that reproduces images based on inhomogeneities within the skin [31]. An optical beam of low coherence light is focused into tissue, and the echo time of delay of light reflected from internal microstructures is measured by interferometry [32]. These light reflections occur at the boundary of refractive index mismatches, which are analogous to the acoustic impedance mismatches associated with ultrasound imaging. Polarization-Sensitive Optical Coherence Tomography (PS-OCT) is a modified form of OCT, whereby the light focused into the burn wound is polarized. In addition to obtaining the location of refractive index mismatches, the polarized nature of the light allows thermally denatured collagen to be differentiated from viable collagen. This is because collagen is a normally a birefringent material, but when thermally denatured it loses its birefringence. Using PS-OCT, de Boer et al showed that it was possible to distinguish thermally denatured collagen from viable collagen in ex vivo burned porcine skin [33]. They produced a cross-sectional image with 15 µm lateral resolution and 10 µm axial resolution, at depths of up to 1 mm. Park et al implemented a high-speed PS-OCT device to image burned rat skin in vivo [32]. They achieved high resolution images at depths of up to 1.5 mm. Polarization-sensitive optical coherence tomography has many advantages that other techniques do not possess. It s greatest advantage is the ability to quantify depth with 10 µm resolution. The fact that it can be done non-invasively also makes 20

34 PS-OCT an attractive option. Disadvantages do exist that limit it s capability. As with ultrasound imaging, PS-OCT has the potential to underestimate the full extent of thermal damage because it only takes into consideration the thermal damage to collagen. Also, since the dermis can be up to 5 mm [7], burn depths between 1.5 mm and 5 mm are unable to be quantified because PS-OCT is depth-limited to approximately 1.5 mm. Finally, like many other techniques, the validity of PS-OCT has not been clinically proven [3] Photoacoustic Estimation of Depth It is quite evident that the competing technologies and techniques of Section have their drawbacks. In short, they are unable to provide a highly-resolved, quantitative estimate of burn depth at all possible dermal depths, particularly in superficial dermal and deep dermal burns. Photoacoustic techniques, though, do have the capability to determine the depth of burn wounds of this type. Photoacoustic detection and imaging, a technique commonly referred to as laserinduced ultrasound, combines many advantages of the other techniques to make it an excellent modality for burn depth determination. Essentially, photoacoustics is the deposition of laser energy into an optically absorbing medium to produce an acoustic signal. Theoretically, photoacoustic methods have the ability to detect burn depth with 20 µm resolution at most skin thicknesses [34, 35]. Unlike OCT, which relies on the information contained in photon propagation, photoacoustic techniques use photons as the energy source but obtain information about the tissue of interest from the propagated acoustic signal. Acoustic signals, unlike optical signals, are much more robust and better at preserving their signal in optically scattering media such as the dermis. A thorough description of photoacoustic theory and pertinent past work is presented in Chapter 2. Recent attempts have been made to apply photoacoustic methods to the estimation of burn depth. Sato et al used a photoacoustic technique to determine the depth of damage by detecting the increased blood volume within the vasculature of the zone of hyperemia in burned rats [36]. They were fairly successful in distinguishing superficial dermal, deep dermal, and full thickness burns from one another based 21

35 on the signal depth. However, the technique was not perfect. Overlapping regions of signal depth existed between each depth classification, signaling the need for more information before an accurate diagnosis of burn depth can be made. Zhang et al induced burns in pigs in vivo, which they later excised and imaged using photoacoustic microscopy [37]. Their technique, like that of Sato et al, determines burn depth by locating the zone of hyperemia. Using photoacoustic microscopy, they were able to provide cross-sectional images of the burn wounds with the zone of hyperemia clearly defined. Unfortunately, though, the technique requires removal of the affected tissue to perform the imaging. As with punch biopsies, the removal of tissue from the wound, let alone the entire wound, is not clinically practical. The purpose of this thesis is to build upon and improve the capabilities of photoacoustic methods in the estimation of burn depth. In Chapter 3 a reflectancemode measurement technique is utilized to make photoacoustic depth estimation a clinically feasible modality. Most importantly, though, Chapter 3 employs a dualwavelength method to better differentiate viable from necrotic tissue within the burn wound. Essentially, it shows how it would be possible to photoacoustically distinguish the viable and necrotic layers of a burn wound based on the type of blood within the layers, either viable or thermally coagulated. Such a technique allows more information about the wound to become available, which helps to formulate a more accurate diagnosis. The optical properties of viable and thermally coagulated blood are important to the implementation of a dual-wavelength photoacoustic technique for burn depth determination or any optical technique related to burn depth determination, for that matter. Since little is known about the optical properties of thermally coagulated blood, a novel photoacoustic method is implemented in Chapter 4 to help elucidate them. In that chapter, the optical absorption coefficient of thermally coagulated blood is derived for the wavelength range from 580 to 700 nm. 22

36 Chapter 2 Photoacoustics 2.1 Introduction Photoacoustics is the blossoming field of science that exists at the intersection of optics and acoustics. More specifically, it involves the production, detection, and analysis of sound from light. In recent years, photoacoustics has become increasingly popular as a tool for biomedical imaging and detection. Unlike many other optical imaging and detection techniques, the signal detected in a photoacoustic system is acoustic and not optical. Instead, light is used as the energy source to produce the acoustic signal. Since acoustic signals traveling within biological tissue are less susceptible to attenuation and scattering than optical signals, the information contained within them is much less likely to be lost. This has made photoacoustics a robust and attractive modality for imaging beyond the range possible that exists for all-optical techniques. This chapter presents a background on the history and recent application of photoacoustics to medical diagnostics, the goal of which, is to highlight its use as an attractive imaging modality for burns. In addition to the review, simple photoacoustic theory will be explained. Lastly, the experimental setup for light delivery and the detection of photoacoustic waves will be discussed. 2.2 History The photoacoustic effect was first discovered by Alexander Graham Bell in He observed that when modulated light was incident on a block of selenium, it caused it to emit a sound with a perfectly perceptible musical tone [38]. The phenomena was soon confirmed by other scientists of the day [39]. However, it s discovery did not stimulate widespread interest in the photoacoustic effect. It was not until microphones, precision instrumentation, and lasers became available, that 23

37 photoacoustics became a burgeoning field. Early improvements in acoustic detection allowed photoacoustics to be used in measuring the IR absorption of gases [39]. Later on, with the advent of laser technology, the related field of photoacoustic spectroscopy was born. Photoacoustic spectroscopy involves the pulsed laser irradiation of condensed matter, which excites a surrounding gas, producing a detectable and characterizable sound [39]. Besides gases, the photoacoustic effect was also studied in liquids. Carome et al used a Q- switched ruby laser to generate a thermoelastic expansion of an optcally absorbing liquid, resulting in a propagated acoustic wave [40]. In his work, he defined the condition of stress confinement, which is necessary to induce the thermoelastic expansion of the liquid. The theory behind the thermoelastic formation and propagation of acoustic waves in liquid was later formalized by Sigrist et al [41]. 2.3 Biomedical Photoacoustics Photoacoustics in Biological Tissue Before discussing the use of photoacoustics in specific biomedical applications, it is important to discuss the study of photoacoustic interactions in biological tissue and tissue-like materials. Since most biological tissue contains approximately 75% water, it can be treated as a liquid medium. Because of this assumption, the previous work by Carome et al and Sigrist et al can be used to describe the photoacoustic effect that occurs within biological tissue. Many investigators have studied the photoacoustic production and detection of acoustic waves in biological tissue. Oraevsky et al showed that it was possible to determine the optical absorption coefficient, µ a, and effective attenuation coefficient, µ eff, of both clear and turbid media, using a photoacoustic method, which they referred to as the time-resolved stress detection (TRSD) technique [42]. Their experiments were performed by irradiating homogeneous biological tissue with a Q-switched Nd:YAG laser and then analyzing the acoustic signal detected with a lithium niobate acoustic transducer. To better simulate the heterogeneous nature of biological tissue, Oraevsky et al expanded the use of the TRSD technique to a turbid, 24

38 two-layer collagenous gel tissue phantom with different optical properties [43]. The TRSD technique was, again, successful at determining both the absorption coefficient and the effective attenuation coefficient of each layer within the tissue phantom. Based on the speed of sound within the tissue phantom, they were also able to determine the thickness of each layer of the phantom. Viator et al performed similar measurements on layered, absorbing soft materials. They were able to extract information about the absorption coefficient of the sample as a function of depth, with a resolution of 20 µm [35]. Unlike the previously mentioned studies that used piezoelectric acoustic transducers located on the side of the sample opposite to the laser irradiation, Paltauf et al used an optical acoustic transducer that made it possible to detect the acoustic wave on the same side as the laser irradiation when determining the absorption coefficient of an absorbing sample [44]. This technique was favorable because it allowed the acoustic wave to be detected where the signal was strongest, at the laser irradiated surface. Scattering is also an important optical property that affects photoacoustic signal generation and detection within turbid biological media. Zhao et al have studied the effects of scattering on acoustic signal propagation [45, 46]. In addition to determining the absorption coefficient and reduced scattering coefficient, µ s, as the studies of the previous paragraph were able to do, they developed a technique to approximate the scattering coefficient, µ s of a weakly absorbing, highly scattering sample. Zhao et al also noted that the effect of scattering was less degradative towards photoacoustic signals than signals from other optical measurement techniques. However, scattering is still strong enough that within most biological tissues the use of photoacoustic methods is limited to approximately 1 cm, albeit with sub-millimeter resolution Medical Imaging The experiments of the previous section, along with other related experiments, set the foundation for photoacoustic tissue imaging. Most importantly, they showed how photoacoustic methods could be used to derive tissue optical properties as a function of depth. Optical properties differ enough between tissue types that they can be considerably contrasting, especially in soft tissue, which x-rays have a tough 25

39 time differentiating. Moreover, since photoacoustics involves non-ionizing radiation it is less harmful than x-rays. However, unlike x-rays, the effect of scattering in biological tissues limits photoacoustics to shallow tissues. Regardless, the optical contrast and sub-millimeter resolution of photoacoustics, have led investigators to focus considerable attention on its use as an imaging modality. To do this, they have sought to develop photoacoustic techniques such that they can spatially resolve varying tissue types within two, if not three dimensions. The following sections describe the efforts made to develop photoacoustics into an imaging modality. Microvasculature Imaging Considerable focus has been given to the photoacoustic imaging of microvasculature, particularly within the skin. Imaging of microvasculature is important to the diagnosis and subsequent treatment of conditions such as Port Wine Stain (PWS) lesions, hemangiomas, telangiectasia, and burns. The skin, being a shallow, soft tissue, is adequately suited to the photoacoustic imaging modality. Pilatou et al performed 3 D in vitro photoacoustic imaging on a vascular cast, created by perfusing the vasculature of a Wistar rat with a red epoxy resin [34]. The vascular cast was removed from the rat and suspended within varying concentrations of Intralipid solutions, creating vasculature phantoms with varying levels of background scattering. At physiological levels of scattering and a laser wavelength of 532 nm, they were able to image to a depth of 7 mm with µm axial resolution and µm lateral resolution. Kolkman et al used photoacoustic methods to image artificial and in vivo blood vessels in two dimensions [47, 48]. They used a narrow aperture sensor, which obfuscated the need for a complex reconstruction algorithm. Two-dimensional images could be produced from individual A-scans (1 D measurements) without having to integrate the signals of multiple A-scans, a reconstruction technique commonly referred to as backprojection imaging. Their sensor also housed the optical fiber that delivered the light to the tissue. This setup allowed for the reflection mode detection of photoacoustic signals, which maximized the signal quality of the in vivo measurements. Using this same sensor, Kolkman et al were also able to find the diameter 26

40 and produce cross-sectional images of artificial blood vessels [49]. Zhang et al used photoacoustic microscopy to produce functional 3 D images of in vivo microvasculature [50, 51]. In particular, they were able to distinguish arterioles from venules based on hemglobin saturation, SO 2, within the blood of the vessels. They achieved this discrimination using a multi-wavelength method. Imaging was done to a depth of 3 mm with an axial resolution of 15 µm and a lateral resolution of 45 µm. Tumor Imaging Cancerous tumors are often marked by their hypervascularization. This hypervascularization, causes an increased presence of blood within tumors, making them relatively strong optical absorbers over much of the visible spectrum, when compared to their surrounding tissue. Because of this optical contrast, investigators have sought to image and detect tumors within soft biological tissues. Of particular interest, is the detection of tumors within the breast, as these can sometimes go unnoticed by traditional mammography. Viator et al created a tumor phantom by embedding 2 mm diameter absorbing spheres within clear and turbid media [52]. Photoacoustic signals were induced within the absorbers, time of flight measurement were obtained, and a convolution algorithm was used to localize them within two dimensions. They were able to locate the absorbers to within 5% of the true location. Esenaliev et al imaged 2 mm diameter gelatin absorbing spheres embedded within turbid gelatin slabs up to depths of 6 cm [53]. In addition, they imaged a 3 x 2 x 0.6 mm piece of liver tissue embedded within chicken breast tissue up to a depth of 8 cm. Khokhlova et al imaged breast tumor phantoms using a 64-element focused array transducer [54]. The phantom was created by immersing a 3 mm piece of bovine tissue within a diluted milk medium. The piece of liver was imaged with 2 mm resolution at depths of up to 3.3 cm. Oraevskyet al used their Laser Optoacoustic Imaging System (LOIS) in a clinical study to detect breast cancer in five patients [55]. Using a 32-element arc-array transducer, they were able to successfully image all tumors in two dimensions with 27

41 400 µm axial resolution, 1,000 µm lateral resolution, and high contrast to the surrounding tissue. However, ultrasound imaging did provide better localization of the tumors and x-rays provided better detail of the tumor structure. Prior to performing a clinical study, Manohar et al used their device, the Twente Photoacoustic Mammoscope (PAM), to image breast tissue phantoms [56]. A 590-element flat array detected tumor-like absorbers in the phantom with resolutions between 3 to 3.5 mm. Manohar et al later performed a clinical pilot study, which they found to be successful at imaging cancerous masses [57]. Rat Brain Imaging The ability to image the rat brain in vivo is of great significance to the scientific community. Rats are often used as models for the study of basic neurological function, which involves the monitoring of structural and functional changes in the brain. Unlike humans, rat brains are adequately suited to optical imaging. This is because their relatively thin skull and small brain size allow light to penetrate the skull and brain quite easily. Ku et al used photoacoustics to produce tomographic images of tumor angiogenesis in rat brains in vivo [58]. Photoacoustic signals were achieved using a Q-switched Nd:YAG laser, operating at a wavelength of 532 nm, which was able to penetrate the skull, yet still be absorbed by blood. Tumor angiogenesis was characterized by an increased level of absorption due to the increased blood volume resulting from hypervascularization. Wang et al added a nanoshell contrast agent when they performed photoacoustic tomography of a rat brain in vivo. Adding the contrast agent increased the absorption in the brain vessels by 63%, which allowed them to produce better images of the vasculature than without the contrast agent. Wang et al showed that photoacoustic methods could be used to produce functional imaging of the rat brain in vivo. They accurately mapped the rat brain structures, with and without lesions, and the functional cerebral hemodynamic changes associated with areas of the brain that were responsive to whisker stimulation. Furthermore, they were able to image the functional level of SO 2 in hyperoxia- and 28

42 hypoxia-induced cerebral hemodynamic changes Other Applications Besides the imaging methods detailed in the previous section, others have devised novel uses of photoacoustics for other biomedical applications. Weight et al designed a system for detecting metastatic circulating melanoma cells in vitro within a flow cell [59]. Laser irradiation of the optically absorbing melanoma cells induces a detectable photoacoustic signal that denotes the presence of metastatic melanoma cells. Laufer et al has developed a method for spatially resolving the SO 2 levels within blood vessels using photoacoustic spectroscopy [60]. The authors envision the incorporation of this technique with photoacoustic imaging to produce 3 D maps of SO 2 levels in microvasculature at centimeter depths with µm resolution. Zhao et al and Kinnunen et al have studied the applicability of photoacoustics to the noninvasive in vivo monitoring of blood glucose levels [61, 62]. In vitro, Kunninen et al found that the photoacoustic amplitude of laser-irradiated pig blood increases with increasing levels of glucose. However, in vivo measurements on human skin by Zhao et al do not show such a consistent relationship and, therefore, more research must be done Motivation for Burn Imaging Based on the previously cited work, it has been shown that photoacoustic methods have the potential to image soft tissue structures with high contrast and resolution at centimeter depths. This is ideally suited for burn imaging since the skin is composed of soft tissue and can reach depths close to 1 cm. Other modalities are either unable to image to this depth or lack the contrast and resolution required to produce adequate images for the quantitative determination of burn depth. As mentioned in Section 1.3.2, several authors have investigated the use of photoacoustics as a modality for burn depth imaging. Their methods are strictly reliant upon the detection and location of viable blood within the zone of hyperemia. However, tissue above the zone of hyperemia can still be viable and should be recognized as such. An estimate of burn depth based solely upon the location of hyperemia could cause the unnecessary excision of the precious viable tissue that might pre- 29

43 side above it. We intend to show how a photoacoustic method that could detect and locate the thermally coagulated blood within the zone of necrosis, in addition to detecting the viable blood within the zone of hyperemia, would better delineate the boundary between the necrotic and viable tissue layers, thereby improving the quantification of burn depth and sparing the unnecessary removal of viable tissue. 2.4 Photoacoustic Theory Five different mechanisms exist that can be responsible for the production of optically induced acoustic waves [63]. These mechanisms are dielectric breakdown, vaporization or material ablation, thermoelastic process, electrostriction, and radiation pressure. Of concern with this thesis, is the thermoelastic production of sound in a planar, optically absorbing medium. In essence, thermoelastic sound is produced by the transient heating of a restricted volume by light energy. This condition is referred to as stress confinement and it requires that the optical irradiance, often done with a laser, occurs over a very short amount of time [40]. More specifically, the laser pulse duration must be less than the amount of time needed for acoustic energy to propagate out of the area of absorption. In a planar medium, this relationship between the laser pulse duration, τ p, the speed of sound in the medium, c s, and the optical absorption depth, δ, is outlined in equation (2.1) and represented graphically in Figure 2.1. τ p < δ (2.1) c s The optical absorption depth is dependent upon the material being irradiated. It is the depth at which the optical energy delivered into the medium is attenuated by 1/e. For a purely absorbing medium, δ = 1/µ a, while in an absorbing and scattering medium, δ = 1/µ eff. Although the tissues under investigation within this thesis consist of scattering elements, the photoacoustic theory presented here will only be concerned with the generation of acoustic waves in purely absorbing media and therefore the optical absorption coefficient, µ a, will be used. The effects of scattering and the use of µ eff will be discussed further in the diffusion theory realm experiments 30

44 Laser Pulse, τ p < δ c s δ = 1 µ a Transmitted Pressure Wave Absorbing Medium Figure 2.1: A laser pulse of duration, τ p < δ/c s, is incident upon a planar absorber. Stress confinement is achieved, a thermoelastic expansion occurs, and a pressure wave is formed. of Chapter 4. If the condition for stress confinement is met and if the fluid is stationary with isotropic acoustic properties, the wave equation for acoustic pressure, p, generated within the laser-irradiated volume is 2 p 1 2 p c 2 s t = β H (2.2) 2 C p t where c s is the speed of sound in the medium, β is the volume thermal expansivity, C p is the constant pressure specific heat capacity, and H is the heat per unit volume and per unit time deposited in the fluid [64]. In general, p and H will depend on the position r = (x, y, z) and time t. When the condition of stress confinement is met by essentially instantaneous laser irradiation (denoted by the impulse function, δ(t), in Equation 2.3), the heating function can be modeled as H(r, t) = H(r)δ(t) (2.3) 31

45 where H(r) is the volumetric heat density. Since the heat energy is created by absorbed optical energy it can be related to the laser fluence, Φ(r), and optical absorption coefficient, µ a (r), of the volume by the following equation H(r) = µ a (r)φ(r) (2.4) In accordance with Beer s Law, the laser fluence decreases exponentially as a function of µ a and depth, z, in a purely absorbing medium. Given the incident radiant exposure, Φ 0, and µ a the energy density can described as a function of depth with following equation H(z) = µ a (z)φ 0 e µaz (2.5) The initial pressure distribution, p 0 (z), immediately following the laser pulse is directly proportional to the energy density function as follows p 0 (z) = Γµ a (z)φ 0 e µaz (2.6) where Γ denotes the fraction of absorbed light energy that is converted to mechanical (acoustic) energy. The value of Γ can range from 0.11 to in biological media [42]. Since the pressure wave radiates in all directions, the acoustic energy can be split evenly between the initial reflected pressure distribution, p 0 (z) trans, and the initial transmitted pressure distribution, p 0 (z) refl (Equation 2.7). The propagation of the initial transmitted pressure distribution is away from the source of laser irradiation. p 0 (z) trans = p 0 (z) refl = 1 2 p 0(z) = 1 2 Γµ a(z)φ 0 e µaz (2.7) The speed at which these acoustic waves propagate is determined by the speed of sound within the medium. Since most biological tissue contains a high percentage of water, the speed of sound within biological tissue is often given the same value as the speed of sound in water, 1.5 mm/µs. Temporal measurement of these propagated acoustic waves is achieved with acoustic transducers, which convert the pressure signal into a voltage signal that is quantifiable by an oscilloscope. Acoustic transducer theory will be discussed in 32

46 greater detail in Section Often times the acoustic transducer is placed opposite the side of the absorber that is being subjected to laser irradiation. This form of measurement is considered a transmission mode measurement. If the measurement takes place on the same side as the laser irradiation, it is considered to be a reflection mode measurement. It is important to note, though, that both the initial reflected and transmitted pressure distributions contribute to the measured signal in either the transmission mode and reflection mode. This combination is a result of acoustic mismatches that often occur at the boundaries of absorbing mediums, which is evident in a transmission mode, acoustic measurement of a laser irradiated absorber that is acoustically mismatched at the surface (Figure 2.2). The absorber and the air above it have different acoustic impedances, which creates a free boundary. At this free boundary, acoustic energy is reflected back into the medium with a negative amplitude. Air Initially Reflected Pressure Wave Laser Pulse, τ p InitiallyTransmitted Pressure Wave Transmitted Pressure Wave Absorbing Medium Acoustic Transducer Figure 2.2: An initially transmitted pressure distribution, p 0 (z) trans, heads downward towards the acoustic transducer. The initially reflected pressure distribution, p 0 (z) refl, heads away from the transducer before being reflected back towards the transducer at the free boundary that exists between air and the absorbing medium. The time-resolved pressure signal of such a measurement is shown in Figure

47 The letter T represents the pressure wave resulting from the initially transmitted pressure distribution, p 0 (z) trans. Since it takes a direct a direct path to the acoustic transducer, it appears closer in time. The initially reflected pressure distribution, p 0 (z) refl, is denoted by the letter R. It has a negative amplitude due to the free boundary and appears later in time because its direction is away from the acoustic transducer prior to being reflected back towards it. The maximum of the T wave and the minimum of the R wave represent the surface of the absorbing medium, the point at which the light energy absorbed and the initial pressure were greatest. Greater depths are characterized by the portions of the pressure waveform to the left of the T maximum and to the right of the R minimum. T Pressure (bar) 0 R Time (µs) Figure 2.3: Time-resolved, transmission mode pressure signal from planar absorber. The T portion of the waveform represents the initially transmitted pressure distribution, p 0 (z) trans. The R portion of the waveform represents the initially reflected pressure distribution, p 0 (z) refl. Photoacoustic signal analysis is primarily concerned with determining the distance from the absorber to the transducer, in addition to determining the optical 34

48 properties of the absorber. Important optical properties to find are the absorption coefficient, µ a, scattering coefficient, µ s, anisotropy factor, g, and the refractive index, n. Certain tissue types, and the pathologies that manifest within them, have characteristic optical properties. Photoacoustic tissue identification is therefore possible if the optical properties of an absorbing tissue can be derived and matched to the optical properties of a known tissue type or pathology. With planar absorbing materials, this is relatively easy. Determining the distance from the transducer to the absorber, particularly the absorber surface, is done by converting the pressure vs time signal into a pressure vs distance signal (Figure 2.4). This conversion is made using the speed of sound approximation in liquids and biological tissues of 1.5mm/µs. The distance located at the maximum of the T wave is taken be the distance from the transducer to the absorber surface. T Surface Distance 1.81 cm Pressure (bar) 0 R Distance (cm) Figure 2.4: Distance-resolved, transmission mode pressure signal from planar absorber. The surface of the absorber is located approximately 1.81 cm above the transducer. 35

49 Determining the optical properties of a planar absorbing medium reduces to determining the absorption coefficient because scattering and anisotropy are not applicable and the refractive index can be estimated or found using other optical methods. The absorption coefficient can be found in two ways, the easiest of which, is an exponential curve fit. Both the T and the R waves can be used to perform this curve fit, since their shapes are governed by an exponential decay (Equation 2.7). However, the T wave is usually more robust and, therefore, used to perform the curve fit, which is done by reversing the waveform at the T maximum on a pressure vs. distance graph (Figure 2.4). This reversed T waveform (Figure 2.5) then represents the relationship given in Equation 2.7. The variables preceding the exponential term reduce to a constant and the waveform can be exponentially curve-fitted to find the value of µ a. The absorption coefficient of a planar medium can also be found from the amplitude of the initial pressure distribution. However, this method requires exact knowledge of the Grunëisen coefficient, laser fluence, and volt/bar ratio of the acoustic transducer, which is hard to obtain in practice. Amplitude measurements are still useful for making relative measurements of absorption strength over different wavelengths because the three previously mentioned items can be held constant. Photoacoustic theory is straightforward, provided the material is a planar absorber. It is quite manageable to find the absorption coefficient of an absorber, along with it s distance from the transducer. However, in real life, the absorbing mediums that we detect photoacoustic signals from have a number of optical properties that define them. Furthermore, these absorbers often have a morphology that is non-planar, which complicates the equation governing pressure distribution and propagation. The non-planar morphology, coupled with the fact that absorbers of interest, such as blood vessels and tumors, are embedded within tissue that has its own set of optical properties, makes the determination of absorber location difficult. Because of the complexity surrounding the theory behind the optical property and location determination of this type of absorber scenario, the photoacoustic theory presented in section will not go beyond that presented for the case of a planar absorber. Minor additions to this theory will be presented in subsequent chapters when 36

50 p 0 (z) trans = 1 2 Γµ ae µ a *z Pressure (bar) Distance, z (cm) Figure 2.5: Initial pressure distribution within the absorbing medium as a function of depth, z, into the medium. deviations from the planar absorbing case are encountered. 2.5 Photoacoustic System Laser System As discussed in the previous section, photoacoustics involves the production of sound from the absorption of light energy. The light used to produce the photoacoustic effect in the experiments of subsequent chapters and in most other photoacoustic applications, is that of a laser light. The coherence, wavelength specificity, and submicrosecond pulsing capability of lasers make their use preferable in photoacoustics. The laser system used for our experiments was a frequency-tripled, Q-switched, Neodymium-doped Yttrium Aluminum Garnet (Nd:YAG) laser (Quantel, LES ULIS cedex, France) pumping an optical parametric oscillator (Vibrant 355 II, Opotek, Carlsbad, CA) (Fig. 2.6). The Nd:YAG laser emits sub-microsecond pulsed light 37

51 at a wavelength of 1064 nm. Pulsing of the laser light is achieved using an optical switch, which is referred to as the Q-switch. Before reaching the optical parametric oscillator (OPO), the laser light passes through both a second harmonic generator and third harmonic generator, which act to reduce the frequency from 1064 nm to 355 nm. Once at the OPO, the 355 nm laser light can be converted to wavelengthspecific light over the range from 410 to 2400 nm. The energy associated with pulsed laser light emitted from the OPO decreases with increasing wavelength. An energy of 25 mj/pulse can be expected at a wavelength of 410 nm, while only 1 mj/pulse can be expected at a wavelength of 2400 nm. Delivery of the pulsed laser light from the OPO to a sample for photoacoustic measurements can be achieved in two ways, either by coupling the light into an optical fiber or by redirecting the beam to the sample with right angle prisms. Both methods have been used in the experiments of the subsequent chapters and will be discussed in greater detail when those experiments are described. Mirror Second Harmonic Generator Third Harmonic Generator Dichroic Mirror 355nm 532nm Polarizer Lens 1064nm Mirror Q-Switched Nd:Yag Laser Optical Parametric Oscillator Figure 2.6: The laser system is shown here. An Nd:YAG laser produces 1064 nm pulsed laser light, which is reduced to a frequency of 355 nm, via a second and third harmonic generator, before reaching the OPO. The OPO can convert the 355 nm laser light to a user-specified wavelength over a range from 410 to 2400 nm. (Photograph courtesy of R.M. Weight) 38

52 2.5.2 Acoustic Detection Transduction of the optically-induced pressure waves into an analyzable electrical signal was achieved using a piezoelectric device. The piezoelectric device was constructed from polyvinylidene fluoride (PVDF) copolymer film (Ktech Corp., Albuquerque, NM). Piezoelectricity is conferred to PVDF film in the manufacturing process, whereby it is stretched and a strong, electrically polarizing field is applied across it [65]. The result of this treatment is a film that is composed of randomly ordered polymer molecules, which causes it to exhibit no net polarization (nonpolarized). However when subjected to mechanical stress, such as that from a pressure wave, polymer molecules orient themselves in a manner that creates a net polarization. The PVDF film becomes polarized across it thickness and a voltage can be measured across. The strength of polarization, or the magnitude of the voltage, is proportional to the strength of the pressure wave that is incident upon the PVDF film. Electrical contacts, connected to an oscilloscope, are situated on both sides of the PVDF film so that changes in the voltage, resulting from mechanical stress, can be recorded and analyzed. Two different PVDF film, piezoelectric devices were created for the experiments of the subsequent chapters. The specifics regarding the construction of these devices is detailed in those chapters. 39

53 Non-Polarized State PVDF Film Pressure Wave Polarized State Figure 2.7: The activity of PVDF film is shown here. When PVDF film is subjected to a mechanical stress, such as that from a pressure wave, a net polarization occurs that is proportional to the stress. The net polarization creates a voltage, which can be measured using an oscilloscope. 40

54 Chapter 3 Photoacoustic imaging of viable and thermally coagulated blood Discriminating viable from thermally coagulated blood in a burn wound can be used to profile burn depth, thus aiding the removal of necrotic tissue. In this study, we used a two wavelength photoacoustic imaging method to discriminate coagulated and non-coagulated blood in a dermal burn phantom. Differences in the optical absorption spectra of coagulated and non-coagulated blood produce different values of the ratio of peak photoacoustic amplitude at 543 and 633 nm. The absorption values obtained from spectroscopic measurements indicate that the ratio of photoacoustic pressure for 543 and 633 nm for non-coagulated blood was 15.7:1 and 1.6:1 for coagulated blood. Using planar blood layers, we found the photoacoustic ratios to be 13.5:1 and 1.6:1, respectively. Using the differences in the ratios of coagulated and non-coagulated blood, we propose a scheme using statistical classification analysis to identify the different blood samples. Based upon these distinctly different ratios, we identified the planar blood samples with an error rate of 0%. Using a burn phantom with cylindrical vessels containing coagulated and non-coagulated blood, we achieved an error rate of 11.4%. These results have shown that photoacoustic imaging could prove to be a valuable tool in the diagnosis of burns. 3.1 Introduction Early and accurate determination of burn depth is crucial in deciding what steps are taken to treat a burn wound [4]. Currently, clinical observation, an inexact This chapter, originally titled Photoacoustic discrimination of viable and thermally coagulated blood using a two-wavelength method for burn injury monitoring, was published in Physics in Medicine and Biology, April 2007, Volume 52. The website for this journal can be found at The abstract for this article can be found at 41

55 science, is the standard method for determining burn depth. Although it is an accurate predictor of full-thickness burns, it is only 50% accurate in the diagnosis of partial thickness burns [66]. A method that could objectively determine the extent of burn damage would provide clinicians with a valuable tool in the monitoring and diagnosis of burn wounds. Furthermore, if a depth profile of the wound were available such that necrotic tissue was differentiated from reversibly damaged or viable tissue, early and accurate excision of the burn wound would be possible, an important factor in the treatment of dermal burns [67]. Many attempts have been made to improve the accuracy of burn depth determination. One such method, laser Doppler imaging (LDI), is a non-invasive technique that determines burn depth by monitoring vascular perfusion in the burn wound. However, this technique gives no measure of absolute depth [68, 69]. Because the technique bases burn depth on the amount of vascular perfusion, factors such as patient heart rate and room temperature must be taken into consideration when performing the measurement. Indocyanine green (ICG) fluorescence and spectral methods have also been attempted but they, too, provide no quantitative measure of burn depth [21, 29]. Polarization sensitive optical coherence tomography (PS-OCT) is a powerful technique that has been used to determine burn depth based on changes in the optical birefringence of skin due to thermally denatured collagen [68, 70]. Although the method can provide absolute burn depths of up to 1.5 mm with 10 µm axial resolution, it is not able to image all burns, as the thickness of skin can be up to 5 mm deep [7]. Photoacoustic techniques, however, may overcome these limitations inherent in the previously discussed methods and have the potential to provide high resolution images at greater depths. Unlike OCT, which relies on information carried by scattered photons, photoacoustic techniques use photons as the energy source but obtain information about the tissue of interest from the propagated acoustic signal. Acoustic signals, unlike optical signals, are much more robust and are better at preserving their signal in optically scattering media such as the dermis. Generation and propagation of photoacoustic waves were studied in gas and liquid media [41, 63]. More recently, the study of photoacoustic generation and propagation 42

56 has been expanded to absorbing media that have been created to mimic human tissue [44]. Though photoacoustic waves can be generated by different mechanisms, photacoustic waves due to thermoelastic expansion are the result of the absorption of short laser pulses under the condition of stress confinement [40]. Stress confinement is achieved when the laser pulse duration is less than the amount of time needed for acoustic energy to propagate out of the area of absorption. This relationship between the laser pulse duration, τ p, and the optical absorption depth, δ, is given by τ p < δ/c s. For dermal tissue, the the speed of sound, c s, is approximately equal to the speed of sound in water, 1500 meter/second. The resulting pressure wave can be analysed to determine the absorbed energy distribution within the sample and obtain the optical and morphological properties of the sample [42]. The burn wound can be modeled as a layered medium, whereby a highly absorbing perfused viable layer lies below a relatively low absorbing, highly scattering, necrotic dermal layer containing an amount of thermally coagulated blood. A photoacoustic study conducted to determine the ability of a probe to detect photodynamically treated esophageal cancer cells found that it was able to detect an absorber lying under a necrotic dermal layer at depths up to 5 mm [71]. Depth resolutions of 20 µm have been obtained in the imaging of layered absorbing media using photoacoustic techniques [35]. PA techniques have been used to image vasculature, in vitro, at depths of up to 7 mm with a depth resolution of µm [34, 58]. Although, photoacoustic techniques do not exhibit the same resolution as PS-OCT in the mm range, they do have the ability to penetrate to greater depths. The goal of this study, however, is not to achieve depth resolved images of burns. Rather, we intend to show that a two wavelength photoacoustic method can be used to discern viable from thermally coagulated blood in a burn phantom. A multiwavelength photoacoustic method has previously been used to determine burn depth in rats [72]. However, it was not based on the presence or absence of viable and coagulated blood, but based on the relative amounts and location of deoxy-hemoglobin and oxy-hemoglobin within the blood of the dermal tissue. Visual inspection and preliminary spectroscopic studies in our laboratory have shown that the optical absorption 43

57 spectra of viable and thermally coagulated blood differ considerably and have the potential to be discriminated more accurately and more easily than analysing the relative levels of deoxy-hemoglobin and oxy-hemoglobin within the blood of a burn wound. The brown colour of thermally coagulated blood indicates that its optical absorption spectrum is relatively featureless, while the optical absorption spectrum of the red, viable blood should exhibit a sharp drop off as the longer wavelengths of the visible spectrum are approached. Due to the linear relationship between the photoacoustic signal produced and the optical absorption characterstics of the medium, the photoacoustic spectrum will exhibit nearly identical characteristics to the optical absorption spectrum. Obtaining photoacoustic measurements at two different wavelengths, one from an area of the spectrum before the absorption of viable blood exhibits a sharp decrease and the other after the decrease, we show that a relationship between the two measurements can lead to the identification of type of blood present within a burn wound. In this paper we develop a classification rule for differentiating between viable and thermally coagulated blood both spectroscopically and photoacoustically. Blood from 18 swine was collected and analysed in this study. The blood was first analysed spectrally to obtain the ratios that define viable and thermally coagulated blood. Photoacoustic ratios were then obtained for the 18 samples of blood and compared with the spectroscopic ratios. We developed a scheme to classify the measurements as coagulated or viable, hereafter referred to as non-coagulated. We then performed a raster scan of a burn phantom and produced an image of it showing the distinction between the coagulated and non-coagulated blood. 3.2 Statistical Method for Classifying Coagulated and Non-coagulated Blood The classification rule we propose begins by assuming there exists two distinct populations of blood as defined by the ratios of their photoacoustic response at 543 and 643 nm. Then classification of observations into one of these populations proceeds by using standard multivariate classification techniques [73] that only rely 44

58 on the first two sample moments (i.e. mean and variance) of the photoacoustic ratios on coagulated blood. In order to develop a classification rule we create a training set of photoacoustic measurements on known samples. We begin by assuming that the probability distribution functions (pdf s) associated with the two populations, coagulated and non-coagulated blood, are Gaussian with equal variances. The assumption of equal variance is used to simplify the classification rule. However, we acknowledge that this assumption may not be satisfied. Nonetheless, the rule we propose is shown empirically to be robust to departures from this assumption. Following the notation and exposition of Johnson and Wichern (1998) [73], we begin by stipulating the photoacoustic ratios come from two distinct populations, Π 1 the coagulated blood, and Π 2 the non-coagulated blood. Let f 1 (x) and f 2 (x) denote the pdf s associated with the random variable x of the two populations, where x denotes the ratio of photoacoustic response. Further, let p(2 1) be the conditional probability of classifying a ratio as belonging to Π 2 when it belongs to Π 1 and p(1 2) be defined similarly. Finally, let c(2 1) denote the cost of classifying a ratio as belonging to Π 2 when it comes from Π 1 and c(1 2) be defined analogously. In the case of burn injury, it may be reasonable to consider coagulated tissue misclassified as healthy (non-coagulated) to be more costly than the opposite case, since a skin graft on a substrate of necrotic tissue will result in a failed graft. The consequences of this case would include an unnecessary surgery and several days of useless treatment. On the other hand, to classify healthy tissue as coagulated results in a deeper excision. Though this case is not optimal as it may damage epithelial regions needed for healing, it will not significantly impede treatment. In fact, the current state of burn care for grafting includes excision of the necrotic tissue and stopping just after the healthy tissue has been reached. The cost of misclassification can be defined by a cost matrix, Π 1 Π 2 Π 1 0 c(2 1) Π 2 c(1 2) 0 where the rows denote the true population and the columns denote how the sample was classified. Clearly, if an element of Π i is classified as Π i, where i = 1, 2, there is 45

59 no cost, as the classification was correct. Now for any classification rule the average or expected cost of misclassification (ECM) is given by ECM = c(2 1)p(2 1)p 1 + c(1 2)p(1 2)p 2, where p i (i = 1, 2) is the prior probability of Π i and p 1 + p 2 = 1. A reasonable classification rule would seek to minimize ECM. Let Ω denote the sample space of all possible observations, R 1 the set of all x observations from Π 1 and R 2 = Ω R 1. The rule that minimizes ECM is then given by ( ) ( ) f 1 (x) c(1 2) R 1 : f 2 (x) p2, c(2 1) p 1 ( ) ( ) f 1 (x) c(1 2) R 2 : f 2 (x) < p2. c(2 1) In the case that we that we assume equal cost of misclassification and p 1 = p 2 = 1/2, this rule simplifies to R 1 : R 2 : f 1 (x) f 2 (x) 1, f 1 (x) f 2 (x) < 1. As in the case of burn injury, it may be more appropriate to specify a rule where it is more costly to misclassify in one direction than the other. For example, c(2 1) = kc(1 2) where k is a constant equal to the ratio of the costs. In the case we consider, k may equal 10 if it is postulated that grafting over necrotic tissue is 10 times more costly than excising healthy tissue. In order to implement these rules we evaluate the density function ratio at a new value x 0. Furthermore, since we don t know the population first and second moments of the pdf s we substitute the sample quantities to establish a classification rule. Specifically, σ1 2 σ2 2 yields the rule ( ) s2 ln + 1 [ (x0 x 2 ) 2 (x 0 x 1 ) 2 ] s 1 2 s 2 2 s 2 1 ( ) s2 ln + 1 [ (x0 x 2 ) 2 (x 0 x 1 ) 2 ] s 1 2 s 2 2 s p 1 [ c(1 2) ln c(2 1) [ c(1 2) < ln c(2 1) ] p 2 p 1 ] p 2 p 1 x 0 Π 1, (3.1) x 0 Π 2, (3.2)

60 for allocating the photoacoustic ratio to the coagulated or non-coagulated populations. Assuming σ 2 1 = σ 2 2, (3.1) and (3.2) can be simplified to (x 0 x 2 ) 2 (x 0 x 1 ) 2 [ c(1 2) ln 2s 2 c(2 1) (x 0 x 2 ) 2 (x 0 x 1 ) 2 2s 2 < ln where s 2 is the pooled variance defined by [ c(1 2) c(2 1) (n 1 1)s 1 + (n 2 1)s 2. n 1 + n 2 2 ] p 2 p 1 ] p 2 p 1 x 0 Π 1, (3.3) x 0 Π 2, (3.4) Here, n i is the total number of measurements for i = 1, 2. Assuming equal cost of misclassification, equal variance and equal prior probabilities produces a rule where if (x 0 x 2 ) 2 (x 0 x 1 ) 2 2s 2 0 then the measurement was performed on blood in Π 1, the coagulated blood. The point x 0 is the ratio of the 543 and 633 nm photoacoustic amplitudes, and should not be confused with the ratio of the pdf s evaluated at x Measure of Future Performance The performance of our classification scheme can in principal be assessed by calculating the actual error rate (AER) AER = f 1 (x)dx + R 2 f 2 (x)dx, R 1 where R 1 and R 2 are determined from sample sizes n 1 and n 2 respectively [73]. However, in practice this quantity cannot be explicitly calculated, since we do not know the actual pdf s, f 1 and f 2. Thus, instead, we determine the APER defined as the fraction of observations in the training sample that are misclassified by the sample. The APER can be easily described by the confusion matrix, Π 1 Π 2 Π 1 n 1c n 1m = n 1 n 1c Π 2 n 2m = n 2 n 2c n 2c 47

61 where the columns denote predicted membership and the rows are actual membership. Thus, APER = n 1 m + n 2m n 1 + n 2. Evaluating our classifying scheme on the same sample used to develop our rule would result in a bias. That is, it would reduce the APER and make our rule appear to perform better than it does in actuality. To alleviate this bias we evaluate our rule using cross validation leave-one-out. Specifically, we leave out one observation from the training set, develop our classification rule and then evaluate the performance on the left out observation. We repeat this process for all the items in the sample and report this figure as our estimated APER. 3.3 Materials and Methods Spectral Analysis of Coagulated and Non-coagulated Blood Since the peak pressure within a photocaoustic wave is linearly related to the optical absorption coefficient, µ a we performed spectral analysis to determine the expected ratio at 543 and 633 nm, S 543/633, based upon µ a at those wavelengths, which could be compared with the experimentally determined ratio, P 543/633, obtained from the photoacoustic point measurements generated by 543 and 633 nm laser irradiation. The blood was analysed over the visible spectrum using a spectrophotometer (HR2000, Ocean Optics, Dunedin, FL). Analysis of the optical absorption spectra also required that the red blood cells (RBC s) be lysed to prevent sedimentation. The non-coagulated blood was analysed by placing a small portion of blood between two microscope slides separated by 125 µm. Integrating spheres were not used for this experiment because the effect of scattering was negligible for such a thin sample. The analysis on the coagulated blood sample was done by placing non-coagulated blood between two slides, separated by a distance of 25 µm, and sealing the edges with epoxy resin. The sealed sample was then baked in a 100 degree Celsius oven (10GC, Quincy Lab, Chicago, IL) for 5 minutes to produce thermally coagulated 48

62 blood between the slides. The absorbance was then measured using a spectrophotometer. A smaller sample thickness, 25 µm, was necessary to analyse the thermally coagulated samples because of the higher absorption coefficient. Although methods for heating the blood differed, both methods coagulated the blood to the point where they exhibited nearly identical optical properties upon visual inspection. Each of the spectra was averaged over 100 measurements with a 14 msec integration time Experimental Setup for Photoacoustic Measurements The experimental setup consisted of a frequency-tripled, Q-switched, Nd:YAG laser pumping an optical parametric oscillator (Vibrant 355 II, Opotek, Carlsbad, CA), operating at wavelengths of 543 and 633 nm. The repetition rate was 10 Hz with a pulse duration of 5 nsec. Pulse energy and laser fluence varied for the photoacoustic measurements being taken and are provided in the relevant sections. The photoacoustic signals were sensed by an acoustic transducer and sent to an oscilloscope. The oscilloscpe (TDS 2024, Tektronix, Wilsonville, OR) had a bandwidth of 200 MHz and an input impedance of 1 MOhm. The oscilloscope was triggered to record the waveforms by a photodiode (DET-210, Thorlabs, Inc., Newton, NJ) with a risetime of 1 nsec. Waveforms were averaged over 128 pulses. Two separate photoacoustic methods were used in this study, one to determine the ideal photoacoustic ratios of coagulated and non-coagulate blood and one to perform a pre-clinical imaging protocol on a blood vessel phantom. Planar Blood Samples In order to determine the ideal photoacoustic amplitude ratios for 543 and 633 nm light, we irradiated planar samples of coagulated and non-coagulated blood. These samples were planar in order to reduce the amount of acoustic diffraction sensed at the transducer. These experiments were performed in transmission mode, as this type of measurement is simpler to perform and more robust than in reflection mode. The laser spot size was approximately 2 mm in diameter. The pulse energy was in the range of mj at 543 nm and mj at 633 nm. Polyacrylamide cylinders were formed to contain a small volume of blood (Fig- 49

63 ure 4.2.5). Polyacrylamide was used to maintain an acoustically matched path from the irradiated blood to the transducer. The cylinders had a height of 20 mm and a diameter of 42 mm. For non-coagulated blood, a well 17 mm in diameter and 10 mm deep was formed in the center of the cylinder. The blood sample was approximately 7 mm deep within the well. For the coagulated sample a 2 mm thick planar layer was formed by enclosing non-coagulated blood in a glass cuvette that was submerged in a 65 degree Celsius water bath. This coagulated layer was placed on the polyacrylamide cylinder. The laser light, as a free beam, was guided to the blood samples Non-Coagulated Blood Coagulated Blood Laser Beam Laser Beam Polyacrylamide Gel Well Polyacrylamide Gel Well PVDF Film Transducer Blood Sample PVDF Film Transducer Blood Sample Oscilloscope Oscilloscope Figure 3.1: The transmission set up was performed by placing blood samples in polyacrylamide cylinders placed on an acoustic transducer. via right angle prisms. The acoustic transducer was housed in an aluminum cylinder and was made from 25 µm thick polyvinylidene fluoride (PVDF) with an active area of 2 2 mm. We irradiated 36 samples, 18 of which were coagulated and 18 were non-coagulated. The measured photoacoustic ratios, P 543/633, for each sample and type of blood, were calculated using P 543/633 = ( ) ( ) P 543 p p P 633 p p ( P 633 p p + ɛ ) 2 (3.5) Simple division of the peak-to-peak photoacoustic amplitude at 543 nm, P 543 p p, by the peak-to-peak amplitude at 633 nm, P 633 p p, results in a singularity when the value of P 633 p p is equal to zero. This equation, referred to as a divide by zero smoothly 50

64 algorithm [74], was used to prevent those singularities from occuring. The value of ɛ was set to To account for a difference in energy at the two wavelengths, the P 543/633 values were normalized by multiplying by an appropriate factor depending on the ratios of energies at the two wavelengths. Blood Vessel Phantom Imaging The blood vessel phantom experiments were performed in reflectance mode, as this type of set up is required for any eventual clinical implementation. The set up for these experiments is shown in Figure 3.2. The imaging was performed on two X-Y Scanning Stage Water Transducer Optical Fiber Nd:YAG Laser 1%Intralipid/Direct Red Polyacrylamide Gel Computer Coagulated Vessel Viable Vessel Oscilloscope Figure 3.2: The experimental set-up used in the photoacoustic analysis of coagulated and non-coagulated blood. A burn phantom containing coagulated and noncoagulated blood was irradiated with 543 and 633 nm light from a optical parametric oscillator. A piezoelectric probe detected the photoacoustic pressure wave as it propagated away from the upper surface of the phantom. cylindrical vessels, one with coagulated and one with non-coagulated blood. The phantom construction is described more fully in the following section. For the imaging scan, the laser beam was coupled into a 1,500 µm optical fiber that delivered light to the phantom surface. The laser spot was 2.2 mm in diameter. The pulse energy was mj at 543 nm and mj at 633 nm. The resulting photoacoustic signals were received by a piezoelectric probe that consisted of a circular piece of 25 µm thick polyvinylidene fluoride (PVDF) film (Ktech Corp., Albuquerque, NM) 51

65 mounted to the terminal end of a piece of semi-rigid coaxial cable (UT-141A, Micro- Coax, Pottstown, PA) [75]. The PVDF film transduced the acoustic signal into an electrical signal, which was transmitted back to an oscilloscope where the acoustic waveform was displayed. As shown in Figure 3.3, the coagulated and non-coagulated blood vessels within the phantom were 1.60 mm in diameter. They were aligned in parallel and located 6.80 mm apart, from center to center. A raster scan was done in a manner such that seven passes, each separated by 1/4 mm, were made perpendicular to the blood vessels. The scan was performed using an motorized x-y stage (X-Y Stage, Sherline, Inc., Vista, CA) using stepper motors controlled by a computer interface (Motion Assistant, National Instruments, Austin, TX). To begin, measurements were taken every 1 mm of the cross-section until a 543 nm photoacoustic signal greater than 4.00 mv was achieved. At this point, the transducer and optical fiber were sent back to the previous measurement point and the scan continued back in the forward direction with measurements taken every 1/4 mm until the 543 nm signal dropped below 4.00 mv. Such a scheme allowed us to minimize time where there were no blood vessels, yet still achieve good resolution at the vessels themselves. The blood vessel phantom contained two artificial blood vessels, one of coagulated blood and the other of non-coagulated blood. Due to the sedimentation of RBC s within whole blood, we were unable to use it in the whole form to represent noncoagulated blood. Instead, we lysed the RBCs by freezing the samples of blood and allowing them to thaw. This procedure had no effect on the portion of spectrum of interest to us, as the optical absorption of the hemoglobin molecule was unaffected. The burn phantom was designed to mimic coagulated and non-coagulated blood vessels embedded within the dermis. A polyacrylamide gel slab containing 1% Intralipid and a minute amount of Direct Red (Sigma Chemical, St. Louis, MO) was made containing two 1.60 mm diameter hollow cylinders, located 1.75 mm below the surface (Figure 3.3). To create a coagulated vessel within the phantom, one of these cylinders was filled with non-coagulated blood, sealed at the ends with silicone, and then immersed in 65 degree Celsius water for 10 minutes. To create a non-coagulated blood vessel, this procedure was repeated without the heating step. Intralipid, a 52

66 Transducer Oscilloscope Oscilloscope Scanner Arm Optical Fiber Transducer Coagulated Vessel Optical Fiber Water Coagulated Blood 1% Intralipid/Direct Absorbing Polyacrylamide Gel Viable Blood Viable Vessel 1.75 mm 1% Intralipid/Direct Red Polyacrylamide Gel 1.60 mm 6.80 mm (b) (a) Figure 3.3: (a) Schematic depicting the burn phantom and photoacoustic measurement setup. (b)photograph of the experimental set-up. The coagulated and noncoagulated blood vessels have been highlighted to indicate their presence just under the surface of the Intralipid/Direct Red solution. purely scattering solution at the relevant wavelengths, was used to mimic the optical scattering properties of the dermis while the Direct Red was added to mimic the mild optical absorption characteristics of the dermis [75]. Optical absorption properties attributable to the epidermis, such as melanin, were not taken into account during this investigation because the epidermal layer is sloughed off in severe burn wounds. Both the transducer and optical fiber were mounted together and rigidly attached to a scan arm. The ends of both the optical fiber and the transducer were submersed in water just above the phantom. This was done to ensure that the acoustic signal generated in the phantom was effectively coupled to the transducer Classification Analysis of Coagulated and Non-coagulated Blood The classification rules described in Section 3.2 were implemented in Matlab R 7.3 (Mathworks, Inc., Natick, MA). We used (3.3) and (3.4) to classify coagulated and non-coagulated blood, respectively. For the planar blood samples, we used the 53

67 leave one out cross validation scheme on 18 randomly chosen samples, 9 from the coagulated set and 9 from the non-coagulated set as the total training set. We then tested the remaining 18 samples using the two classification rules. For the planar samples and the blood vessel phantoms, we determined the APER. We report the ratio in terms of 543:633. For the blood vessel phantom, we used the same rules on each scan point. Initially, we used the training set mentioned above. We also developed a second training set on three of the seven cross sections and classified the remaining four cross sections. This second set was done since the variance of the planar measurements was smaller than the variance of the cylindrical measurements, suggesting that they may be a better and more relevant training set. Subsequent to implementing our statistical classification scheme, we used a deterministic method to assign appropriate pixel values. We plotted the scans using Matlab with a custom colormap, in which the values (0, 1/2) were assigned shades of brown to denote coagulated blood and [1/2, 1] were assigned shades of red to denote non-coagulated blood. Values of zero were assigned to pixels in which the 543 nm signal did not reach the threshold signal of 4 mv and were thus considered background. These pixels were colored white. For each pixel, we used the following function to assign an appropriate color. This function takes a photoacoustic ratio and assigns it a value in the range [0, 1/2]. x x 0 g(x 0 ) = 2 max x x where x 0 is the photoacoustic ratio and x is the mean ratio of the coagulated or noncoagulated distribution, whichever is relevant. If the classification rule determined that the pixel was coagulated, the function g(x 0 ) would be given the pixel, with a possible range of (0, 1/2]. If the classification rule determined that the pixel was non-coagulated, the pixel value would be g(x 0 ) + 1 with a possible range of (1/2, 1]. 2 Since the Matlab colormap uses values from [0, 1], this color map assigned shades of brown to coagulated pixels and shades of red to non-coagulated pixels. For the planar and cross section training sets, we produced images of the vessels and determined the APER. 54

68 3.4 Results Spectral Analysis of Coagulated and Non-coagulated Blood The optical absorption spectra for both coagulated and non-coagulated blood of the 18 different samples were averaged and plotted as shown in Figure 3.4. The ratio, S 543/633, for non-coagulated blood was 15.7:1, while the S 543/633 for thermally coagulated blood was 1.6:1. Absoroption Coefficient, µ a (cm -1 ) Coagulated Blood Viable Blood Wavelength (nm) Figure 3.4: Average optical absorption spectra for coagulated and non-coagulated blood. Arrows indicate the location of 543 and 633 nm Photacoustic Discrimination of Planar Blood Samples The resulting waveforms from 543 and 633 nm laser excitation of coagulated and non-coagulated blood vessels were plotted together to show the relationship between the two. A sample of the data obtained for one of the phantoms is shown in Figure 3.5. From the graphs, we took our peak-to-peak pressure amplitude for the corresponding wavelength, Pp p, λ at the initial peak of the pressure wave. The mean P 543/633 value for the 18 samples of non-coagulated blood in the training set 55

69 Amplitude (mv) Amplitude (mv) P P 543 P P (a) Time (μs) (b) Time (μs) Figure 3.5: Pressure waveforms resulting from 543 and 633 nm laser excitation for (a) non-coagulated blood and (b) coagulated blood. Peak-peak measurement points are indicated. was 13.46:1. The mean P 543/633 value for the 18 samples of coagulated blood in was 1.6:1. The APER for this planar set was 0% Photoacoustic Imaging Images of the 2-dimensional scan were created using the pixel values, P ij, at each measurement point (Figure 3.4.3(a) and (b)). The image with the planar training set was similar to the vessel training set on reconstruction of the coagulated vessel, though the non-coagulated vessel was reconstructed better with the vessel training set. The APER for the planar training set was 15.7%, while the APER for the vessel training set was 11.4%. 3.5 Discussion Photoacoustics offers a new tool in the diagnosis of burn injury, as the contrast in optical properties of injured and healthy tissue, coupled with the robust 56

70 Planar Training Set Vessel Training Set Figure 3.6: Two-dimensional images produced to highlight different blood types types. (a) Viable blood was highlighted in the image using P expected 543/633 = (b) Thermally coagulated blood was emphasized using P expected 543/633 = 1.6. signal carrying properties of acoustic waves in skin, provides information for noncontact evaluation of burn injury. While burn injury is dynamic and comprises many physiological processes of wound healing, we propose that the necrotic layer is represented by coagulated blood in contrast to non-coagulated blood perfusing inflamed or healthy tissue. This study showed that the two aforementioned states of blood can be distinguished using ratios of the photoacoustic pressure induced at 543 and 633 nm with good discrimination in a burn phantom. Though this method is still preliminary and needs refinement prior to taking it to clinical application, the ability to classify the two blood states allows non-invasive, spatial distinction between viable and necrotic tissue. Further work needs to be done, however, before photoacoustics can be considered for clinical management of burn injury. Primarily, the depth resolution and depth limit need to be determined. The necrotic layer, containing coagulated blood, may have an optical thickness too great for sufficient light penetration to the viable layer if the burn is too deep. We expect, however, 57