Optical Clearing of Tissues and Blood (SPIE Press Monograph Vol. PM154)

Bellingham, Washington USA

Library of Congress Cataloging-in-Publication Data Tuchin, V. V. (Valeriæi Viktorovich) Optical clearing of tissues and blood / Valery V. Tuchin. p. cm. Includes bibliographical references and index. ISBN 0-8194-6006-0 1. Tissues--Imaging. 2. Tissues--Optical properties. 3. Imaging systems in medicine. 4. Medical physics. I. Title. QP88.T66 2005 612'.028--dc22

2005022126

Published by SPIE—The International Society for Optical Engineering P.O. Box 10 Bellingham, Washington 98227-0010 USA Phone: +1 360 676 3290 Fax: +1 360 647 1445 Email: [email protected] Web: http://spie.org Copyright © 2006 The Society of Photo-Optical Instrumentation Engineers All rights reserved. No part of this publication may be reproduced or distributed in any form or by any means without written permission of the publisher. The content of this book reflects the work and thought of the author(s). Every effort has been made to publish reliable and accurate information herein, but the publisher is not responsible for the validity of the information or for any outcomes resulting from reliance thereon. Printed in the United States of America.

Contents

Preface

vii

Introduction: Brief Review

ix

Chapter 1 1.1 1.2 1.3 1.4

Tissue and Blood Optical Immersion by Exogenous Chemical Agents Tissue Structure and Scattering Properties Principles of Optical Immersion and Chemical Agent Diffusion Thermodynamics of Water Transport Through a Semipermeable Membrane Tissue Swelling and Hydration

1 1 12 18 19

Chapter 2 Optical Clearing of Fibrous Tissues 2.1 Spectral Properties of Immersed Eye Sclera 2.1.1 Monte Carlo modeling 2.1.2 In vitro measurements 2.2 In vitro Frequency-Domain Measurements of Eye Sclera 2.3 In vivo Measurements of Eye Sclera 2.4 Dura Mater Immersion and Agent Diffusion Rate 2.5 Conclusions

23 23 23 32 52 54 58 60

Chapter 3 Optical Clearing of Skin 3.1 Introduction 3.2 In vitro Speckle and Spectral Measurements 3.3 In vivo Spectral Reflectance Measurements 3.4 In vivo Frequency-Domain Measurements 3.5 OCT Imaging

61 61 63 70 75 78

Chapter 4 Optical Clearing of Gastric Tissue 4.1 Spectral Measurements 4.2 OCT Imaging

83 83 84

Chapter 5 Tissue Structural Properties Studies at Optical Immersion 5.1 Polarization Measurements 5.2 Confocal Microscopy 5.3 Fluorescence Measurements 5.3.1 Transmittance of fluorescent radiation 5.3.2 In vivo measurements for intact human skin

87 87 94 95 95 97

v

vi

Optical Clearing of Tissues and Blood

5.3.3 Two-photon scanning fluorescence microscopy 5.4 The Second Harmonic Generation

98 103

Chapter 6 Optical Clearing of Blood, Cells, and Microorganisms 6.1 Blood Flow and Microvessel Network Imaging 6.2 OCT Measurements of Blood Optical Clearing 6.3 Theoretical Analysis of Blood Optical Clearing 6.4 Optical Immersion of Cells and Microorganisms

107 107 112 125 133

Chapter 7 Chemical Agent Delivery 7.1 Introduction 7.2 Cosmetic Preparations and Skin Permeation 7.3 Diffusion of Macromolecules in Tissues 7.4 Enhanced Diffusion of Molecules Through Tissues 7.4.1 Occlusion 7.4.2 Chemical enhancers 7.4.3 Physical enhancers 7.5 Skin Reservoirs for Topically Applied Clearing Agents

147 147 148 151 155 155 156 157 162

Chapter 8 Applications 8.1 Glucose Sensing 8.1.1 Introduction 8.1.2 Tissue and blood scattering spectroscopy 8.1.3 OCT measurements 8.2 Speckle Technologies 8.2.1 Speckle topography and tomography 8.2.2 Diffusion-wave spectroscopy and functional imaging 8.3 Precision Tissue Photodisruption and Tattoo Removal

167 167 167 167 170 175 175 186 188

Chapter 9 Other Methods of Tissue Optical Properties Control 9.1 Tissue Compression and Stretching 9.2 Temperature Effects and Tissue Coagulation 9.3 Tissue Whitening

193 193 205 211

Conclusion

213

References

217

Index

245

Preface

This book describes an optical clearing method based on reversible reduction of tissue scattering due to refractive index matching of scatterers and ground matter, which was of great interest for research and application in the last decade and is a promising technique for future developments in the fields of tissue imaging, spectroscopy, phototherapy, and laser surgery. The basic principles, recent results, advantages, limitations, and future of the optical immersion method applied to clearing of the naturally turbid biological tissues and blood are overviewed in the book. The refractive index matching concept for enhancement of in-depth light and laser beam penetration into tissues and blood is discussed on the basis of in vitro and in vivo studies using optical spectroscopy, polarization, and coherence domain techniques. The optical properties of tissues with basic multiple scattering, which are transformed to a low scattering mode, are analyzed. It is shown that light reflection, transmission, scattering, and polarization can be effectively controlled. The diagnostic abilities of the method based on contrasting of abnormalities, on in-depth profiling of tissue and blood, and on monitoring of endogenous and exogenous matter diffusion within tissue are demonstrated. The author is grateful to Ms. Sharon Streams, who has supported the idea to publish the book, for her assistance in editing and production. The author is very thankful to attendees of the short courses “Coherence, Light Scattering, and Polarization Methods and Instruments for Medical Diagnosis,” “Tissue Optics and Spectroscopy,” and “Tissue Optics and Controlling of Tissue Optical Properties,” which the author has given during the SPIE Photonics West Symposia, SPIE/OSA European Conferences on Biomedical Optics, and OSA CLEO/QELS Conferences over the last ten years, for their stimulating questions, fruitful discussions, and critical evaluations of the presented materials. Their responses were very valuable for the preparation of this edition. The original part of this work was supported within both Russian and international research programs by grant N25.2003.2 of the President of Russian Federation “Supporting of Scientific Schools,” grant N2.11.03 “Leading ResearchEducational Teams,” contract No. 40.018.1.1.1314 “Biophotonics” of the Ministry of Industry, Science and Technologies of RF, grant REC-006 of CRDF (U.S. Civilian Research and Development Foundation for the Independent States of the Former Soviet Union) and the Russian Ministry of Education, the Royal Society grant for a joint project between Cranfield University (UK) and Saratov State University, grants of National Nature Science Foundation of China (NSFC), and by Palomar Medical Technologies Inc. vii

viii

Optical Clearing of Tissues and Blood

The author greatly appreciates the cooperation, contribution, and support of all his colleagues from the Optics and Biomedical Physics Division of the Physics Department and Research-Educational Institute of Optics and Biophotonics at Saratov State University, especially that of Profs. Yu. P. Sinichkin, D. A. Zimnyakov, V. I. Kochubey, I. L. Maksimova, and E. I. Galanzha, and Drs. A. N. Bashkatov, E. A. Genina, A. B. Pravdin, G. V. Simonenko, K. V. Larin, I. V. Meglinsky, S. P. Chernova, and I. V. Yaroslavsky. The author would like to thank all of his numerous colleagues and friends all over the world for collaboration. He is especially grateful to Dr. B. Chance and Dr. A. Yodh from the University of Pennsylvania (USA) for collaboration in the framework of the CRDF Research Program, to Dr. R. K. Wang and Dr. X. Xu from Cranfield University (UK), Dr. Q. Luo from HUST University (China), and Stoyan Tanev from Vitesse Re-Skilling Canada Inc. (Canada). The author expresses gratitude to his wife, Natalia, and to all of his family, especially to his daughter Nastya and grandchildren Dasha, Zhenya, and Stepa, for their indispensable support, understanding, and patience during the writing of this book. Valery Tuchin September 2005

Introduction: Brief Review

Reflection, absorption, scattering, and fluorescence in living tissues and blood can be effectively controlled by various methods.1−197 Staining (sensitization) of biological materials is used extensively to study mechanisms of interaction between their constituent components and light, and also for diagnostic purposes and selective photodestruction of individual components of living tissues. This approach underlies the diagnosis and photodynamic therapy (PDT) of malignant neoplasm, UV-A photochemotherapy of psoriasis and other proliferative disorders, angiography in ophthalmology, and many other applications in medicine. In the visible and NIR spectrums, tissues and bioliquids are low absorbing, but highly scattering media.7 Scattering defines spectral and angular characteristics of light interacting with living objects, as well as its penetration depth; thus, optical properties of tissues and blood may be effectively controlled by changes of scattering properties. The living tissue allows one to control its optical (scattering) properties using various physical and chemical actions such as compression, stretching, dehydration, coagulation, UV irradiation, exposure to low temperature, and impregnation by chemical solutions, gels, and oils. All these phenomena can be understood if we consider tissue as a scattering medium that shows all optical effects that are characteristic to turbid physical systems. It is well known that the turbidity of a dispersive physical system can be effectively controlled by providing matching of refractive indices of the scatterers and the ground material. This is a so-called optical immersion technique. Another possibility for controlling the optical properties of a disperse system is to change its packing parameter and/or scattered sizing.29,37,46 Control of optical properties tissues in vivo is very important for many medical applications. A number of laser surgery, therapy, and diagnostic technologies include tissue compression and stretching used for better transportation of the laser beam to underlying layers of tissue. The human eye compression technique allows one to perform transscleral laser coagulation of the ciliary body and retina/choroid.35,41,57 The possibility of selective translucence of the upper tissue layers should be very useful for developing of the eye globe imaging techniques and for detecting local inhomogeneities hidden by a highly scattering medium in functional tomography. Results on the control of human sclera optical properties by tissue impregnation with osmotically active chemicals such as Trazograph (x-ray contrast), glucose, and polyethylene glycol (PEG), as well as hypaque-60 (x-ray contrast), have been reported.6,11,21,24,31,32,36,38,82,84–89,162,163,174 In general, the scattering coefficient µs and scattering anisotropy factor g of a tissue is dependent on the refractive index mismatch between cellular tissue comix

x

Optical Clearing of Tissues and Blood

ponents: cell membrane, cytoplasma, cell nucleus, cell organelles, melanin granules, and the extracellular fluid. For fibrous (connective) tissue (eye scleral stroma, corneal stroma, skin dermis, cerebral membrane, muscle, vessel wall noncellular matrix, female breast fibrous component, cartilage, tendon, etc.), index mismatch of the interstitial medium and long strands of scleroprotein (collagen-, elastin-, or reticulin-forming fibers) is important. The refractive index matching is manifested in the reduction of the scattering coefficient (µs → 0) and increase of single scattering directness (g → 1). For skin dermis and eye sclera µs , reduction can be very high.32,36,45,90,198 For hematous tissue such as the liver, its impregnation by solutes with different osmolarity also leads to refractive index matching and reduction of the scattering coefficient, but the effect is not so pronounced as for skin and sclera, due to cells changing size as a result of osmotic stress.79,80 It is possible to achieve a marked impairment of scattering by means of the intratissue administration of appropriate chemical agents. Conspicuous experimental optical clearing in human and animal sclera; human, animal, and artificial skin; human gastrointestinal tissues; and human and animal cartilage and tendon in the visible and NIR wavelength ranges induced by administration of x-ray contrast agents (Verografin, Trazograph, and Hypaque-60), glucose, propylene glycol, polypropylene glycol-based polymers (PPG), polyethylene glycol (PEG), PEG-based polymers, glycerol, and other solutions as has been described in Refs. 6–8, 11, 21, 24, 31–40, 45, 52–54, 82–98, 101, 102, 104, 105, 107–110, 116, 117, 127, 138, 139, 141–144, 161–175, 179–183, 186–189, 192, 196, and 197. Coordination between refractive indices in multicomponent transparent tissues showing polarization anisotropy (e.g., cornea) leads to its decrease.5,10 In contrast, for a highly scattering tissue with a hidden linear birefringence or optical activity, its impregnation by immersion agents may significantly improve the detection ability of polarization anisotropy due to reduction of the background scattering.8,33,34,174,175,192,197 Concentration-dependent variations in scattering and transmission profiles in α-crystalline suspensions isolated from calf lenses are believed to be related to osmotic phenomenon.74 Osmotic and diffusive processes that occur in tissues treated with Verografin, Trazograph, glucose, glycerol, and other solutions are also important.32 Osmotic phenomena appear to be involved when optical properties of biological materials (cells and tissues) are modulated by sugar, alcohol, and electrolyte solutions. This may interfere with the evaluation of hemoglobin saturation with oxygen or identification of such absorbers as cytochrome oxidase in tissues by optical methods.79,80 Experimental studies on optical clearing of normal and pathological skin and its components (epidermis and dermis) and the management of reflectance and transmittance spectra using water, glycerol, glycerol-water solutions, glucose, sunscreen creams, cosmetic lotions, gels, and pharmaceutical products were carried out in Refs. 7, 8, 19, 45, 54, 83, 85–92, 94–96, 101, 102, 104, 105, 108, 110, 118, 142, 167, 168, 172, 179, 180, 186–189, 192, 193, 196, and 197. The control of skin optical properties was related to the immersion of refractive indices of scatterers

Introduction: Brief Review

xi

(keratinocytes components in epidermis and collagen, and elastin fibers in dermis) and ground matter, and/or reversible collagen dissociation.101 In addition, some of the observed effects appear to have been due to the introduction of additional scatterers or absorbers into the tissue or, conversely, to their washing-out. A marked clearing effect through the hamster45 and the human86,90,92,102,105,142 skin, the human and rabbit sclera,21,84 and rabbit dura matter143 was occurred for an in vivo tissue within a few minutes of topical application (eye, dura matter, skin) or intratissue injection (skin) of glycerol, glucose, propylene glycol, Trazograph, and PEG and PPG polymers. Albumin, a useful protein for index matching in phase contrast microscopy experiments,65,75–77 can be used as the immersion medium for tissue study and imaging.20,28 Proteins smaller than albumin may offer a potential alternative because of relatively high scattering of albumin. Sometimes medical diagnosis or contrasting of a lesion image can be provided by the enhancement of a tissue’s scattering properties by applying, for instance, acetic acid, which has been successfully used as a contrast agent in optical diagnostics of cervical tissue.20,28,128–133 It has been suggested that the aceto-whitening effect seen in cervical tissue is due to coagulation of nuclear proteins. Therefore, an acetic acid probe may also prove extremely significant in quantitative optical diagnosis of precancerous conditions because of its ability to selectively enhance nuclear scatter.20,28 Evidently, the loss of water by tissue seriously influences its optical properties. One of the major reasons for tissue dehydration in vivo is the action of endogenous or exogenous osmotic liquids. In in vitro conditions, spontaneous water evaporation from tissue, tissue sample heating at a noncoagulating temperature, or its freezing in a refrigerator push tissue to loose water. Typically in the visible and NIR spectrums, far from water absorption bands, the absorption coefficient increases by a few dozen percent, and the scattering coefficient by a few percent due to closer packing of tissue components caused by its shrinkage. However, the overall optical transmittance of a tissue sample increases due to the decrease of its thickness at dehydration.43,44 Specifically, in the vicinity of the strong water absorption bands, the tissue absorption coefficient decreases due to less concentration of water in spite of higher density of tissue at its dehydration. It is possible to significantly increase transmission through a soft tissue by its squeezing (compressing) or stretching it.73 The optical clarity of living tissue is due to its optical homogeneity, which is achieved through the removal of blood and interstitial liquid (water) from the compressed site. This results in a higher refractive index of the ground matter, whose value becomes close to that of scatterers (cell membrane, muscle, or collagen fibers). Closer packing of tissue components at compression makes the tissue a more organized system—what may give less scattering due to cooperative (interference) effects.37,46 Indeed, the absence of blood in the compressed area also contributes to altered tissue absorption and refraction properties. Certain mechanisms underlying the effects of optical clearing and changing of light reflection by tissues at compression and stretching were proposed in Refs. 21, 22, 35, 37, 46, 57, 145, and 198.

xii

Optical Clearing of Tissues and Blood

Long-pulsed laser heating induces reversible and irreversible changes in the optical properties of tissue.43,44,127 In general, the total transmittance decreases and the diffuse reflectance increases, showing nonlinear behavior during pulsed laser heating. Many types of tissues slowly coagulated (from 10 min to 2 hrs) in a hot water or saline bath (70–85◦ C) exhibit an increase of their scattering and absorption coefficients. UV irradiation causes erythema (skin reddening), stimulates melanin synthesis, and can induce edema and tissue proliferation if the radiation dose is sufficiently large.18,56,121,122 All these photobiological effects may be responsible for variations in the optical properties of skin, and need to be taken into consideration when prescribing phototherapy. Also, UV treatment is known to cause color development in the human lens. Natural physiological changes in cells and tissues are also responsible for their altered optical properties, which may be detectable and, thus, used as a measure of these changes. For example, measurements of the scattering coefficient allow one to monitor glucose52–54,153,154,156–159 or edema152 in the human body, as well as blood parameters.16,17 Many papers report optical characteristics of blood as functions of hemoglobin saturation with oxygen. The alterations of the optical properties of blood caused by changes of the hematocrit value, temperature, and flow parameters can be found in Refs. 7, 16, and 17. As a particle system, whole blood shows pronounced clearing effects that may be accompanied by induced or spontaneous aggregation and disaggregation processes, as well as RBC swelling or shrinkage at application of biocompatible clearing agents with certain osmotic properties.8,47,64,111–115,142,160,177,178,184,185,192,197

1 Tissue and Blood Optical Immersion by Exogenous Chemical Agents

1.1 Tissue Structure and Scattering Properties Soft tissue is composed of closely packed groups of cells entrapped in a network of fibers through which interstitial liquid percolates. At a microscopic scale, the tissue components have no pronounced boundaries, thus tissue can be considered as a continuous structure with spatial variations in the refractive index. To model such a complicated structure as a collection of particles, it is necessary to resort to a statistical approach.199 The tissue components that contribute most to the local refractive index variations are the connective tissue fibers (either collagen or elastin forming, or reticulin forming) that form part of the noncellular tissue matrix around and among cells, and cell membrane; cytoplasmic organelles (mitochondria, lysosomes, and peroxisomes); cell nuclei; and melanin granules.28,78–80,199–204 Figure 1 shows a hypothetical index profile formed by measuring the refractive index along a line in an arbitrary direction through a volume of tissue and corresponding to the statistical mean index profile. The widths of the peaks in the actual index profile are proportional to the diameters of the elements, and their heights depend on the refractive index of each element relative to that of its surroundings. This is the origin of the tissue-discrete particle model. In accordance with this model, the index variations may be represented by a statistically equivalent volume of discrete particles having the same index, but different sizes. The refractive indices of tissue structure elements, such as the fibrils, the interstitial medium, nuclei, cytoplasm, organelles, and the tissue itself can be derived using the law of Gladstone and Dale, which states that the resulting value represents an average of the refractive indices of the components related to their volume fractions:29 n¯ =

N


ni fi ,




fi = 1,

(1)

i

i=1

where ni and fi are the refractive index and volume fraction of the individual components, respectively, and N is the number of components. The statistical mean index profile in Fig. 1 illustrates the nature of the approximation implied by this model. According to Eqs. (1), the average background index is defined as the weighted average of the refractive indices of the cytoplasm and 1

2

Optical Clearing of Tissues and Blood

Figure 1 Spatial variations of the refractive index of a soft tissue. A hypothetical index profile through several tissue components is shown, along with the profile through a statistically equivalent volume of homogeneous particles. The indices of refraction labeling of the profile are defined in the text.199 Reprinted with permission.

the interstitial fluid, ncp and nis , as n¯ 0 = fcp ncp + (1 − fcp )nis ,

(2)

where fcp is the volume fraction of the fluid in the tissue contained inside the cells. Since approximately 60% of the total fluid in soft tissue is contained in the intracellular compartment, in accordance with Refs. 20, 21, 27, 28, 199, 205–210, ncp = 1.37 and nis = 1.35, it follows that n¯ 0 = 1.36. The refractive index of a particle can be defined as the sum of the background index and the mean index variation, n¯ s = n¯ 0 +

n,

(3)

which can be approximated by another volume-weight average,

n = ff (nf − nis ) + fnc (nnc − ncp ) + for (nor − ncp ).

(4)

Here, subscripts f, is, nc, cp, and or refer to the fibers, interstitial fluid, nuclei, cytoplasm, and organelles, which are identified as the major contributors to the index variations. The terms in parentheses in this expression are the differences between the refractive indices of the three types of tissue components and their respective backgrounds. The multiplying factors are the volume fractions of the elements in the solid portion of the tissue. The refractive index of the connectivetissue fibers is about 1.47, which corresponds to the approximately 55% hydration of collagen, its main component.32 The nucleus and the cytoplasmic organelles in mammalian cells that contain similar concentrations of proteins and nucleic acids, such as the mitochondria and the ribosomes, have refractive indices that fall within a relatively narrow range (1.38–1.41).20,28 The measured index for the nuclei is nnc = 1.39.209,210 Accounting for this and supposing that nor = nnc = 1.39, the

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

3

mean index variation can be expressed in terms of the fibrous-tissue fraction ff only:

n = ff (nf − nis ) + (1 − ff )(nnc − ncp ).

(5)

Collagen and elastin fibers comprise approximately 70% of the fat-free dry weight of the dermis, 45% of the heart, and 2–3% of the nonmuscular internal organs.199 Therefore, depending on the tissue type, ff may be as small as 0.02 or as large as 0.7. For nf − nis = 1.47 − 1.35 = 0.12 and nnc − ncp = nor − ncp = 1.39 − 1.36 = 0.03, the mean index variations that correspond to these two extremes are

n = 0.03–0.09. For example, the nucleus and cell membrane of fibroblasts have an index of refraction of 1.48, the cytoplasm has an index of 1.38, and the averaged index of a cell is 1.42.211 The collagenous fibrils of cornea and sclera have an index of refraction of 1.47, and the refraction index of the ground matter is 1.35.206 The relative index of human lymphocytes in respect to plasma varies from 1.01 < m < 1.08.212 Additional information on the refractive indices of biological cells and tissues may be found in Refs. 1, 7, 29, 213, and 214. The matter surrounding the scatterers (intercellular liquid and cytoplasm), the so-called ground substance, is composed mainly of water with salts and organic components dissolved in it. The ground matter index is usually taken as n0 = 1.35– 1.37. The scattering particles themselves (organelles, protein fibrils, membranes, protein globules) exhibit a higher density of proteins and lipids in comparison with the ground substance and, thus, a greater index of refraction, ns = 1.39–1.47. This implies that the simplest way to model tissue is to consider the binary fluctuations in the index of refraction of the various tissue structures. The refractive index variation in tissues, quantified by the ratio m ≡ ns /n0 , determines the light scattering efficiency. For example, in a simple monodisperse tissue model, such as dielectric spheres of equal diameter 2a, the reduced scattering coefficient is215   2πa 0.37  2 µs ≡ µs (1 − g) = 3.28πa ρs (m − 1)2.09 , (6) λ where µs = σsca ρs is the scattering coefficient, σsca is the scattering cross section, ρs is the volume density of the spheres, g is the scattering anisotropy factor, and λ is the light wavelength in the scattering medium. This equation is valid for noninteracting Mie scatterers, g > 0.9, 5 < 2πa/λ < 50, 1 < m < 1.1. For example, epithelial nuclei can be considered as spheroidal Mie scatterers with refractive index nnc , which is higher than that of the surrounding cytoplasm ncp . Normal nuclei have a characteristic diameter d = 4–7µm. In contrast, dysplastic nuclei can be as large as 20 µm, occupying almost the entire cell volume. In the visible range, where the wavelength λ0  d, the Van de Hulst approximation can be used to describe the elastic scattering cross section of the nuclei:216,217     2 sin δ 2 2 sin δ 1 2 + , (7) σsca (λ, d) = πd 1 − 2 δ δ

4

Optical Clearing of Tissues and Blood

with δ = 2πd(nnc − ncp )/λ0 ; λ0 is the wavelength of the light in vacuum. This expression reveals a component of the scattering cross section, which varies periodically with inverse wavelength. This, in turn, gives rise to a periodic component in the tissue optical reflectance. Since the frequency of this variation (in inverse wavelength space) is proportional to particle size, the nuclear size distribution can be obtained from the Fourier transform of the periodic component. Absorption for most tissues in the visible region is insignificant, except for the absorption bands of blood hemoglobin and some other chromophores.7,19,218 The absorption bands of protein molecules are mainly in the near-UV region. Absorption in the IR region is essentially defined by water contained in tissues. The ns /n0 ≡ m ratio determines the scattering coefficient. For example, in a simple monodisperse model of scattering dielectric spheres (Mie theory), the reduced scattering coefficient µs is defined by Eq. (6), where µs ∼ (m − 1)2 . It follows from Eq. (6) that only a 5% increase in the refractive index of the ground matter (n0 = 1.35 → 1.42), when that of the scattering centers is ns = 1.47, will cause a sevenfold decrease of µs . In the limit of equal refractive indices for nonabsorbing particles and background material, m = 1 and µs → 0. In a living tissue, the relative refractive index is a function of the tissue’s physiological or pathological state. Depending on the specificity of the tissue-state refractive index of the scatterers and/or the background change (increase or decrease); therefore, light scattering may correspondingly increase or decrease. Light scattering and absorption of particles that compose tissue or blood can be calculated by Mie theory.219–221 The relevant parameters are the size (radius a) of the particles, their complex refractive index 

ns (λ0 ) = ns (λ0 ) + ins (λ0 ),

(8)

the complex refractive index of the dielectric host (ground material in tissues or plasma in blood) n0 , and the relative refractive index of the scatterers and the ground materials, m = ns /n0 . The imaginary part of the complex refractive indices is responsible for light losses due to absorption. Mie theory yields the absorption and scattering efficiencies and the phase function from which the absorption and scattering coefficients, µs = ρσsca and µa = ρσabs , and the scattering anisotropy g are calculated; ρ is the scatterers’ (particles’) density. The corresponding scattering and absorption cross sections σsca and σabs , scattering phase function p(θ), and g factor are described by219 ∞

  2π
σsca = 2 (2n + 1) |an |2 + |bn |2 , k

(9)

n=1 ∞

σabs =

   2π
(2n + 1) Re(an + bn ) − |an |2 + |bn |2 , 2 k

(10)

n=1

 1  p(θ) = 2 2 |S1 |2 + |S2 |2 , k r

(11)

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

5


∞   2n + 1 4π g= 2 Re an bn∗ n(n + 1) k σsca n=1

+

∞
n(n + 2) n=1

n+1

   ∗ ∗ Re an an+1 + bn bn+1 ,

(12)

where k = 2π/λ is the wave number; an and bn are the Mie coefficients, which are functions of the relative complex refractive index of the particles, m = ns /n0 , and size parameter, 2πan0 /λ0 ; θ is the scattering angle; r is the distance from the scatterer to the detector; S1 and S2 are the elements of the amplitude scattering matrix, and the asterisk indicates that the complex conjugate is to be taken. The coefficients an and bn are  Sn (y)Sn (x) − mSn (y)Sn (x)     an = S  (y)ζ (x) − mS (y)ζ (x) , n n n n (13)  (y)S (x) − S (y)S  (x)  mS n n  n n  ,  bn = mSn (y)ζn (x) − Sn (y)ζn (x) where 2πan0 , λ0 2πans , y= λ0 ns m= , n0 x=

(14)

where ζn and Sn can be written in terms of Bessel functions:   0.5 πz    Jn+0.5 (z), Sn (z) =   2  ζn (z) = Sn (z) + iCn (z),   0.5   πz    Cn (z) = − Nn+0.5 (z), 2

(15)

where Jn+0.5 (z) is the Bessel function of the first kind and Nn+0.5 (z) is the Bessel function of the second kind. The derivatives of Sn and Cn can be obtained through   0.5  0.5 π πz      Jn+0.5 (z) + Jn+0.5 (z),  Sn (z) = 8z 2  0.5  0.5  π πz      Cn (z) = − Nn+0.5 (z) − Nn+0.5 (z). 8z 2

(16)

6

Optical Clearing of Tissues and Blood

The introduction of the specific scattering and absorption coefficients extrapolated to a volume fraction of 100% is useful for describing the scattering and absorption properties of tissues with controlled structural properties.150,151 In that case, and when the particles are sufficiently distant to prevent dependent scattering, the scattering, transport scattering, and absorption coefficients are proportional to the dimensionless volume fraction of the scatterers cs  µs = cs σ¯ sca , µs = cs σ¯ sca 1 − g(λ0 , a) , µa = cs σ¯ abs , (17) where the specific scattering and absorption coefficients σ¯ sca and σ¯ abs are expressed in cm−1 . For some tissues, the size distribution of the scattering particles may be essentially monodispersive, and for others it may be quite broad. Two opposite examples are transparent eye cornea stroma, which has a sharply monodispersive distribution, and turbid eye sclera, which has a rather broad distribution of collagen fiber diameters.29,213 The optical parameters of broad-banded particle size distributions are values averaged over the distribution, weighted by the volume fractions of particles with different diameters (2a). The resulting specific optical coefficients are the averaged values and can be defined using the concept of combination of photondiffusion theory and particle representation of a tissue.199 A microoptical model developed by the authors of Ref. 199 explains most of the observed scattering properties of soft tissue. The model treats the tissue as a collection of scattering particles whose volume fractions are distributed according to a skewed log-normal distribution, modified by a packing factor p to account for correlated scattering among densely packed particles. To account for the interparticle correlation effects, which are important for systems with volume fractions of scatterers higher than 1–10% (dependent on particle size), the following expression is valid for the packing factor ωp of a medium filled with a volume fraction fs of scatterers with different shapes:219 ωp =

(1 − fs )p+1 , [1 + fs (p − 1)]p−1

(18)

where p is a packing dimension that describes the rate at which the empty space between scatterers diminishes as the total density increases. The packing of spherical particles is described well by the packing dimension p = 3. The packing of sheet-like and rod-shaped particles is characterized by a p that approaches 1 and 2, respectively. Since the elements of tissue have all of these different shapes and may exhibit cylindrical and spherical symmetry simultaneously, the packing dimension may lie anywhere between 1 and 5. When one calculates optical coefficients at high concentrations of particles, the size distribution η(2a) should be accounted for. There is no universal distribution size function that would describe all tissues with equal adequacy. In optics of dispersed systems, Gaussian, gamma, or power size distributions are typical.222 Polydispersion for randomly distributed scatterers can be accounted for by using the

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

7

gamma distribution or the skewed logarithmic distribution of the scatterers’ diameters, cross sections, or volumes.21,29,32,199,223,224 In particular, for turbid tissues such as eye sclera, the gamma radii distribution function is applicable:21,32 η(a) = a µ exp(−µβ),

(19)

where σ/am = 2.35µ−0.5 , β = a/am , σ is the halfwidth of the distribution, and am is the more probable scatterer radius. For epithelial cells and their nuclei scattering structures, log-normal size distributions of spherical or slightly prolated ellipsoidal particles are characteristic:225   1 [ln(a) − ln(am )]2 η(a) = √ exp − . 2σ2 aσ 2π

(20)

In particular, for epithelial cells and their nucleus components, two log-normal size distributions for small and big spherical scatterers with the following parameters were found in a certain line of rat prostate carcinoma cells:225 am1 = 0.012 µm, σ1 = 1.15 µm, and am2 = 0.59 µm, σ2 = 0.43 µm. For description scattering characteristics of a particle with a complex shape, differ from sphere or long cylinder, some special procedures, for example, the method of T-matrices, can be applied.220,221,225 Complexly shaped scatterers, like cells themselves, may be modeled as aggregates of spherical particles. To account for the interparticle correlation effects, the distribution size function, described, for example by Eqs. (19) and (20), should be replaced by the correlation-corrected distribution199 η (2a) =

[1 − η(2a)]p+1 η(2a). [1 + η(2a)(p − 1)]p−1

(21)

Most of the observed scattering properties of soft tissue that are explained in the model treat tissue as a collection of scattering particles, whose volume fractions are distributed according to a skewed log-normal distribution modified by a packing factor, to account for correlated scattering among densely packed particles.199 Thus, the scattering centers in tissue have a wide range of dimensions and tend to aggregate into a complex form suggestive of fractal objects. The skewed logarithmic distribution function, which is the most plausible on physical grounds, is used extensively in particle-size analysis. The skewed logarithmic distribution function for the volume fraction of particles of diameter 2a has the view:199   Fv [ln(2a) − ln(2am )]2 3−Df , η(2a) = (2a) exp − Cm 2σ2 where   √ (4 − Df )2 σ2 4−Df Cm = σ 2π(2am ) exp 2

(22)

8

Optical Clearing of Tissues and Blood

is the normalizing factor and 

∞

Fv =

η(2a)d(2a) 0

is the total volume fraction of the particles; and the quantities 2am and σ set the center and width of the distribution, respectively; Df is the (volumetric) fractal dimension. At the limit of an infinitely broad distribution of particle sizes η(2a) ≈ (2a)3−Df .

(23)

For 3 < Df < 4, this power-law relationship describes the dependence of the volume fractions of the subunits of an ideal mass fractal on their diameter 2a. These size distributions expand the size distributions described by Eqs. (19) and (20) to account for the fractal properties of tissues. For calculations of the optical coefficients at a high concentration of particles, the size distribution η(2a) [Eqs. (22) and (23)] should be replaced by the correlation-corrected distribution, described by Eq. (21).199 Scatterers in the epidermal layer of the skin also exhibit a log-normal size distribution, whereas the spatial fluctuations in the index of refraction of dense fibrous tissues, such as the dermis and many other tissues, follow a power law.224,226 Assuming that the waves scattered by the individual particles in a thin slice of the modeled tissue volume add randomly, then the scattering coefficient of the volume can be approximated as the sum of the scattering coefficients of the particles of a given diameter,199 µs =

Np


µs (2ai ),

(24)

i=1

where µs (2ai ) =

η(2ai ) σs (2ai ); vi

Np is the number of particle diameters, η(2ai ) is the volume fraction of particles of diameter 2ai [see Eqs. (18)–(23)], and σs (2ai ) is the optical cross section of an individual particle with diameter 2ai and volume vi . The volume-averaged phase function p(θ) (and scattering anisotropy parameter g) of the tissue slice is the sum of the angular-scattering functions pi (θ) (and anisotropy parameters gi ) of the individual particles weighted by the product of their respective scattering coefficients Np p(θ) =

i=1 µs (2ai )pi (θ) ; Np i=1 µs (2ai )

(25)

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

9

Np g=

i=1 µs (2ai )gi (2ai ) . Np µ (2a ) s i i=1

(26)

The reduced scattering coefficient as usually defined as µ = µ(1 − g). Mie theory predicts that scattering introduced by spherical micrometer-sized particles is strongest if the particle radius and wavelength are of the same order of magnitude. Mie theory is strictly applicable only to particles of particular regular shapes, but results are still useful if the shape is irregular. In that case, the oscillatory structure of the scattering coefficient and anisotropy factor as a function of particle size, which is observed with spherical particles (Fig. 2), is averaged out.219

(a)

(b)

Figure 2 Scattering properties of nonabsorbing particles at a wavelength of 633 nm calculated by Mie theory.150,151 The transport scattering coefficient (a) strongly depends on both the particle size and the relative refractive index. This graph is approximately symmetric. The axis of symmetry is at ns /n0 = 1. While the transport scattering coefficient equals zero there, in (b) the scattering anisotropy factor is maximal. In some parts of the range shown, the functions are not monotonic, but rapidly oscillating.

The transport scattering coefficient increases strongly with the ratio of the real part of scatterer index and the background medium index, ns /n0 . In turn, the scattering anisotropy factor is maximal when this ratio approaches 1 (Fig. 2). For fully matched refractive indices of scatterers and background material, the scattering coefficient goes to zero, which means that only absorption is now responsible for the light beam extinction. However, as it follows from Mie theory, absorbing particles suspended in an index matched medium cause strongly forward-directed resonance scattering. Light absorption by such particles is smaller than expected from their bulk absorption coefficient.150,151 For the 1 µm diameter particles with ns = 1.6 and bulk absorption coefficient of their material equal to 104 cm−1 in an index matched medium, the particle system absorption coefficient µa = cs × 4120 cm−1 .

10

Optical Clearing of Tissues and Blood

(a)

(b)

Figure 3 Wavelength dependencies of scattering by nonabsorbing particles at partial refractive index matching, ns /n0 = 1.07, calculated by Mie theory.150 The specific transport scattering coefficient (a) of particles with very small or very big diameter is almost independent of the wavelength over the whole visible spectral range. For particles with intermediate diameters, it increases strongly toward shorter wavelengths. The scattering anisotropy factor (b) for very small and very big particles scarcely depends on the wavelength, but depends very much on the particle size.

However, in practice, the total index matching cannot always be provided, thus other mechanisms of tissue clearing may be essential. Sometimes, the action of hyperosmotic chemical agents or strong mechanical compression may lead to reversible or irreversible change in the scatterers’ size. The wavelength dependencies of scattering parameters for systems of partially matched refractive indices of scatterers and background (ns /n0 = 1.07) are shown in Fig. 3. Such level of matching is typical for many normal connective and cell-structured tissues. The spectral variation of the relative index has been neglected in calculations, but may be relevant in practice. If particle size and ratio of refractive indices are fixed, the wavelength dependencies are caused by the spectral variation of the ratio of the wavelength to the particle size. For particles with a refractive index close to that of the host (see Fig. 3), the scattering coefficient of the particle systems with very small or very big diameters of particles is almost independent of the wavelength in the range from 400 to 800 nm, while that of the system with the intermediate diameters of particles decreases with wavelength. The same tendency in the wavelength dependence (no dependence for very small and very big scatterers, and decreased for intermediate diameters) is expected for the scattering anisotropy factor. There are plateaus if the particles are much smaller (isotropic scattering) or larger (highly anisotropic scattering) than the wavelength, with a steep increase between.150 It follows from this consideration that reduction of scattering may be associated not only with the refractive index matching, but also with the changes of the scattering system sizing. Both aggregation to big-sized particles and disaggre-

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

11

gation to small-sized particles will lead to scattering damping, but the scattering anisotropy properties of the newly formed system should be quite different. The last can be used in the understanding of the tissue clearing mechanisms associated with particle sizing and refractive index matching. Conceptually, for many situations, the leading mechanism of tissue clearing might be a refractive index matching, because the equalizing of refractive indices of scatterers and surrounding media always takes place at tissue immersion, dehydration, or compression, and the sensitivity of the scattering properties to refractive index matching is very high. The transmission of collimated light by a tissue layer of thickness d is defined by Bouguer–Beer–Lambert law:1–7 Tc =

I (d) = exp(−µt d), I0

(27)

where I0 and I (d) are the intensities of the incident and detected light, respectively; µt = µa + µs is the attenuation coefficient. For the human sclera at the wavelength λ = 800 nm, the absorption coefficient µa ∼ = 1.6 cm−1 and the reduced scattering  −1 7 ∼ coefficient µs = µs (1 − g) = 38 cm . For g = 0.9, µs ∼ = 380 cm−1 . Measurement of diffusely reflected light is often used to infer bulk tissue optical properties for the aims of in vivo tissue spectroscopy and imaging. To provide such measurements, an adequate calculating algorithm should be derived. The diffusion equation solved subject to boundary conditions at the interfaces is the basis for the calculation algorithm.227–231 These boundary conditions are derived by considering Fresnel’s laws of reflection, and balancing the fluence rate and photon current crossing the interface. For the source term modeled as a point scattering source at a depth of one transport mean free path, l t , and extrapolated boundary approach satisfying the boundary condition, the spatially resolved steady-state reflectance per incident photon R(r sd ) is expressed as230,231 R(rsd ) =

   FU 1 exp(−µeff r1 ) lt µeff + 4π r1 r12    1 exp(−µeff r2 ) + (lt + 2zb ) µeff + r2 r22   FF exp(−µeff r1 ) exp(−µeff r2 ) + , − 4πD r1 r2

(28)

where r sd is the distance between the light source and the detector at the tissue   2 ; r = (l + 2z )2 + r 2 ; surface (source-detector separation), cm; r1 = lt2 + rsd 1 t b sd zb = 2AD is the distance to the extrapolated boundary, where A = (1 + Reff )/(1 − Reff ), Reff is the effective reflection coefficient, which can be found by integrating the Fresnel reflection coefficient over all incident angles,229 and D ∼ = 1/(3µs + µa )

12

Optical Clearing of Tissues and Blood

is the photon diffusion coefficient. The parameters FU and FF represent the fractions of the fluence rate and the flux that exit the tissue across the interface. These values are obtained by integration of the radiance over the backward hemisphere230 and depend on refractive index mismatch on the boundary.231 The measurement of the intensity of a back-reflected light from a tissue for different source-detector separations rsd is the basis of the spatially resolved technique, which allows one to evaluate the absorption and the scattering coefficients using, for example, the analytical expression (28), valid for a highly scattering thick tissues.

1.2 Principles of Optical Immersion and Chemical Agent Diffusion Let us consider the principles of the optical immersion technique based on the impregnation of a tissue or dilution of blood by a biocompatible chemical agent, which also may have some osmotic properties. Any connective (fibrous) tissue can be effectively impregnated by a liquid agent or its solution. As an example of fibrous tissue, human sclera will be analyzed. A model of the human sclera in a local region can be represented as a slab with a thickness d that is filled by thin and long dielectric cylinders (collagen fibers) with average diameter ∼100 nm and refractive index nc = 1.474 (see Figs. 4 and 5).32,39,40,232 The cylinders are located in planes that are in parallel to the slab surface, but within each plane their orientations are random (see Fig. 4). The space between collagen fibers is filled by homogeneous ground substance with refractive index n0 = 1.345. A considerable refractive indices mismatch between collagen fibers and ground substance turns the system turbid, i.e., causes the multiple scattering and poor transmittance of propagating light. The refractive index of the background is a controlled parameter and may be changed in the range from 1.345 to 1.474, which is what transits the system from multiple to low-step and even single-scattering mode. For nc = n0 = 1.474, the medium becomes totally homogeneous and optically transparent if absorption of scatterers is small. The described model of tissue is applicable to any fibrous tissue including skin dermis and muscle. Indeed, refractive indices and fibers’ diameters and their spacing should be adequately chosen for a specific tissue type. Owing to the fibrous structure of the sclera, it is quite reasonable to assume that the dynamics of fluid diffusion within the tissue could be approximated by free diffusion.32,161,233 Therefore, to describe the dynamics of the refractive index change and corresponding decrease of the scattering coefficient when a chemical agent diffuses within the interfibrillar substance of a tissue, we use the model of free diffusion with the approximate solution of the diffusion equation32,233 ∂Cf (x, t) ∂ 2 Cf (x, t) = Df , ∂t ∂x 2

(29)

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

13

Figure 4 Collagen fibrils in human sclera.232 Scleral collagen fibrils display various diameters. They are much larger than those in the cornea. Mf is the microfibril (×18,000, scanning electron microscopy). Copyright Association for Research in Vision and Ophthalmology. Reprinted with permission.

14

Optical Clearing of Tissues and Blood

Figure 5 Schematic representation of human scleral sample and the geometry of light irradiation.32

where Cf (x, t) is the fluid concentration, Df is the coefficient of diffusion, and x is the spatial coordinate. This equation is applicable in cases when the rate of the process is not limited by membranes, such as the diffusion of substances in the interfibrillar space or when a substance in solution has a high rate of permeation through membranes.233 For a plane slab with a thickness d, which is placed at the moment t = 0 in a solution with the initial concentration of the agent Ca0 (the initial concentration of the agent within the slab is equal to 0, i.e., t = 0, 0 ≤ x ≤ d, Ca (x, 0) = 0, and the boundary conditions are Ca (0, t) = Ca (d, t) = Ca0 ), Eq. (29) has the following solution describing the time-dependent distribution of agent concentration within a sample:32,233           t πx 1 9t 3πx  4    exp − sin + exp − sin  1 − π τ d 3 τ d      Ca (x, t) = Ca0 , 25t 5πx 1     exp − sin + · · · +   5 τ d (30) where τ= Da is the agent diffusion coefficient.

d2 , π2 Da

(31)

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

15

The ratio of the amount of dissolved matter mt at the moment t to its equilibrium value m∞ is defined as233 d

Ca (x, t)dx Ca0 d         1 9t 1 25t t 8 + exp − + exp − + · · · . (32) = 1 − 2 exp − τ 9 τ 25 τ π

mt = m∞

0

This ratio in turn defines the volume-averaged concentration of an agent Ca (t), which in the first-order approximation has a view32,117,161 1 Ca (t) = 2



d 0

   t ∼ . Ca (x, t)dx = Ca0 1 − exp − τ

(33)

Equations (30)–(33) allow one to find the time-dependent concentration of chemical agents with a relatively low molecular weight at depth x within a tissue sample, or time variations of the total amount of these agents in the sample if the diffusion coefficient Da of these molecules in the tissue is known. On the other hand, measurements of Ca (t) make it possible to estimate the Da value of implanted molecules in the interstitial fluid of the tissue. For low molecular weight compounds, the values of their diffusion coefficients in their own media are about 10−5 cm2 s−1 ; for water, Da = 2.5 × 10−5 cm2 s−1 , and for saccharose, Da = 0.5 × 10−5 cm2 s−1 .233 When an agent is administrated through only one sample surface, such a situation also may take place for in vivo agent applications, Eq. (33) is still valid, but with another expression for the characteristic diffusion time:179 τ=

d2 . Da

(34)

Equations (30)–(34) were established for diffusion through a homogeneous slab. Due to its fibrous structure, a tissue can be presented as a porous material that leads to modification of the chemical agent diffusion coefficient Da =

Dai . p

(35)

Here, Dai is the chemical agent diffusion coefficient within the interstitial fluid and p is the porosity coefficient defined as: p=

V − VC , V

(36)

where V is the volume of the tissue sample, and VC is the volume of collagen fibers.

16

Optical Clearing of Tissues and Blood

To describe the diffusion of bigger molecules, the theory of hindered diffusion through a permeable membrane should be used.32,161,233 Based on Fick’s law, which limits the flux of matter J (mol/s/cm2 ) to gradient of its concentration: J = −Da

dC . dx

(37)

For stationary transport of matter through a thin membrane, we have233 J = Pa (C1 − C2 ),

(38)

where Pa = Da /d is the coefficient of permeability, and C1 and C2 are the concentrations of molecules in two spaces separated by a membrane. Using Eqs. (37) and (38), it is possible to find the variation in concentration of molecules inside a closed space with a volume V , surrounded by a permeable membrane, and with an area S by using the following equation:233 dC2 Pa S = (C1 − C2 ). dt V

(39)

For a large external volume when C1 can be considered as a constant, Eq. (39) has an approximate exponential solution in a form similar to Eq. (33),32,161,233 with C2 = Ca , C1 = Ca0 , and τ described by Eq. (34), that indicates that in experiments with tissue plane slabs (see Fig. 6.), V = S × d, where S and d are the area and thickness of the sample. The form of this equation is the same as that for free diffusion, but the values of the diffusion coefficient for free and hindered diffusion can be significantly different.

Figure 6 Schematic representation of diffusion of the immersion agent into a tissue sample and light transmittance and scattering.117

In tissue impregnation by a chemical agent, the refractive index of the background (interfibrillar) media n0 is a time-dependent function of the agent concentration, which penetrates into a sample Ca (t) and is defined by Eq. (32). The timedependent volume fraction of the agent within the tissue sample fa is proportional

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

17

to its concentration Ca , thus using the law of Gladstone and Dale [see Eq. (2)], we can write n0 (t) = n0 (t)f0 (t) + na fa (t),

(40)

where f0 (t) + fa (t) = 1. For application of nonosmotic or low-osmotic agents, the initial refractive index of the interfibrillar space can be considered as being independent of time, n0 (t) ∼ = n0 (t = 0). The expression for the scattering coefficient, derived for a system of noninteracting thin cylinders with a number of fibrils per unit area ρs has a form:32,161  5 4 3   π a n0 2 2 2 ∼ (m − 1) 1 + 2 , µs = ρs (m + 1)2 λ30

(41)

where ρs = fcyl /πa 2 , fcyl is the surface fraction of the cylinders’ faces, a is the cylinder radius, m = ns /n0 is the relative index of refraction of the cylinders (scatterers) to the background (interfibrillar space), and λ0 is the wavelength in vacuum. As a first approximation, it is reasonable to assume that the radii of the scatterers (fibrils) and their density cannot be significantly changed by chemicals (no tissue swelling or shrinkage take place), the absolute changes of n0 are not very high, and variations in µs are caused only by the change in the refractive index of the interstitial (interfibrillar) space in respect to the refractive index of the scatterers. Then, accounting for the fact that for a majority of tissues m ≈ 1, the ratio of the scattering coefficients at a particular wavelength as a function of the refractive index ratio m can be written in the following form32,161   m2 − 1 2 ∼ µs2 = µs1 . m1 − 1

(42)

Indeed, this relation describes the tissue scattering properties’ change due to refractive index match or mismatch caused by changes of refractive indices of the scatterers or the background, or the both. The similar equation for Mie spherical particle systems follows from Eq. (6). Due to square dependence, the sensitivity to indices matching is very high, for instance, for m1 = 1.1 and m2 = 1.01, µs2 ∼ = 0.01µs1 . For the immersion technique, the refractive index of the scatterers ns is usually kept constant during tissue impregnation by an agent. Thus, we can use Eq. (40) to rewrite Eq. (42) in a form that is specific for the tissue impregnation by an agent with a weak osmotic strength: 2 ns −1 n0 (t) µs (t) = µs (t = 0) ×  2 . ns −1 n0 (t = 0) 

(43)

18

Optical Clearing of Tissues and Blood

It should be noted that a more rigorous approach to calculating the scattering coefficient must be based on a consideration of light scattering by densely packed systems of thin dielectric cylinders or spherical particles with a particular size distribution [see Eqs. (18)–(26)]. To estimate changes of tissue collimated transmittance caused by the agent diffusion into a sample (see Fig. 6) Eqs. (27), (33), (40), and (43) should be used together. For the spatially-resolved reflectance experiments, Eq. (27) should be replaced by Eq. (28). Usually, immersion agents do not have strong absorption bands within the wavelength range of interest; thus, the absorption coefficient may often be considered as a constant value. Indeed, the diffuse transmittance and reflectance as well as differential scattering characteristics (angular dependent scattering) for a tissue sample can be calculated if the behavior of the scattering anisotropy factor g at the optical immersion is known. For Mie particles, the analysis of the g-factor behavior due to refractive index matching can be done using Eq. (12); the corresponding calculations are presented in Figs. 2 and 3.

1.3 Thermodynamics of Water Transport Through a Semipermeable Membrane Water balance in living tissues is one of the important features of tissue condition. At tissue interaction with external or internal molecules’ diffusion (proteins, sugars, alcohols, polymers, etc.), tissue water should be involved in molecular displacement processes.32 Water may be transported through a membrane (a certain tissue layer) by an increased concentration of dissolved substance in one of two parts of the system. This happens for membranes that are more permeable for water than for dissolved material and the process is called osmosis.233 The simplest case of water transport is when a membrane is permeable for water and totally impermeable for molecules of dissolved substances. However, in general, biological membranes are permeable for both water and dissolved substances, but the degree of permeability for each can be quite different. This is the most complicated case to describe, but the situation becomes simpler when water and dissolved substances permeate through the same paths inside a membrane (such as interfibrillar spaces (pores) in fibrous tissues, which are filled by the interstitial fluid containing water). In that case, the fluxes of water and dissolved substances interact and each flux is dependent on the degree of interaction. Such interaction between the stationary fluxes can be well described within the framework of irreversible thermodynamics.233 Assuming that in a system there is only one type of dissolved molecule (i.e., two fluxes move through a membrane—the water flux JW and a dissolved matter JS , which are proportional to the gradients of the chemical potential of water and dissolved matter), the volumetric flux is defined as233 JV = JW V W + JS V S ,

(44)

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

19

where V W and V S are the partial mole volumes, in the form JV = Lp (
p − σRT
CS ).

(45)

The flux of the dissolved matter can be expressed as233 JS = RT ω
CS + C S (1 − σ)JV .

(46)

Here, in Eqs. (45) and (46), Lp is the phenomenological coefficient indicating that the volumetric flux can be induced by rising hydrostatic pressure
p; σ is the reflection coefficient (σ = −(Lpd /Lp ), where Lpd is the phenomenological coefficient indicating, on the one hand, the volumetric flux that can be induced for the membrane by the osmotic pressure RT
CS , and on the other, the efficiency of the separation of water molecules and dissolved matter); ω = (LD − Lp σ2 )C S , where LD is the phenomenological coefficient characterizing the interchange flux induced by osmotic pressure RT
CS ; and C S is the average concentration of dissolved matter in two interacting solutions. For the ideal partially permeable membrane, σ = 1. For membranes that are permeable for molecules of dissolved matter, 0 < σ < 1. Equations (45) and (46) are valid for solutions with a low concentration. It should be noted that the volumetric flux for a partially permeable membrane described by Eq. (45) has the same mechanism of creation for both hydrostatic and osmotic pressure. So for porous (fibrous) materials (such as sclera, derma, muscle, dura mater), it is expected that osmotic pressure induces the flux of water due to increasing hydrostatic pressure, but not through independent diffusion of water molecules caused by their concentration gradient, because this entails considerably more resistance.

1.4 Tissue Swelling and Hydration When applying a chemical agent, a change in the pH level of the environment is very important because swelling or shrinkage of tissue is expected.234 The swelling or shrinkage of a fibrous tissue is caused not only by the increase (decrease) of collagen (elastin) fibril size, but also by the increase (decrease) of the sample volume due to the rise (diminution) of the mean distance between fibrils. It is well known that the change of the environmental pH to the more acid or more alkaline side from a colloid isoelectric point increases the degree of swelling. It is explained by the appearance of a positive or negative charge of colloid particles and, therefore, by the increase of hydration degree. In general, the initial pH condition of the tissue under study and the acid or alkaline nature of the impregnated solution may lead to different dependencies of tissue thickness or volume on chemical agent concentration (or time of impregnation) due to changes of pH. Such behavior of a tissue sample should be taken into account when optical measurements are used for estimation of tissue properties. For example, the swelling or shrinkage was watched for different initial conditions of a scleral tissue sample preparation and solutions used.32,84,85,88,162,163

20

Optical Clearing of Tissues and Blood

A detailed study of the swelling of bovine sclera and cornea as a function of pH and ionic strength of the bathing medium, using an equilibration technique that prevents the loss of proteoglycans during swelling is presented in Ref. 235. X-ray diffraction was used to measure the intermolecular spacings (IMS), fibril diameters and D periodicity, and interfibrillar spacings (IFS) of collagen as functions of pH, ionic strength, and tissue hydration. Hydration H was defined as H=

Weightwet − Weightdry Weightdry

.

(47)

It was found that both tissues swelled least near pH 4 (the isoelectric point), that higher hydrations were achieved at low ionic strengths, and that sclera swelled about one-third as much as cornea under most conditions. The IMS in both tissues decreased as the ionic strength was increased; for scleral hydration H ∼ = 2.5 and pH 7.0, the IMS changed from 1.71 to 1.61 nm at a 33-fold increase of ionic strength. The IMS causes virtually no change in the hydration when H > 1; H = 3.2 is physiological hydration. The corresponding mean value for the cornea was 1.75 ± 0.04 nm (n = 12) and for the sclera it was 1.65 ± 0.02 nm (n = 9) at pH 7.4. For dehydrated tissues (H = 0), the mean spacing was 1.25 ± 0.02 nm (n = 2) for the cornea and 1.27 ± 0.01 nm for the sclera. The packing of fibrils, defined as (IFS)2 , is another important parameter that determines the control of tissue light scattering. For bovine cornea at physiological pH 7.4, the squared IFS decreased linearly from approximately 9.2 × 103 nm2 for a highly hydrated tissue (H = 9.7) to approximately 2.1 × 103 nm2 at tenfold lesser hydration, and was equal to 4.2 × 103 nm2 at physiological hydration, H = 3.2. Both fibril diameters [mean value 39.0 ± 1.1 nm (n = 6)] and the D periodicity [mean value 64.50 ± 0.35 nm (n = 6)] of corneal collagen were essentially independent of hydration and pH when hydration was above unity. This means that the fibrils preferentially absorb the initial water and then remain at a relatively constant diameter. The remaining unchanged state of the D periodicity with hydration indicates no significant changes in the dimensions along the axis of the fibrils during swelling. The same tendencies are expected for the sclera as a collagen-based media. Following Ref. 235, let us consider the distribution of water within swollen fibrous tissue. The volume of a tissue at a given hydration may be expressed in terms of the dry volume. Tissue hydration, defined by Eq. (47), can be written as H=

VW , VT ρ T

(48)

where VW and VT are the volumes occupied by the water and the dry tissue, respectively. The density of water is taken as 1 g/ml, and ρT is the density of the dry tissue. Because the dried fibrous tissue can be considered to consist of two parts,

Tissue and Blood Optical Immersion by Exogenous Chemical Agents

21

fibrils (collagen, elastin, reticulin) and extrafibrillar material, ρT can be expressed as ρT =

MF + Mext ρF VF + ρext Vext = , VF + Vext VF + Vext

(49)

where MF and Mext are the mass per unit volume of the fibrils and the extrafibrillar material in the dried tissue, respectively; VF , ρF and Vext , ρext are the volume fractions and the densities of the fibrils and the extrafibrillar material in the dried tissue, respectively. Using Eq. (48), we can find the volume of the hydrated tissue as VH = VT + VW = VT (1 + ρT H ).

(50)

To assess the effect of added water on the interfibrillar spacings, the fraction of cellular water of volume VCW must be deduced from the total water in the tissue. So, on the assumption that the volume of the cellular water changes proportionally with that associated with each fibril, Eq. (50) may be rewritten as     VCW H . (51) VH = VT + VW = VT 1 + ρT 1 − VW At physiological hydration of the cornea, the cellular water is approximately 15% of the total water in the tissue; thus, VCW /VW can be taken as 0.15. ρT can be evaluated from Eq. (49) if VF , ρF , Vext , and ρext are known. For cornea ρF = 1.41 g/ml and ρext = 1.06 g/ml,235 VF and Vext can be calculated on the assumption that the dry collagen fibrils are packed with a hexagonal arrangement, and that they are distributed evenly throughout the tissue. In this case, the volume of tissue associated with each fibril (corresponding to the volume of a “unit cell”) is i 2 /1.15, where i is the nearest-neighbor distance (see Ref. 235). Thus, VF can be expressed as VF =

1.15πr 2 , i2

(52)

where r is the radius of the dry collagen fibril. At hydration H = 0 of bovine cornea, i = 34 ± 2 nm and r can be calculated to be ∼ 13.5 nm according to the percentage reduction in the intermolecular spacings obtained between H = 3.2 (physiological hydration) and H = 0 (dried state).235 This leads to the values VF = 0.555 and Vext = 1 − VF = 0.445. Using these values in Eq. (51), the expression describing the volume of the cornea at its hydration can be written as VHyd = VT (1 + 1.066 · H ).

(53)

Equations (51) and (53) should apply equally to any volume of tissue, i.e., to the volume associated with each fibril—i 2 /1.15 for a hexagonal arrangement or

22

Optical Clearing of Tissues and Blood

i 2 /1.12 for a “pseudohexagonal” arrangement.235 For bovine cornea, i 2 /1.12 = 1032 nm2 ; thus, at physiological hydration H = 3.2, the total volume associated with each fibril (the volume of dry material and the volume of water) is equal to 1032 nm2 × (1 + 1.066 × 3.2) = 4552 nm2 . This theoretical value correlates well with the experimental one of squared interfibrillar spacings, equal to 4.2×103 nm2 . The swelling of scleral tissue follows similar principles as for cornea with the same isoelectrical point around pH 4, but at a lower level of swelling. As it was noted in Ref. 235, there are several reasons for the low hydrations of the sclera: the low concentration of proteoglycans; a high collagen content and larger fibrils with a smaller combined surface area than in the cornea; and structural peculiarities connected with the fibrils arranged in bands, which may branch and interweave with each other. It was found by the authors of Ref. 235 that the light scattering (less transmittance) of bovine cornea increased more rapidly with hydration (even below physiological hydration) at pH values around the isoelectric point. For example, at pH 5 transmittance was approximately 98% for H = 2, 87% for H = 3.2, and only 12% for H = 6. In contrast, the light scattering at higher pH values (6–8) changed slowly with the hydration: transmittance was higher than 90% for each level of hydration from 1 to 7, with the local maximum of transmittance of 98% for H = 4. According to current models,7,29,37 corneal transparency at a given wavelength depends on certain structural parameters such as fibril diameters, the density of fibril packing, the position of each fibril relative to its neighbors, and the refractive indices of the collagen and the interfibrillar matrix, and changes in one or more of these parameters may be sufficient to decrease or increase light scattering. From these studies, it follows that to improve corneal transparency caused by stromal edema, hypertonic drops extracting enough water from the tissue may be used. As it was shown in Ref. 235, sodium chloride should be better than other hypertonic preparations for the treatment of corneal edema, because it may also reduce the swelling pressure in the stroma and decrease the fibril diameter if used frequently. The connection between the hydration H (mg water per mg dry tissue weight) and corneal thickness d (in mm) is described by the following empirical formulas:164,165,236 for rabbit cornea H = 10 · d − 0.42,

(54)

H = 5.3 · d − 0.67.

(55)

and for bovine cornea

2 Optical Clearing of Fibrous Tissues

2.1 Spectral Properties of Immersed Eye Sclera 2.1.1 Monte Carlo modeling Normally, eye sclera is a turbid medium that is nontransparent in the visible range (see Fig. 7).35,198 The origin of scleral spectra formation can be understood on the basis of light scattering by a system of polydispersive irregularly arranged collagen cylinders immersed in the ground substance with a lower refractive index29 and strong absorption bands.35,36,198 With natural thickness of 0.6–0.8 mm, this tissue shows multiple scattering and looks like a white matter. The transition from a multiple to a low-step/single scattering can be provided by the drying of a tissue sample36,198 [Fig. 7(c)] or its impregnation by an immersion liquid.32,36,39,161 Figure 5 is a schematic representation of the human scleral sample structure and geometry of light irradiation. Analytical approaches for describing the propagation of light in the sclera are valid only when strongly simplifying assumptions are used, which make the model substantially less adequate. Thus, the direct simulation of photon migration in a medium using a Monte Carlo simulation was used for calculating spectral characteristics and photon statistics.32,39 The Monte Carlo simulation of the sclera transmission and reflection spectra was carried out using the probability function for the free photon path l.1–7 The ordering of scatterers (thin dielectric cylinders) was taken into account, using the experimental radial distribution function g(r) obtained from electron micrographs of human sclera39,232 (Figs. 4 and 8). Light propagation in a densely packed disperse system can be analyzed using the radial distribution function g(r), which statistically describes the spatial arrangement of particles in the system. The function g(r) is the ratio of the local number density of the fibril centers at a distance r from a reference fibril (at r = 0) to the bulk number density of the fibril centers.238 It expresses the relative probability of finding two fibril centers separated by a distance r; thus, g(r) must vanish for values of r ≤ 2a (a is the radius of a fibril—fibrils cannot approach each other closer than touching). The radial distribution function of scattering centers g(r) for a certain tissue may be calculated on the basis of tissue electron micrographs (see Fig. 4). The technique for the experimental determination of g(r) involves counting the number of particles placed at a specified spacing from an arbitrarily chosen initial particle, followed by statistical averaging over the whole ensemble. In a two-dimensional case, the particle number
N at the spacing from r to r +
r 23

24

Optical Clearing of Tissues and Blood

(a)

(b)

(c) Figure 7 Transmittance spectra of human sclera:35,198 (a) Total transmittance for three samples. (b) Axial (collimated) transmittance for the same three samples. (c) Total transmittance of dry sclera sample.

Optical Clearing of Fibrous Tissues

25

Figure 8 Histograms of radial distribution functions g(r) obtained from electron micrographs of human sclera.237 Reprinted with permission.

is related to function g(r) by the following equation:
N = 2πρg(r)r
r,

(56)

where ρ is the mean number of particles for a unit area. The radial distribution function g(r) was first found for the rabbit cornea by Farrell et al.238 Similar calculations for several regions of the human eye sclera237 are illustrated in Fig. 8. Electron micrographs from Ref. 232, averaged for 100 fibril centers, were processed. The function g(r) for the sclera was obtained on the basis of the spatial distribution of the fibril centers, neglecting the discrepancy in their diameters. Some noise is due to the small volume of statistical averaging. The obtained results present evidence of the presence of a short-range order in the sclera, although the degree of order is less pronounced than in the cornea. The function g(r) = 0 for r ≤ 100 nm, which is consistent with the mean fibril diameter of ≈100 nm derived from the electron micrograph (see Fig. 4).232 The first peak in the distribution gives the most probable separation distance, which is approximately 285 nm. The value of g(r) is essentially unity for r ≥ 750 nm, indicating a short-range order in the system. The short-range order, being characterized by a ratio of this specific distance (decay of spatial correlation) to the most probable particle separation distance, (750/285) ≈ 2.7, is smaller than the similar ratio for the cornea, (170/50) = 3.4. It was assumed that the ordering affects only each individual event of the interaction between a photon and a particle. The angular dependence of the scattered

26

Optical Clearing of Tissues and Blood

light intensity by a particle was taken in a form corresponding to the angular dependence for a fiber with a diameter equal to the modal value of the size distribution, i.e., 120 nm. The effect of the multiple scattering is included automatically in the procedure of simulation of the photon path by the Monte Carlo method, in that part of the simulations where the spatial distribution of scattering centers is assumed completely arbitrary. This approximation is valid if the dimension of the region of local ordering of scattering particles is far smaller than the mean free path of a photon in the medium, which is the case for the sclera. The results of such modeling for a scleral sample (1 mm thickness, 120 nm mean fibril diameter, and 285 nm mean separation between collagen fibrils, their refractive index, nc = 1.474; and initial refractive index of interfibrillar space, n0 = 1.345) are presented in Figs. 9–12. The collimated transmittance represents a fraction of photons leaving a tissue layer in a direction that differs from the direction of the incident radiation by no more than by ±0.5◦ , which corresponds to their entering into the aperture of a spectrometer. Total transmittance and diffuse reflectance spectra accounting for a real geometry (losing of some amount of the scattered light) of the integrating spheres used in experiments also were calculated. The calculations were performed for different values of the refractive index of the background substance, from 1.345 to 1.450, corresponding to different levels of indices matching. It is clearly seen that the model describes the major features of normal and immersed tissue spectra in the visible range. A comparison of calculated and experimental spectra (see Fig. 13) shows that refractive index matching can be considered as the main mechanism responsible for tissue-enhanced translucence. The Monte Carlo simulation technique allows one to describe the transition of the tissue scattering mode from complete diffusion to coherent (ballistic photons dominate), caused by refractive index matching. Such transition is well illustrated by the histograms in Figs. 10–12.39,237 The numbers of back- and forwardscattered photons collected by integrating spheres were calculated. These histograms show that for sclera with unmatched or slightly matched refractive indices (n0 = 1.370), there is a broad distribution of the number of scattering events (with the mean value of 25–30 collisions for NIR) that forward-traveling photons undergo; no ballistic photons (the coherent part of transmitted light) are seen. For fairly matched refractive indices, there are ballistic photons that come into being. In particular, for moderately matched refractive indices (n0 = 1.410), the unscattered and low-step scattered photons dominate in both directions—forward and back— with the mean number of collisions for the forward-traveling NIR photons of 3–4 and a rather big ballistic component. For strongly matched indices (n0 = 1.450), the ballistic component dominates and both scattering components in the forward and backward directions are small. In the NIR, the optical clearing of tissue and transformation of the scattering mode from multiple to low or even single steps begins much earlier than for visible light. A strong ballistic component formed at the tissue clearing gives perspective to coherent-domain diagnostic methods being more widely used in biomedicine.

Optical Clearing of Fibrous Tissues

27

(a)

(b)

(c) Figure 9 Monte Carlo modeling. Collimated (a) and total (b) transmittance spectra as well as diffuse reflectance (c) spectra of human sclera of 1 mm thickness calculated for different refractive indices matching with a geometry very close to the experimental one (see Fig. 13); refractive index of collagen fibrils nc = 1.47 and interfibrillar material n0 = 1.35, 1.37, 1.39, 1.41, 1.43, and 1.45.39,237 Reprinted with permission.

28

Optical Clearing of Tissues and Blood

(a)

(b)

(c) Figure 10 Monte Carlo modeling with a double-integrating sphere system. Distributions of the number Nph of forward (F) and backscattered (B) photons undertaken a definite number of collisions Ns before escaping the human scleral slab of 1 mm thickness (double-integrating sphere geometry) for slightly matched refractive indices of collagen fibrils and interfibrillar material (nc = 1.474, n0 = 1.370): (a) λ = 400 nm; (b) λ = 600 nm; (c) λ = 800 nm.39,237 Reprinted with permission.

Optical Clearing of Fibrous Tissues

29

(a)

(b)

(c) Figure 11 Monte Carlo modeling with a double-integrating sphere system. Distributions of the number Nph of forward (F) and backscattered (B) photons undertaken a definite number of collisions Ns before escaping the human scleral slab of 1 mm thickness (double-integrating sphere geometry) for partly matched (midlevel) refractive indices of collagen fibrils and interfibrillar material (nc = 1.474, n0 = 1.410): (a) λ = 400 nm; (b) λ = 600 nm; (c) λ = 800 nm.39,237 Reprinted with permission.

30

Optical Clearing of Tissues and Blood

(a)

(b)

(c) Figure 12 Monte Carlo modeling with a double-integrating sphere system. Distributions of the number Nph of forward (F) and backscattered (B) photons undertaken a definite number of collisions Ns before escaping the human scleral slab of 1 mm thickness (double-integrating sphere geometry) for strongly matched refractive indices of collagen fibrils and interfibrillar material (nc = 1.474, n0 = 1.450): (a) λ = 400 nm; (b) λ = 600 nm; (c) λ = 800 nm. 39,237 Reprinted with permission.

Optical Clearing of Fibrous Tissues

31

(a)

(b)

(c) Figure 13 Experimental spectra of the human scleral samples measured for different time intervals of administration of Trazograph-60. Spectra were measured 1 min after the sample was immersed in solution and then at 2 min intervals. The measurement time for an individual spectrum, scanning from higher to lower wavelengths, was about 85 s. (a) Collimated transmittance Tc ; sample thickness, 0.6 mm. (b) Total transmittance Tt ; sample thickness, 0.7 mm. (c) Diffusion reflection Rd ; sample thickness, 0.7 mm (heavily pigmented tissue).32,161

32

Optical Clearing of Tissues and Blood

2.1.2 In vitro measurements The total transmittance, diffuse reflectance, and collimated transmittance were measured in the 200–2,200 nm wavelength range using commercially available Varian Cary 5E, 500, or 2415 spectrophotometers with an internal integrating sphere.32,45,97,98,161 To reconstruct the absorption and reduced scattering coefficients of a tissue from such measurements, the inverse adding-doubling (IAD) method239 or inverse Monte Carlo (IMC) method was applied.84 For in vitro and especially in vivo studies of tissue optical clearing, fiberoptic grating-array spectrometers such as LESA-5, -6, and -7 (BioSpec, Russia) and PC1000, PC2000, and USB2000 (Ocean Optics Inc., USA) are suitable due to fast spectra collection in a course of immersion agent action.56,84,90,108,110,116,117,143,172,174 Typically, the spectral range of interest is from 400 to 1000 nm and the spectrometer fiber probe consists of seven optical fibers. One fiber transmits light to the object, and six fibers collect the reflected radiation. The mean distance between the irradiating and receiving fibers is about 200 µm for PC1000 and LESA-6, and about 2 mm for LESA-5. The spectrometers are calibrated using a white slab of BaSO4 with a smooth surface. Spectra were measured in vitro with human sclera samples.32,161 The sclera was carefully purified from ciliary body and retina, washed, and cut into pieces of area 10 × 10 mm. The sclera sample was placed into a cell of volume 1 ml filled with osmotic liquid or physiological solution. Three different types of chemical agents were used for scleral optical clearing in Refs. 32 and 161. The main parts of the experiments were performed using the x-ray contrast agent Trazograph (a derivative of 2-, 4-, 6-triiodobenzene acid) with molecular weight of about 500; 60% and 76% solutions in water. Some measurements were performed for two osmolytes with quite different molecular weights, such as glucose (∼180) and PEG (6000 or 20,000). At room temperature, measured by refractometer, Abbe refractive indices of some of the used agents were the following: Trazograpth-60, n = 1.437; Trazograph-76, n = 1.460; PEG (6000) solutions, n = 1.368 (0.4 g/ml), 1.394 (0.6 g/ml), 1.403 (0.8 g/ml), and 1.469 (1.0 g/ml); glucose solutions, n = 1.363 (0.2 g/ml), 1.378 (0.3 g/ml), 1.391 (0.4 g/ml), and 1.415 (0.54 g/ml) (see Table 1). For the glucose-water solutions, the refractive index nglw at any wavelength in the visible and NIR spectrums, where glucose has no strong absorption bands, can be calculated using the expression:32,52,53 nglw = nw + 1.515 × 10−6 × Cgl ,

(57)

where Cgl is the glucose concentration in mg/dl and nw is the refractive index of water59 (λ in nanometers): nw (λ) = 1.31848 +

6.662 λ[nm] − 129.2

6878 1.132 × 109 1.11 × 1014 ∼ + . = 1.3199 + 2 − λ λ4 λ6

(58)

Optical Clearing of Fibrous Tissues

33

For example, nw can be found for 820 nm from Eq. (58) as nw = 1.328; thus, nglw values for this specific wavelength can be calculated from Eq. (57) with dependence on the used concentration of glucose Cgl . It should be noted that when using proteins as immersion agents or modeling the behavior of the refractive index of tissues, blood, and their components, one may use a remarkable property of proteins, which is that equal concentrations of aqueous solutions of different proteins all have approximately the same refractive index npw .65 Moreover, the refractive index varies almost linearly with concentration Cp npw (λ) = nw (λ) + βp (λ) · Cp ,

(59)

where nw is the refractive index of water and βp is the specific refractive increment; Cp is measured in grams per 100 ml (g/dl). A similar equation can be also used for the estimation of the refractive index of glucose and other solutions at a specific wavelength, βgl = 0.00143 at 589 nm246 (see Table 1). The typical transmission spectra Tc (λ) and Tt (λ), and diffusion reflection spectra Rd (λ), measured by the integrating sphere spectrophotometer for different time intervals of Trazograph-60 administration, are presented in Fig. 13.32,161 It is easily seen that the untreated sclera is poorly transparent for visible and NIR light. Trazograph administration makes this tissue highly transparent—up to 70–75% at 600–800 nm for the sample kept in solution for 7–10 min. In addition, its spectral reflectivity decreased from 35–40% to 13–15% in this wavelength range. In general, for many of the measured samples, it can be concluded that for untreated scleral samples, transmittance was less than 1–2% in the range 400–500 nm and increased up to 6–30% for NIR wavelengths, depending on the sample thickness and pigmentation. Trazograph or other agent administration not only leads to increased transmittance, but changes the form of the spectral curve: on average, for the short wavelengths, collimated transmittance increased from 1–2% to 20% (10– 20 times), and for the long wavelengths, from 20–30% to 50–80% (∼2.5 times). For the optically cleared sclera, the collimated light makes the main contribution to transmittance. Direct measurements performed for a scleral sample of 0.75 mm thickness treated with Trazograph-60 for 40 min showed that transmittances for the detector acceptance angle of 30◦ were T30 ∼ = 35% at 400 nm and ∼ ∼ ∼ T30 = 85% at 840 nm, and Tc = 27% at 400 nm and Tc = 85% at 840 nm. It also follows from the CCD image of a laser beam transmitted through the sclera at different levels of optical clearing (see Fig. 14), that the process of formation of a ballistic group of photons (see the center of the pattern) at reduction of scattering multiplicity can be observed. These images also qualitatively support the calculated statistics of photon interactions at tissue clearing (Figs. 10–12). The efficiency of tissue clearing depends on the concentration and temperature of the solutions used. For bovine sclera at room temperature (18.5◦ C), the maximum collimated transmittance at 450 nm is in the range Tc max = 13% (Trazograph60), 22% (glucose, 45%), 39% (Trazograph-76), and 46% (PEG (6000), 80%); and

1.5863

93.1

88% (vol/vol) 71% (vol/vol)

1.454 1.430

100%

20% 50%

100%

100% 10% 25% 30% 50% 100%

Concentration 100%

1.4747

Aniline

92.1

Glycerol

1.4770 1.354 1.365 1.369 1.396 1.470

1.455 1.460

78.1

DMSO (dimethyl sulfoxide)

RI 1.4326

Glycerol + DMSO (%)

MW, Da 76.1

Immersion agent 1,2-Propylene glycol (PPG)

Mixes well with organic solvents, dissolved in water (6%), is used in production of drugs, antioxidants240–243

RA, RT

RA, RT; mixes well in any concentration with water, ethanol, methanol, acetone; absorbs moisture well from air (up to 40% in mass); glycerolwater solutions freeze at low temperatures; applied for skin softening; in cosmetics, medicine, and the food industry.240–243 “Ophthalgan” is a glycerol-based ophthalmic solution clinically used to decompress swollen corneas; contains not more than 1% water and 0.55% chlorobutanol (chloral derivative)165 WS, RA, RT; 88% glycerol is available from a common drug store

WS, RA, RT; a natural substance derived from wood pulp; has the unique capability to penetrate living tissues and to associate with water, proteins, carbohydrates, nucleic acid, ionic substances, and other constituents of living systems; possesses hygroscopic and antiinflammatory properties; FDA approved as a preservative of transplanting organs and for interstitial cystitis treatment244

Comments A dihydric alcohol, is used as a vehicle and pharmaceutical solvent in medicine, cosmetics, and the food industry; it mixes well with water and many organic solvents, possesses hygroscopic and antimicrobial properties182,240–243

Table 1 Biocompatible immersion agents: MW—molecular weight; RI—refractive index; WS—water solution; RA—refractometer Abbe (λ = 589 nm); RT—room temperature.

34 Optical Clearing of Tissues and Blood

182.2

58.5 267.8

Mannitol

Sodium chloride

Iodomethylene

Methyl salicylate

182.2 Hydrate 200

MW, Da 180.2

Sorbitol

Immersion agent Glucose

1.535–1.537

1.7425 100%

100%

50%

1.374 βNaCl = 0.00160

0.16 g/ml

25% 50%

Concentration 0.18 g/ml 0.20 g/ml 0.30 g/ml 0.40 g/ml 0.54 g/ml 7M 40% 45% 72%

1.357

1.375 1.420

βgl = 0.00143

RI 1.360 1.363 1.378 1.391 1.415 1.460 1.390 1.398 1.442

Dissolves well in alcohol and ether, but badly in water (0.07%); fragrant substance used as a perfume component, analgetic or antiinflammatory substance in medicine240–243

Is used in immersion analysis; irritant of mucosa; has a narcotic effect243

Diuretic used at injection to treat edemas; in neurosurgical practice is used as hyperosmotic agent for reduction of intracranical pressure and tissue swelling; it works as a hydroxyl radical scavenger and improves cerebral blood flow;245 WS, RA, RT117 βNaCl is the specific refractive increment at the wavelength of 589 nm [Eq. (59)]246

Carbohydrate with sweet taste, used by diabetic patients as a substitute for sugar and for intravenous nutrition, diuretic162 WS, RA, RT

βgl is the specific refractive increment at the wavelength of 589 nm [Eq. (59)]246

RA, RT104 WS, RA, RT, pH 3.5117,162 WS, RA, RT117,162

Comments WS, RA, RT117,162 In medicine, is used for intravenous nutrition and as a component of blood substitutes243

Table 1 (Continued).

Optical Clearing of Fibrous Tissues 35

1.47 1.47

PEG-based prepolymers

1.370 1.405 1.368 1.394 1.403 1.469

PPG-based prepolymers

1,550 6,000

1.459–1.467

150 – 40,000

Polyethylene glycol (PEG)

1.437 1.460

1.485

1.437 1.460

RI 1.568

1.460

∼500 ∼500

∼500 ∼500

MW, Da

IV contrast

Hypaque-60 Hypaque-76

Verografin-76

Trazograpth-60 Trazograph-76

Immersion agent Benzyl benzoate (Ascabiol)

100%

100%

25% 50% 0.4 g/ml 0.6 g/ml 0.8 g/ml 1.0 g/ml

100%

100% 100%

100%

100% 100%

Concentration 100%

Table 1 (Continued).

Hydrophilic102

Lipophilic102

Up to MW 400—liquid; dissolves well in many organic solvents and water (up to MW = 600); is used in pharmaceutics and cosmetics240–243 WS, RA, RT32,161,169

X-ray contrast agent, Ioxaglate Meglumine and Ioxaglate Sodium (Hexabrix, Mallinckrodt Inc.), OCT, group refractive index, λ = 1300 nm111

X-ray contrast agent, Diatrizoate Meglumine (Princeton, Nycomed Inc.)36,38

RA, RT; x-ray contrast agent162–164

Comments Oily aromatic liquid used in dermatology as a lotion at topical application for scabies and pediculosis treatment240–243 RA, RT; x-ray contrast agent, a derivative of 2,4,6triiodobenzene acid32,161,169

36 Optical Clearing of Tissues and Blood

Oils: Paraffin Coconut Olive Sesame Sunflower Turpentine Cotton Poppy Almond Beechnut Nut Castor Linseed Sandal Cedar

Methylcellulose (Celevac)

Immersion agent Dextrans

13,000–140,000

473,000

65,500

MW, Da 10,500

1.440 1.455 1.460–1.469 1.468 1.470 1.470 1.4730 1.463–1.4742 1.475 1.475 1.4795 1.478–1.480 1.485 1.508 1.516

1.337

RI 1.345(0.001) 1.353(0.002) 1.345(0.001) 1.353(0.002) 1.339(0.001) 1.341(0.001) 1.346(0.002) 1.52 10%

Concentration 5 g/dl 10 g/dl 5 g/dl 10 g/dl 0.5 g/dl 2 g/dl 5 g/dl 0.25 g/ml

Table 1 (Continued).

RA, RT76,240,241

Absorbs water well; used in medicine for internal application as a purgative remedy; usually has no any side effects240–243 WS, RA, RT

OCT, group refractive index, λ = 1300 nm111

Comments Dextrans of MW 40,000–80,000 are used as blood plasma substitutes243 OCT, group refractive index, λ = 820 nm112

Optical Clearing of Fibrous Tissues 37

1.343

1.345

1.360 1.342 βp = 0.00179 0.00183 0.00181 0.00181 0.00181 0.00170(1) 0.00179 0.00184 0.00187

Ovalbumin + PBS

Bovine serum albumin (BSA) + PBS

Native white-egg

5%-agar-agar

Proteins: Total serum (human) Euglobulin Pseudoglobulin Total albumin Recrystallized albumin Lipoprotein Hemocyanin Helix Octopus Carcinus

RI 1.532–1.538 1.550–1.560 1.601 1.332

MW, Da

Phosphate buffer saline (PBS)

Immersion agent Cloves Anise Cinnamon

Concentration

Table 1 (Continued).

βp is the specific refractive increment at the wavelength of 589 nm [Eq. (59)]; RA, RT65

RA, RT76

BSA was added to PBS, mixed and then filtered through 0.4 and 0.2 µm filters to remove undissolved protein; the principle of minimum deviation was used to measure refractive index208

Ovalbumin was added to PBS, mixed, and then filtered through 0.4 and 0.2 µm filters to remove undissolved protein; the principle of minimum deviation was used to measure refractive index208

pH 7.4208

Comments

38 Optical Clearing of Tissues and Blood

Human α2 -globulin Human β2 -globulin Hemoglobin

Human γ-globulin

Egg albumin

Bovine serum albumin

Horse serum albumin

βp = 0.00181 0.001862 0.001833 0.001830 0.001912 0.001844 0.001924 0.001854 0.001901 0.001883 0.001820 0.001876 0.001875 0.00188 0.00183 0.00185 0.001988

Human serum albumin

RI 0.001813 0.001945 0.00178 0.00193(5) βp = 0.0014(5) 0.0014(5) 0.0016(20)

MW, Da

Carbohydrates Lipoids Nucleic acid compounds

Immersion agent Egg albumin Sheep CO hemoglobin Ox globin Ox CO hemoglobin

Concentration

Data from Ref. 76; White light λ = 578.0 nm λ = 589.3 nm; crystal White light λ = 436.0 nm λ = 546.0 nm λ = 436.0 nm λ = 546.0 nm λ = 578.0 nm λ = 436.0 nm λ = 546.0 nm λ = 578.0 nm λ = 578.0 nm λ = 589.3 nm λ = 589.3 nm λ = 589.3 nm

Comments

Table 1 (Continued).

Optical Clearing of Fibrous Tissues 39

1.337 1.339

1.340

Ultrasound gel

Aloe vera gel (aloe vera, triethanolamine, D-alpha tocopherol (natural vitamin E), carbomer, tetrasodium EDTA, methylparaben, imidazolidinyl urea)

Aloe vera (skin gel) (99% aloe vera gel, pure water, propylene glycol, carbomer, triethanolamine, DMDM hydantoin, diazolidinyl urea)

1.360–1.420
n = 0.002–0.005

Water-glycerol gelatinous gels

RI 0.001892 0.001818 0.001890 0.001822 0.00188 1.360–1.420
n = 0.002–0.005

MW, Da

Water solutions of albumin

Fibrinogen

β-lactoglobulin

Immersion agent Lactoglobulin

100%

100%

100%

Concentration

Skin moisturizer; provides temporary relief of minor skin irritations such as sunburns and dry, chapped skin; RA, RT, commercially available

Forms a protective barrier that helps retain moisture and promotes healing; nonoily moisturizer provides effective relief from sunburns, minor burns, skin irritation, chafing, itching, and dry skin; RA, RT, commercially available

RA, RT, commercially available

Gel kits are prepared by dilution of concentrated salt-free and purified gelatinous gels with turbidity of 0.5 × 10−3 cm−1 in 0.2% sterile solution of glycerol in 0.5% solution of NaCl with pH 7.0–7.2; other compositions with 1% glycerol, 1% glycerol, and 0.5% glucose or 10% sasharose also can be used; RA, RT76

Fifth fraction of bovine or human serum albumin; contains total mass of serum albumin and about 3% of α-globulin and less than 0.5% of β-globulin; RA, RT76

Comments λ = 436.0 nm λ = 546.0 nm λ = 436.0 nm λ = 589.3 nm

Table 1 (Continued).

40 Optical Clearing of Tissues and Blood

Immersion agent Clear cold cream (water, butylene glycol, cyclomethhicone, PEG-32, glycerol, diglicerol, isopentyldiol, PEG-6, hydrogenated polyisobutene, sodium chloride, decyl polyglucose, dimethiconol, isododecane, panthenol, cetyllactate, witch hazel distillate, tocopheryl acetate, glycolic acid, dimethicone copolyol, tributyl citrate, fragrance, phenoxyethanol, benzylalcohol, PG, methylparaben, propylparaben, ethylparaben, butylparaben, butylmethoxydibenzoylmethhane)

MW, Da

RI 1.403

Concentration 100%

Table 1 (Continued).

Comments Oil-free formula for skin cleaning; use in conjunction with skin moisturizers; RA, RT, commercially available

Optical Clearing of Fibrous Tissues 41

42

Optical Clearing of Tissues and Blood

Figure 14 Specklegrams recorded at two different moments of time in the far-field zone for a sample of translucent human eye sclera irradiated with a focused beam of a He:Ne laser (633 nm). Enhanced translucence was provided by Trazograph-60 in a cuvette during (a) 2.5 and (b) 10 min.85

at 700 nm, Tc max = 73% (glucose, 45%), 76% (Trazograph-60), 99% (Trazograph76), and PEGs (6000 and 20,000), 80%).169 The maximal transmittance is achieved at 15–30 min. At physiological temperature, this time interval is considerably shortened. For example, for a PEG 20,000 solution (80%), the time interval for maximal tissue clearing changed from 27 min at 18.5◦ C to 12 min at 38◦ C. The time-dependent collimated transmittance of a scleral sample measured at 633 nm concurrently with Trazograph-60 administration is presented in Fig. 15. It shows the dynamics of tissue clearing. Similar characteristics were measured for glucose and PEG administration. The registration of the dynamic response of the intensity transmission can be used to estimate diffusion coefficients of the interacting fluids: water and agent (Trazograph, glucose, glycerol, PEG, etc.). Based on the theoretical background given earlier, we can estimate the coefficient of diffusion of the agent assuming that water and agent have the same paths for diffusion. The fol-

Optical Clearing of Fibrous Tissues

43

lowing set of equations gives the simple algorithm for diffusion coefficient determination: Eqs. (27), (31), (33), (40), and (43). More sophisticated algorithms accounting for tissue swelling and shrinkage, and appropriate to the measuring procedure of inverse optical problem solving (IAD, IMC), are also available.84,116,117,155,162

Figure 15 Time-dependent collimated transmittance (dots) of a 0.5 mm thick scleral sample measured at 633 nm concurrently with administration of Trazograph-60.32 The solid and dashed lines represent the mean value and the upper and lower limits of Tc obtained by calculation [Eqs. (27), (30), (37), and (40)] of Tc using experimental data: DT = (1.46 ± 0.19) × 10−5 cm2 /s (see Table 2, sample # 1).

Data for diffusion coefficients’ values of different samples of human sclera are collected in Table 2. It is easily seen that estimated values of DT , calculated using about 30 magnitudes of Tc measured for different time intervals for each sample, have quite reasonable rms errors and differences in mean values from sample to sample. As can be seen from Fig. 15, rms values include the low-frequency oscillations of Tc (t), which can be caused by spatial-temporal fluctuations of the agent diffusivity on interacting with tissue structure. On average, the DT values are not far from the values of Da for diffusion of low-weight molecules in water.233 It should be noted that for the osmolytes, fluid transport within tissue is more complicated because there are at least two interacting fluxes, so the model for describing these processes should be more complicated and should include monitoring of additional measurement parameters, such as the refractive index of the chemical agent, tissue weight and/or thickness, and osmotic pressure in a process of tissue clearing. Such monitoring of the refractive index of Trazograph-60 in the bath during a process of scleral clearing gave 1.4370 (t = 0), 1.4321 (12 min), 1.4222 (20 min), and 1.4025 (40 min). Measurements of tissue samples’ weights before and after administration of the agents gave the following values: Trazograph-60 (sample 5 × 8 × 0.6 mm3 ), 54 mg (t = 0) and 51 mg (34 min); glucose (40%, pH 3.5) (sample 10 × 11 × 0.5 mm3 ), 82 mg (t = 0) and 66 mg (20 min); PEG (6000) (1 g/ml) (sample 8 × 10 × 0.5 mm3 ), 65 mg (t = 0) and 48 mg (60 min). Thus, the relative decrease of the sample weight is 5.5% for Trazograph-60, 15.5% for 40% glucose, and 28% for PEG (6000). Both experiments with refractive index and weighting show differences in osmotic properties of the used agents and

3.5 3.5 – 3.5

– 7.2

40% 40% 72% 40%

0.72 g/ml 100% 100%

Propylene glycol

Trazograph-60

–

6.0 3.5 3.5

0.3 g/ml 0.4 g/ml 40%

pH 6.3

Glucose

Human sclera

Concentration 0.18 g/ml

Immersion liquid

Tissue (0.93 ± 0.15) × 10−6

(2.39 ± 0.58) × 10−6 (2.47 ± 0.08) × 10−6 (5.61 ± 0.75) × 10−6 (9.19 ± 2.67) × 10−6 (2.06 ± 0.37) × 10−6 (2.44 ± 0.86) × 10−6 (2.76 ± 0.10) × 10−6 (1.85 ± 0.13) × 10−6

(0.57 ± 0.09) × 10−6

(1.47 ± 0.36) × 10−6 (1.52 ± 0.05) × 10−6 (3.45 ± 0.46) × 10−6 (5.62 ± 1.64) × 10−6 (1.29 ± 0.23) × 10−6 (1.50 ± 0.53) × 10−6 (1.7 ± 0.06) × 10−6 (1.14 ± 0.08) × 10−6 (2.05 ± 0.10) × 10−8

#2 (0.40 ± 0.11) × 10−5 #3 (0.53 ± 0.08) × 10−5 #4 (1.20 ± 0.10) × 10−5

#1 (1.46 ± 0.19) × 10−5

Da , cm2 /sec, mean ± sd, at 37◦ C

Da , cm2 /sec, mean ± sd, at 20◦ C

Table 2 Diffusion coefficients of chemical agents in tissues.

d = 1.00 ± 0.01 mm, fixed tissue174 d = 0.5 mm (#1 and #2) and 0.4 mm (#3 and #4)32

d1 = 1.10 ± 0.01 mm, d2 = 0.52±0.01 mm, extrapolated to 37◦ C, Ref. 174

d = 0.51, see Fig. 20; extrapolated to 37◦ C, Ref. 162 d = 0.5; extrapolated to 37◦ C, Ref. 173 d = 0.5; extrapolated to 37◦ C, Ref. 84

Sample thickness, d = 0.50 ± 0.01; 0.51 ± 0.02, and 0.48 ± 0.02 mm, respectively; extrapolated to 37◦ C, Ref. 116

Comments

44 Optical Clearing of Tissues and Blood

Glycerol Glucose Glucose

Human skin (in vivo)

40%

Glucose

Rat skin (in vivo)

88%

Glycerol

Rat skin

40%

88% 40%

0.16 g/ml 0.2 g/ml

Mannitol Glucose

Human dura mater

–

80%

PEG (6000) D2 O

45%

100%

Concentration

Glucose

Verografin-76

Immersion liquid

Human eye lens

Bovine sclera

Tissue

3.5

6.32 3.5

3.5

6.32

6.05 5.99

–

–

–

7.2

pH

–

Ref. 180 Ref. 179 Ref. 90

(2.56 ± 0.13) × 10−6

Extrapolated to 37◦ C, averaged for six samples, both skin sides, epidermal and dermal, were interacted with the agent172 d1 = 0.73 mm skin with fat layer; d2 = 0.70 mm and d3 = 0.57 mm skin with removed fat 61,179

Extrapolated to 37◦ C, Ref. 117

Confocal Raman spectroscopy, samples #1–4 intact, and #5 fixed155

Average for 5 samples; extrapolated to 37◦ C, Ref. 237 Average for 4 samples237

d = 0.49 ± 0.01 mm, fixed tissue174

Comments

(1.16 ± 0.03) × 10−6 (1.10 ± 0.15) × 10−6

(8.31 ± 3.70) × 10−7

(5.12 ± 2.27) × 10−7

(1.9 ± 0.1) × 10−6 (3.1 ± 0.1) × 10−6 (3.3 ± 0.7) × 10−6

(2.08 ± 0.65) × 10−6 (2.59 ± 0.46) × 10−6

(3.54 ± 2.44) × 10−5

Da , cm2 /sec, mean ± sd, at 37◦ C

#1 3.88 × 10−6 #2 0.90 × 10−6 #3 3.57 × 10−6 #4 3.53 × 10−6 #5 0.75 × 10−6 (1.31 ± 0.41) × 10−6 (1.63 ± 0.29) × 10−6

(1.94 ± 1.52) × 10−5

(2.18 ± 1.50) × 10−5

(0.32 ± 0.01) × 10−6

Da , cm2 /sec, mean ± sd, at 20◦ C

Table 2 (Continued). Optical Clearing of Fibrous Tissues 45

46

Optical Clearing of Tissues and Blood

their tissue dehydration abilities, which are in the range: low (Trazograph-60), midlevel (40% glucose), and high [PEG (6000)]. It follows that at optical clearing of the sclera by Trazograth-60 the process of replacement of the interfibrillar fluid by the agent dominates. The rather big reduction in the refractive index in the bath with the sample means that water effectively escapes tissue, and the small loss of the sample weight says that tissue water is replaced by Trazograph-60. Thus, we may assume that in the system there are two approximately equal fluxes moving through a tissue layer: the water flux JW directed out of a layer, and a dissolved matter JS directed into a layer, which are proportional to the gradients of the chemical potential of water and the dissolved matter [see Eq. (45)].233 For glucose and especially for PEG, dehydration plays an important role due to the inequality of the two fluxes: the water flux JW out of a tissue layer is stronger than a dissolved matter JS flux into a layer. Thus, structural changes of collagen fibrils and interfibrillar spacings caused by tissue dehydration and described in Section 1.4235 should be accounted for in the tissue clearing model based on tissue dehydration. The interaction of an osmotically active agent penetrated inside a tissue with collagen fibrils may be responsible for quasi-periodic low-frequency (3–4 min of period) oscillations of the light transmittance that are well seen in Fig. 15. The oscillating character of the tissue response may be explained as a multistep origin of fluid diffusion.32,85 The first-step osmolyte penetration into the tissue leads to refractive index matching of the interstitial fluid and hydrated fibril collagen—the significant translucence of tissue growth. The second step is characterized by the interaction of osmolytes, contained within the renovated interfibrillar liquid, with fibril collagen, which leads to collagen dehydration and the consequent growth of its refractive index that slightly breaks down optical matching and causes a slight decrease of transmittance. The subsequent imbalance of water-osmolyte concentrations leads in turn to penetration of an additional amount of osmolyte into the sample, which causes reestablishment of the refractive index matching and a corresponding light transmittance—this is the origin of the third step. The inertia of each of the considered processes may cause the establishment of quasi-periodic oscillations, with the periods and amplitudes depending on the parameters of the nonlinear system. Measured values of osmotic pressure for Trazograph-60 were equal to 4.3 MPa and 7.1 MPa for Trazograph-76.32 For untreated sclera, the value of osmotic pressure was equal to 0.74 MPa, and it increased after administration of Trazograpth-60 for 30 min—up to 5.02 MPa. On one hand, the osmotic pressure causes the flows’ generation and their intensities [see Section 1.3 and Eqs. (45) and (46)], but on the other hand, rather strong osmotic pressure may destroy tissue structure. A direct histological study showed that there are no serious irreversible changes in the cellular and fibrous structure of human sclera for a rather long period of osmolyte administration.170 For example, for Trazograph-60, this time is at least about 30 min, and rather minor changes of tissue structure, which are characterized by a moderate swelling of tissue components, were seen. The reversibility of tissue structure change with an osmolyte administration is also demonstrated by the data in Fig. 16,85,171 which show that the multiple-single

Optical Clearing of Fibrous Tissues

47

scattering transition (i.e., optical translucence, improvement of linear polarization) is reversible when the osmolyte bath is replaced by a physiological solution, and vice versa when the osmolyte is administered again.

Figure 16 Time dependencies for the values of mean speckle intensity Is  and its polarization components I⊥  and I||  measured in the paraxial region of a sample of human sclera with a thickness of 0.4 mm and averaged over the scanning trajectory (1.5 mm): 1, 2, 3 sequential measurements with a sample that was first placed in Trazograph-60, then in a physiological solution (0.9% NaCl), and then again in Trazograph-60; λ = 633 nm.85

The theoretical and experimental results show that administration of osmolytes to the sclera affects the refractive index matching of the collagen fibrils and interfibrillar fluid, leading to dramatic changes (reductions) in the scattering properties of the sclera. For different osmolytes, refractive index matching can be implemented in different ways: (i) water can leave the interfibrillar space and exit the sample (a dehydration); (ii) the administered fluid can enter the tissue and fill up the interfibrillar space, and water can partly leave the interfibrillar space and partly exit the sample. The first mechanism is characteristic only for ideal osmolytes, the second, for nonideal ones. For fibrous tissue similar to sclera, all tested chemical agents can be considered as nonideal osmolytes, because molecule sizes of these agents were much less than the mean cross section of the interfibrillar space, which is about 185 nm, when the diameter of the biggest molecule of PEG (20,000) should be less than 5 nm. Indeed, the structure of the interfibrillar matrix and the molecular structural properties may also have their influence on diffusion; thus, the diffusivity of foreign molecules and corresponding strength of water flux is different for different agents used. A stronger water flux in the sclera was induced by PEGs, a midlevel flux by 40% glucose, and a small flux by Trazograph-60 (see weight measurements).32 The dynamics of tissue optical clearing using osmolytes is defined by a characteristic time response of about 3 to 10 min. This is in good agreement with results obtained by Rol,35 but he used a pointwise mechanical stress or local heating

48

Optical Clearing of Tissues and Blood

induced by a laser beam. Actually, as follows from Eq. (45), osmotic pressure and hydrostatic pressure caused, for example, by mechanical stress, have the same mechanism for inducing the fluid flux, and the time response is defined by water diffusion through the interfibrillar space. So, optical clearing using mechanical stress should be somewhat equivalent to the action of an ideal osmolyte, because a local stress picks up water from the compressed site and diminishes the tissue layer thickness. In practice, optical clearing with osmolytes may be more preferable to apply than a mechanical stress, because there are more possibilities of controlling the time response and efficiency using various chemical agents; in addition, the function of these agents may be combined (tissue clearing and treatment). These results are general and can be used to describe many other fibrous tissues. It should be noted that human sclera can be considered as a living scattering etalon in the visible range, like a white glass (see diffuse reflectance spectra in Fig. 13). For example, due to the simpler structure and stable and controllable parameters of sclera in comparison with skin, light scattering technologies of glucose monitoring designed for skin measurements52–54,156–159 may be more effective in application to sclera. In this connection, it is interesting to analyze a change in the sclera color during its clarification.34,39 The quantitative estimation of this change from transmission and reflection spectra in Fig. 13 was done by calculating the chromaticity coordinates for the MKO 1964 color system. From the calculated color triangles, it follows that the native sclera has a reddish tint in the transmitted light; however, this does not appreciably change the visual perception because of a very low transmission coefficient. During sclera clarification, its color becomes whiter. In the diffuse reflectance, the native sclera is white, as is visually observed. Upon clarification, the sclera color in the reflected light slightly shifts from white to bluish. To more precisely study time-dependent transmittance and reflectance of a tissue, a fiber-optic photodiode array or CCD spectrometer providing a fast collection of spectra should be used. This is especially important for diffusion coefficient determination in in vitro studies and in in vivo monitoring of tissue clarification. Using a fiber-optic photodiode array spectrometer shown in Fig. 17, detailed in vitro measurements for human sclera were done on tissue impregnation by various solutions such as glucose, Trazograph, Verografin, and propylene glycol, which do not have strong absorbing bands within the spectral range of interest, 400–800 nm.116 In in vitro study, the conjunctiva and ciliary body, as well as the retina with choroid were removed. The mean thickness of samples was about 0.5 mm. They were fixed on a plastic plate with a square aperture of 5 × 5 mm2 (effective impregnation by a chemical agent via both surfaces of the sample was provided) and placed in a 5 ml cuvette filled with the solution under study. To understand the mechanisms of scleral tissue optical clearing, the collimated transmittance spectra and change of the scleral sample weight were measured concurrently with the administration of glucose solution. Figures 18, 19, and 20 illustrate the dynamics of transmittance spectra and typical weight change. It is easily seen that the untreated sclera is a poorly transparent media for visible light. Glucose administration makes this tissue highly transparent. As it follows from Fig. 19, the characteristic time response of sclera optical clearing is about 5 min.

Optical Clearing of Fibrous Tissues

49

Figure 17 Experimental setup for in vitro and in vivo measurements of the collimated light transmittance and the reflectance spectra.84 In vitro measurements: 1, optical fiber; 2, aluminum jacket; 3, neutral filters; 4, sclera sample; 5, osmotically active solution; 6, 0.5 mm pinhole; 7, cuvette.

Figure 18 Time-dependent collimated transmittance spectra of the human sclera sample impregnated by a 40% glucose: (1) 10 s; (2) 1 min; (3) 2 min; (4) 3 min; (5) 4 min; (6) 5 min; (7) 6.5 min; and (8) 8.5 min after the scleral sample was immersed in a 40% glucose.84

Based on these measurements and accounting for the fact that commercially available 40% glucose from the drugstore, which has a low pH (3.5), was used, the following model of action of osmotically active liquid on a fibrous tissue seems to be adequate. At the first stage, which takes place approximately 5 minutes after a sample is placed in glucose solution, substantial optical clearing was accompa-

50

Optical Clearing of Tissues and Blood

Figure 19 Time-dependent collimated transmittance of the human sclera sample measured at 420 nm (squares); 589 nm (circles); 630 nm (up triangles); and 700 nm (down triangles) concurrently with administration of 40% glucose.84

Figure 20 Calculated time-dependent thickness of the human sclera sample (Monte Carlo simulation as the best fit to the experimental data for the collimated transmittance shown in Figs. 18 and 19).61 Reprinted with permission.

nied by a sample thickness decrease. Thus, we may suppose that optical clearing occurs due to two main mechanisms: (i) refractive index matching between collagen fibers and penetrated glucose, and (ii) glucose osmotic action that dehydrates tissue resulting in up to a 25% decrease of thickness. In the late stage of glucose administration, to the seventh minute, the optical clearing process saturates, due to equilibration of fluid (glucose, water, proteins, salts, etc.) concentrations in the system, and the thickness increases somewhat. From the seventth to fifteenth minute, the inclusion of the thickness change (increase to its initial thickness) in the optical clearing is well seen on the background of the inclusion of the saturated molecular fluxes—collimated transmittance goes down a little bit, but is still very high.

Optical Clearing of Fibrous Tissues

51

The further tissue swelling with time of up to 20% of the initial thickness to the twenty-first minute does not seriously influence tissue transmittance. It is important that in spite of complex behavior of tissue thickness with the administration of this specific chemical agent (40% glucose with pH 3.5), thickness variations do not strongly affect optical clearing. Such nonmonotonous behavior of tissue thickness (first shrinkage and later swelling) can be explained using results of Ref. 235, where it was experimentally shown that for bovine sclera, hydration (swelling) may change from H = 2.48 for pH 7, close to physiological, to H = 7.15 for pH 3. In our case, this means that at the first stage when tissue pH, which is close to the physiological one, is not seriously affected by glucose (a small amount is penetrated into the sclera), the dehydration of tissue dominates due to the osmotic action of the glucose: but in the late stages of glucose administration, due to a large amount of glucose penetrated in and in the bath the pH of the whole system—the tissue/glucose bath is going down and swelling takes place. It should be stressed again that the discussed effects of tissue shrinkage and swelling are important but not dominant in glucose action; thus, the experimental data for the collimated transmittance (Figs. 18 and 19) and the time-dependent measurements of tissue sample thickness changes under immersion solution action (Fig. 20) can be used to estimate the glucose diffusion coefficient in sclera.37,61 The detailed model of glucose transport in fibrous tissue is described in Ref. 117. Equations (27), (33), (40), and (43) are the basis for this model, which can be used for reconstruction of the diffusion constant. The estimated average value of the diffusion coefficient of 40% glucose transport in the scleral sample is DG = (3.45 ± 0.46) × 10−6 cm2 s−1 at temperature 20◦ C (see Table 2). This value is not far from the values of Da for diffusion of low-weight molecules (such as sucrose, glucose, etc.) in water at zero concentration (3.6–5.2) ×10−6 cm2 s−1 at 12 − 15◦ C.233,240,241 When agents with a strong osmotic strength are used, the diffusion coefficient should be close to that of water diffusion in a tissue, because this is the main flux in the system. With dependence on tissue structure, this value should be equal or above the value of the diffusion coefficient of water in water, DW = 2.5 × 10−5 cm2 s−1 . The diffusion coefficient is a function of dimension and form of the diffusing molecule:233 Da = const × M −S .

(60)

For small molecules, S = 1/2, and for spherical molecules diffusing in water (large proteins), S = 1/3. In general, the parameter S for diffusion in water is in the range 0.3–0.5, and for diffusion through a biological membrane, it is about 3.5. For example, changing the molecule’s weight M from 45 to 122 with diffusion in water changes the diffusion coefficient from 1.6 × 10−5 cm2 s−1 to 0.8 × 10−5 cm2 s−1 (twofold), and for the same molecule’s diffusion through a plasmatic membrane, from 1.4 × 10−8 cm2 s−1 to 2.0 × 10−10 cm2 s−1 (70-fold).233

52

Optical Clearing of Tissues and Blood

2.2 In vitro Frequency-Domain Measurements of Eye Sclera The dynamic response of optical properties (modulation depth and phase shift of intensity modulation of the backscattered light) of human eye sclera in respect to the interval of a chemical agent (solution, gel or oil) administration can be measured using a photon-density wave (frequency-domain) technique.86 When the intensity of the light source is modulated at a frequency ω, a photon-density wave is induced in a scattering medium:1,3,4,7,29 A(r) = Adc + Aac exp[−i(ωt − θ)],

(61)

where Adc , Aac , and (ωt − θ) are the dc and ac components of the amplitude of the photon-density wave and its phase, respectively. Photon-diffusion theory provides independent determination of the absorption and reduced scattering coefficients from the measurements at a single modulation frequency. The expressions for the measured quantities of the phase delay θ and ac amplitude Aac have been presented elsewhere.29 These expressions depend on the source-detector separation rsd , reduced scattering coefficient µ s , and absorption coefficients µa . The intensity and phase of photon-density waves generated by an NIR optical source were measured at several source-detector separations. The light source was a laser diode with a wavelength 786 nm and 4 mW power at the end of coupled multimode fiber (core diameter 62.5 µm).86 The intensity modulation depth of approximately 80% at a frequency of 140 MHz was provided by modulation of the injection current of the laser diode. The experimental setup was designed at the University of Pennsylvania (Fig. 21). A multifiber detection system with small source-detector separations together with a Dicon multichannel fiber-optic switcher has been used for immersion experiments on human sclera in situ for a whole eyeball.86 Data shown in Fig. 22 are for the temporal changes of ac amplitude during Trazograph-60 administration for three different source-detector separations and two different techniques of immersion solution administration—by injection and by drops. The clearing of scleral tissue was observed during the first 3 min of Trazograph-60 administration by injection. For small source-detector separations (about 1–2 mm) and a relatively large one (3.5 mm), the temporal dependencies are quite different. Keeping in mind that at the first 3 min after injection of the chemical agent, the positive time scale corresponds to a decrease of scattering due to tissue immersion, the opposite tendencies of considered dependencies can be understood as the following. For the small source-detector separation close to the exact back-reflectance geometry, the intensity of reflected light decreases along with the scattering decrease; and for rather large separations, when lateral photon diffusion effects are important, intensity at first goes up with decreased scattering, but if scattering continues to decrease, intensity will lessen. That is why a local maximum on a curve for a separation of 3.5 mm was observed. At the third minute after chemical agent injection, due to its diffusion into neighboring tissue regions, amplitudes for all separations have a tendency to go to their initial values. Another

Optical Clearing of Fibrous Tissues

53

Figure 21 The homodyne frequency domain I&Q measuring system:86 intensity modulated at 140 MHz diode laser, 786 nm, eliminates the sample via fiber, the scattered light is collected by the fiber or the multifiber probe and detected by the APD (Hammatsu, C5331-3); after filtration, such signals as Idc and Qdc , are registered; DC, I , and Q components are digitized by the A/D board; the phase and amplitude caused by light transport in a tissue 2 }1/2 . are found from the equations:
 = tan−1 (Qdc /Idc ), A = {Q2dc + Idc

(a)

(b)

Figure 22 Frequency-domain measurements. The time-dependent changes in amplitude of the optical signal from a human eyeball in situ after (a) Trazograph-60 injection and (b) Trazograph-60 drops in the vicinity of detector fiber tip.86

technique of chemical agent administration by drops shows the same tendencies for small and large separations as for injection, but essential changes of the amplitudes happen momentarily after chemical agent drops are applied, and then the

54

Optical Clearing of Tissues and Blood

amplitudes slowly change in opposite directions. Such behavior depends on the specific features of a chemical agent application, which are: (i) superficial impregnation (underlines the importance of the surface immersion effect) and (ii) continuous renovation of the chemical agent on the tissue surface (many drops during the measurement interval). This study, which was performed under circumstances that are very close to in vivo measurements, also shows that the impregnation of the eye sclera by osmotically active chemical agents affects the reversible refractive indices matching of the collagen fibrils and interstitial media that leads to dramatic reduction of the tissue scattering ability—up to 60% in ac signal change for 10–12 min with Trazograph60 application.

Figure 23 In vivo time-dependent reflectance spectra of rabbit eye sclera measured concurrently with administration of 40% glucose solution: (1) 1 min; (2) 4 min; (3) 21 min; (4) 25 min; and (5) 30 min after drop of glucose into the rabbit eye.84

2.3 In vivo Measurements of Eye Sclera In vivo measurements were done for a rabbit eye using the experimental setup presented in Fig. 17. Experimental spectra and dynamic response on selected wavelengths are shown in Figs. 23 and 24. The surface temperature of the rabbit eye was ∼38◦ C. 40% glucose was used as a chemical agent for the scleral optical clearing, administered in the form of eye drops. A significant decrease of the reflectance during the first 5 min of glucose administration is seen. Dips appearing at 420, 530, and 570 nm are caused by blood perfusion. The lower reflectance at 420 nm is caused by the strong absorption of blood. Evidently, faster decay at this wavelength reflects blood perfusion dynamics due to eye conjunctiva and sclera inflammation caused by light irradiation and osmotic action of the glucose. Because blood has less influence at 630 to 700 nm, the measured dynamic responses can be used for in vivo estimation of the diffusion coefficient for glucose in scleral tissue.

Optical Clearing of Fibrous Tissues

55

Figure 24 In vivo time-dependent reflectance of rabbit eye sclera measured concurrently with administration of 40% glucose at 420 nm (down triangles), 630 nm (squares), and 700 nm (up triangles).84

From experimental data presented in Figs. 23 and 24, one can see that for in vivo measurements reflectance decreases up to twofold; such a value of decrease is comparable with in vitro studies for Trazograph-60 immersion [see Fig. 13(c)]. Transmittance measurements are difficult to provide in in vivo experiments; thus, to estimate the translucent efficiency at optical immersion, we may use the results of in vitro measurements of collimated transmittance that show up to a 26-fold increase in transmittance. Using Monte Carlo modeling based on experimental data and the arrangement used for in vivo and in vitro measurements, a correct comparison of in vivo and in vitro clearing efficiency can be done.173 Results of such a comparison are summarized in Table 3. The calculated ratios of maximal (untreated tissue) to minimal (well-treated tissue) diffuse reflectance Rd for in vitro and in vivo measurements show the same tendency of their change with the wavelength. Differences in absolute values of this ratio, which are higher for the in vivo case, can be explained by a multilayered structure of the living tissue consisting of the conjunctiva, the Tenon’s capsule, the sclera itself, the ciliary muscle, and the ciliary pigmented epithelium (singly or in combination)—some of which are extremely absorbing. The living tissue seems to be more effectively controlled by an immersion phenomenon due to the stronger influence of the absorbing layers, which reduces the fluence rate of the backscattered photons as the light penetrates more deeply inside the tissue (due to reduction of scattering coefficient), where absorption is maximal for this specific tissue. Less scattering causes shorter photon migration paths and less probability for photons to be absorbed. Other reasons for more effective control are the blood perfusion and metabolic activity of leaving tissue, which cause more effective impregnation of tissue at physiological temperatures, in spite of the fact that the agent was applied only to the exterior surface of the sclera. In general, the rate of agent diffusion in a tissue (see Table 2) increases from the minimal rate for the fixed tissue samples, when

56

Optical Clearing of Tissues and Blood

additional bonds between the protein molecules hindering the agent transport are formed, to the midlevel for the fresh tissue samples, and the highest for in vivo immersion. It is also important that the total transmittance for in vivo case is three- to sixfold more effectively controlled by tissue immersion than that for separated scleral samples measured in vitro. The total transmittance of the anterior eye layers measured at the posterior interface of the sclera determines the laser energy applied to the ciliary body, when its coagulation is needed. The collimated transmittance, in turn, determines the efficiency of laser irradiation through the sclera at some local area of the eye bottom to destroy a tumor, for example.

Figure 25 Time-dependent light absorption fractions for different layers of a rabbit eye at 700 nm calculated using Monte Carlo simulation for tissue impregnated by 40% glucose. Squares correspond to conjunctiva, circles to sclera, up triangles to ciliary muscle, and down triangles to ciliary pigment epithelium.173 Reprinted with permission.

The time-dependent light absorption fractions for different layers of the rabbit eye at 700 nm calculated using Monte Carlo simulation for tissue impregnated by a 40% glucose is shown in Fig. 25. The initial values of the scattering and absorption coefficients for various tissue layers were taken from Refs. 41 and 42. The time-dependent diffuse reflectance, the total transmittance, and the light absorbed fractions at tissue immersion were calculated using the in vivo studies discussed above. From the graphs of Fig. 25, it follows that due to a significant translucence of the upper layers of the rabbit eye, the lower-absorbing layers of the eye membrane, such as ciliary body components, are well irradiated, and thus absorb light

420 630 700

Wavelength, nm

Experimental values Rmax /Rmin Tc max /Tc min in vivo in vitro 2.09 14.6 1.44 25.8 1.38 17.8

Monte Carlo modeling Rd max /Rd min Rd max /Rd min in vivo in vitro 1.66 1.13 1.87 1.18 1.99 1.22

Tt max /Tt min in vivo 6.35 3.28 3.03

Tt max /Tt min in vitro 1.12 1.31 1.10

Table 3 Measured and calculated ratios of maximal and minimal values of optical parameters of human (in vitro) and rabbit (in vivo) sclera at administration of 40% glucose.173 Experimental values correspond to data in Figs. 19 and 23.

Optical Clearing of Fibrous Tissues 57

58

Optical Clearing of Tissues and Blood

well. It is found that as much as the light absorption fraction in the conjunctiva and the sclera is decreased, in the ciliary body, it is considerably increased. This confirms the declared possibility of using osmotically active substances to enhance efficiency of transscleral selective photodestruction of the ciliary body.21,35,36

2.4 Dura Mater Immersion and Agent Diffusion Rate Optical clearing of human dura mater is important for cerebral optical diagnostics, phototherapy, and laser surgery. Dura mater is a typical fibrous tissue and demonstrates the same behavior of optical clearing as eye sclera, cornea, or skin dermis and muscle, but has its own diffusion coefficient, and characteristic time and degree of clearing, defined by its structure. The first results from in vitro experimental study of human and rabbit dura mater optical clearing under action of mannitol, glucose, and glycerol solutions at various concentrations are presented in Refs. 62,94,117,143, and 163. Figure 26 illustrates the dynamic changes in the rabbit dura mater turbidity after application of glycerol.143 A resolution target was placed under a sample. After the treatment of glycerol for 1 min, the target, which was not visible under the native dura mater [Fig. 26(a)], was seen through the specimen [Fig. 26(b)]. Results of the measurement of the optical properties [Fig. 26(c)] confirm the visually observed reduction in scattering. Figure 26(c) shows the increase in transmittance within the wavelength range of 400–750 nm as a function of the time the sample was soaked in glycerol. The hemoglobin absorption became much more prominent after

Figure 26 Visual changes in in vitro turbid rabbit dura mater and the measured optical changes before and after epidural application of glycerol. (a) Native dura mater placed over the resolution target, bar = 1 mm. (b) One-minute application of glycerol, bar = 1 mm. (c) Transmittance spectra for native dura mater measured at application of glycerol for 1, 2, and 10 min.143 Reprinted with permission of Springer Science and Business Media.

Optical Clearing of Fibrous Tissues

59

the application of glycerol [Fig. 26(c)]. This indicates that the quality of images received by techniques based on the detection of hemoglobin absorption spectra can be significantly improved through reduction of scattering of the upper tissue layers. In vivo studies of glucose and glycerol action on rabbit dura mater at the open cranium and epidural agent application also confirm the concept of effective optical clearing of fibrous tissue.143 Total optical clearing was achieved very fast (in 50 s) after tissue treatment by glycerol.

(a)

(b) Figure 27 (a) The collimated transmittance spectra and (b) corresponding temporal dependencies measured for the human dura mater sample in a course of administration of 40% glucose solution in a bath.94,117

Figure 27 presents the collimated transmittance spectra and temporal dependencies of the spectral components for human dura mater samples impregnated by glucose solution. It is well seen that glucose is also a very effective agent for dura mater clearing. Using such measurements for glucose and mannitol, and the algorithm described in Refs. 84 and 117, the diffusion coefficients for a 40% glucose and mannitol solution (0.16 g/ml) were found: DG = (5.43 ± 0.88) × 10−6 cm2 s−1 and DM = (1.67 ± 0.21) × 10−6 cm2 s−1 (see Table 2).117

60

Optical Clearing of Tissues and Blood

2.5 Conclusions It is shown that administration of osmolytes to a fibrous tissue allows one to effectively control its optical characteristics. The dynamics of tissue optical clearing are characterized by a time response of about 5–10 min, which is defined by the diffusivity of an immersion agent in a tissue, tissue condition (intact or fixed), and tissue thickness. The tissue shrinkage and swelling may play an important role in the tissue clearing process. At a prolonged time of some chemicals’ administration (for example, glucose at pH 3.5), tissue shrinkage at the first step of clearing may be replaced by a swelling, which in turn may cause saturation or even slight reduction of tissue optical transmittance. Dynamic optical characteristics can be used for the determination of the diffusion coefficient of endogenous (metabolic) and exogenous (chemical agent) fluids in human tissues. Obtained values for the diffusion coefficients of glucose, mannitol, Trazograph, and PEG (6000) in intact human sclera and dura mater correspond well with values of diffusion coefficients for small molecules diffusing in water.

3 Optical Clearing of Skin

3.1 Introduction Skin has a very complicated structure, as schematically presented in Fig. 28. It possesses a protective function preventing penetration of pollutions and microorganisms inside the body. The outermost cellular layer of skin is epidermis, which consists of stratum corneum (SC) (mostly dead cells) and four layers of living cells. Stratum corneum is a lipid-protein biphasic structure, having a thickness of only 10–20 µm on most surfaces of the body. Due to cell membrane keratinization, tightly packing of cells, and lipid bridges between them, SC is a dense medium with a poor penetration for foreign molecules.247–250 The excellent diffusional resistance of the SC makes the transdermal delivery of immersion agents and water lost by skin difficult. To understand the transport and barrier functions of the skin, it is important to have knowledge of the distribution of water and ions within the different layers.249,250 Water content is known to influence: various physical characteristics, such as brittleness, elasticity, tensile strength, and viscoelasticity; barrier characteristics; electrical resistance; thermal conductivity; and appearance. The SC receives water from within the body, but water also may be taken up from the environment. From within the body, water reaches this tissue from the sweat glands and by diffusion from underlying tissues. In vivo, the diffusion of water across the SC is a passive process that can be modified with the application of osmotic agents. The water content of the innermost layer of the SC is in equilibrium with the adjacent moist granular layer. The outside cell layer, however, is in equilibrium with the environment and it is certainly drier than the innermost cornified layer. Thus, there exists a concentration gradient causing transepidermal water lost (TEWL). No significant difference was found for the diffusion across epidermis and SC. The diffusion coefficient D of the flow of water through a stationary macromolecular gel (the tissue) corresponds to viscous flow through a very fine porous medium. As has been determined in strongly hydrated SC, D is about four orders of magnitude less than the self-diffusion coefficient in water.249 The diffusivity D of water in SC increases from ∼3×10−10 to 10−9 cm2 /s as humidity H increases from 46% to 81%. It should be noted that the hydration of the dermis is not significantly different from that of the viable cell layers of the epidermis.250 The average water content of the SC as measured in Ref. 250 is 54%, while other authors arrived at a water content as low as 15–40% in the same layer. The rate of diffusion of molecules with a molecular weight of 119 in the SC of volunteers is in the range from ∼10−10 to 3.5 × 10−10 cm2 /s.248 61

62

Optical Clearing of Tissues and Blood

Figure 28 Human skin structure.

Dermis is the next thicker layer of the skin, which is mostly fibrous tissue well supplied by blood, and thus can be easily impregnated by exogenous or endogenous liquids (immersion agents). Subcutaneous tissue contains a big portion of a fat cellular layer, which is much less penetrative for diffusing molecules than dermis. This specific structure of skin defines the methodology of its effective optical clearing, which is related to the matching of refractive indices of scatterers (keratinocyte components in epidermis, and collagen and elastin fibers in dermis) and ground matter.19,45,83,90,101,251 Experimental studies of optical clearing of skin using water, glycerol, glycerol-water solutions, glucose, propylene glycol, oleic acid, DMSO, sunscreen creams, cosmetic lotions, gels, and pharmaceutical products were carried out in Refs. 19, 45, 75, 83, 85–90, 94, 101, 102, 104, 105, 167, 171, 180, 189–193, 196, 251, 252.

Optical Clearing of Skin

63

3.2 In vitro Speckle and Spectral Measurements Table 4 and Fig. 29 illustrate the efficiency of the different immersion agents’ action on the transmittance of human skin stripped samples (30–40 µm in thickness) taken from volunteers using glass substrate-glue technology.83 Because of the small thickness of the sample, with only a few layers of dried and living kerationcytes, and the agent supply through the living cell layer, the rate and efficiency of immersion were very high. Table 4 Efficiency of the immersion agent action on skin stripping (30–40 µm thickness) expressed as a ratio of mean transmitted intensities after (IA ) and before (IB ) lotion application; n is the index of refraction of the lotion used.83

Immersion agent

Glycerol-water-urea solutions

n IA /IB

1.449 12.8

1.380 3.7

1.356 4.9

1.354 5.9

1.348 4.1

DMSO 50%

Ultrasound gel

1.396 7.9

1.337 5.3

Figure 29 shows the spatial series of the speckle intensity fluctuations of the transversely scanned focused laser beam transmitted by the human skin sample before and after immersion agent application and the corresponding empirical probability density function of the speckle intensity fluctuations.83 The speckle patterns (see Fig. 14) are produced as a result of interference of a large number of wavelets with random phases that arise when a coherent beam interacts with a scattering medium (a tissue sample).7 The average size of a speckle in the far-field zone is estimated as dav ∼

λ , ϕ

(62)

where λ is the wavelength and ϕ is the angle of observation. The tissue sample is characterized by the standard deviation σh of the altitudes (or refractive index) of inhomogeneities and the correlation length Lc of these inhomogeneities (random relief). Speckles either distort the results of measurements, and should be eliminated, or provide new information about the structure or the motion of cells or tissue components. The spatial or temporal intensity fluctuations are arose when: (i) the observation point over a screen is scanned, (ii) the scanning of a laser beam over a sample is provided (Fig. 29), and (iii) the sample itself or its components are moveable. The first-order statistics describe the properties of the speckle pattern point by point by using the intensity probability density distribution function p(I ) and the contrast VI =

σI , I 

σ2I = I 2  − I 2 ,

(63)

64

Optical Clearing of Tissues and Blood

(a)

(b)

(c)

(d)

Figure 29 Modification of transmitted speckles at optical immersion of a thin tissue sample (the human skin stripped sample of 30–40 µm thickness).83 The spatial series of the speckle intensity fluctuations of the focused He:Ne (633 nm) laser beam transmitted by the sample measured (a) before and (b) after immersion agent (glycerol-water-urea solution) application; X is the scanning distance of the focused laser beam along the sample surface, I is the intensity in the arbitrary units. The corresponding empirical probability density functions p(I) of the speckle intensity fluctuations (c) before and (d) after immersion agent application.

where I  and σ2I are the mean intensity and the variance of the intensity fluctuations, respectively. For ideal conditions, when the complex amplitude of the scattered light has Gaussian statistics, the contrast is VI = 1 (developed speckles), and p(I ) is represented by a negative exponential function: p(I ) = I 

−1

 I . exp − I 


(64)

This means that the most probable intensity value in the speckle pattern is equal to zero—total destructive interference is much more probable. As it follows from the measured spatial series of the speckle intensity fluctuations [Figs. 29(a) and 29(b)] at the immersion of a skin sample (mostly SC) by glycerol-water-urea solution, speckles become bigger in size and brighter; thus, the mean intensity increases (see Table 4). The corresponding empirical probability

Optical Clearing of Skin

65

density function p(I ) of the speckle intensity fluctuations transfers from a negative exponential function, described by Eq. (64), to a function characterized by a more uniform distribution with a higher probability of the higher-transmitted intensities. An in vitro study of rat dorsal skin impregnated by anhydrous glycerol, when the agent was applied to the dermal side of the skin sample,45 it showed a power wavelength dependence of the reduced scattering coefficient in the wavelength range from 500 to 1200 nm, described by the expression224,226 µs ∝ λ−h ,

(65)

was shown. For the normal skin, the reduced scattering coefficient at 500 nm was µs ≈ 50 cm−1 and h = 1.12.45 With increased time of the skin sample being in glycerol, µs (500 nm) and h were subsequently decreased (mostly due to dehydration effect); µs ≈ 30 cm−1 and h = 1.09 for 5 min in glycerol, µs ≈ 20 cm−1 and h = 0.85 for 10 min, µs ≈ 12 cm−1 and h = 0.52 for 20 min, and µs ≈ 23 cm−1 and h = 0.9 for the rehydrated sample kept in physiological phosphate buffered saline solution for 20 min.45 A 60% decrease in hydration was estimated on the basis of changes in the water absorption peaks and a 21.5% corresponding decrease in thickness was found going from the native tissue to the tissue treated with glycerol for 20 min. The rehydration process caused the thickness and turbidity of the sample to go back toward the initial states, but during the course of rehydration, which lasted 20 min, the turbidity µs did not reach the initial state. Accounting for the relatively short period of time (∼20 min) for optical clearing of the skin samples of about 1 mm in thickness in this experiment and the high viscosity of glycerol, its action as a hyperosmotic agent should have mostly drew interstitial water out of the tissue and, at a slower rate, should have replaced the water and salts of the ground substance. The twenty minutes of rehydration are also not enough for water to reenter all of the cells and collagen fibers in the tissue; thus, the scattering coefficient and spectral power parameter h for rehydrated tissue are somewhat less than their initial values. More prolonged administration of glucose (up to 6 hr) and glycerol (up to 45 min) into the fresh rat skin samples at room temperature in the course of tissue collimated transmittance measurements was also done.172,179,180 These studies were performed to clarify the mechanisms of the skin optical clearing and to optimize the technique. To avoid tissue damage and to provide less viscosity of the chemical agent, a glycerol-water solution (88%) and 40% glucose (both available in a drug store) were used as immersion agents. Skin samples were 0.57–0.90 mm in thickness and 1 × 1 cm2 in area, some of them contained whole skin including epidermis, dermis, and hypodermic fatty layer, and for others the fatty layer was removed. Hairs were removed by tweezers, and the immersion agent was applied for both sides of the sample in a bath. Figures 30 and 31 illustrate the typical collimated transmittance spectra and optical clearing dynamics. It is well seen that the untreated rat skin is poorly transparent for the visible light. Both glucose and glycerol administration make this tissue highly transparent; the 15-fold increase of

66

Optical Clearing of Tissues and Blood

the collimated transmittance for glucose [Fig. 31(a)] and tenfold increase for glycerol [Fig. 31(c)] at 700 nm for the samples with a fatty layer kept in solution for 45 min are seen. The efficiency is substantially greater with the removed fatty layer [Fig. 31(b)], and about a 50-fold transmittance increase is seen for the glucose solution at the same wavelength during the same time interval; further increase of transmittance and its saturation happened for more prolonged time intervals.

Figure 30 Collimated transmittance spectra of the rat skin sample measured in a course of glycerol (88%, vol/vol) administration in a bath at different time intervals (sample thickness of 0.9 mm).172

It is evident that the penetration rate of the agent into the skin is much slower than that for the sclera or dura mater, which takes only 8–10 min to be saturated by Trazograph or glucose solutions. In comparison with the sclera and dura mater, no saturation of the clearing process was seen for up to six hours if the fatty layer is not removed. This phenomenon can be connected with the low permeability of the epidermal and fat tissue cellular layers for any molecules, which slows down both fluxes—the water flux from the tissue and the immersion agent flux from the outside to the inside of the tissue. Saturation of the optical transmittance can be expected when the equilibrium state in the immersion agent/water diffusion process is achieved, i.e., when concentrations of water and immersion agent inside and outside tissue are approximately equal. For skin with epidermis and fatty layer, such saturation was not reached even after six hours of glucose administration; but without the fatty layer, saturation was achieved in one hour. Using the earlier discussed algorithm described by Eqs. (27), (33), (40), and (43) and experimental data (see Fig. 31), the diffusion coefficient of water in the skin with glycerol action can be estimated. Such estimation is valid for the agents

Optical Clearing of Skin

67

(a)

(b)

(c) Figure 31 Time-dependent collimated transmittance of the rat skin samples (1 hr after autopsy, hairs were removed using tweezers) measured at different wavelengths in a course of administration of immersion solution in a bath. (a) Sample thickness 0.73 mm, with hypodermic fatty layer, immersion agent 40% glucose; (b) sample thickness 0.57 mm, with removed hypodermic fatty layer, immersion agent 40% glucose; (c) sample thickness 0.9 mm, with hypodermic fatty layer, immersion agent glycerol-water solution (88%).179,180

68

Optical Clearing of Tissues and Blood

with a strong osmotic strength, because the water flux dominates in the system. The mean value of the diffusion coefficient averaged for the six rat skin samples at 20◦ C for glycerol-water solution penetration, mostly from the dermal side of the skin, is equal to (5.12 ± 2.27) × 10−7 cm2 s−1 (see Table 2), which is about two orders less than the diffusion coefficient of water in water, DW ∼ = 10−5 cm2 s−1 (see Ref. 241); or one order less than water diffusion in an intact human lens, DW ∼ = 3.0 × 10−6 cm2 s−1 (see Ref. 155). For a subcutaneous fat-free sample, a 40% glucose action is characterized by a higher diffusion rate, D = (3.1 ± 0.1) × 10−6 cm2 s−1 (see Table 2), which may be due to the more effective penetration of glucose into a tissue. Using near-infrared spectroscopy (800–2200 nm), mass and water loss measurements, and transdermal skin resistance measurements, such enhancers of skin permeability as dimethyl sulfoxide (DMSO) and oleic acid, a monounsaturated fatty acid, were compared in a propylene glycol (PG) application onto the epidermal surface of samples of fresh porcine skin with a thickness of 1.52 ± 0.18 mm.167 It was shown that, when compared with DMSO as an enhancer, oleic acid has a similar synergetic optical clearing effect. Accounting for clinical safety reasons, oleic acid could be an optimal choice as an enhancer for optical clearing of skin because it is recognized as a safe agent (free of side effects), and DMSO has some potential toxicity. After application of the oleic acid solution (0.1 M of oleic acid and PG-40), the total transmittance measured on the wavelength 1278 nm of the skin sample increased by 41% and 58%, respectively, for 30 and 60 min treatments, while diffuse reflectance decreased by 39% and 47%, respectively. Diffuse reflectance spectra were transformed into apparent absorbance (A = log 1/Rd ). The difference in apparent absorbance between two wavelengths, 1936 nm and 1100 nm, was adopted to monitor the change in water content.97,98,141,142,166,167 It is mportant that the oleic acid solution provided the greatest water loss in comparison with the other tested solutions—37% and 46% after 30 min and 60 min treatments, respectively. As for DMSO-50, water loss was 15% and 20%; PG-80, 20% and 29%; and PG-80 + DMSO-50, 34% and 44%, after 30 min and 60 min treatments, respectively. But the mass loss with the oleic acid solution application was the minimal among the tested solutions—after 30 min, PG-80 provided 10.9% mass loss, PG-80 + DMSO-50 provided 6.4%, and oleic acid (0.1 M) + PG-40 provided 6.3%. More mass loss was obtained after 60 min of these agents’ application—PG-80, 14.2%; PG-80 + DMSO-50, 9.9%; and oleic acid (0.1 M) + PG-40, 8.3%. The comparison of water and mass loss data gives a nice confirmation of the basic conception of the optical clearing; that refractive index matching is achieved by two main diffusing processes—water flux from tissue (dehydration) and agent flux into tissue (replacement of interstitial water by the agent). A method of accelerating penetration of the index matching compounds by enhancing skin permeability by creating a lattice of microzones (islets) of limited thermal damage in the SC was recently proposed.196 A combination of a flashlamp system (EsteLux, Palomar Medical Technologies, Inc.) and a specially designed

Optical Clearing of Skin

69

appliqué with a pattern of absorbing centers (center size ∼75 µm, lattice pitch ∼450 µm) has been used to create the lattice of islet damage (LID). Several index matching agents, including glucose and glycerol, have been tested. The essential optical clearance of a full-thickness pig, rat, and human skin in vitro sample has been demonstrated with 40% glucose and 88% glycerol solutions after creating a LID with a few optical pulses (fluence 14–36 J/cm2 , 20 ms pulse duration). One example for in vitro pig skin of 1 mm of thickness is shown in Fig. 32.

(a)

(b) Figure 32 The accelerating penetration of the index matching compound by enhancing skin permeability by creating a lattice of microzones (islets) of limited thermal damage in the stratum corneum.196 (a) The white light transilluminating image of the pig skin specimen of 1 mm thickness at topical application of 40% glucose, 2 pulses of 27 J/cm2 (20 ms) with island damage mask and 2 pulses of 27 J/cm2 (20 ms) after mask removing; optical clearing is more effective in the area of the lattice of islet damage. (b) Time-dependent spectral transmittance on 470 and 650 nm within the area of maximal translucence. Copyright © 2005, Wiley-Liss Inc., a subsidiary of John Wiley & Sons, Inc. Reprinted with permission.

70

Optical Clearing of Tissues and Blood

3.3 In vivo Spectral Reflectance Measurements In vivo topical application of glycerol, glucose, Trazograph (x-ray contrasting substance), propylene glycol, and cosmetic lotions and gels also made human skin more translucent, within a time periods from a few minutes to a few hours.86,88,89,168,186,188 Water loss or increase by means of moisturizing substances seriously influence skin optical properties.187–190,248–250,253 NIR reflectance spectroscopy is used as a method to directly determine changes in free, bulk, and protein-bound water and to assess scattering effects in skin, for the evaluation of skin care products.188 The following spectral bands are associated with water: free water, 1879 nm; bulk water, 1890 nm; and protein-bound water, 1909 and 1927 nm. The effect of increases in ambient humidity is associated with increased levels of free water in the skin, while moisturizers containing hydroxyethyl cellulose, propylene glycol, dipropylene glycol, and glycerol contribute to a decrease in the light scattering.188 The water observed in such experiments is primarily in the SC since only a small part of the reflected light comes from the epidermis or below. Noninvasive measurement of the SC hydration can be performed using attenuated total reflectance Fourier transform infrared (ATR FTIR) spectroscopy.189,248,253 Three absorption bands are relevant for determining water content in the SC: 3300 cm−1 (3030 nm), O H and N H vibrations; 1645 cm−1 (6079 nm), amide I band; and 1545 cm−1 (6472 nm), amide II band. The amide I band intensity is pronounced in the presence of water due to the strong absorption of water at 1645 cm−1 and the changes in the carbonyl absorption under the influence of water, while the amide II band intensity is due to protein alone. The intensity ratio of the amide I/amide II bands, also called the moisture factor, is assumed to be a relative measure of the SC hydration.189 Various SC moisturizers based on glycerol, propylene glycol, sodium lactate, natural moisturizing vegetal compounds, liposomes, butylene glycol, polyglycerylmethacrylate, and urea were used for an in vivo SC hydration study.189 Depending on the composition and concentration, the maximal SC hydration could be reached in 0.5–2 hr after application of the substance on the skin surface. For some substances, a considerable moisturizing effect was detectable up to eight hours following application. Dual wavelength (1300 and 1450 nm) optical coherence reflectance measurement is a prospective technique for depth profiling water absorption within the skin.190 To enhance clearing agent permeation through the SC, a number of specific procedures such as heating, electrophoresis, and sonophoresis are usually applied.180,186,247 To increase the efficiency of the topical application of the clearing agents, gels containing clearing agents (Verografin, glycerol, or glucose) were designed.186 The diffusion rate of the agents within the gel layer can be rather high, and this along with a comparatively large volume of the gel provided the constant concentration of the clearing agent, equal to the agent content in the gel, at the skin surface. For intact skin of a volunteer, the best dynamics, i.e., the rate and the degree of clearing (17%), was observed in the case of Verografin gel [Fig. 33(a)]; after 40 min of observation, clearing still proceeds with marked rate, while for

Optical Clearing of Skin

71

glycerol gel, after 27 min the curve flattens out; no additional clearing was observed in 40 min of glucose gel application. As a barrier function of the skin is associated mainly with SC, the measurement was carried out on the skin after 30–50 µm epidermal glue stripping [Fig. 33(b)]. Application of glucose gel to the skin without upper epidermis gave a rapid 10% drop of reflected light intensity. Glycerol gel gave better results; over the time of observation, the decrease of reflected signal ranged up to ∼20%, which was twice that attained for intact skin. Surprisingly, no clearing effect of Verografin gel was observed. The electrophoretic applicator and gel with twice the content of gelatin were also applied to human skin optical clearing.186 In Fig. 33(c), the results for glycerol gel are shown. When the active electrode was connected as an anode, the reduction of optical signal by ∼20% was attained. This value is comparable to the results with stripping, but the time of attainment of minimum signal is nearly halved. When the active electrode was connected as a cathode, an increase of back-reflectance was observed over the whole duration of measurement. The effect was attributed to the development of erythema. It may be concluded that for the topical application of glycerol gel and glucose gel, the employment of epidermal stripping and the electrophoresis technique does lead to the enhancement of the dynamics of in vivo optical clearing of human skin. The best characteristics were obtained with electrophoretic administration of glycerol from an anode. In the case of glucose, stripping and electrophoresis from a cathode give results close to each other, but the application of glucose should be terminated after 10–12 min because of the risk of deterioration of clearing by erythema development. The effective in vivo human skin optical clearing due to accelerated skin permeability for refractive index matching agents at creating a lattice of islet thermal damage in the SC was recently demonstrated.196 The administration of glucose or glycerol by intradermal injection into rat or hamster skin causes a decrease of reflectance and the corresponding increase of tissue transmittance.45,94,108,110,172,179,180 This effect was observed at all the wavelengths during the 15–18 min period after glucose injection.108,110 The highest degree of tissue reflectance change is found at the wavelengths from 580 to 750 nm, where scattering dominates. At the 16th min, the reflectance of the skin was minimal (maximal transmittance), it decreased about 3.5-fold at 700 nm, then the tissue went slowly back to its normal state—at the 76th min, only a twofold reduction of reflectance was observed. It was shown that glycerol injection causes a more prolonged effect on tissue optical clearing, but reflectance decreased a little bit less than for glucose injection. This can be explained by the higher viscosity of glycerol and by its mostly indirect action through tissue dehydration. The reaction of the rat skin on injection of distillate water, as a model of a nonosmotic agent, also was studied. The reduction of reflectance was observed only for a short period (the first minute) after injection. It happened due to a much higher transmittance of the injected water in respect to the surrounding tissues. At the second minute,

72

Optical Clearing of Tissues and Blood

(a)

(b)

(c) Figure 33 Back reflectance at 830 nm measured by a fiber-optic probe (rsd = 1.2 mm, source-fiber of 0.2 mm and detector fiber of 1 mm in diameter) perpendicular to the skin surface at clearing gel (glucose: 3.9 ml 40% glucose and 0.2 g gelatin; glycerol: 1.3 ml glycerol, 0.1 g gelatin, and 2.6 ml distillate water; verografin: 2.6 ml Verografin-60, 0.1 g gelatin, and 1.3 ml distillate water) application on (a) intact skin; (b) on skin after glue epidermal stripping; (c) at electrophoretic application of gelatin gel with glycerol (1.3 ml glycerol, 0.2 g gelatin, and 2.5 ml distillate water).186

Optical Clearing of Skin

73

water diffuses into bulk tissue and the transparency of tissue was decreased—the reflectance spectrum was elevated gradually to its initial value. The injection of water does not cause the immersion clearing of the skin. The virtual transparent window (VTW) of ∼4 mm diameter in the skin is created with the living time of ∼30 min for 40% glucose and more than 60 min for 75% glycerol. Such a window allows one to clearly identify visually, by the naked eye, blood microvessels in the skin.108,110 The swelling white ring (edema) appeared around the VTW after agent injection. The images of skin after intradermal injection of glucose, glycerol, and water were recorded by a digital video camera. The diameters of the swelling area (DS ) and VTW (DT ), and their ratio (DS /DT ), were measured (Fig. 34).108 For glucose injection, the diameter of the VTW was registered at the first t minute after injection. At the second minute, the diameter was slightly decreased. For the next 15 min, this diameter and the diameter of the swelling area were not changed. After the 20th min, the significant reduction of the VTW was observed. For glycerol injection, the diameter of the VTW was approximately the same, but the swelling ring was bigger, and both transmittance and swelling were seen for a longer time than for glucose injection [Fig. 34(b)]. The injection of distillate water causes only the appearance of swelling at the site of the injection. The diameter of the swelling area is decreased gradually and the swelling disappears up to the 30th min after injection. Figure 35 shows the reflectance spectra and the corresponding time-dependent reflectance for a few spectral components measured, for a healthy human volunteer, with intradermal 40% glucose solution.90 The reflectance spectra show a scattering background determined by the diffusion reflection of the skin layers with the well-pronounced bands caused by blood optical absorption. Within one hour after glucose injection, the skin reflection coefficient decreases on average by a factor of 3.8 and then exhibits a slow increase, which indicates that glucose is eliminated from the observation area and that the skin reflectance tends to restore itself to the initial level. Based on this result and the proposed skin clearing model, we may suggest that the main contribution to clearing in the first stage (first hour) is due to the refractive index matching between collagen fibrils of the dermis (n = 1.46) and the interstitial space (initially n = 1.36) to which glucose (n = 1.39) diffuses. Estimated from the experimental data [Fig. 35(b)], the diffusion coefficient of glucose in dermis is DG = (2.56 ± 0.13) × 10−6 cm2 /s; this value is 3.6-fold less than for glucose diffusion in water at 37◦ C, DG ≈ 9.2 × 10−6 cm2 /s, and reflects the character of dermis permeability for glucose. For applications, it is important that skin preserves transparency (low reflectance) for a few hours after injection, which is defined by glucose diffusion along the skin surface, because upper and lower layers of the skin—epidermis and fat—have much lower permeability than dermis. For the optical clearing effect to continue to be seen, glucose should diffuse at least at the distance l = 1.25– 1.75 mm for the fiber probe used in the experiments (Fig. 17), i.e., the diffusing (optical clearing) time τ ≈ l 2 /DG ≈ 1.7–3.3 hrs (which well corresponds to the experimental data) [Fig. 35(b)].

74

Optical Clearing of Tissues and Blood

(a)

(b) Figure 34 The changes of skin reaction on injection of (a) 40% glucose and (b) 75% glycerol: Q—diameter of virtual transparent window (VTW) (DT ); 2—diameter of swelling area around VTW (DS ); ∗ , +—ratio DS /DT ; !—diameter of swelling area at injection of distillate water.108,110

As is well seen from Fig. 35(a), at dermis clearing (reduction of scattering) the contrast of the hemoglobin absorption bands is significantly higher than for the control, but for prolonged immersion (curve 3), contrast is again not very high. This result is very important for contrasting of tissue abnormalities (tumors) associated with hemoglobin or other probe absorbers concentration (for instance, indocyanine green dye). Therefore, there is an optimal immersion time interval (for human skin at glucose injection is of 60 min) that allows one to see skin absorbers and localize them more precisely at reduced scattering. Indeed, for prolonged immersion, contrast goes down due to less light interactions with absorbers at low-step scattering.

Optical Clearing of Skin

75

(a)

(b) Figure 35 (a) The reflectance spectra and (b) the time-dependent reflectance at three wavelengths (420, 500, and 700 nm) of human skin measured at hyperdermal injection of 0.1 ml of 40% glucose into the internal side of the forearm of the male volunteer for different time intervals; (1) intact skin, (2) at 23 min, and (3) at 60 min after injection.90

3.4 In vivo Frequency-Domain Measurements The dynamical response of optical properties of human skin treated by a chemical agent can be measured using a photon density wave (frequency-domain) technique.86 The intensity and phase of photon-density waves generated by an NIR optical source (786 nm) were measured at several source-detector separations (see Fig. 21).86,88 For the small (1–3 mm) source-detection separation measurements allowing for thin tissue layer examination, a special multichannel fiber-optic probe was designed. It was used together with a Dicon multichannel fiber-optic switcher. The dynamical response of optical properties (modulation depth and phase shift of intensity modulation of the backscattered light) was measured for human skin via

76

Optical Clearing of Tissues and Blood

the interval of a chemical agent administration. The measurement for each separation was done during 10 s and averaged, corresponding to one point in Fig. 36. The relative amplitude (normalized to the initial amplitude) and the phase changes (the current phase minus the initial phase) during 20 min of glycerol topical application are shown in Fig. 36. Only scattering changes must be considered due to the extremely low absorption of glycerol at the measuring wavelength. The observed amplitude and phase changes are small, reflecting minor permeation of epidermal cell layers to any chemical agent. Nevertheless, these measurements show enough sensitivity of the frequency-domain method to the small changes of the scattering coefficient of the skin.

(a)

(b)

Figure 36 Frequency-domain back-reflectance measurements for the small source-detector separations.86,88 The time-dependent changes of the (a) amplitude and (b) phase shift of the signal for several source-detector separations (1.14, 2.08, and 3.5 mm) for in vivo study of a human arm under 20 min glycerol administration.

For large (2.5 cm) source-detector separation studies, the source and detector fiber tips have been mounted in the rubber pad and fastened to the surface of a human forearm to avoid random moving artifacts. The cosmetic gel with refractive index n = 1.403 (cold cream, see Table 1) has been put on the surface of the arm and the phase and ac amplitude measurements have been provided continuously. One sampling point corresponds to one second. The results of measurements during 30 min of gel administration are shown in Fig. 37(a). The observed temporal quasi-periodic fluctuations in the phase and the amplitude of the optical signal are caused mainly by the heartbeats. Results of the calculations of tissue optical parameters are shown in Fig. 37(b). The initial values of µs and µa for human skin were taken from Ref. 7, and the relative changes of these parameters were calculated with a running-average procedure for every 5 s interval in order to exclude the influence of heartbeats. Corresponding temporal evolutions of the scattering and absorption coefficients have been

Optical Clearing of Skin

77

(a)

(b) Figure 37 Frequency-domain back-reflectance measurements for the large source-detector separation (2.5 cm).86,88 (a) The raw experimental data of the phase and ac amplitude of the optical signal and (b) calculation of the absorption and the scattering coefficients. The cosmetic gel with refractive index n = 1.403 (cold cream) was used.

received. This study shows that there are no noticeable changes in the absorption during the gel administration trial. A slight increase in the absorption probably can be explained by the increase of the water content in the skin due to the moisture effect of the applied gel. The selected source-detector separation (2.5 cm) and corresponding measuring volume are too large to make the matching effect a useful procedure for topical application of the gel. Only about a 6% reduction of the scattering coefficient averaged over the large measuring volume was observed. This means that the scattering coefficient of the upper (superficial) layers of the skin changed more effectively. Refractive index matching of the fiber tips and tissue surface is also important.

78

Optical Clearing of Tissues and Blood

In vivo frequency-domain measurements for immersed tissues show that the refractive index matching technique provided by the appropriate chemical agent or cosmetic preparation application can be successfully used in tissue spectroscopy and imaging when reduction of scattering properties is needed.

3.5 OCT Imaging The typical optical coherence tomography (OCT) fiber-optic system employs a broadband light source (a superluminescence diode), delivering light at the central wavelength of 820 nm or 1300 nm with a bandwidth of 25–50 nm (Fig. 38). Such an OCT system provides 10–20 µm of axial and transverse resolution in free space with a signal-to-noise ratio up to 100 dB.14,254

Figure 38 The OCT system employs a broadband light source, delivering an output power of 1 mW at the central wavelength of 820 nm with a bandwidth of 25 nm, yielding 12 µm axial resolution in air that determines the imaging axial resolution.142 Back-reflectance measurements were performed by directing low-coherence light at the tissue and detecting the reflections from the front and in-depth interfaces between tissue layers by the use of an optic-fiber integrated system: CLs, collimating lens; D1 , D2 , detectors; FC2 , FC3 , fiber couplers; PCs, polarization controller; OL, objective lens. Polarization fibers are used to provide the maximum obtainable resolution for the given spectral width of the source. The transverse resolution was measured at 16 µm. Reprinted with permission of Springer Science and Business Media.

The result of the OCT study is the measurement of optical backscattering or reflectance, R(z), from the tissue versus axial ranging distance, or depth, z. The reflectance depends on the optical properties of the tissue, i.e., the absorption µa and scattering µs coefficients, or total attenuation coefficient µt = µa + µs . The relationship between R(z) and µt is, however, highly complicated because of the high and anisotropic scattering of the tissue. But for optical depths less than 4,

Optical Clearing of Skin

79

the reflected power can be approximately proportional to −µt z on an exponential scale, according to the single scattering model,142 i.e., R(z) = I0 α(z) exp(−2µt z),

(66)

where I0 is the optical power launched into the blood sample and α(z) is the reflectivity of the tissue sample at depth z. The factor 2 in the exponential accounts for the light passing through the tissue sample twice after it is backscattered. Optical depth is a measure of depth in terms of number of mean free path lengths, i.e., µs z. α(z) is linked to the local refractive index and the backscattering property of the blood sample. If α(z) is kept constant, µt can be obtained theoretically from the reflectance measurements at two different depths, z1 and z2 :   R(z1 ) 1 µt = ln , 2(
z) R(z2 )

(67)

where
z = |z1 − z2 |. As noise is inevitable in the measurement, a final result should thus be obtained by use of a least-squares fitting technique to improve the accuracy of the determined value of µt . Optical clearing (enhancement of transmittance)
T by an agent application can be estimated using the following expression: 

R a − Rs
T = Rs

 × 100%,

(68)

where Ra is the reflectance from the backward surface of the sample impregnated by an agent, and Rs is that of a control sample. Multiple scattering is a detrimental factor that limits OCT imaging performances: imaging resolution, depth, and localization. To improve the imaging capabilities, the multiple scattering of tissue must be reduced. The immersion technique in the application of biocompatible agents is expected to be a prospective technique for OCT, because the depth of OCT images and their contrast can very easily be essentially improved at immersion.14,30,45,63,91,93,97,99,102,138,142,144,168,254 OCT imaging combined with agent immersion is a useful technology for skin disease diagnosis and monitoring. To illustrate the dynamics of skin optical clearing after the application of glycerol, a set of OCT images (820 nm) of a rat skin sample was recorded at regular time intervals over a period of 40 min (Fig. 39).91 Both the index matching effect, leading to the enhanced depth capability, and the localized dehydration effect, leading to the improvement of imaging contrast, are clearly evident. Analogous results were received for fresh porcine and chicken skin at imaging on 1300 nm by 50% and 80% glycerol solutions by R. K. Wang. The OCT image of human skin with psoriatic erythrodermia acquired in 60 min after application of glycerol [Fig. 40(b)] differs from the initial image [Fig. 40(a)] in greater penetration depth and better contrast.63 These image improvements facilitate identifying important morphological phenomenon of acanthosis.

80

Optical Clearing of Tissues and Blood

Figure 39 Dynamic OCT images (λ = 820 nm) at (a) 0, (b) 3, (c) 10, (d) 15, (e) 20, and (f) 40 min after a topical application of 80% glycerol solution onto rat skin.91 Images performed right after the rat was sacrificed; all the units presented are in millimeters, and the vertical axis presents the imaging depth. Reprinted with permission.

(a)

(b)

Figure 40 OCT images of skin with psoriatic erythrodermia: (a) before topical application of glycerol; (b) 60 min after application of glycerol.63 Reprinted with permission of Springer Science and Business Media.

The possibility of in vivo diagnostics of malignant melanoma, observation of subepidermal blisters, and control of the scattering properties of skin through the

Optical Clearing of Skin

81

saturation of skin with glycerol by its topical application was demonstrated on the basis of OCT imaging of human skin both in vitro and in vivo.168 According to the estimates done by the authors of Ref. 168, the scattering coefficient for the SC with glycerol is reduced by a factor of two (Fig. 41). For epidermis and upper dermis, the coefficient of scattering decreases by 20%. For deeper dermis layers, the coefficient of scattering lowers only by 5%. The effect on enhancement of both the imaging depth and contrast was found in in vivo studies of human skin optical clearing at topical application of the 50% propylene glycol solution.92

Figure 41 OCT signals (in intensity-depth coordinates) obtained for in vivo probing of human fingertip skin. The upper curve corresponds to skin saturated with glycerol. The scattering coefficient for the skin with glycerol is reduced: for the SC by a factor of two; for epidermis and upper dermis by 20%; and for deeper dermis layers, only by 5%.168

The OCT images captured from the skin site of a volunteer at hyperdermal injection of 40% glucose allows one to estimate the total attenuation coefficient [see Eq. (67)].92 The attenuation initially goes down and then goes up with time. Such behavior well correlates with the spectral measurements shown in Fig. 35 and also reflects the index matching induced by the glucose injection. The light beam attenuation in tissue, I /I0 ∼ exp(−µt ), for intact skin (0 min) was found from OCT measurements as I /I0 ∼ = 0.14, and for immersed skin at 13 min as I /I0 ∼ = 0.30, i.e., the intensity of transmitted light increased 2.1-fold. That value also correlates well with the spectral measurements. It should be noted that high sensitivity of an OCT signal to immersion of living tissue by glucose allows one to monitor its concentration in the skin at a physiological level.156–159 Although glycerol and glucose are effective optical clearing agents when injected into the dermis,45,92,110 they normally do not penetrate so well into intact

82

Optical Clearing of Tissues and Blood

skin. In recent OCT experiments with human skin in vivo at topical application during 90–120 min of the combined lipophilic polypropylene glycol-based prepolymers (PPG) and hydrophilic polyethylene glycol (PEG)-based prepolymers, both with indices of refraction of 1.47 that closely match that of skin scattering components in the SC, epidermis and dermis, it was shown that the polymer mixture can penetrate intact skin and improve OCT images to see dermal vasculature and hair follicles more clearly.102 This composition may have some advantages in skin optical clearing due to a hydrophilic component that may more effectively diffuse within living epidermis and dermis; less osmotic strength may also have some advantages, but the optical clearing depth could not be improved radically in comparison with topical application of other clearing agents, such as glycerol, glucose, x-ray contrast, and propylene glycol, because of the principle limitations of chemical agent diffusion through intact cell layers. Thus, to provide fast and effective optical clearing of skin, the appropriate well-known or newly developed methods of enhanced agent delivery should be applied. Some of them are discussed in Sections 3.3 and 7.4.

4 Optical Clearing of Gastric Tissue

4.1 Spectral Measurements Transmittance and diffuse reflectance measurements were recently performed over a range from 800 to 2200 nm for previously frozen and fresh native porcine stomach cardiac and pyloric mucosa sections of 1.2–1.6 mm thickness.97,98,141,142,166,167 Mucosa consists of moist epithelium and the connective tissue immediately beneath it. Mucosal structure is somewhat identical to skin with cell epidermal and fibrous dermal layers. The absence of the dead cell layer, such as the SC of skin, makes normal mucosa more permeable for chemical agents. Immersion solutions (glycerol/DMSO/water) of different concentrations were topically applied onto the epithelium surface of the sample and then spectra were acquired at time intervals of 5, 10, 20, and 30 min. The difference in apparent absorbance (extracted from the diffuse reflectance) between two wavelengths, 1936 nm and 1100 nm, was used to estimate water content. Some results are presented in Fig. 42. It can be seen from Figs. 42(a) and 42(b) that, over the whole wavelength range investigated, the transmittance was increased with time, and diffuse reflectance was decreased over the range of 800–1370 nm. The greatest increase in transmittance was at 1278 nm, and the greatest decrease in reflectance was at 1066 nm. It is found that there is a strong correlation between optical clearing and water desorption.97,98,141,166,167 At 30 min after the treatment, 80% glycerol caused 15% water loss, whereas 50% glycerol and 50% DMSO caused 9% and 7% water loss, respectively. The patterns of optical clearing are similar to those of water desorption. The water loss was maximal (∼19%) and optical transmittance at 1278 nm was also maximal (∼30%) for the mixture of 50% glycerol and 30% DMSO (synergetic effect). Reduction of scattering and water absorption allows one to get more pronounced signatures of absorbing bands of tissue components. In particular, this is demonstrated by apparent absorbance spectra (1400–2200 nm) in Fig. 42(c), measured at 50% DMSO solution application. The major features of these spectra are the bands near 1450 nm and 1936 nm, corresponding to the first overtone of OH stretch in water and the combination mode of OH stretch and HOH bend in water, respectively. DMSO application significantly changes the absorbance spectrum of the tissue. The peaks of 1730 nm and 1675 nm appeared 5 min after DMSO administration, i.e., with the water loss (scattering and water absorption reduction) and may correspond to resolvable CH groups in lipids and proteins. 83

84

Optical Clearing of Tissues and Blood

(a)

(b)

(c) Figure 42 Optical changes for porcine stomach fresh pyloric mucosa before and after application of immersion solution measured by spectrophotometer with an integrating sphere at the time intervals of 0, 5, 10, 20, and 30 min [(a) and (b) from bottom to top, and (c) from top to bottom, respectively]. (a) Total transmittance and (b) diffuse reflectance over the range 800–2200 nm after topical application of 80% glycerol onto the epithelium surface of the sample of thickness 1.6 ± 0.2 mm.97,142 (c) Apparent absorbance spectra calculated from diffuse reflectance measurements over the range 1400–2200 nm after application of 50% DMSO, the sample of thickness was 1.15 ± 0.12 mm.98

4.2 OCT Imaging In vitro studies of optical clearing of gastrointestinal tissues, such as stomach, esophagus, and colonic mucosa, were recently performed using the OCT imaging technique.92,93,97,144 Figure 43 shows two OCT images of intact and treated

Optical Clearing of Gastric Tissue

85

(by 80% propylene glycol solution) normal fresh human stomach tissue (fundus). A more cleared image with the excellent differentiation of epithelium, isthmus, lamina propia, and muscular tissue is achieved with the agent action.92,93

(a)

(b) Figure 43 OCT images of a normal fresh human stomach tissue (fundus): (a) without and (b) with topical application of 80% propylene glycol solution. E, epithelium; LP, lamina propia; MM, muscularis mucosae.92 Copyright 2002. Reprinted with permission from IOS Press.

Figure 44 The time course of repeated OCT A-scans of porcine stomach tissue with the application of glycerol. The horizontal and vertical axes present the time (min) and the imaging depth (mm), respectively; the registration of OCT signal starts at the time about 0.5 min after the agent application.144 Reprinted with permission.

86

Optical Clearing of Tissues and Blood

Figure 44 illustrates the M-mode OCT images obtained from the repeated A-scans of porcine stomach with the application of glycerol.144 From the image, it is clearly seen that the penetration depth increases gradually with the increase of time duration. There is a slope to the surface of the tissue. The downward trend of the tissue surface is attributed to the tissue dehydration induced by the chemical agent. It should be pointed out that the above experiments were performed on in vitro biological tissues. The dynamic optical clearing effect induced by the chemical agent would differ from that of an in vivo case. Because of the cell self-regulation and blood circulation, the living tissues would have less dehydration after the application of hyperosmotic chemical agent.

5 Tissue Structural Properties Studies at Optical Immersion

5.1 Polarization Measurements Many biological tissues are optically anisotropic.7−11,15,21−24,33−35,37,200−204,209, 238,255−261 Tissue birefringence results primarily from the linear anisotropy of fibrous structures, which form extracellular media. The refractive index of a medium is higher along the length of a fiber than along the cross section [Fig. 45(a)]. A specific tissue structure is a system composed of parallel cylinders that create a uniaxial birefringent medium, with the optic axis parallel to the cylinder axes. This is called birefringence of form. A large variety of tissues, such as eye cornea, tendon, cartilage, eye sclera, dura mater, testis, muscle, nerve, retina, bone, tooth, myelin, etc., exhibit form birefringence. All of these tissues contain uniaxial and/or biaxial birefringent structures. For instance, in bone and tooth, these are mineralized structures originating from hydroxyapatite crystals, which play an important role in hard tissue birefringence. In particular, dental enamel is an ordered array of such crystals surrounded by a protein/lipid/water matrix.7,15,37,213

(a)

(b)

Figure 45 Models of tissue birefringence: (a) system of long dielectric cylinders, (b) system of thin dielectric plates.37,197

Tendon consists mostly of parallel, densely packed collagen fibers. Interspersed between the parallel bundles of collagen fibers are long, elliptical fibroblasts. In general, tendon fibers are cylindrical in shape with diameters ranging from 20 to 400 nm.262 The ordered structure of collagen fibers running parallel to a single axis 87

88

Optical Clearing of Tissues and Blood

makes tendon a highly birefringent tissue. Arteries have a more complex structure than tendons. The medial layer consists mostly of closely packed smooth muscle cells with a mean diameter of 15–20 µm. Small amounts of connective tissue, including elastin, collagenous, and reticular fibers, as well as a few fibroblasts, are also located in the media. The outer adventitial layer consists of dense fibrous connective tissue. The adventitia is largely made up of collagen fibers, 1–12 µm in diameter, and thinner elastin fibers, 2–3 µm in diameter. As with tendon, the cylindrical collagen and elastin fibers are ordered mainly along one axis, thus causing the tissue to be birefringent. Myocardium, on the other hand, contains fibers oriented along two different axes. Myocardium consists mostly of cardiac muscle fibers arranged in sheets that wind around the ventricles and atria. In pigs, the myocardium cardiac muscle is comprised of myofibrils (about 1 µm in diameter) that, in turn, consist of cylindrical myofilaments (6–15 nm in diameter) and aspherical mitochondria (1–2 µm in diameter). Myocardium is typically birefringent since the refractive index along the axis of the muscle fiber is different from that in the transverse direction.262 Form birefringence arises when the relative optical phase between the orthogonal polarization components is nonzero for forwardly scattered light. After multiple forward scattering events, a relative phase difference accumulates and a delay (δoe ) similar to that observed in birefringent crystalline materials is introduced between the orthogonal polarization components. For organized linear structures, an increase in phase delay may be characterized by a difference (
noe ) in the effective refractive index for light polarized along, and perpendicular to, the long axis of the linear structures. The effect of tissue birefringence on the propagation of linearly polarized light is dependent on the angle between the incident polarization orientation and the tissue axis. Phase retardation (δoe ) between the orthogonal polarization components is proportional to the distance (d) traveled through the birefringent medium15,256 δoe =

2πd
noe . λ0

(69)

A medium of parallel cylinders is a positive uniaxial birefringent medium (
noe = (ne − no ) > 0) with its optical axis parallel to the cylinder axes [Fig. 45(a)]. Therefore, a case defined by an incident electrical field directed parallel to the cylinder axes will be called “extraordinary,” and a case with the incident electrical field perpendicular to the cylinder axes will be called “ordinary.” The difference (ne − no ) between the extraordinary index and the ordinary index is a measure of the birefringence of a medium comprised of cylinders. For the Rayleigh limit (λ
cylinder diameter), the form birefringence becomes15,37
noe = (ne − no ) =

f1 f2 (n1 − n2 )2 , f1 n1 + f2 n2

(70)

Tissue Structural Properties Studies at Optical Immersion

89

where f1 is the volume fraction of the cylinders, f2 is the volume fraction of the ground substance, and n1 and n2 are the corresponding indices. For a given index difference, maximal birefringence is expected for approximately equal volume fractions of thin cylinders and ground material. For systems with large diameter cylinders (λ  cylinder diameter), the birefringence goes to zero. For a system of thin plates [Fig. 45(b)], the following equation is valid:15,37 n2e − n2o = −

f1 f2 (n1 − n2 ) , f1 n21 + f2 n22

(71)

where f1 is the volume fraction occupied by the plates, f2 is the volume fraction of the ground substance, and n1 and n2 are the corresponding indices. This implies that the system behaves like a negative uniaxial crystal with its optical axis aligned normally with the plate surface. The dynamics of the polarization structure of the tissue image at immersion can be easily observed using an optical scheme with a “white” light source and a tissue sample placed between two polarizers, either in parallel or crossed. Figure 46 illustrates the evolution of polarization images during scleral optical clearing.37,83 In such experiments, a tissue layer “works” as a phase plate (or number of phase plates175,255 ), whose linear birefringence is spatially and temporally dependent. As scattering goes down with time due to refractive index matching, the birefringence of the fibrillar structure of the sclera affects the transmittance of the optical system. The spatial inhomogeneities of images may be due to spatial variations of the sample thickness and structure, which both may influence the efficiency of chemical impregnation and corresponding phase shift between the orthogonal optical field components.34 At reduction of scattering, the degree of linearly polarized light propagating in sclera is improving. This is clearly seen from the experimental graphs in Figs. 16 and 47.174 As far as tissue immersed the number of scattering events decreases correspondingly depolarization efficiency goes down, and the residual polarization degree of transmitted linearly polarized light increases. As a result, the dynamics of tissue average transmittance and polarization degree are similar (see Fig. 16). It follows from Figs. 16 and 47 that tissue optical clearing leads to the increase of the depolarization length.15,256–261 Due to less scattering of the longer wavelengths, the initial polarization degree is the highest for these. Polarization imaging is a useful tool for detection of subsurface lesions, but it is effective only at depths smaller than the depolarization length.258–260 Optical clearing may give a possibility to substantially increasing the depth of polarization imaging. Reduction of the scattering at optical immersion makes it possible to detect the polarization anisotropy of tissues more easily, and to separate the effects of light scattering and intrinsic birefringence on the tissue polarization properties. It is also possible to study birefringence of form with optical immersion, but when the immersion is strong, the average refractive index of the tissue structure is close to the index of the ground media, and the birefringence of form may be too small to

Figure 46 Polarization images of a scleral sample (white light source, crossed polarizers). Images from left to the right correspond to 4, 5, 6, 7, 8, 9, 9.5, and 10 min of tissue impregnation by Trazograph-60. The supporting wires of the sample are seen for the translucent tissue.37,83

90 Optical Clearing of Tissues and Blood

Tissue Structural Properties Studies at Optical Immersion

91

Figure 47 Time-dependent polarization degree, (I|| − I⊥ )/(I|| − I⊥ ), of collimated transmittance measured in vitro at different wavelengths for rabbit eye sclera at administration of 40% glucose.174

see because both phenomena based on the refractive index mismatch—scattering due to irregular refractive index variations, and birefringence due to regular ones [see Eqs. (6), (70), and (71)]. At a midlevel of refractive index matching,34 when n0 = 1.39, the measured birefringence of a scleral sample is
noe ∼ = 10−3 . The image contrast C(t) = B(t)/Bmax , where B(t) is the current sample brightness and Bmax is the maximal one, characterizing transmittance of linear polarized light through a tissue sample, was used for a quantitative evaluation of the diffusion process of an agent in a tissue. A white-light video-digital polarization microscope is suitable for the measurements.33 Sections of the various connective and vascular tissues of a thickness of 0.1–1.5 mm were studied. The immersion solution was heated up to (36–40)◦ C and simply dropped on the tissue sample surface. Both Figs. 48 and 49 show different rates of tissue optical clearing for different tissues when a tissue turns from initially turbid (multiple scattering) at t = 0 to a less depolarized and more transparent state (less scattering), C(t) → 1. Evidently, such a difference depends on tissue structure, which defines the initial (natural) turbidity and an efficiency of a chemical agent interaction with a tissue. For example, vein and aorta samples having approximately the same initial turbidity and degree of linear depolarization quite differently interact with the immersion agent (see Fig. 48)—denser aorta is less penetrative for the agent than vein; therefore, its action on aorta can be seen only in a few hours, whereas for vein sample about 10 min is enough to complete clearing.

92

Optical Clearing of Tissues and Blood

Figure 48 Experimental temporal dependencies for image contrast of linear polarized light transmitted through vascular tissue sections (aorta and vein-vena cava inferior ) measured by a white-light video-digital polarization microscope at the application of Trazograph-60.33,89

Figure 49 Experimental temporal dependencies for image contrast of linear polarized light transmitted through tissue sections measured by a white-light video-digital polarization microscope at the application of Trazograph-60.33,89

At reduction of scattering, tissue birefringence can be measured more precisely; in particular, birefringence of form and material can be separated. For example, for the translucent human scleral sample by its impregnation with a highly concentrated glucose solution (about 70%), the measured optical anisotropy34
n = (ne − no ) was equal to ≈10−3 . This is 1.5- to 4.5-fold less than for other

Tissue Structural Properties Studies at Optical Immersion

93

birefringent tissues,15 and mostly explained by the reduction of inclusion of birefringence of form at optical immersion. The additional measurements of the collimated transmittance allows one to estimate the refractive index of the ground substance of the translucent tissue n2 using the expressions from radiative transfer and Mie theories [see Eqs. (27) and (42)]. For human scleral sample impregnated by 70% glucose solution, n2 was evaluated to 1.39. Using this value and the value of the refractive index of hydrated collagen, n1 = 1.47 and
n = 10−3 , the collagen volume fraction f1 was calculated from Eq. (70) as f1 ∼ = 0.32, which well correlates with an estimation made in Ref. 37. Figure 16 illustrates the reversibility of the polarization immersion effect. A polarization-speckle microscope working in transmittance mode was used to carry out these measurements.85,171 The sample was irradiated by a linear polarized focused laser beam that was scanned along the trace of 1.5 mm on the sample surface to average the speckle modulation in the far zone, where the analyzer and photodetector were placed. Two orthogonal linear polarized components of the transmitted light were detected. It can be seen that the sample initially has poor transmittance with the equal intensity components I|| = I⊥ and that multiple scattering takes place. When the immersion agent acts in the 14th min, I|| prevails substantially over I⊥ , and the tissue becomes less scattering. The subsequent action of the physiological solution, which washes out the immersion agent, returns the tissue to its normal state, and it becomes turbid again after 22 min with no measured difference between the intensities of the orthogonally polarized components. The secondary application of the immersion agent again makes the tissue less scattering and more polarization sensitive with a maximum reached at 28 min. Figure 50 shows the reversible loss of turbidity and birefringence in rodent tail tendon observed at glycerol (13 M) application. The dark background in each of the images demonstrates the extinction of illuminating light at crossed polarizers in the polarized light microscope used in the measurements. Characteristic banding patterns observed in the tendon sample indicate ordered fibril organization. The distribution of pattern brightness corresponds to the distribution of a phase shift between the orthogonal optical field components [see Eq. (69)], and the background smooth brightness corresponds to light scattering. Loss of transmittance at the sample edges and appearance of bright spots in the middle of the sample in the course of glycerol action indicate refractive index matching of collagen fibers (not seen in the image due to their small diameter). The complete refractive index matching at the edge region happens earlier than in the middle of the sample, and turns tissue to lose scattering and birefringence completely in this region. In the middle region of the sample, refractive index matching is not completed and mostly scattering is reduced (loss of turbidity); thus, bright and dark areas corresponding to a certain phase shift are well seen. Tissue shrinkage at glycerol action due to tissue dehydration as well as hypothesized by authors of Ref. 101 the reversible dissociation of collagen fibers at glycerol action both may have some influence on the pattern formation. The rehydration of the tissue sample in saline makes fully visible its banding structure in the crossed polarizers due to resumption of the tissue birefringence and turbidity approximately to the initial states.

94

Optical Clearing of Tissues and Blood

Figure 50 Reversible loss of turbidity and birefringence in rodent tail tendon following glycerol (13 M) application observed by polarized light microscopy at crossed polarizers. (A) Before glycerol application: banding patterns observed in tendon indicate ordered fibril organization. (B) At glycerol application: loss of transmittance at the sample edges and bright spots in the middle indicate refractive index matching of collagen fibers; the complete refractive index matching at the edge region turns tissue to lose scattering and birefringence, as in the middle sample region refractive index is not completed and mostly scattering is reduced (loss of turbidity). (C) The tissue sample after rehydration in saline. (Figure was kindly presented by Alvin T. Yeh and Bernard Choi.)

Practically all healthy connective and vascular tissues show the strong or weak optical anisotropy typical of either uniaxial or biaxial crystals.33,34,175,255 Pathological tissues show isotropic optical properties.258–260

5.2 Confocal Microscopy The increase of the upper tissue layers’ transparency can improve the penetration depth, image contrast, and spatial resolution in confocal microscopy as well.99,100 Through Monte Carlo simulations of the point-spread function, it was shown that the signal spatial localization offered by a confocal probe in the skin tissues during their clearing is potentially useable for reticular dermis monitoring (Fig. 51).99 The results of the simulation predict that after 20 min of the chemical agent diffusion

Tissue Structural Properties Studies at Optical Immersion

95

after intradermal glycerol or glucose injection, signals from the tissues located twice as deep in the skin can be detected.

Figure 51 Point spread function for confocal skin probing during the 40% glucose diffusion (after intradermal injection). (a) 5 min diffusion; (b) 10 min diffusion; (c) 20 min diffusion.99

A significant improvement of the confocal microscopy signal at glycerol administration is well seen from theoretical axial profiles of a detected signal calculated for three different in-depth focusings (Fig. 52).100

5.3 Fluorescence Measurements 5.3.1 Transmittance of fluorescent radiation Recently, the improvement of the detected fluorescence signal traveling through skin in in vitro and in vivo experiments at topical application of hyperosmotic clearing agents, such as anhydrous glycerol (13 M, index n = 1.47) and pure DMSO (14 M, index n = 1.47), and a highly concentrated glucose (7 M, index n = 1.46), was demonstrated.104 Fluorescence measurements were performed for hamster dorsal skin with clearing agent applied to the subdermal side of the skin and rhodamine fluorescent film placed against the same skin side. Fluorescence was induced by a dye laser pulse at 542 nm, delivered to the skin epidermal side

96

Optical Clearing of Tissues and Blood

(a)

(b) Figure 52 The axial profile of a detector signal distribution (a) before and (b) 20 min after glycerol administration (intradermal injection) predicted by the numerical Monte Carlo simulation for a confocal microscope focusing at (1) 300 µm, (2) 600 µm, and (3) 900 µm into the skin. Confocal probe parameters are: lens diameter 5 mm and focal length 10 mm; pinhole diameter is 10 µm; the height of the lens above the surface is 9.7 mm.100

by a fiber bundle, and was detected by a collection fiber bundle from the epidermal surface at wavelengths longer than 565 nm. The skin flap window preparation in an area void of blood vessels was used for in vivo studies. Approximately equal enhancement of transmitted fluorescence was achieved for in vitro and in vivo measurements (Fig. 53). On average, up to a 100% increase in fluorescence intensity is seen after 20 min of glucose and glycerol application, and up to a 250% increase for DMSO. A significantly bigger increase in the case of DMSO is associated with its twice greater osmolarity than for glycerol and glucose of the concentrations used.

Tissue Structural Properties Studies at Optical Immersion

97

Figure 53 Comparison of the percent increase in fluorescent signal due to 100% glycerol, 100% DMSO, and 7M glucose in vivo.104

A significant enhancement of both the sensitivity (∼fivefold) and the spatial resolution (∼threefold) for low-level light-emitting probes (a broadband violetblue chemiluminescence with a center wavelength of 425 nm) was demonstrated in in vitro experiments with a 3 mm thick fresh porcine skin sample at topical application of 50% glycerol for a period of 30 min.107 A higher efficiency of luminescent light transportation through the skin at immersion in that case is connected with a higher initial scattering and absorption of skin at the shorter wavelengths. Refractive index matching effectively damps light scattering and thus absorption of light also is accomplished, due to a smaller number of photons circulating within a tissue. In a recent theoretical study,106 it was shown that by use of refractive index matching at the skin interface and a fiber-optical fluorescence probe, one can improve the contrast and spatial resolution of the shallow sampling volume. 5.3.2 In vivo measurements for intact human skin Both model experiments described above well demonstrated changes of tissue layer transmittance at optical immersion for light from a fluorescent source placed behind a tissue layer. However, fluorophores are more often distributed within a tissue layer, or even a multilayered tissue structure may contain a number of different fluorophores. In that case, the behavior of the fluorescence signal at tissue immersion is not so evident, because the cross section for fluorescence emission depends on the amount of the absorbed light by fluorescent centers. Such a cross section decreases as multiple scattering decreases. Thus, at tissue optical clearing, instead of enhancement of a fluorescence signal, one can see its damping. Evidently, that depends on the depth, where the fluorophore is, and what layer of the tissue is optically cleared. Figure 54 illustrates that fluorescence can be damped at tissue scattering reduction due to refractive index matching.105 These data were received in vivo for the

98

Optical Clearing of Tissues and Blood

Figure 54 In vivo dynamics of human skin autofluorescence intensity (λexc = 337 nm) after intradermal injection of 0.1 ml of Trazograph-76. Intensity values at different emission wavelengths are normalized by correspondent initial values.105

human skin at intradermal injection of the immersion liquid, Trazograth-76. Such behavior of the autofluorescence signal means that the main fluorophore (collagen) is in the dermis, where the immersion agent was inserted. However, with time due to more in-depth penetration of the exciting light and less attenuation of the induced fluorescence by the upper layers of the skin, fluorescence intensity is going up. Tissue optical clearing can be a helpful technology in looking for endogenous or exogenous fluorophores distribution within a tissue, and their differentiation. 5.3.3 Two-photon scanning fluorescence microscopy A new direction in laser spectroscopy of tissues and cells is associated with multiphoton (two-, three-photon) fluorescence scanning microscopy, which makes it possible to image functional states of an object or, in combination with autocorrelation analysis of the fluorescence signal, determine the intercellular motility in small volumes.263–266 The two-photon technique employs both ballistic and scattered photons at the wavelength of the second harmonic of the incident radiation coming to a wide-aperture photodetector exactly from the focal area of the excitation beam. A unique advantage of two-photon microscopy is the possibility of investigating three-dimensional distributions of chromophores excited with ultraviolet radiation in thick samples. Such an investigation becomes possible because chromophores can be excited (e.g., at the wavelength of 350 nm) with laser radiation whose wavelength falls within the range (700 nm) where a tissue has a high transparency. Such radiation can reach deeply lying layers and produces less damage in tissues. Fluorescent emission in this case lies in the visible range (>400 nm) and comparatively easily emerges from a tissue and reaches a photodetector, which

Tissue Structural Properties Studies at Optical Immersion

99

registers only the legitimate signal from the focal volume, without any extraneous background. In a two-photon excitation process, the rate of excitation is proportional to the average squared photon density. This quadratic dependence follows from the requirement that the fluorophore must simultaneously absorb two photons per excitation process. To demonstrate that a multiphoton excitation process has occurred, it is necessary to measure the intensity of the fluorescence as a function of the intensity of the excitation light. A two-photon excitation process is characterized by a slope of 2 on a log-log plot of measured intensities; a three-photon excitation process is characterized by a slope of 3. The rate of two-photon excitation is expressed as the number of photons absorbed per fluorophore per pulse and is a function of the pulse duration, the pulse repetition rate, the photon absorption cross section, and the numerical aperture of the microscope objective that focuses the light.263–266 The laser light in a two-photon excitation microscope is focused by the microscope objective to a focal volume. Only in this focused volume is there sufficient intensity to generate appreciable excitation. The low photon flux outside the focal volume results in a negligible amount of fluorescence signal. The origin of the optical sectioning capability of a two-photon excitation microscope is due to the nonlinear quadratic dependence of the excitation process and the strong focusing capability of the microscope objective. Most specimens are relatively transparent to near-infrared light. The focusing of the microscope objective results in twophoton excitation of ultraviolet absorbing fluorophores in a small focal volume. It is possible to move the focused volume through the thickness of the sample and thus achieve optical sectioning in three dimensions. Thus, the optical sectioning in a two-photon excitation microscope occurs during the excitation process. Investigations of tissues and cells by means of two-photon microscopy are characterized by the following typical parameters of laser systems: the wavelength ranges from 700 to 1000 nm, the pulse duration is 100–150 fs, the pulse repetition rate is 76–80 MHz, and the mean power is less than 10 mW (typically about 3 mW). Such parameters can be achieved with mode-locked dye lasers pumped by a Nd:YAG laser or with titanium sapphire lasers pumped by an argon laser. Diodepumped solid-state lasers also hold much promise for the purposes of two-photon microscopy. However, it has been shown that the effect of light scattering in multiphoton fluorescence scanning microscopy is to drastically reduce penetration depth to less than that of the equivalent single-photon fluorescence while largely leaving the resolution unchanged.266,267 This happens mostly due to excitation beam defocusing (distortion) in the scattering media. Although some improvement in the penetration depth of two-photon microscopy can be obtained by optimizing the pulse shape and repetition rate for the sample under investigation,268 reduction of scattering is believed to be more effective in penetration depth and image contrast improvement.269 Two-photon fluorescence microscopy provides high-resolution images of human skin in vivo.270,271 Evidently, the technique is applicable for many other tissues, but normally its penetration depth is limited to 20–30 µm.

100

Optical Clearing of Tissues and Blood

The first demonstration of two-photon in-depth signal improvement using the optical immersion technique with hyperosmotic agents, such as glycerol, propylene glycol, and glucose, was done by the authors of Ref. 269 in ex vivo experiments with human dermis. Thick (150 µm) slices of dermis excised during plastic surgery were imaged within the same day. Images were collected in stacks, each comprising four images of 100 µm2 area taken at depths of 20, 40, 60, and 80 µm from the surface of the sample. Before data acquisition, the sample was immersed in 0.1 ml of phosphate buffered saline (PBS) in order to prevent drying and shrinkage. Then, the sample was immersed in 0.5 ml of an optical clearing agent (OCA) and one image stack was acquired every 30 s for 6–7 min. Finally, the OCA was removed and the sample was immersed again in 0.1 ml of PBS, in order to observe the reversibility of the clearing process (see Fig. 16). Glycerol and propylene glycol were both used in anhydrous form, and glucose as a concentrated aqueous solution (5 M). The upper limit of tissue shrinkage was estimated as 2% in the course of 6–7 min of OCA application. The average contrast in each image and relative contrast (RC) were defined as269 Contrast =

N lines
i,j =1

  Iij − Iij ,

RC(%) = 100

Contrast[OCA] − Contrast[PBS] , Contrast[PBS]

where Iij is the mean intensity of the nearest eight pixels and Nlines = N − 2, with N = 500; Contrast [OCA] and Contrast [PBS] are calculated for OCA and PBS immersion, respectively. Contrast, as defined here, is linearly dependent on the fluorescence intensity and varies according to structures in the image. Hence, its usefulness is primarily to enable comparison between images of the same sample at the same depth, maintaining the same field of view. Normalization to the total intensity would be required in order to compare different images. The relative contrast RC also serves for the purpose of comparison. Figure 55(A) shows two typical image stacks: the first received for a sample immersed in PBS and the second received 7 min after application of glycerol. The images show connective tissue in human dermis, which is primarily composed of collagen and elastin fibers. The enhancement of the contrast, as well as increasing of penetration depth (from 40 to 80 µm) and total intensity [i.e., the intensity summed over all pixels, Fig. 55(B)], are clearly seen from the images. The corresponding absolute and relative contrast levels are plotted in Figs. 55(A) and 55(D). The RC has a value of 215% at 40 µm and dramatically increases with increasing depth. The effect on deeper layers is greater because of the cumulative effect of the reduction in scattering in the superficial layers of the tissue sample, which provides less attenuation of the incident and detected fluorescent light. The contrast is also dependent on fluorescence intensity, which is proportional to squared intensity of the excitation intensity and mostly dependent on excitation beam focusing ability. Better focusing (less focused beam distortion) is achieved in less scattering media.

Tissue Structural Properties Studies at Optical Immersion

101

Figure 55 Two-photon microscopy of human skin ex vivo by use of glycerol as optical clearing agent.269 (A) Image stackes for a skin dermis sample immersed in PBS (upper) and after immersion for 7 min in glycerol (lower). The corresponding total intensity, contrast, and relative contrast are plotted in (B), (C), and (D), respectively. Reprinted with permission.

The contrast verses time and depth for all three OCA used in the experiments is presented in Fig. 56. In each case, the sequence PBS-OCA-PBS was followed with the measuring interval of 30 s per step. This figure shows that application of each OCA resulted in a contrast enhancement with varying degrees of efficiency and saturation. The dynamics and the final contrast level attained depend on the OCA and on the depth. Saturation of contrast occurs most rapidly in superficial layers of the sample. This is consistent with a diffusion model for the penetration of the agent from the surface into the tissue [see Eqs. (29), (33), and (34)], i.e., if the contrast is proportional to agent concentration, then the saturation time at a given depth will be proportional to the depth. As it follows from Fig. 56,∗ glycerol is the most efficient with respect to saturation level (RC = 49.7% at 20 µm depth, ∼304% at 40 µm depth, ∼1900% at 60 µm depth, and ∼9260% at 80 µm depth), but also the slowest. Propylene glycol provides RC ∼ 64% at 20 µm depth, ∼1090% at 40 µm depth, ∼5640% at 60 µm depth, and ∼447% at 80 µm depth. Whereas, glucose (5M) is the worst with RC = 10.9% at 20 µm depth, ∼134% at 40 µm depth, ∼471% at 60 µm depth, and ∼406% at 80 µm depth, but diffuses ∗ Numerical data for RC values are provided by D. D. Sampson.

102

Optical Clearing of Tissues and Blood

Figure 56 Two-photon microscopy of human skin ex vivo by use of optical clearing agents.269 Contrast versus incubation time and depth for immersion in (A) glycerol, (B) propylene glycol, and (C) glucose, preceded and followed by application of PBS. In the right column, the results obtained at 60 and 80 µm depths are presented on magnified scales. Reprinted with permission.

Tissue Structural Properties Studies at Optical Immersion

103

three times faster than glycerol and five times faster than propylene glycol. Diluted agents gave similar tendencies in contrast enhancement and increase of penetration depth, providing higher efficiency in both characteristics with OCA concentration increase. These data illustrate that, as well as in linear spectroscopy, refractive index matching is the leading mechanism in reduction of tissue scattering and two-photon signal improvement. In contrast to in vivo single-photon fluorescence spectroscopy (see Fig. 54), where fluorescence intensity may decrease at multiple scattering decrease, the two-photon tomography signal is always increased due to less distortion of the focused beam and less attenuation of a two-photon fluorescence signal by superficial optically cleared tissue layers. However, some specificity in action of three different OCAs was found. The results presented in Fig. 56 show, for propylene glycol and glucose, a slowing in the rate of contrast increase following addition of PBS rather than a decrease as it is seen for glycerol. Such behavior may be associated with a lower inclusion of the dehydration mechanism in optical clearing for propylene glycol and glucose, and higher amount of these agents diffused into a tissue in comparison with glycerol.

5.4 The Second Harmonic Generation Optical clearing seems to be a promising technique for improvement of detected signals in multiphoton microscopy, nonlinear spectroscopy, and imaging of tissues.101,191 On the other hand, these techniques might be useful in the understanding of molecular mechanisms of tissue optical clearing at immersion and dehydration. Collagen, as a main component of connective tissues, has an appreciable nonlinear susceptibility for second-harmonic generation (SHG). The helix of collagen secondary structure is noncentrosymmetric, satisfying a condition for SHG, which self-assembles into higher-order structures. Collagen has been shown to have a dominant uniaxial second-order nonlinear susceptibility component aligned along the fiber axis.272,273 In such multicomponent tissues as skin, SHG light is generated mostly within dermis, not in cellular layers like epidermis or subcutaneous fat. SHG techniques have a number of advantages connected with incident wavelength dividing and selectivity to tissue structure, which allow one to easily reject surface reflection and multiple scattering of the incident light in the epidermis layer without any gating technique. In addition, SHG polarimetry is an effective tool to probe collagen orientation in the tissues.191 Evidently, due to optical clearing, less scattering in the epidermis for incident long-wavelength light (800 nm) and especially for backward SHG shortwavelength light (400 nm) may improve SHG images of dermis collagen structures. At 100% glycerol application to rodent skin dermis and tendon samples, as well as to engineered tissue model (raft), a high efficiency of tissue optical clearing was achieved in the wavelength range from 400 to 700 nm, but the SHG signal was

104

Optical Clearing of Tissues and Blood

significantly degraded in the course of glycerol application—and it was returned back to the initial state after tissue rehydration by application of saline.101 The loss of SHG signal in Ref. 101 is associated with the collagen fibers’ reversible dissociation and corresponding loss of fibril organization at glycerol action. Such an explanation is somewhat contradictory, because less organization of collagen fibers will lead to higher scattering.37 Since the significant effect of optical clearing at glycerol application is tissue dehydration, the explanation following from the data of Ref. 191 seems to be more adequate. Using reflection type SHG polarimetry, it was shown in Ref. 191 that SHG polarization signal (SHG radar graphs) for chicken skin dermis was almost unchanged (Fig. 57) and the SHG intensity was decreased to about a fourth at tissue dehydration. Authors have hypothesized that the decrease of the SHG intensity results in a change in linear optical properties, i.e., scattering efficiency (reduction of the SHG photon recycling), rather than that of efficiency of SHG radiation in the tissues. As it follows from Fig. 57, the tissue fixation process also indicates almost unchanged SHG polarization radar graphs while SHG intensity was slightly increased. Since formalin fixing induces crosslinking of collagen in tissues, this result may imply that the cross-linking does not

Figure 57 SHG radar graphs received for native samples of chicken dermis;191 (a) and (b) are SHG signal distributions before and 5 hr after formalin fixation, (c) and (d) are those before and 13 hr after air drying.

Tissue Structural Properties Studies at Optical Immersion

105

affect collagen orientation, but essentially contributes to the efficiency of the SHG signal.191 These two examples well illustrate the dependence of the SHG signal on light scattering of the sample, which was decreased at tissue dehydration and increased at tissue fixation. Thus, to study tissue structure (collagen orientation) using SHG, one of the methods providing light scattering suppression may be applied, such as SHG polarimetry191 or the optical immersion technique.

6 Optical Clearing of Blood, Cells, and Microorganisms

6.1 Blood Flow and Microvessel Network Imaging Small blood microvessels can be clearly identified visually by the naked eye in in vivo study of hamster45,109 and rat110 skin, when a transparent window in the skin was created by glycerol drops applied to the subdermal side of a native hamster dorsal skin flap window preparation,109 or by intradermal injection of glycerol45,110 or 40% glucose110 (see Fig. 34). In in vitro study of fresh human fat tissue with a topical application of a propylene glycol (PG) solution (50–80% with pure water), blood vessels are also seen.138 In order to have a quantitative sense of fat tissue optical clearing, the corresponding OCT images were registered (see Fig. 58). Two images of a fat sample without and with topical application of PG solution were captured. It is well seen that OCT techniques are very capable of visualizing the fat cells because of their sizes, which are within the limits of the OCT system spatial resolution. By comparison of images of Figs. 58(a) and 58(b), the light penetration enhancement is evident. Particularly, the blood vessel wall is clearly visualized after the topical application of PG solution, whereas without application of chemical agents, the light is almost blocked by the blood vessel. This is due to the fact that the blood inside the vessel highly scatters the incoming light, which degrades the imaging performance. Besides more precise visualization of the vessel network, immersion agents may have influence on blood microvessels’ functioning45,109,110,143,183 that gives the possibility to control functioning of tissue within the area of agent action. Functioning of microvessels of rat mesentery under the topical action of glycerol and glucose was described.110 A topical application of 75% glycerol during the initial period of 1–3 s led to a slowing down of blood flow in all microvessels (arterioles, venules, and capillaries). After 20–25 s, the stasis appeared and vessels were dilated by 30% on average, intravascular hemolysis took place; to 1 min after agent application, diameters of vessels were increased still more, to 40%. To the 6th min, stasis was maintained in all vessels, but the diameters of vessels were slightly decreased. Such changes of microcirculation were only local within the area of glycerol application. The topical application of glucose also decreased blood flow velocity in microvessels. For example, at the introduction of 40% glucose on a venule with a diameter of 11 µm and with initial an flow rate of 1075 µm/s, the flow rate decreased to 510 µm/s at 3 s after the glucose application, and to 202 µm/s at 5 s. Similarly with the action of glycerol, there were dilation and stasis of blood 107

108

Optical Clearing of Tissues and Blood

(a)

(b) Figure 58 OCT images of the sample of fresh human fat tissue where the blood vessel is clearly visible: (a) before and (b) after topical application of propylene glycol solution.138 Measurements were done using the OCT system presented in Fig. 38.

flow in all vessels (arterioles, venules, capillaries, and shunts) within 20–30 s, but no intravascular hemolysis was found and only RBC aggregates in the lumen of microvessels were seen. The strength of vessel dilation was more than that for glycerol: the mean diameter increased by 30% to the 30th s after glucose application, but to the 4th min it rose on average by 2.5 times. From the 3rd to 5th min, blood flow appeared again in a few microvessels, the velocity of reflow was markedly slower than in the control. The changes in blood flow were also local, but with a larger area than for glycerol, approximately 1 × 1 cm2 ; there were no disturbances in the functioning of blood microvessels in the other parts of the mesentery. Evidently, the decrease in glucose concentration and corresponding loss of the agent hyperosmotic property led to a more gentle glucose action on blood circulation; in particular, no blood stases were observed for 20% glucose, blood flow in all vessels after 3–4 min of glucose application was not significantly different from the initial one. The vasculature under the dura mater also became visible after the treatment of glycerol in in vivo rabbit experiment (Fig. 59).143,183 The reflectance decreased as a function of time of glycerol action, which proved the visual observation. The dura mater nearly recovered to native condition after 1 min. Velocity images of in vivo cerebral blood flow (CBF) under the effect of glycerol are shown in Fig. 60. Glycerol was applied around the exposed area. When glycerol diffused in brain tissue and influenced CBF under the dura mater, CBF in the exposed area would also change. Figure 60 illustrates the spatio-temporal characteristics of CBF changes under the treatment of glycerol. Under the action

Optical Clearing of Blood, Cells, and Microorganisms

109

Figure 59 Visual changes of image contrast of blood vessels for in vivo rabbit dura mater before and after treatment with glycerol epidurally: (a) native dura mater; (b) 50 s application of glycerol, bar = 1 mm.143 Reprinted with permission of Springer Science and Business Media.

of glycerol, blood flow first decreased while the blood vessels underneath the dura mater became increasingly visible. Then blood flow increased to near baseline; at the same time the turbidity of the dura mater returned. Figure 61 gives the time course of changes in four different vessels, which are expressed as the ratio of the measured velocity in the conditions of treatment with glycerol to that of the control condition. Vessel 2 is an arteriole. Vessels 1, 3, and 4 are veinules. Blood flow in vessel 2 (arteriole) began to decrease after 20 s of glycerol application, while that in the other vessels (veinules) decreased immediately after application with glycerol. The blood flow in vessel 1 decreased slower than that in the other vessels, which suggested that blood flow in the arteriole had a different response from that in the veinules. Blood flow in all vessels decreased to 70%–80% of baseline after treatment with glycerol. An example of the subdermal side of a native hamster dorsal skin flap window preparation is shown in Fig. 62(a). The main arteriole (A) is 97 ± 8 µm in diameter (lumen) and the main venule (V) is 188 ± 21 µm in diameter. The diameters of the branches, a and v, are 92 ± 18 µm and 181 ± 21 µm, respectively. Figure 62(b) shows the blood vessels in the same window preparation 10 min after the application of 100% glycerol. The smallest branches of the arterioles and venules can now be seen in the image. This is likely due to the increased clearing of the tissue overlying the vessels and could also occur with vasodilation. The venule branch, v, is dilated to 259 ± 19 µm. The main vessels and the arteriole branch (a), however, are not noticeably dilated. After 20 min, the main venule branch in the window preparation appears very dark and is occluded [Fig. 62(c)]. The diameters of the main vessels are 97 ± 18 mm (A), and 189 ± 20 mm (V), and the diameters of the branches are 141 ± 17 mm (a) and 259 ± 21 mm (v). For glycerol as a clearing agent, these findings were confirmed in experimental studies performed using laser speckle imaging of blood flow on golden Syrian hamsters with surgically installed dorsal skin fold window.274 However, surprisingly, the authors of Ref. 274 have found that such a clearing agent as a mixture

110

Optical Clearing of Tissues and Blood

Figure 60 Blood-flow images following the epidurally applied glycerol around the exposed area of in vivo dura mater. (a) The white-light image of the interest area. (b)–(h) Blood flow maps expressed as measured velocity, which is proportional to the blood flow velocity, during the treatment with glycerol and represented by images at the time points shown in Fig. 61. (b) Imaged blood flow before the application of glycerol (control), four vessels are indicated. (c) 10 s application of glycerol, no obvious change in blood flow was observed. (d) 20 s application of glycerol, blood flow began to decrease. (e) 30 s application of glycerol, the blood vessels underneath the dura mater began to be clear. (f) 40 s application of glycerol, blood flow decreased and the transparency of the surrounding dura mater increased. (g) 50 s application of glycerol, more blood vessels could be seen through the dura mater and the blood flow decreased significantly. (h) 70 s application of glycerol, blood flow increased and the dura mater became turbid again. Bar = 1 mm.143,183 Reprinted with permission of Springer Science and Business Media.

Optical Clearing of Blood, Cells, and Microorganisms

111

Figure 61 The time course of change in relative blood flow in vessels 1, 2, 3, and 4, which are indicated in Fig. 60(b), before and after the application of glycerol epidurally. After 20 s, the blood flow in vessel 2 (arteriole) began to decrease, while blood flow in the other vessels (veinules) decreased immediately after the application of glycerol. Decreases of blood flow in these vessels were 20–30% of the baseline. b, c, d, e ,f, g, and h denote the time points of corresponding images in Figs. 4.50 (b), (c), (d), (e), (f), (g), and (h).143,183 Reprinted with permission of Springer Science and Business Media.

Figure 62 Images of the subdermal side of in vivo hamster dorsal skin flap window preparation.109 (a) Native, the main arteriole (A) is 97 ± 18 µm in diameter (lumen) and the main venule (V) is 188 ± 21 µm in diameter; the diameters of the branches, a and v, are 92 ± 18 µm and 181 ± 21 µm, respectively. (b) 10 min after the application of glycerol (100%); (c) 20 min after the application of glycerol, the diameters of the main vessels are 97 ± 18 mm (A), and 189 ± 20 mm (V), and the diameters of the branches are 141 ± 17 mm (a) and 259 ± 21 mm (v). Scale bar: 0.25 cm.

of prepolymers PPG/PEG, in spite of its hyperosmotic properties, does not induce blood flow alterations. Optical clearing of vascularized tissue may have some important biomedical applications connected with investigation of vascular system structure and functioning, including relation of diameters of arterioles and venules, capillary density, bifurcation angles, etc. These parameters can be important in physiology and therapy for the diagnosis and treatment of some diseases (vascular disease, cancer, etc.). On the other hand, the optical clearing effect coupled with temporary and local cessation of blood flow in microvessels in the area of treatment may help vascular photothermal therapy significantly.109

112

Optical Clearing of Tissues and Blood

6.2 OCT Measurements of Blood Optical Clearing Refractive index mismatch between erythrocyte cytoplasm and blood plasma causes strong scattering of blood that, for example, prevents getting high quality images of intravascular structures through whole blood. The refractive index of erythrocyte cytoplasm is mostly defined by hemoglobin concentration (blood hematocrit).16 Hemoglobin oxygenation67 and glycation68,69 may have some role in refractive index mismatch. The scattering properties of blood are also dependent on erythrocytes’ volume and shape, which are defined by blood plasma osmolarity,16,176 and aggregation or disaggregation ability.47,112,177,178,275–277 Recently, the feasibility of index matching as a method to overcome the limited penetration through blood when obtaining OCT tissue images has been demonstrated for circulating, steady-state, or sedimenting blood.47,111,112,177,178 Glucose, low- and high-molecular dextrans, x-ray contrast, glycerol, and some other biocompatible agents were used to increase the refractive index of blood plasma closer to that of the erythrocyte cytoplasm to improve penetration depth of OCT images. The OCT and other noninvasive imaging techniques, based on backreflectance, polarization-sensitive, fluorescence, multiphoton, and Raman spectroscopies, which already have witnessed widespread and exciting application in biomedical diagnostics,14,15,29,30,213,254,256,278–285 may have additional advantages in the early diagnosis of vascular disease through optically clarified blood. Normal human blood is a scattering system that consists of about 43% by volume of scattering particles (99% red blood cells (RBC), 1% leukocytes and thrombocytes) and about 57% by volume of plasma.16,286 Under normal physiological condition, hematocrit ranges from 36.8% to 49.2%.16 Propagation of light in such a medium can be studied within the model of light scattering and absorption by an individual particle, taking into account the interparticle correlation effects and polydispersity. As it was noted, the blood plasma osmolarity is an important factor in changes in the scattering properties of blood.16,47,49 The effects of glucose, glycerol, Trazograph, and propylene glycol, which are hyperosmotic agents, led to significant change of blood plasma osmolarity. The change in osmolarity induces a variation of the RBC volume due to water exchange and, therefore, has an impact on the hemoglobin concentration within the RBC and consequently on their refractive index. It was demonstrated that glucose solution with concentration less than 20% led to an increase of blood scattering due to the osmotic dehydration of erythrocytes.49 Significant optical clearing was obtained at glucose concentrations higher than 40%, but such concentrations can cause erythrocytes aggregation.47 The result of the OCT study is the measurement of optical backscattering or reflectance, R(z), from the RBCs versus axial ranging distance, or depth, z. The reflectance depends on the optical properties of blood, i.e., the absorption µa and scattering µs coefficients, or total attenuation coefficient µt = µa + µs . The relationship between R(z) and µt is, however, highly complicated because of the high and anisotropic scattering of blood; but for optical depths less than 4, reflected

Optical Clearing of Blood, Cells, and Microorganisms

113

power can be approximately described by Eq. (66). Optical depth is a measure of depth in terms of the number of mean free path lengths, i.e., µs z.α(z) is linked to the local refractive index and the backscattering property of the blood sample. If α(z) is kept constant, as for a laminar blood flow for circulating blood or measurements before sedimentation process begins for noncirculating blood should be provided, µt can be obtained theoretically from the reflectance measurements at two different depths, z1 and z2 [see Eq. (67)]. Optical clearing (enhancement of transmittance)
T by an agent application can be estimated using Eq. (68), where Ra is the reflectance from the backward surface of the vessel within a blood sample with an agent, and Rs is that with a control blood sample (whole blood with saline). The 1300 nm OCT system was used for taking images of the reflector through circulating blood in vitro.111 As immersion substances, dextran (group refractive index, 1.52) and IV contrast (group refractive index, 1.46) were taken. The system allows the blood to be circulated in vitro through transparent tubing to reproduce coronary flow. Human whole blood was obtained from discarded blood bank samples and diluted to a hematocrit of approximately 35% with normal saline. Blood was pumped through a closed system of tubing by a perfusion pump (Sigma). The flow rate was 200 ml/min, which is approximately the peak flow in the coronary artery. The diameter of the tubing was 6 mm, approximately the diameter of a normal adult coronary. A reflector was placed in the tubing; the section of the reflector imaged is approximately 2 mm below the inner surface of the tubing. Once blood was introduced into the system and circulated, OCT imaging was performed on the reflector. The total intensity of the signal off the reflector is used to represent penetration [see Eq. (68)]. The more light that is scattered by blood, the less the signal is off the reflector. After baseline data had been obtained with blood, test substances were added to the blood. The test substances were dextran (0.25 g/ml in normal saline), IV contrast, or normal saline.111 Hematocrit and RBC concentrations were measured before and after the experiments. All substances added had a volume of 40 ml, which was added to a total volume of 260 ml. For the saline control, a 7 ± 3% increase in signal intensity was noted, which was not a statistically significant effect. A 69 ± 12% increase in
T was noted for dextran, which was statistically different from the saline control (p < 0.005). For the IV contrast, a 45 ± 4% increase was noted, which was also significantly different from the control (p < 0.001). By OCT imaging in the presence of saline, blood (Hct 35%), or lysed blood (Hct < 1%), it was directly demonstrated that RBC intracellular/extracellular mismatch, and not membranes or hemoglobin absorption, is the main source of nearinfrared attenuation by blood. In the presence of blood, the reflector was difficult to locate. However, when the RBCs were lysed, signal intensity returned to values not significantly different from saline. The fact that the cell membrane is not the major source of scattering is not surprising since it is too small relative to the wavelength to significantly scatter (see Figs. 2 and 3).

114

Optical Clearing of Tissues and Blood

In the case of dextran, the effect was consistent with index matching. With the IV contrast, a small but significant decrease in RBC volume was noted by a decrease in hematocrit, but not the RBC number relative to the saline control. Therefore, some improvement in penetration may be due to a reduction in cell volume. The lack of improved penetration with the addition of normal saline (40 ml) is consistent with data of Ref. 16 that suggested dilution of the hematocrit to below 10% was necessary before significant improvement in penetration was seen. Studies of blood scattering reduction by the immersion technique using various osmotically active solutions, which are biocompatible with blood, like saline, glucose, glycerol, propylene glycol, Trazograph (x-ray contrast), and dextrans were also described.47,112,177,178 The 820 and 1310 nm OCT systems were applied for taking images of the reflector through a layer of uncirculating fresh whole blood. One of the experimental setups is presented in Fig. 57. The OCT system used yields 12 µm axial resolution in free space. This determines the imaging axial resolution, which is comparable with the dimensions of RBCs or small aggregates. It was shown that for uncirculating blood, the sedimentation may play an important role in blood clearing using the immersion technique and OCT allows for precise monitoring of blood sedimentation and aggregation. Venous blood was drawn from healthy volunteers and stabilized by 9NC coagulation sodium citrate 3.2% or by K2E EDTA K2. The major blood samples were prepared immediately after blood taking by gently mixing blood and agent (for agents in the liquid state) or agent-saline solution (for agents in the solid state) with low-rate manual rotating for 1 min before each OCT measurement. A few samples were stored before measurements up to 24 hr after blood taking. Four groups of the blood samples with various hematocrit values were investigated in the study (see Table 5).47,112,177,178 A few different glass vessels of 0.2–2 mm thickness were used as blood sample holders. For some holders, to enhance reflection from the bottom interface, a metal reflector was used. The sample holder was mounted on a translation stage at the sample arm and was placed perpendicular to the probing beam. The amplitude of reflected light as a function of depth at one spatial point within the sample was obtained. The result is the measurement of optical backscattering or reflectance, R(z), from the RBCs versus axial ranging distance, or depth, z, described by Eq. (66). The total attenuation coefficient µt and optical clearing (enhancement of transmittance)
T by an agent application were estimated using algorithms described by Eqs. (67) and (68), respectively. Averaging for a few tenths of z-scans was employed. The scattering µs and reduced scattering coefficient µ
s of blood depend on mismatch of the averaged refractive indices of blood plasma and erythrocyte cytoplasm. The ratio nRBC /nbp ≡ m determines the scattering coefficient, where nRBC is the mean refractive index of erythrocyte cytoplasm and nbp is the mean refractive index of the blood plasma. For the model of the RBC ensemble as a monodisperse system of noninteracting scattering dielectric spheres of radius a irradiated at a NIR wavelength λ, when 5 < 2πa/λ < 50, the anisotropy scattering factor, g > 0.9, and 1 < m < 1.1, µ
s is described by Eq. (6).

Saline Dextran A Dextran B Dextran C

56.5% (control) 6.5% + 50% saline

Saline Trazograph-60 Propylene glycol Glycerol Glucose Dextran A Dextran B Dextran C

56.5% (control) 1.62 g/dl

1.62 g/dl

35% (control) 2.43 g/dl

Concentration (volume % or g/dl) 13% (control) 13% 0.325 g/dl

Saline Dextran A Dextran B Dextran C

Agent Saline Glycerol Dextran

17.4

17.4

26

Hema-tocrit (%) 35

36.5 29.5 30.0 32.5

42 26 26 20 57 43 54 58

37.1 (1.3) 38.2 (2.4) 29.7 (3.6) 31.2 (1.8)

µt (cm−1 ) 61 (3) 51 (5) 55 (5)

52.1 110.6 150.5

–

– 45.3 77.2 117.1 45.3 47 44.6 20.5

– 11.9 (8.3) 100.1 (20.2) 86.7 (29.1)


T (%) – 13.7–95.0 20.2–78.4

Male volunteer, 35 yr old47

Female volunteer, 35 yr old47,177

Male volunteer, 24 yr old172

Comments From 1 to 24 hr after taking blood; dextran sulfate, M ≈ 500,000177

Table 5 Total attenuation coefficient and enhanced transmittance
T (%) of blood samples diluted by saline and added agents; pH for all solutions was approximately 7.5; dextran A (M = 10,500); dextran B (M = 65,500), and dextran C (M = 473,000).

Optical Clearing of Blood, Cells, and Microorganisms 115

Saline Dextran A Dextran A Dextran A Dextran B Dextran B Dextran B Dextran C Dextran C

Agent Saline Dextran A Dextran B Dextran C

80% (control) 1 g/dl 5 g/dl 10 g/dl 1 g/dl 5 g/dl 10 g/dl 0.5 g/dl 5 g/dl

Concentration (volume % or g/dl) 56.5% (control) 1.62 g/dl

8

Hema-tocrit (%)

13.5 17.5 (0.9) 14.3 (1.2) 12.2 (1.8) 14.2 (1.5) 13.0 (2.8) 11.5 (1.3) 10.0 (1.6) 13.3 (0.7)

µt (cm−1 ) 25.6 (1.6) 22.5 (2.4) 19.0 (3.8) 14.3 (4.3) 11.4 (6.2) 11.3 (3.3) 49.4 (12.1) 21.1 (5.4) 49.0 (26.2) 76.8 (21.2) 106.3 (39) 67.0 (5.8)

–


T (%) – 20.5 (4.2) 44.5 (3.4) 47.0 (9.7)

Table 5 (Continued).

Male volunteer, 36 yr old, hemoglobin: initial—175 g/l, diluted—37 g/l, Ref. 112

Comments Male volunteer, 23 yr old112

116 Optical Clearing of Tissues and Blood

Optical Clearing of Blood, Cells, and Microorganisms

117

The blood plasma contains up to 91% of water, 6.5–8% (about 70 g/l) various proteins, and about 2% of low molecular compounds. Because of the low concentration and relatively low refractive index of low molecular chemical compounds, the mean blood plasma (background) index can be estimated as the weighted average of the refractive indices of water (92%), nw , and proteins (8%), np : nbp = fw nw + (1 − fw )np ,

(72)

where fw is the volume fraction of water contained in plasma; nw = 1.329 at 800 nm and the index of proteins can be taken as np = 1.470.65 Since approximately 92% of the total plasma is water, it follows from Eq. (72) that nbp = 1.340. The following empirical formula can be used for estimation of the blood plasma index in the wavelength range from 400 to 1000 nm:49,64 nbp (λ) = 1.3254 + 8.4052 × 103 λ−2 − 3.9572 × 108 λ−4 − 2.3617 × 1013 λ−6 . (73) The refractive index of erythrocyte cytoplasm, defined by the cell-bounded hemoglobin solution, can be found from Eq. (59). As it follows from Eq. (6), about a tenfold reduction of the scattering coefficient µ
s is expected when the refractive index of the blood plasma is changed from nbp = 1.340 to 1.388 and the refractive index of RBC cytoplasm is kept constant, nRBC = 1.412 (for hemoglobin concentration in cytoplasm of 400 g/l).16 For slightly diluted blood (hematocrit ∼35%), optical clearing was found only for dextran of molecular weight M = 500,000, with concentration in a blood sample of 0.325 g/dl and glycerol with volume fraction of 13%. Values of
T , characterizing optical clearing, were from 20.2% to 78.4% for dextran and from 13.7% to 95% for glycerol, depending on time of blood sample storage. The minimal and maximal values have been found for blood samples that were stored after taking the blood for a short (1–3 hr) and for a long (24 hr) time interval, respectively. For the time interval of 4–6 hr of blood storage,
T = 46.5% for dextran and 74.5% for glycerol. Evidently, at high concentrations of RBC in a sample, interaction of used agents with blood significantly depends on the physicochemical parameters of blood, which may change after prolonged storage. Thus, all other measurements were done as fast as possible after taking the blood. For 56.5% diluted blood by a saline, the blood samples with Trazograph-60, propylene glycol, and glycerol had a lower total attenuation coefficient than the control (Table 5). Optical clearing
T was from 45.3% to 117.1% as measured immediately after mixture, when sedimentation is not critical for the optical properties of the blood layer. The minimal attenuation (approximately one half of that for the control) and the maximal enhancement of transmittance (
T = 117.1%) were found for the application of glycerol. Propylene glycol is also a good enhancer of blood transmittance (
T = 77.2%). The effects of glucose, glycerol, Trazograph-60, and propylene glycol, which are hyperosmotic agents, on the scattering properties of human and animal tissues have been demonstrated in in vitro and in vivo studies, and presented in this chapter. The blood study under discussion

118

Optical Clearing of Tissues and Blood

also shows the possibility of increasing the transmittance and reducing the scattering coefficient in blood after application of the agents, assuming that the agents cause no change in the absorption coefficient. The hypothesis that the agents reduce light scattering primarily as a result of better index matching with ground substance is supported by the results presented in Table 5. Similar effects of increase in transmittance and decrease in scattering were demonstrated by use of dextrans of various molecular weights. Table 5 shows that all three used dextrans; A, B, and C reduced the amount of the attenuation (scattering) coefficient in blood with respect to saline. Optical clearing
T was in the range from 52.1% to 150.5%. The dextran with the highest molecular weight appeared to have a much stronger effect on the increase in transmittance immediately after mixing. A blood sample mixed with an agent of higher refractive index, for example, dextran C, had a higher reflectivity from the metal surface than did agents such as saline (control) and dextran A with lower refractive indices. The results support well the hypothesis that the refractive index matching effect is important for clearing of 50% diluted blood. It can be seen from Table 5 that dextrans C and B at concentration of 2.43 g/dl in 35% diluted blood are effective agents to decrease the light attenuation of blood compared to the saline control, with the total attenuation coefficient decreased from 37.1 cm−1 for the saline control to 31.2 cm−1 and 29.7 cm−1 , respectively. The optical clearing capability
T was approximately 90% and 100% for dextran C and B, respectively. It is interesting to note that dextran C, providing a higher refraction, had less effect than that of dextran B at the same concentration. Moreover, the increase in concentration (refraction power) cannot always achieve higher optical clearance. The 0.5 g/dl dextran C had a stronger effect than 5 g/dl in 20% blood with 80% saline samples. The changes in the scattering property brought above by the addition of dextran solution may first be explained by the refractive index matching hypothesis. It can be seen that scattering can be reduced when the refractive index of plasma is increased. The refractive index of dextran saline solution was increased with concentration in all molecular weight groups (see Table 1). The measured indices of blood samples with dextrans were in good agreement with the theoretical values calculated according to the equation n = cb nb + (1 − cb )nsaline , where cb is the volume fraction (20%) of whole blood in the diluted sample and nsaline is the index of saline with or without dextrans. As expected, the refractive index of blood with dextran increases as the concentration of the added dextran increases due to an increase of the index of the ground matter of the sample. It should be noted that the total attenuation coefficient for glucose and dextrans was not changed significantly with respect to the control; nevertheless, transmittance enhancements of 45% for glucose and of 52–150% for dextrans were found. The concurrent increase of attenuation and transmittance by dextran B and C, relative to A, shows not only that refractive index matching is important for blood layer optical clearing, but also that RBC aggregation, which defines the scattering indicatrix, may be substantial. Dextran macromolecules are neutral polymers. The

Optical Clearing of Blood, Cells, and Microorganisms

119

(a)

(b)

(c)

(d)

Figure 63 Blood-smear microscopy: blood with low-molecular-weight dextran A (M = 10,500) (a) immediately after the application of the agent for which there are no aggregates and (b) 15 min after application of the agent (some small aggregates are seen). Blood with high-molecular-weight dextran C (M = 473,000) (c) immediately after application of the agent (rouleaux are clearly seen) and (d) 15 min after application of the agent (the network of aggregates is seen).47

high molecular weight dextrans are used artificially to induce RBC aggregation by bridging surfaces of adjacent cells after adsorption on their surfaces. The low molecular weight dextrans prevent normal blood aggregation. From blood smear microscopy (see Fig. 63), we can see that rouleaux have occurred in the blood diluted with dextran C, but that no aggregates were produced in the blood mixed with low molecular weight dextran A.153,154,177,178 The lower sedimentation rate of blood with dextrans B and C and the higher sedimentation rate of blood with dextran A relative to the rates for whole blood or for blood diluted with saline also reflect the aggregation abilities of various dextrans.47 The high molecular weight dextran C has greatly changed the scatter (RBCs and aggregates) morphology and size. Normal RBCs are biconcave disks of 8 µm diameter and 2 µm thickness when they are in an isotonic solution. It is known that the RBCs sizes, shapes, and orientations contribute to the properties of blood backscattering.275–277 Aggregation results in a decrease in diffusing surfaces, which in turn leads to a decrease of the backscattered signal.276 It can be concluded that the greater transmittance enhancement of dextran C is governed strongly by the scattering changes (refractive index matching) accompanied by RBC aggregation. Some discrepancy between µt and
T can be also explained by the fact that different algorithms are used to estimate them: µt is defined as a single scattering parameter and
T is defined as an experimental value that accounts for multi-

120

Optical Clearing of Tissues and Blood

ple scattering, which is why it is more sensitive for reduction of scattering. It has been shown that immersion leads at first to a reduction in the number of scattering events, and only then to the appearance of ballistic photons (see Figs. 10–12).39 For example, for the scattering system described in Ref. 39, which normally has for unmatched indices of scaterers ns = 1.47 and ground material n0 = 1.35 a total transmittance at 800 nm of 10%, after index matching (n0 = 1.41) the total transmittance rises to 45%, but the number of scattering events (as many as 10–15) remains high [see Fig. 11(c)]. It should be noted that for fresh erythrocyte concentrates flowing at a physiological velocity at an oxygen saturation of 98%, the change of hematocrit from 0.4 to 0.2 causes a reduction of the transport scattering coefficient µ
s from 16.8 to 8.8 cm−1 at 633 nm, measured by an integrating sphere technique.16 The corresponding changes in scattering coefficient µs and anisotropy factor g are µs from 850 to 800 cm−1 and g from 0.980 to 0.989. More than a tenfold decrease in the total attenuation coefficient measured by OCT (see Table 5) relative to the value of the scattering coefficient measured by an integrating sphere technique for blood samples with approximately equal hematocrit also show the limitations of a single scattering algorithm to extract a proper value of the total attenuation coefficient. Another reason for such strong discrepancy between µt and
T data may be caused by spatial variations of the reflectivity of the blood sample α(z) associated with the variations of the local refractive index and the backscattering property of the blood sample. Sedimentation increases the transparency of a blood layer because there is less bulk scattering as the RBCs fall. As expected, the undiluted blood sample has the lowest reflectance from the metal plate because it has a higher concentration of scatterers (RBCs) and a lower sedimentation rate.47,177,178 For the slightly diluted blood by a saline and at addition of the low molecular weight dextran A, reflectance from the metal plate in the depth of the blood sample increases, because both dilution by saline and addition of dextran A cause more intense sedimentation. Dextran B, which has a mean molecular weight, permits higher reflectance than the control (i.e., blood diluted by saline) only during the first 4 min. This result reflects the competition between two processes: refractive index matching, which is important at the beginning especially for dextran, and sedimentation, which is more important for the control sample after time has elapsed. The high molecular weight dextran C permits an increase in metal plate reflectance compared with that of the control for only a short period at the beginning of sedimentation, when only the refractive index matching effect dominates. Such behavior shows that after some time interval RBC sedimentation may be more important for increasing reflectance than the refractive index matching effect, which is provided by dextrans with higher molecular weight. This result is clearly seen from the in-depth reflectance profiles presented in Fig. 64, which show the three main evolution peaks in time: the first peak is independent of time and is induced by reflectance at the glass-blood interface; the second peak, which is broad and has some structure, is caused by reflectance at the RBC-plasma interface (within this peak, aggregates can be seen); and the third

Optical Clearing of Blood, Cells, and Microorganisms

121

peak is caused by the metal reflector. Qualitatively, the intensity of the third peak shows changes in the blood layer’s transmittance, and the second peak defines the attenuation coefficient of this layer.

Figure 64 OCT in-depth reflectance profiles measured for slightly diluted whole blood (13% volume fraction of saline, hematocrit ∼35%, concentration of dextran 3.25 g/dl in a blood sample) (a), (d), (g) at the beginning of the sedimentation process; (b), (e), (h) at 5 min; and (c), (f), (i) at 10 min for (a)–(c) saline only, (d)–(f) dextran A added, and (g)–(i) dextran C added. The first peak is induced by reflectance at the glass-blood interface; the second peak, which is broad and has some structures, is caused by reflectance at the RBC-plasma interface; and the third peak is caused by the metal reflector.47

The histograms in Fig. 65 illustrate the dynamic changes in transmission of light after application of saline and dextrans in terms of the changes in backscattered intensity from the metal surface at the bottom of the experimental vessel for mildly diluted and highly diluted sets of samples, showing the same relative values of refractive index and sedimentation clearing of a blood layer as was discussed for slightly diluted blood. The difference lies only in the critical time interval when sedimentation clearing is greater than refractive index matching clearing. For mildly diluted blood [Fig. 65(a)], this happens within 7.5 min of the beginning of sedimentation; for highly diluted blood, at ∼7.5 min [Fig. 65(b)]. As shown above, for the slightly diluted blood the critical time is ∼4 min.

122

Optical Clearing of Tissues and Blood

(a)

(b) Figure 65 Dynamic changes in reflectance from the metal surface in blood caused by addition of agents (saline and dextrans A, B, and C): (a) blood dilution by 35% of saline (dextrans concentration 2.43 g/dl) and (b) blood dilution by 56.5% of saline (dextrans concentration 1.62 g/dl).47

Therefore, these results have shown that, because of sedimentation, low molecular weight dextran A may have a much greater ability to enhance transmittance than do high molecular weight dextrans B and C. From blood smear microscopy, the rouleaux in blood induced by the highest molecular weight dextran C at the beginning of mixing interact further with other rouleaux to form rouleaux networks with time, i.e., the degree of aggregation is much enhanced with time [Fig. 63(d)]. However, the degree of aggregation in the blood with the lowest molecular weight dextran A was much less [Fig. 63(b)], even less than for blood diluted by saline only.47,112 To clarify the role of RBC aggregation on optical clearing, and accounting for the fact that the aggregation process is time dependent, the blood sample was allowed to sediment after the addition of dextrans and before measurements were done. Figure 66 shows the summary of the effect of dextrans compared to the

Optical Clearing of Blood, Cells, and Microorganisms

123

Figure 66 A summary of the effects of dextrans A, B, and C (D × 10, D × 70, and D × 500, respectively) compared to that of the saline control on light transmission after 10 min sedimentation. Lower-concentration dextran C and dextran B had significant effects in enhancing light transmission. Efficiency of higher concentration dextrans was much lower than that of the saline control112 (Copyright ©Institute of Physics Publishing).

saline control on light transmission for the sample with 20% blood and 80% saline after 10 min sedimentation. It can be seen from Fig. 66 that the influence of dextran on the light transmission was different from that at the beginning of mixing dextrans in blood. The lower concentration (0.5 g/dl) dextran C still had the strongest effect on reducing the scattering of light in blood, with a 2.8-fold stronger effect than that of the saline control. However, enhancement by the highest concentration of dextran C (5 g/dl) and dextran B (10 g/dl) was dramatically lower than that of the saline control. At the beginning, they both had a very high blood optical clearing capability with 67.5% and 76.8%
T , respectively. In addition, the effect was decreased with the increase of dextran in blood within all three groups, contrary to the expectation of the refractive index matching hypothesis. The decreased aggregation capability of dextran with concentration explained well that light transmission decreased less with the increase of dextran for both types (midmolecular and large-molecular). Over a range of concentrations, dextran C and B induced RBC aggregation. However, dextrans have been known to exert a biphasic effect on RBC aggregation; they induce aggregation at low concentration, and disaggregation at high concentration.287 For example, with dextran B, the maximal aggregation size is obtained at approximately 3%, above which the size decreases. In OCT measurements of Ref. 112, 2 g/dl dextran C and 5 g/dl dextran B in 20% blood with 80% saline appeared to be the critical concentration to affect RBC aggregation. Their aggregation parameters became smaller than those of 0.5 g/dl dextran C and 1 g/dl dextran B. When the concentration increased to 5 g/dl for dextran C and 10 g/dl for dextran B, they played a role of disaggregation. That is the reason why the cells are much less packed than with the saline control, accounting for the reduced light transmission. Although refractive index matching suggested a higher light transmission, it can be seen that the aggregationdisaggregation effects are now dominant.

124

Optical Clearing of Tissues and Blood

The behavior of RBC in flow is dependent on the processes of aggregationdisaggregation, orientation, and deformation. For normal blood, rouleaux are easily decomposed to their individual cell constituents as blood flow (shear) increases. In some pathological cases, however, the capillary circulation is seriously affected because nonseparable rouleaux are formed. Increased RBC aggregability has been observed in various pathological states, such as diabetes and myocardial infarction, or following trauma.288 The aggregation and disaggregation properties of human blood can be used for the characterization of the hemorheological status of patients suffering different diseases.277 In this connection, optical clearing methodology for the controlling of optical properties of blood using molecules with a specific action on the RBC and plasma may be useful in the monitoring of blood parameters in flow. It is obvious that refractive index matching is not the only factor to affect transmittance in these experiments. The amount of the aggregation certainly has an effect, but other factors, such as the hematocrit, the manner in which RBCs are packed, the shape of the aggregates, the variance in size of these aggregates, and the fluctuation of all these parameters in time and space, may all contribute. It should be noted that the blood plasma’s osmolarity is also an important factor in changes in the scattering properties of blood and, therefore, in the control of blood clearing and the improvement of contrast of OCT images obtained within or behind a layer of blood. In the discussed experiments, the osmolarity of the plasma was different for each of the added agents. Variation in plasma osmolarity led to changes in the shape of the erythrocytes; RBC shrinking (acanthocytes) took place when the plasma was hyperosmotic, and RBC swelling (spherocytes) took place when the plasma was hypo-osmotic. For the diluted blood samples (hematocrit of 7.5%), the scattering coefficient µs showed a slight decrease (∼10%) with increasing osmolarity in the range from 225 to 450 milliosmol/l; also, the anisotropy factor g decreases from 0.995 to 0.991, and correspondingly the reduced (transport) scattering coefficient µ
s increased linearly with osmolarity up to 70%.16 Such a strong effect on the scattering properties of the blood solution was caused not only by changes in cell shape but also by the variation with osmolarity of the refractive index of the cell-bounded hemoglobin solution. The refractive index of the cell-bounded hemoglobin solution can be estimated from Eq. (59). Assuming a mean erythrocyte volume of 90 µm3 and an inner cell hemoglobin concentration of 350 g/l for isotonic conditions, the values of refractive indices and sphere-equivalent diameters were calculated (see Table 6). This table clearly shows that the osmolarity of the blood solution can substantially change the scattering properties of the blood layer. Refractive index matching is easier to achieve in conditions of low osmolarity, but for OCT imaging of RBC or their aggregates, the hypertonic conditions are preferable.

Optical Clearing of Blood, Cells, and Microorganisms

125

6.3 Theoretical Analysis of Blood Optical Clearing From the above analysis of experimental data, it follows that to describe theoretically light transport in the immersed blood, we have to consider blood as a turbid medium with multiple scattering, defined by scattering and absorption properties of individual particles (erythrocytes) and by concentration effects and polidispersity of the cell suspension. The erythrocyte size and complex refractive indices (n
, n

) of erythrocytes and blood plasma define the absorption µa and scattering µs coefficients, and scattering anisotropy factor g. The size, shape, and optical parameters of blood cells as well as optical properties of blood suspension are described in Refs. 16, 277, and 289–291, and are presented in Table 6. The erythrocyte mean volume at isotonic medium is 94 ± 14 µm3 and the volume distribution is in the range from 30 to 200 µm3 . The hemoglobin concentration in hemolyzed blood is between 134 and 173 g/l. The volume fraction of cells within whole blood is defined as the hematocrit (Hct). Each erythrocyte contains approximately 29 pg of hemoglobin. The hemoglobin concentration within an erythrocyte ranges from 300 to 360 g/l. The real part of the refractive index of the red blood cell is very close to 1.4 in the wavelength range from 400 nm to 1200 nm.289,290,292,293 Table 6 RBC parameters found from osmolarity of the blood solution.16

Osmolarity, milliosmol/l 250 (hypotonic) 300 (isotonic) 400 (hypertonic)

Hematocrit, RBC RBC hemoglobin Refractive Equivalent % volume, concentration, g/l index sphere µm3 at 589 nm diameter, µm 8.1 7.5 6.6

96.7 90.0 78.6

325 350 400

1.397 1.402 1.412

5.70 5.56 5.32

The phase function and scattering cross section of an individual erythrocyte depend on its orientation.294 However, light-scattering characteristics of a large number of randomly distributed nonspherical particles are very close to light-scattering characteristics of a system of randomly distributed spherical particles with equal volume.16,295 Therefore, calculations can be done for a model of homogeneous spheres with the volume equal to the volume of real erythrocytes. Such a model provides simpler calculations than that for a rigorous theory accounting for particle nonsphericity220 and allows one to account for particle polydispersity in the simplest way; in particular, on the basis of the data presented in Ref. 290 (see Fig. 67). The presence of big particles in the distribution can be associated with small aggregates of RBC. Evidently, the hemoglobin concentration CHb in an erythrocyte correlates with its volume VRBC . In accordance with the data of Ref. 289, such dependence is defined by: CHb = 0.72313 − 0.00451VRBC ,

(74)

126

Optical Clearing of Tissues and Blood

Figure 67 Volume (size) distribution function of RBC in a whole blood.290 The volume fraction of RBC (hematocrit) is 45%, which corresponds to venous blood of an adult male. Reprinted with permission.

Figure 68 Spectral dependencies of the index of refraction (real part) for: (1) RBCs, 100% oxygenation of hemoglobin and its mean concentration in RBC of 340 g/l;290,294 (2) oxygenated (100%) hemoglobin; and (3) deoxygenated hemoglobin at concentration of 140 g/l,294 which corresponds to the mean concentration of hemoglobin in whole blood.67 Copyright 2004 by the American Physical Society. Reprinted with permission.

Optical Clearing of Blood, Cells, and Microorganisms

127

Figure 69 Absorption spectra of deoxy- and oxyhemoglobin. —2— εdλ , pH = 5.5–9.5; —"— εoλ , pH = 5.5–10.0.286

where CHb is the hemoglobin concentration in g/ml, and VRBC is the RBC volume in µm3 . The spectral dependence for the real part of the RBC refractive index is presented in Fig. 68,290,294 and the spectral dependence for the imaginary part can be calculated using data presented in Fig. 69. Both refractive index components are proportional to hemoglobin concentration in RBC; the real part is defined by Eq. (59) with βHb = 0.001942 dl/g for 589 nm16 and βHb = 0.00284 dl/g at 640 nm,294 and the imaginary part by the following expression: n

= αHb CHb ,

(75)

where αHb is the spectrally dependent coefficient equal to 1.477 × 10−6 dl/g at 640 nm.294 Because the concentration of salts, sugars, and other organic components in RBC cytoplasm is negligible, hemoglobin can be considered to be dissolved in water only [see Eq. (59)]; thus, the spectral dependence of the refractive index of the medium in which hemoglobin is dissolved is defined by water. For a more precise description of the refractive index of this medium n0 (λ) when organic components are accounted for, instead of nw (λ), n0 (λ) = nw (λ) + 0.007 may be used.184 The spectral dependence of the real part of the refractive index of blood plasma can be described by the empirical Eq. (73) and, because blood plasma contains up to 91% water, and only 6.5–8% (about 70 g/l) proteins (hemoglobin, albumin, and globulin), and about 2% of low-molecular compounds, its imaginary part is negligible and can be ignored in calculations. For further calculation of scattering and absorption coefficients and the scattering anisotropy factor, the Mie theory valid for a homogeneous spherical particle is used. The corresponding equations for scattering and absorption cross sections and the anisotropy factor are given by Eqs. (9), (10), and (12). For a densely packed

128

Optical Clearing of Tissues and Blood

polydisperse particle system, which whole blood is, the absorption and scattering coefficients and scattering anisotropy factor are defined by [see Eqs. (24) and (26)]199 µa =

N
RBC

(76)

Ni σai ,

i=1

µs = F (Hct) NRBC

N
RBC

(77)

Ni σsi ,

i=1

µsi gi , g = i=1 NRBC i=1 µsi

(78)

where F (Hct) is the packing function of RBC292,296,297 and accounts for the interparticle correlation effects, Hct is the hematocrit, NRBC is the number of RBC diameters (volume fractions), Ni = fRBCi /VRBCi is the number of RBC in a unit volume of blood, fRBCi is the volume fraction of RBCs with volume VRBCi = (4/3)πai3 (see Fig. 67), and ai is the radius of an individual equivalent volume spherical particle. At glucose application as an immersion agent, the spectral dependence of the index of refraction of blood plasma corrected by the added glucose-water solution should be accounted for [see Eqs. (57), (58), and (73)] as nbp+gl (λ) = nbp (λ) + 0.1515 × Cgl ,

(79)

where nbp (λ) is the refractive index of blood plasma defined by Eq. (73), and Cgl is the concentration of glucose in g/ml. Because glucose has no strong absorption bands within the spectral range from 400 to 1000 nm, its absorption may be neglected. It can be also hypothesized that glucose molecules do not bounding with proteins in blood plasma and hemoglobin in RBC during the limited time (a few minutes, maximum) of their interaction. As we already discussed, the RBC is very sensitive to changes in blood plasma osmolarity (see Table 6). At osmolarity increase due to cell dehydration, the RBC volume decreases, hemoglobin concentration within the cell increases, and the index of refraction increases. Glucose injection in blood causes the linear increase of plasma osmolarity with glucose concentration, up to 6000 mOsm/l at glucose concentration in blood plasma of 1.0 g/ml. Indeed, for patients, such large glucose concentrations may be applied only locally in the vicinity of a vessel wall site under spectroscopic study or optical imaging. Using data of Ref. 16, the following empirical relation was suggested to describe RBC volume change with osmolarity:49 



osm VRBC (osm) = VRBC (300) 0.463 + 1.19 exp − 376.2

 ,

(80)

Optical Clearing of Blood, Cells, and Microorganisms

129

where VRBC (osm) is the RBC volume in µm3 at the given osmolarity, expressed in mOsm/l, and VRBC (300) is the RBC volume at isotonic osmolarity osm = 300 mOsm/l. At glucose injection, the local hematocrit (Hct) decreases. If Hct before the injection of glucose was 45% at osm = 300 mOsm/l; then at Cgl = 0.05 g/ml, osm = 580 mOsm/l and Hct = 32%; at Cgl = 0.1 g/ml, osm = 850 mOsm/l and Hct = 26%; and at Cgl = 0.2 g/ml, osm = 1400 mOsm/l and Hct = 22%. At further increase of glucose, concentration (0.3–1.0 g/ml) hematocrit is constant and Hct ∼ = 21%, in spite of a linear increase of blood plasma osmolarity (2000–6000 mOsm/l). Results of modeling of the scattering properties control for a whole blood at its immersion (local intravessel injection) by a glucose solution at different concentrations accounting for RBC packing function in the form F (Hct) = (1− Hct),292,297 polidispersity (150 volume fractions of volume (size) distribution, Fig. 67), osmolarity, and hematocrit effects are presented in Figs. 70 and 71. The scattering coefficient and scattering anisotropy factor were calculated. The scattering coefficient behavior with concentration and wavelength (Fig. 70) is defined by: (i) the change in blood plasma osmolarity (increase of scattering for all wavelengths far from the Soret band caused by RBC shrinkage and increase of the refractive index for low concentrations of glucose, see Table 6); (ii) reduction of blood hematocrit (plays some role in the scattering decrease for glucose concentrations, less than 0.3 g/ml); (iii) refractive index matching, the main effect (a significant reduction of scattering for glucose concentration from 0.5 to 0.7 g/ml, depending on the wavelength); and (iv) dispersion of hemoglobin absorbing bands (within a strong Soret band (415 nm) it does not allow a significant reduction of scattering and modifies slightly the position of the dip and the depth of scattering damping for other lower-absorbing hemoglobin bands, 542 and 575 nm). The maximal damping of the scattering corresponds to 900 nm, where the influence of hemoglobin bands’ dispersion is minimal, but the highest glucose concentration of 0.7 g/ml is needed in that case. The scattering coefficient increase at higher glucose concentrations is caused by refractive index mismatch, when the refractive index of the RBCs becomes less than that for blood plasma modified by adding glucose. The same factors define the behavior of the scattering anisotropy parameter (Fig. 71): RBC shrinkage causes a decrease of the g factor for small glucose concentrations, refractive index matching for moderate concentrations causes its increase, and further refractive index mismatch causes its reduction. For applications, especially when OCT endoscopy is used, concurrent reduction of scattering and increase of the g factor at immersion agent administration is important. The transport scattering coefficient, µ
s = (1 − g)µ, decreases more dramatically; thus, transport free path length for a photon, lt ∼ = 1/µ
s , will also be increased dramatically. Thus, a greater amount of photons, which carry information about the hidden object (for example, thin-wall plaques in the coronary arteries) can be detected. From data in Figs. 70 and 71, it follows that for the wavelength 900 nm, the scattering coefficient of blood is changed from 1200 cm−1 to approximately 50 cm−1 and g factor from 0.991 to 0.994 at glucose immersion; thus, the transport free path

130

Optical Clearing of Tissues and Blood

Figure 70 Theoretical modeling of blood optical clearing at glucose intravessel injection. The calculated damping curves for light scattering at glucose concentration.184 The initial blood hematocrit of 45% and RBC hemoglobin concentration of 322 g/l; 150 volume fractions of RBC with different volume (size) in accordance with the RBC volume distribution function, presented in Fig. 67, were used in the modeling.

Figure 71 Theoretical modeling of blood optical clearing at glucose intravessel injection. The calculated curves for the scattering anisotropy factor at glucose concentration.184 The initial blood hematocrit of 45% and RBC hemoglobin concentration of 322 g/l; 150 volume fractions of RBC with different volume (size) in accordance with the RBC volume distribution function, presented in Fig. 67, were used in the modeling.

Optical Clearing of Blood, Cells, and Microorganisms

131

length increases more than 35 times. Correspondingly, the depolarization depth of blood,257,298,299 which is proportional to lt , should be much larger at optical clearing. The described method for immersed blood modeling is applicable for any other biocompatible immersion agent administration, such as dextrans and x-ray contrasting agents (see Tables 1 and 5). At blood clearing, there also exists another possibility of blood immersion, using the local blood hemolysis, which can be provided in the vicinity of the fiber-optic endoscopic probe.64 Table 7 shows the expected values of the local changes of blood hematocrit and hemoglobin concentration in plasma at the release of hemoglobin from RBC at blood hemolysis. In that method, free hemoglobin is the immersion agent. To model optical properties, all earlier discussed effects should be taken into account. The local increase of hemoglobin concentration in plasma can lead to local change of plasma osmolarity:49 osm
= osm +

CbpHe , MHb

(81)

where osm is the plasma osmolarity under the physiological condition (280– 300 mOsm/l), CbpHb is the concentration of plasma hemoglobin in g/l, and MHb is the molar mass of hemoglobin (MHb = 66,500 g/M). The expected change of RBC volume calculated using empirical Eq. (80) is not more than 0.1% at hemolysis less than 20%. For simplicity, the polydispersity of RBC can be taken into account on the basis of the six-fraction blood model given in Table 8,300 which correlates with the more complete distribution given in Fig. 67. Table 7 Calculated values of the local blood hematocrit and hemoglobin concentration in plasma (free hemoglobin) at different degrees of hemolysis for the initial hemoglobin concentration within a RBC corresponding to physiological condition, i.e., 33 g/dl.64

Hemolysis, %

Hematocrit, %

Hemoglobin, g/dl

0 5 10 15 20

42 39.9 37.8 35.7 33.6

0 1.15 2.23 3.23 4.91

Table 8 Size distribution of the equivalent spherical particles modeling RBC.300

Volume fraction, % Radius, µm

4

14

30

32

14

6

1.2 ± 0.2

1.7 ± 0.3

2.2 ± 0.2

2.7 ± 0.3

3.4 ± 0.4

4.3 ± 0.5

132

Optical Clearing of Tissues and Blood

Calculations of the absorption coefficient, the scattering coefficient, and the anisotropy factor of whole blood at normal conditions and at local hemolysis have been performed using Eqs. (76)–(78), with the packing function F (Hct) = (1 − Hct)(1.4 − Hct).297 In contrast with the small changes of the absorption coefficient, more significant changes of the scattering properties of blood have been observed at increase of free hemoglobin concentration in plasma (Figs. 72 and 73). Figure 72 shows the evolution of the scattering spectra of blood with a rather smooth decrease (up to 30–40%) of scattering for all wavelengths with free hemoglobin release at hemolysis.

Figure 72 Theoretical wavelength dependencies of the scattering coefficient of whole blood at different degrees of blood hemolysis.64 The expected amount of free hemoglobin release at hemolysis is given in Table 7. Reprinted with permission.

Figure 73 shows the scattering coefficient and g factor for two wavelengths, 633 nm and 820 nm, versus the local hemolysis rate. The maximal decrease of the scattering coefficient for these wavelengths is 40%, while the anisotropy factor, at wavelength 633 nm, increases from 0.9940 to 0.9952—and at 820 nm, from 0.9919 to 0.9929. The described method can be realized not only at blood hemolysis, but also at local free hemoglobin injection. Hemoglobin administration also may serve as a clearing agent for tissue clearing, when clearing is needed in the spectral range far from the strong absorption bands of hemoglobin. On the other hand, the found sensitivity of scattering properties of blood to RBC hemolysis may be used for design-

Optical Clearing of Blood, Cells, and Microorganisms

133

Figure 73 Calculated the scattering coefficient and the scattering anisotropy factor at two wavelengths 633 and 820 nm versus the degree of blood hemolysis.64 The expected amount of free hemoglobin release at hemolysis is given in Table 7. Reprinted with permission.

ing of the effective optical technology for in vivo monitoring of blood hemolysis in vessels.

6.4 Optical Immersion of Cells and Microorganisms The optical immersion method is a valuable technique for study of the refractive and scattering properties of livings cells.48,65,76,77,176,208,211,301,302 For cellular refraction measurements, this technique has been use since the 1950s.48,65,76 It was successfully used in combination with phase refractometry to study water and solids distribution in animal cells (mechanisms of cell cornification); to examine mechanisms of animal cell motility connected with water redistributions; to study cell permeability, damage, and death, and the vitality and growth cycles of bacteria, fungi, yeasts, and spores. Some hematological applications of the immersion technique in cell suspensions studies are discussed earlier in Sections 6.2 and 6.3. In cell examination, the requirements regarding immersion agents should be somewhat different from those in tissue optical clearing, where for many applications only cell damage is critical and the strong osmotical properties of agents is one of the leading mechanisms for tissue clearing. In general, an immersion substance (IS) to be used for the refractometry of living cells should fulfill the following requirements.65,76

134

Optical Clearing of Tissues and Blood

(1) IS should be nontoxic, with no affect on the structure or function of a living cell, i.e., chemically inert and not affecting any chemical components of a cell. (2) It should be isotonic, i.e., does not cause any changes of cell volume. Cell shrinkage or swelling induced by water displacement from a cell into the surrounding medium, or vice versa, are accompanied by corresponding changes of cell refractive index; thus, the measured refractive index will not be a true value. The isotonic property is a biological character of a cell in solution, which is connected with such physical characteristic of a solution as its isoosmotic property, but may be not similar. To provide isoosmotic and isotonic conditions, IS should exert a low osmotic pressure; that is, it should consist of dissolved particles with high molecular weight and dimensions. For example, the osmotic pressure of a 10% water solution of bovine serum albumin (BSA) is equivalent to the pressure of a 0.08% water solution of sodium chloride (NaCl). (3) IS should not penetrate the cell when the refractive index of a whole cell is under study. Otherwise at cell immersion in a medium with the higher refractive index, this substance diffuses inside the cell and equalizes refractive indices inside and outside the cell, making the measured cell refractive index value far from a true value. Thus, for many cases, IS should have a macromolecular structure to prevent cell permeability. However, for some specific cases, when intracellular organelles are under investigation, immersion agents with a controllable permeation can be used. (4) IS should be freely soluble in water so that the refractive index of the solution can be equal or exceed that of the part of the cell to be measured. As for animal cells, the refractive index range is from 1.350 to 1.426 (see Table 9), and for bacterial cells, from 1.360 to 1.420;76 IS refractive index values must be variable in these ranges with the step of 0.002–0.005. The best decision is to find two well-mixing solutions, one with the minimal index and another with the maximal index of the range under study. It is very important that for each of mixed solutions, IS keeps its isotonic properties. Evidently, this condition can be satisfied if both solutions have a low osmotic pressure and their mixing is not accompanied by a specific chemical reaction causing increase of osmotic pressure. (5) IS should be optically transparent and isotropic, i.e., conditions of less absorption and scattering in the measuring wavelength range as well as less linear birefringence and chirality should be provided. (6) IS should be stable in the range from room to physiological temperatures, and should not change their optical properties at prolonged storage. Such requirements are fulfilled almost completely for water solutions of albumin and water-glycerol gelatinous gels.65,76 The fifth fraction of bovine or human serum albumin contains total mass of serum albumin, about 3% of α globulin, and less than 0.5% of β-globulin,76 and its index of refraction has a linear dependence

Optical Clearing of Blood, Cells, and Microorganisms

135

Table 9 Experimental mean values of phase n or group ng refractive indices of tissues, blood, and their compounds measured in vitro and in vivo*; rms values are given in parentheses.

Tissue

λ, nm

n, ng

Comments

Ref. 303

Human aorta: Normal: Intima Media Adventitia Calcified: Intima Media

456–1064 456–1064 456–1064

1.39 1.38 1.36

456–1064 456–1064

1.39 1.53

Human bladder: Mucosa Wall Integral

456–1064 456–1064 456–1064

1.37 1.40 1.38

Human brain: Gray matter White matter White and Gray

456–1064 456–1064 456–1064

1.36 1.38 1.37

Human colon: Muscle Submucosa Mucosa Integral

456–1064 456–1064 456–1064 456–1064

1.36 1.36 1.38 1.36

Female breast tissue: Normal Malignant

800 800

1.403 1.431

Ref. 304

Human esophagus: Mucosa

456–1064

1.37

Ref. 303

Human fat: Subcutaneous Abdominal

456–1064 456–1064

1.44 1.46

Bovine fat

633

1.455

Porcine fat

488 632.8 1079.5 1341.4 632.8

Human heart: Trabecula Myocard

1.510(0.002) 1.492(0.003) 1.482(0.002) 1.478(0.004) 1.493(0.005)

Ref. 214, fiber-optic refractometer (FOR); homogenized tissue Ref. 305, laser refractometer (LR)

Ref. 306, LR Ref. 303

456–1064 456–1064

1.40 1.38

136

Optical Clearing of Tissues and Blood Table 9 (Continued).

Tissue

λ, nm

n, ng

Human femoral vein Kidney: Human Human Canine Porcine Bovine

456–1064

1.39

456–1064 633 633 633 633

1.37 1.417 1.400 1.390 1.390

Ref. 303 Ref. 214, FOR

Liver: Human Human

456–1064 633

1.38 1.367

Ref. 303 Ref. 214, FOR; homogenized tissue

633 633 633

1.380 1.390 1.390

456–1064 633

1.38 1.38

633

1.38

456–1064 633

1.37 1.400

Canine Porcine Bovine Lung: Human Canine Porcine Muscle Human Canine Bovine Ovine||

Ovine⊥

Porcine||

Porcine⊥

Porcine|| Porcine⊥

633 488 632.8 1079.5 1341.4

1.410 1.404(0.003) 1.389(0.002) 1.378(0.004) 1.375(0.003)

488 632.8 1079.5 1341.4 488 632.8 1079.5 1341.4 488 632.8 1079.5 1341.4 632.8 632.8

1.402(0.002) 1.389(0.002) 1.375(0.003) 1.373(0.003) 1.402(0.002) 1.381(0.002) 1.372(0.003) 1.370(0.003) 1.399(0.002) 1.379(0.002) 1.370(0.002) 1.367(0.003) 1.380(0.007) 1.460(0.008)

Comments

Ref. 303 Ref. 214, FOR; homogenized tissue

Ref. 303 Ref. 214, FOR; homogenized tissue Ref. 305, LR, tissue samples labeled as || and ⊥ are the same sample with the tissue fibers oriented in parallel and perpendicular to the interface, respectively

Ref. 306, LR

Optical Clearing of Blood, Cells, and Microorganisms

137

Table 9 (Continued).

Tissue

λ, nm

n, ng

Comments

Human skin: Stratum corneum (SC)*

1300

ng = 1.51(0.02)

Ref. 308, OCT, reference mirror and focus tracking

Epidermis* Dermis SC* (Palm of hand)

1300 1300 1300

ng = 1.41(0.02) ng = 1.34(0.02) (nng )1/2 = 1.47(0.01)

Epidermis* (Palm of hand, Granular layer) Epidermis* (Palm of hand, Basal layer) Epidermis* (Volar side of lower arm) Upper dermis* (Palm of hand) Upper dermis* (Volar side of lower arm) SC* (Dorsal surface of a thumb) Air/skin interface* (Volar side of a thumb) SC/epidermis interface* (Volar side of a thumb) Pigskin

1300

(nng )1/2 = 1.43(0.02)

1300

(nng )1/2 = 1.34(0.02)

1300

(nng )1/2 = 1.36(0.01)

1300

(nng )1/2 = 1.41(0.03)

1300

(nng )1/2 = 1.43(0.02)

Pigskin (treated by a detergent solution) Human stratum corneum SC Rat skin Mouse skin

Ref. 309, OCT, focus tracking by moving of fiber tip/collimating lens

Bifocal OCT refractometer, Ref. 310

Ref. 311

980

(nng )1/2 = 1.50(0.02)

980

(nng )1/2 = 1.56

980

(nng )1/2 = 1.34

1300

(nng )1/2 = 1.415

1300

(nng )1/2 = 1.365

400–700 456–1064 456–1064

1.55 1.42 1.40

Ref. 312, OCT, focus tracking

Ref. 206 Ref. 303

138

Optical Clearing of Tissues and Blood

Table 9 (Continued).

Tissue

λ, nm

n, ng

Comments

Spleen: Human Canine

456–1064 633

1.37 1.400

Ref. 303 Ref. 214, FOR; homogenized tissue

633

1.400

456–1064 456–1064 456–1064

1.39 1.38 1.38

Ref. 303

1.391(0.002) 1.373(0.002) 1.361(0.003) 1.359(0.004)

Ref. 305, LR

400–700

1.335

Ref. 206

400–700 400–700

1.52(0.01) 1.38(0.1)

Ref. 206

850

1.4245 (T 1.4239 (T 1.4223 (T 1.4216 (T 1.4186 (T 1.4027 (T 1.4016 (T 1.4000 (T 1.3986 (T

400–700

1.336

400–700 400–700 400–700

1.376 1.47 1.35

400–700 400–700

1.386 1.406

Porcine Human stomach: Muscle Mucous Integral Porcine small intestine

Human cerebral spinal fluid Rat mesentery Rat mesentery* Rat mesentery

Human eye: Aqueous humor Cornea: Integral Fibrils Ground substance Lens: Surface Center

488 632.8 1079.5 1341.4

= 25◦ C) = 30◦ C) = 35◦ C) = 38◦ C) = 40◦ C) = 42◦ C) = 44◦ C) = 46◦ C) = 48◦ C)

Ref. 66, OCT, several pieces of rat mesenteries, mainly composed of phospholipids bilayers; gelto-liquid phase transition in the range from 38 to 42◦ C

Ref. 206

Optical Clearing of Blood, Cells, and Microorganisms

139

Table 9 (Continued).

Tissue

λ, nm

n, ng

Comments

Vitreous humor Tears Sclera

400–700 400–700 442–1064

1.336 1.3361–1.3379 1.47–1.36

Ref. 303

Cornea:

Human: Fibrils Extrafibrillar material Ox: Fibrils Extrafibrillar material Rabbit: Fibrils Extrafibrillar material Trout: Fibrils Extrafibrillar material

Abbe refractometer measurements and calculations on the basis of from x-ray diffraction data, Ref. 313 589 589

1.411(0.004) 1.365(0.009)

589 589

1.413(0.004) 1.357(0.009)

589 589

1.416(0.004) 1.357(0.010)

589 589

1.418(0.004) 1.364(0.009)

Bovine: Stroma Hydrated fibrils Hydrated extrafibrillar matrix Dry collagen Dry extrafibrillar material Solvent (salt solution) Hydrated stroma:

Data from Refs. 314 and 315 589 589

1.375 1.413

589

1.359

589 589

1.547 1.485

589

1.335

589

1.335 + 0.04/(0.22 + 0.24H)

H = 3–8, H = 3.2— physiological hydration314

140

Optical Clearing of Tissues and Blood

Table 9 (Continued).

Tissue

λ, nm

n, ng

Comments

Calf cornea: Normal

820

ng = 1.380(0.001)

Hydrated (H = 1.5–5): ng (H ) = a + b/(H + 1)

820

a = 1.324(0.002) b = 0.272(0.009)

Ref. 316, OCT, reference mirror method; H = 5.3 · d − 0.67, d is the corneal stroma thickness in mm

Human cornea

550

1.3771

Obstfeld, 1982, datum from Ref. 317

589

1.380(0.005)

Patel et al., 1995, datum from Ref. 317 Ref. 318, OCT Ref. 317, OCT, 21◦ C, Ref. 317, extrapolation of datum for 550 nm Ref. 317, extrapolation of datum for 589 nm Ref. 317, extrapolation of datum for 855 nm

Human tooth: Enamel Enamel Apatite Dentin matrix Enamel

855 1270 1270

ng = 1.3817(0.0021) ng = 1.389(0.004) ng = 1.386

1270

ng = 1.390(0.005)

1270

ng = 1.3838(0.0021)

220 400–700 400–700 Visible

1.73 1.62 >1.623 1.553(0.001)

Ref. 206

856

ng = 1.62(0.02)

Dentin Enamel Dentin

856 850 850

ng = 1.50(0.02) ng = 1.65 ng = 1.54

Ref. 321, OCT

Human nail*

850

ng = 1.51

Ref. 321, OCT

Human hair shaft: Black Brown Red Blond Gray White

850 850 850 850 850 850

ng = 1.59(0.08) ng = 1.58(0.06) ng = 1.56(0.01) ng = 1.57(0.01) ng = 1.58(0.01) ng = 1.58(0.01)

Ref. 322, OCT

Ref. 319, optical immersion method Ref. 320, OCT, reference mirror method

Optical Clearing of Blood, Cells, and Microorganisms

141

Table 9 (Continued).

Tissue Human blood

whole

Human whole blood 0%-solution in water 20%-solution in water 40%-solution in water 60%-solution in water Undiluted blood (extrapolated) Human plasma

blood

Human red blood cells (dry): Healthy patients (n = 7, fixed RBC) Diabetic patients (n = 9, fixed RBC) Healthy patients (n = 7, intact RBC) Diabetic patients (n = 9, intact RBC) Hemoglobin: Oxygenated (from porcine blood)

λ, nm

n, ng

Comments

633 488 632.8 1079.5 1341.4

1.400 1.395(0.003) 1.373(0.004) 1.363(0.004) 1.360(0.005)

Ref. 214, FOR Ref. 305, LR

Ref. 323, the equilateral hollow prism 633

1.34

633

1.35

633

1.35

633

1.36

633

1.38

488 632.8 1079.5 1341.4

1.350(0.002) 1.345(0.002) 1.332(0.003) 1.327(0.004)

Ref. 305, LR

550

1.61–1.66

550

1.56–1.62

Ref. 70, pH = 6–8, Nomarski polarizinginterference microscope

550

1.57–1.61

550

1.61–1.64

800

1.392(0.001)

Ref. 67, OCT, 37◦ C, Hemoglobin of 93 g/l

142

Optical Clearing of Tissues and Blood Table 9 (Continued).

Tissue Deoxygenated (from porcine blood) Glycated (Glucose from 40 to 400 mg/dl) Glycated (Glucose from 400 to 800 mg/dl)

λ, nm

n, ng

Comments

800

1.388(0.002)

820

1.382 → 1.415

820

1.415 → 1.385

Collagen (Type I) Dry Fully hydrated

850 850

ng = 1.53(0.02) ng = 1.43(0.02)

Ref. 324, OCT

Cytoplasm

400–700

1.350–1.367

Ref. 206

Nuclei of cervical epithelium cells: Normal: Basal/parabasal Intermediate Superficial

Far visible/NIR Far visible/NIR Far visible/NIR

Cervical intraepithelial neoplasia (CIN 3): Basal/parabasal Intermediate Superficial

Far visible/NIR Far visible/NIR Far visible/NIR

Refs. 68, 69, (OCT), Hemoglobin of 140 g/l

Ref. 307, histology, cytometry, and modeling; 20 nuclei 1.387(0.004–0.007) were analyzed for 1.372(0.004–0.006) each case;
n is 1.414(0.005–0.008) the refractive index spatial fluctuation n(
n)

1.426(0.008–0.010) 1.404(0.007–0.009) 1.431(0.008–0.011)

on concentration with an increment of βp = 0.00185 [Eq. (59)]. Some discrepancy from linearity is seen for high concentrations that are possibly connected with their relative high viscosity. For microbiological studies, protein solutions are usually prepared with the refractive index range from 1.360 to 1.420 with the interval of 0.002 on the basis of a 0.5–0.6% solution of NaCl in distilled water.76 The water-glycerol gelatinous gels are applicable when cell motility does not allow one to provide precise measurements using protein solutions.76 Such gels fix and immobilize cells, preserving their vitality; they are optically transparent with low birefringence and high stability of optical properties, if correctly exploited, and with low osmotic pressure. The isotonic gel kits for studies are prepared by dilution of concentrated salt-free and purified gelatinous gels with turbidity of 0.5 × 10−3 cm−1 in a 0.2% sterile solution of glycerol in a 0.5% solution of NaCl with pH 7.0–7.2; other compositions with 1% glycerol, 1% glycerol, and 0.5% glucose, or 10% sasharose also can be used. Using protein solutions and gelatinous gels as immersion substances and in phase-contrast microscopy with

Optical Clearing of Blood, Cells, and Microorganisms

143

effective suppression of the background light, refractive indices of numerous bacteria of families, such as Coccaceae, Bacteriaceae, Bacillaceae, Spirillaceae, and Proactinomycetaceae, were measured.76 On the basis of refraction measurements of the concentration of dry materials and water in bacterial cells, their density, bacterial growth-cycle, as well as rehydration of lyophilized bacterial cells and hydration of spores were studied. Differentiation of vital and dead cells in lyophilized cell preparations and the percentage of vital spores were also determined.

Figure 74 Experimentally determined ratio of scattering intensities from cells (rat fibroblast cell clone MR1; ∼105 cells/ml) immersed in media of low and high indices of refraction.208 The media of low index had n = 1.332 (PBS) in both cases. The media of higher index had n = 1.345 (BSA in PBS) for the solid curve and n = 1.343 (ovalbumin in PBS) for the dashed curve. Data at higher angles are not shown because of interference from scattering off of protein molecules dissolved in the media.

One of the examples of cell light scattering analysis using the immersion technique is presented in Fig. 74, where experimentally determined ratios of scattering intensities from cells (rat fibroblast cell clone MR1; ∼105 cells/ml) immersed in media of low and high indices of refraction are shown.208 As a medium with a low index of refraction n = 1.332, phosphate buffer saline (PBS) (see Table 1) was taken in both cases. The media of higher index had n = 1.345 (bovine serum albumin (BSA) in PBS) for the solid curve and n = 1.343 (Ovalbumin in PBS) for the dashed curve (see Table 1). The scattered light intensity at small angles (<20 deg) was significantly greater when the cells are immersed in PBS with a low refractive index than when they are immersed in a protein solution with a higher index. Thus, it may be concluded that there is significant scattering at small angles from cell structures that are in contact with the immersion substance (IS). However, at larger angles (>40 deg) the effect on light scattering of increasing index of refraction of IS is much smaller. Following the estimations of the authors of Ref. 208, the percentage of light scattering from internal cellular structures can be determined.

144

Optical Clearing of Tissues and Blood

Accounting for the fact that the ratio of scattering intensity from cells suspended in IS with low and high indices for angles above 40 deg is 1.3 (Fig. 74), the fraction of scattering intensity from particles internal to the cell can be estimated. The scattering intensity in IS with low refraction is given by Inc + Ic , where Inc and Ic are the intensities of scattering from structures not in contact and in contact with the IS, respectively. In the IS with high refraction, the scattering from the particles in contact with the IS is reduced by about a factor of 2.1 (see Fig. 74, small angles) and the scattering is given by Inc + 0.48Ic . Thus, the relative light scattering from internal cell components when the cells are immersed in PBS, Inc /(Inc + Ic ) ≈ 0.55, because (Inc + Ic )/(Inc + 0.48 × Ic ) ≈ 1.3. The optical clearing effect can be most easily demonstrated by analyzing phase microscope images of bacterial cells containing only a cytoplasm and a membrane.76 If a biological object is homogenous, matching of its refractive index value with that of the host medium will make it optically invisible. In the case of a bacteria containing only a cytoplasm and a membrane, matching of the refractive index of the cytoplasm with that of the extra-cellular fluid will make the image of the cytoplasm disappear and sharply enhance the brightness of the optical image of the membrane. In a case where the refractive index of the extra-cellular fluid is externally controlled by the administration of an appropriate chemical agent, the disappearing of the cytoplasm and the sharp enhancement of the membrane brightness can be used as an efficient method of measuring the refractive index of the cytoplasm and monitoring of cell vitality. Figure 75 illustrates the optical clearing effect.

Figure 75 Effect of the extracellular fluid to cytoplasm refractive index matching on the phase-contrast microscope image of dysentery bacteria cells. There is a sharp enhancement of the membrane brightness. (Ref. 76, p. 70)

The finite-difference time-domain (FDTD) approach was recently suggested as a promising tool for a more detailed study of the optical clearing effect in cells, and its possible applications explored.302 In Ref. 302, the 3D and 2D FDTD simulation results of light transmission through a biological cell containing only cytoplasm and membrane are presented. The calculated 2D distributions and two cross sections of phase of the Ez component of the forward scattered light through the

Optical Clearing of Blood, Cells, and Microorganisms

145

Figure 76 Finite-difference time-domain (FDTD) modeling of light scattering by biological cells in controlled extracellular media.302 The calculated 2D distributions and two cross sections of phase of the Ez light component in the near field of a biological cell containing only cytoplasm and membrane. It is clearly seen that refractive index matching (graphs on the right) significantly enhances the phase contrast of the cell membrane. Cell radius, 0.72 µm; membrane thickness, 0.06 µm; cytoplasm refractive index, 1.36; membrane refractive index, 1.47. Asymmetry is due to the z polarization of the incident light. The simulations were performed by the FDTD solutions software, which is commercially available from Lumerical Solutions Inc., Vancouver, BC, Canada.

biological cell in the near field are shown in Fig. 76. It is clearly seen that refractive index matching (graphs on the right) significantly enhances the phase contrast of the cell membrane as it follows from the experimental data of Fig. 75. The intensity of phase microscope images is directly proportional to the phase accumulated by the light beam after its propagation through the cell. Calculations were done for typical parameters of a microbial cell: the diameter of the cell (1.44 µm) and

146

Optical Clearing of Tissues and Blood

the thickness of the membrane (0.06 µm), the refractive index of the cytoplasm (1.36) and of the membrane (1.47). In calculations, the FDTD cell size was taken as 0.02 µm and the extra-cellular fluid refractive index values were 1.33 for no refractive index matching and 1.36 for refractive index matching conditions.

7 Chemical Agent Delivery

7.1 Introduction In this section, we will continue to discuss physicochemical properties of chemical agents and their delivery through biological membranes. Such biological membranes and tissue components as the human epidermal membrane (HEM), the human cervical mucus (HSM), the living skeletal muscle fibers (SMF), and cell membranes will be considered as the examples. We will concentrate on the discussion of permeation efficiency of chemical agents and biological molecules often used in medicine and cosmetics. Some data on physicochemical properties of immersion agents and physiological analytes and their diffusivity are presented in Tables 1, 2, and 9, 66–70,206,214,303–324 and Fig. 77, which summarizes the refractive indices of biocompatible immersion agents and biological tissues, tissue components, animal cells, bacterial cells, and viruses. These data allow one to find an immersion agent suitable for a certain study or application. In particular, Table 10 presents characteristics (refractive index and osmolarity) of a variety of chemical agents with different optical clearing potentials (OCP), defined as the ratio of the values of tissue reduced scattering coefficient before and after agent action, OCP ≡ µs (before)/µs (after).325 OCP was measured in vitro at agent application to the dermis side of human skin using a Franz diffusion chamber after 20 min application time. It follows from Table 10 that there is no correlation between OCP and refractive index for used agents with indices in the range from 1.43 to 1.48, as well as no correlation with osmolality in a wide range from 1,643 to 26,900 mOsm/kg, but the highest values of OCP (from 2.4 to 2.9) are provided by the agents having both the highest refractive index and osmolality, such as glycerol, 1,4-butanediol, and 1,3-butanediol. Evidently, for effective optical clearing, the diffusivity of an agent through a biological membrane should be important. The transmembrane permeation is defined by two major factors—by the coefficient of agent distribution between the surrounding solution and membrane K=

C1m C2m = , C1 C2

(82)

where C1m and C2m are concentrations inside the membrane in the vicinity of its surfaces; and by the diffusion rate of the agent through the membrane.233 The diffusion coefficient Da that appears in the Eq. (37) is a function of the size and form of a molecule. For the spherical molecules (proteins) diffusing in water, Eq. (60) 147

148

Optical Clearing of Tissues and Blood

Table 10 Chemical agent characteristics and in vitro measured optical clearing potential (OCP) at agent application to dermis side of human skin using a Franz diffusion chamber; OCP is defined as the ratio of values of tissue reduced scattering coefficient before and after agent action, OCP ≡ µs (before)/µs (after), was measured after 20 min application time.325

Hydroxy-terminated chemical agent Glycerol 50% TMP (trimethylolpropanol) 100% TMP 1,3-butanediol 1,4-butanediol Ethylene glycol MPDiol glycol (1,3-diol, 2-methyl-propane) P-0062a

Refractive index

Osmolality (mOsm/kg)

1.47 1.43 1.47 1.44 1.44 1.43 1.44 1.48

14,550 6,830 13,660 22,050 26,900 22,640 23,460 1,643

OCP 2.9 ± 0.8 2.2 ± 0.3 2.1 ± 0.7 2.4 ± 0.7 2.8 ± 0.5 1.9 ± 0.6 2.3 ± 0.2 2.0 ± 0.5

a

P-0062 is a polyethylene glycol based prepolymer developed at the University of California, Irvine.

with S = 1/3 is valid.233 For small molecules—from hydrogen to trisaccharides— a steeper slope with S = 1/2 was found.233 Thus, instead of Eq. (38) the following expression can be written for the agent flux through a membrane: J = Pa K(C1 − C2 ),

(83)

where Pa = Da /d is the coefficient of membrane permeability, and C1 and C2 are the concentrations of molecules in two spaces separated by a membrane. In this section, we will also discuss general principles of designing of cosmetic preparations allowing one to provide deep permeation within skin, mechanisms of diffusion of proteins and other macromolecules in tissue components, and chemical and optical methods of the enhanced molecule delivery through biological membranes.

7.2 Cosmetic Preparations and Skin Permeation The general principles of designing of cosmetic preparations allowing one to provide deep permeation within skin for the improvement of its physiological properties are discussed elsewhere.326 As it was shown earlier, the same cosmetic preparations with, or even without, corrections may serve as the optical immersion compositions. This is the best solution when the immersion composition improves both physiological and optical properties of the skin. However, the excellent diffusional resistance of the SC makes the transdermal delivery of an immersion agent difficult.247 In order to understand the transport and barrier functions of the skin, it is important to have knowledge of the water and ions distribution within the different layers.249,250 Water content is known to influence: various physical characteristics,

Chemical Agent Delivery

149

Figure 77 The scale of refractive indicies of biocompatible immersion agents and biological tissues, tissue components, and animal and bacterial cells and viruses. Adapted from Ref. 76.

such as brittleness, elasticity, tensile strength; barrier characteristics; electrical resistance; and thermal conductivity. The SC receives water from within the body, but water also may be taken up from the environment. From within the body, water

150

Optical Clearing of Tissues and Blood

reaches this tissue from the sweat glands and by diffusion from underlying tissues. In vivo, the diffusion of water across the SC is a passive process. The water content of the innermost layer of the SC is in equilibrium with the adjacent moist granular layer. The outside cell layer, however, is in equilibrium with the environment and it is certainly drier than the innermost cornified layer. Thus, there exists a concentration gradient causing transepidermal water lost (TEWL). No significant difference was found for diffusion across epidermis and SC. The diffusion coefficient D of the flow of water through a stationary macromolecular gel (the tissue) corresponds to viscous flow through a very fine porous medium. As has been determined in strongly hydrated SC, D is about 4 orders of magnitude less than the self-diffusion coefficient in water.249 The diffusivity (D) values of water in SC as a function of its water content (Cm , g/cm3 ), or relative humidity (RH, %) for three subjects are presented in Table 11. The average water content of the SC as measured in Ref. 250 is 54%, while other authors arrived at water content as low as 15–40% in the same layer. It should be also noted that the hydration of the dermis is not significantly different from that of the viable cell layers of the epidermis. Table 11 Diffusivity (diffusion coefficient D) of water in SC as a function of its relative humidity (RH, %), or water content (Cm , g/cm3 ), for three subjects.249

Water content, RH, %

Water content, Cm , g/cm3

D × 1010 cm2 /s, subject A

D × 1010 cm2 /s, subject B

D × 1010 cm2 /s, subject C

46 62 81 93

0.096 0.127 0.194 0.358

3.19 4.30 9.26 8.11

3.19 3.52 9.57 8.34

2.50 2.93 4.01 3.77

Lipids define a high permeability of creams and lotions in upper layers of epidermis and hair follicles.326 Ethers of fatty acids with single-atom spirits like isopropylmiristate, isopropylpalmitate, and isopropyllaurinate are very important chemicals as components of deep penetrating creams and lotions. Technical lecithin (60% natural phospholipids (major phosphatidylcholine), 30–35% plant oil, glycerol, et al.) is a basis for many nourishing (nutritive) creams due to its possibility of penetrating deep into the skin. Silicon wax and oils easily penetrate into hair follicles via friction (rubbing), do not induce inflammation due to low surface tension, and do not influence the thermal balance of the skin. Emulsions are oils in water and water in oils, with the sizes of particles more than 0.1 µm. Emulsions like oils in water are widely used in cosmetics for deep penetration into the skin, as providers for biologically active substances, and other effects. Nourishing (nutritive) creams easily penetrate to the deep layers of epidermis, and prevent transepidermal water loss (TEWL). Skin hydration can be provided by two types of mechanisms—osmotic or physiological. As the hydrating substances,

Chemical Agent Delivery

151

sodium lactate, pyrrolidonecarboxylic acid, derivatives of amino acids and sugars, proteins, and mucopolysaccharides are usually used. As a hygroscopic component, glycerol is often used (usually less than 10% in composition). At present, glycerol is usually replaced by propylene glycol. To understand the transport and barrier functions of the skin, it is important to have knowledge of the water and ions distribution within the different layers.249,250 Currently, in the market of cosmetic products, numerous creams and lotions providing enough deep impregnation of the skin are available. Many cosmetic emulsions, gels, and lotions for skin hydration use gyaluronic acid (the best for TEWL); sea collagen (also good for TEWL); liposomes and nanospheres—fat particles (for transportation of biologically active substances to the deep layers of epidermis and hair follicles). As a rule, creams on the basis of liposomes and nanospheres are used after application of peeling creams, for example, creams containing α-hydroxy acids (AHAs) or abrasive creams, which allow making skin relief more smooth and penetrative for liposomes and nanospheres. Liposomes have been suggested as a vehicle for dermal and transdermal drug delivery, but the knowledge about the interaction between lipid vesicles and human skin is still poor. In Ref. 327, the visualization of liposome penetration into human skin in vitro using a confocal microscope was done. Liposomes were prepared from phospholipids in different compositions and labeled with a fluorescent lipid bilayer marker. Liposome compositions containing DOPE (dioleylphosphatidylethanolamine) were able to penetrate deeper into the SC than that from liposomes without DOPE; the liposomes containing DOPE may fuse or mix with skin lipids in vitro and loosen the SC lipid bilayers. Among the factors not affecting SC penetration were negative charge, cholesterol inclusion, and acyl chain length of the phospholipids. Fusogenicity of the liposome composition appears to be a prerequisite for skin penetration. The liposome sizes determined by the quasielastic light scattering method were in the range 40–76 nm. The penetration depth into the skin in 72 hr was in the 2–38 µm range. It should be noted that effective mixing of liposomes containing DOPE with SC lipid bilayer happens in a few minutes.327

7.3 Diffusion of Macromolecules in Tissues As the SC in skin, the mucus that coats the surfaces of the gastrointestinal, respiratory, and reproductive tracts provides the outermost barrier against pathogens,328 and, thus, may give some difficulty in the delivery of immersion agents. Secreted mucins, the gel-forming component of mucus, are large proteins (10–40 Da, 0.5– 5 µm long) that are highly glycosylated (up to 80% sugar by weight) and that form an entangled viscoelastic gel.328 The mucin fibers with diameter of 3–10 nm are organized in a mesh with spacing of 90–3000 nm, which is much larger than the hydrodynamic diameter of most globular proteins (2–15 nm). In the model in which the effective radius of the mesh spacing is greater than the hydrodynamic radius of

152

Optical Clearing of Tissues and Blood

the diffusing particle, and no interaction between the solute and the polymer gel is assumed, the ratio of diffusion in a gel and diffusion in water is described as.328
   Dg π rs + r f 2 , = exp − Dw 4 rg + rf

(84)

where Dg and Dw are the diffusion coefficients of the particle in the polymer gel and water, respectively; and rs , rf , and rg are the radii of the particle, of the gel fiber, and of the mesh effective spacing, respectively. The results of experiments on the diffusion of macromolecules in human midcycle cervical mucus done using fluorescence recovery after photobleaching (FRAP) and fluorescence microscopy techniques are presented in Table 12. The ratio of diffusion coefficients, Dmuc /DPBS , indicates the speed with which a protein diffuses in mucus compared with a saline buffer. Most proteins (15–650 kDa) diffused as fast in mucus as they did in water (Dmuc /DPBS = 0.84–1.1). Two proteins, human lysozyme and cationized horse ferritin, seemed to be slowed slightly (Dmuc /DPBS = 0.73–0.75). However, two human immunoglobulin molecules, IgM and IgM-Fc5µ , were slowed significantly (Dmuc /DPBS = 0.33–0.41). Two major mechanisms may stop particles from readily diffusing through a mucus gel. Particles can stick to mucin fibers or they can be hindered by the size of the mesh spacing between the mucin fibers. Most of the proteins tested were able to diffuse in the human mucus samples as fast as they diffuse in water, indicating that the size of the mesh spacings did not significantly hinder or block the diffusion of any proteins tested and any other molecule with a comparable size.328 Table 12 Diffusion of proteins in human cervical mucus.328

Proteins

MW, kDa

Human lysozyme Horse myoglobin Chicken lysozyme Porcine pepsin Human immunoglobulin (IgM-Fabµ ) Human α-amylase Human lactoferrin Human immunoglobulin (IgG) Human immunoglobulin (IgA) Human immunoglobulin (IgM-Fc5µ ) Horse ferritin Cationized horse ferritin Human immunoglobulin (IgM)

15 17 14 35 50 59 82 152 160 ∼300 650 650 950

d, nm

Dmuc /DPBS

3.5 3.8 4.1 4.5 6.0 5.5 6.1 11 11 11 12 12 16

0.75 ± 0.23 0.84 ± 0.21 1.1 ± 0.2 0.98 ± 0.02 0.94 ± 0.23 0.90 ± 0.02 0.86 ± 0.21 0.87 ± 0.19 0.85 ± 0.02 0.41 ± 0.12 0.88 ± 0.18 0.73 ± 0.32 0.33 ± 0.11

Microinjection in combination with microspectrophotometry were employed to measure the diffusivity of proteins along the longitudinal axis of single intact

Proteins Cytochrome c Myoglobin Hemoglobin Catalase Ferritin Earthworm hemoglobin

MW, kDa 12.4 17.0 64.5 247.5 450–900 3,700

d, nm 3.1 3.5 5.5 10.2 12.2 30

Soleus Dm (n), 10−8 cm2 /s 13.0± 0.6(27) 12.5± 1.3(12) 6.3± 0.5(11) 2.6± 0.4(12) 0.6± 0.1(10) 0

EDL Dm (n), 10−8 cm2 /s 16.2± 0.6(22) 18.7± 0.8(12) 6.2± 0.4(13) 1.9± 0.2(16) 0.9± 0.1(26) 0

Dw , 10−8 cm2 /s 120 112 74 43 38 13

Soleus Dm /Dw 0.11 0.11 0.08 0.06 0.016 0

EDL Dm /Dw 0.13 0.19 0.08 0.04 0.02 0

Table 13 Diffusion of proteins in muscle (Dm ) and aqueous solutions (Dw ) at 22◦ C.329 Extensor digitorum longus (EDL) and soleus muscle of adult rats were studied; n is the number of experiments, where each fiber was used once for microinjection.

Chemical Agent Delivery 153

154

Optical Clearing of Tissues and Blood

skeletal muscle fibers taken from either the extensor digitorum longus (EDL) or soleus muscle of adult rats.329 Experimental data are summarized in Table 13 and Fig. 78. In general, the D values of the same proteins measured in muscle fibers are 6–63-fold lower than in diluted aqueous solution. This result is quite different from protein diffusion in the cervical mucus (Table 12), where for the same proteins, such as ferritin and myoglobin, the diffusion coefficient is only 0.84Dw and 0.88 Dw , respectively. It is important that the D value of the smaller proteins cytochrome c and myoglobin is significantly higher (p < 0.001) in the fast muscle type (EDL). The diffusion coefficient measured for myoglobin at physiological temperature (37◦ C) in soleus fibers (n = 24, not shown in Table 13) amounts to D = (22.0 ± 1.2) × 10−8 cm2 /s, or 1.76 times higher than at room temperature (22◦ C). No diffusive flux at all of earthworm hemoglobin (d = 30 nm) along the fiber axis could be detected.

Figure 78 Dependency of the diffusion coefficient (D) on molecular mass for diffusion in aqueous solution (empty symbols) and in different muscle fiber preparations (filled symbols) at 20◦ C.329 In the low molecular mass range (<1 kDa), the results obtained for the diffusion of ions, metabolites, and small inert molecules are plotted (filled circles), whereas on the right side of the figure (>10 kDa), the D values found for proteins are shown: filled triangles;329 filled squares [Ashley et al. (1974) and Maughan and Lord (1988)].329 For earthworm hemoglobin (3.7 × 106 Da), only the D for diffusion in water is shown, as it is for eukaryotic ribosomal particles—subunits and complete ribosome (empty squares); ×, translational D value for sarcoplasmic diffusion of myoglobin as derived from the results obtained by Livingston et al. (1983)329 for rotational myoglobin diffusion in bovine heart muscle.

Figure 78 gives a compilation of diffusion data obtained for low-molecularmass molecules (ions, ATP, Pi , phosphocreatine, O2 ) and for proteins in muscle preparations. Relative to their D values in water, intracellular diffusivities of small molecules were reduced by a factor of ∼2. The translational protein [cytochrome c (12.4 kDa), myoglobin (17.0 kDa), aequorin (∼20 kDa), hemoglobin (64.5 kDa), intrinsic muscle proteins (12–144 kDa)] diffusivities (Fig. 78) are

Chemical Agent Delivery

155

considerably more reduced in muscle fibers (∼tenfold for a molecular mass between 10 and 100 kDa). An abrupt decrease of sarcoplasmic diffusion with increasing molecular size is obvious for a molecular mass of >100 kDa. Proteins, such as glycerine-aldehyde-3-P-dehydrogenase (144 kDa), catalase (248 kDa), ferritin (450 kDa), and β-galactosidase (540), diffuse remarkably slowly, and earthworm Hb (3,700 kDa) probably does not diffuse at all in muscle cells. Figure 78 also shows the D values measured in a dilute aqueous solution for earthworm Hb (30 nm) and for both the subunits (38S and 59S) and the complete particle (81S) of eukaryotic ribosomes with dimensions 26 × 32 nm. Authors of Ref. 329 had concluded that (i) sarcoplasmic protein diffusion is strongly impaired by the presence of the myofilamental lattice, which also gives rise to differences in diffusivity between different fiber types; (ii) contractions do not cause significant convection in sarcoplasm and do not lead to increased diffusion transport; and (iii) in addition to the steric hindrance that slows down the diffusion of smaller proteins, diffusion of large proteins is further hindered when their dimensions approach the interfilament distances, and this molecular sieve property progressively reduces intracellular diffusion of proteins when the molecular diameter increases to more than ∼10 nm.

7.4 Enhanced Diffusion of Molecules Through Tissues 7.4.1 Occlusion247 Occlusion enhances the percutaneous absorption of a variety of compounds. The effect is relatively independent of the structure of the compound. For example, hydration of the SC appears to enhance diffusion of water as well the percutaneous absorption of a homologous series of alcohols, phenols, and steroids. Occlusion leads to a threefold increase in the percutaneous absorption of several steroids applied from acetone vehicles. Occlusion reduces or blocks TEWL and the evaporation of volatile solvents or compounds from the skin surface. In turn, this results in a profound (300–400%) increase in the water content of the SC. Most transdermal preparations are partially or completely occlusive. Partial occlusion may also be obtained with formulations based on petrolatum, ointments, or creams, although lotions and powders offer little occlusive activity. In addition, baths act to increase the water content of the SC and enhance percutaneous absorption. The effect of occlusion on the water content of the SC is relatively transitory, and typically returns to “normal” levels within 15 min after removal of an impermeable wrap. Since TEWL also returns to a normal level, it is likely that the reduction in barrier activity is also transitory. Water uptake by the SC under occlusive conditions is primarily localized in the corneocytes. Hydration appears to have very little influence on the structure or properties of the intercellular lipids. Full hydration of the SC by occlusion appears to provoke the formation of water pools associated with rough structures. These structures can be considered as small water channels, which reduce the diffusional

156

Optical Clearing of Tissues and Blood

pathlength and resistance for hydrophilic compounds. However, lipophilic as well as amphiphilic drugs may also profit from such shortened pathways. A possible mechanism of occlusion action is that the swelling of the corneocytes directly alters the skin barrier function. Swelling of the corneocytes may provide an alternative penetration pathway, i.e., by facilitating entry into the corneocytes, increasing the diffusivity of compounds through the corneocytes, or altering the structure of a minor lipid component. 7.4.2 Chemical enhancers In general, skin permeation enhancers act at the level of the SC.247 The molecular basis their activity can be attributed to include: (i) an increase in the partitioning of compounds into the SC; (ii) an increase in the diffusivity of the compound through the SC; and (iii) a change in the penetration pathway. The in vitro studies of passive transport of polar molecules like urea, mannitol, sucrose, and raffinose across intact and two-hour ethanol pretreated human epidermal membrane (HEM) and theoretical analysis of the hindered diffusion showed that permeation pathways of HEM can be characterized by membrane porosity.330 Effective pore radii estimates for intact HEM fell between 1.5 to 2.5 nm while similar estimates fell compactly between 1.5 to 2.0 nm for ethanol pretreated HEM. Thus, an approximately 100-fold increase in permeability for ethanol pretreated HEM relative to intact was explained by increased porosity of HEM at extraction of HEM lipids by ethanol pretreatment, while creating pores with effective radii that are quite small. Up to three orders of permeation decrease was found for large molecules (8000 Da) in comparison with the small ones (∼200 Da).331 The examination of macromolecules’ (up to 18 kDa) permeation through ethanol pretreated (2 hr) HEM yielded estimates of effective pore size for this biological membrane in the range 2.2–5.4 nm.331 Approximately twice larger pores radii at studies with larger molecular size may reflect the existence of a distribution of pore sizes; probe permeants of larger molecular size would then yield a larger average pore size than those determined with smaller molecular permeants. Such behavior was also observed in skin penetration studies where ethanol was topically applied.332–335 In these studies, ethanol reduced the barrier of the SC due to its interaction with the intercellular lipids, resulting in enhanced SC permeation of topically applied substances including aspirin. Effects of ethanol/propylene glycol composition on macroscopic barrier properties of the skin were also analyzed.333 Recently, ethanol evaporation through the skin was measured after oral intake of ethanol (0.30 to 0.52 g per kg of body weight) at skin sites differing in the thickness of the SC and the density of follicles and sweat glands.336 The selective sealing of the skin appendages had no significant influence on ethanol evaporation; this indicates that the evaporation of orally ingested ethanol occurs mostly through the SC lipid layers. Thus, an influence of ethanol on the penetration of topically

Chemical Agent Delivery

157

applied products can be expected. However, in the study of the authors of Ref. 336, orally administered ethanol had no effect on the penetration of a topically applied UV filter substance. Presumably, the available concentration of ethanol within the SC was too small (a theoretical maximum of 1.7 mg per 1 cm2 skin surface) to influence its permeation significantly. Therefore, the effect of topically applied substances should not be influenced by a single ethanol dose of 31.2 g as used in the study.336 Permeability of biological membranes may be induced not only by ethanol; a number of various chemical agents may serve as enhancers of membrane permeation.233,247 For example, such a polyenic antibiotic as Amphotericin B provides a twofold increase in water permeation through a cell membrane, more than 44-fold for glycerol, and more than 200-fold for urea. 233 Dimethylsulfoxide (DMSO), a popular aprotic solvent, is also a good enhancer. This is a natural substance derived from wood pulp; has a unique capability to penetrate living tissues and to associate with water, proteins, carbohydrates, nucleic acid, ionic substances, and other constituents of living systems; possesses hygroscopic and anti-inflammatory properties; and is FDA approved as a preservative of transplanting organs and for interstitial cystitis treatment (Table 1).337 A concentration of approximately 60% is required for activity of DMSO to disrupt the human skin barrier function and enhancement ratios of 20–200 have been reported.247 DMSO provides irreversible disruption of the SC, perhaps due to solubilizing the intercellular lipids and/or denaturing proteins. Unfortunately, DMSO has some side effects such as skin irritation, chemical instability, the degree of damage to the SC, and the characteristic taste/odor of its oxidation products. Some other polar aprotic solvents also have been used experimentally as enhancers, but only one of them, decylmethylsulfoxide, has found its way into practice.247 There are a great variety of surfactants that reduce the skin barrier function, classified as nonionic, anionic, and cationic depending on their charge at physiological pH.247 Certain unsaturated fatty acids increase percutaneous agent absorption by reducing the skin barrier function as well as by increasing the thermodynamic activity of compounds in some formulations.247 The monounsaturated fatty acid oleic acid (C18) is frequently chosen as a good enhancer for a wide variety of polar and moderately lipophilic compounds.167,247 Oleic acid increases TEWL in in vivo by approximately 1.5- to 2-fold, consistent with a general decrease in the skin barrier function. There is a well-established synergy between the enhancer activity of oleic acid and propylene glycol vehicles.167,247 7.4.3 Physical enhancers Occlusive conditions that can be provided by application of an impermeable wrap on the skin surface enhance the percutaneous absorption of a variety of compounds.247 Occlusion reduces or blocks TEWL and the evaporation of volatile solvents or compounds from the skin surface. Partial occlusion may also be obtained with formulations based on petrolatum, ointments, or creams. A possible

158

Optical Clearing of Tissues and Blood

mechanism of occlusion action is that induced swelling of the corneocytes may directly alter the skin barrier function by increasing the diffusivity of compounds through the swelled cells. The stripping of the SC represents the simplest approach to physically enhancing the percutaneous absorption of a compound across the skin [see Fig. 33(b)].186,247 Though it is not always suitable for therapeutic applications because of irritation responses, skin stripping is a very useful scientific tool for evaluating the maximum amount of agent percutaneous absorption that can be expected from a topical application. 338–340 The ultrasound or phonophoresis (sonophoresis) technique provides the enhanced absorption of low-molecular-weight compounds as well as proteins such as insulin.247 A continuous or pulse exposure of ultrasound of frequency from 20 kHz to 10 MHz at intensity of up to 3 W/cm2 can be applied for up to 10 min. The intensity is limited by heat production in the tissue. The enhancement activity of high-frequency ultrasound (5–10 MHz) is connected with the induction of convective pathways through hair follicles and disruption of the intercellular lipid lamellae; the cavitation phenomenon when small air bubbles are formed within the SC plays an important role in its permeation. A low-frequency ultrasound may be particularly suitable for enhancement.247 In vivo application of ultrasound of 20 kHz to skin of hairless rats for 1 hr resulted in a 100-fold increase in TEWL and sufficient delivery of insulin through the skin to reduce the blood glucose levels of the rats. Iontophoresis refers to the enhancement of agent percutaneous absorption by the application of moderate (0.5 V/cm2 ) voltages across the skin [see Fig. 33(c)].186,247 Iontophoresis is not restricted to charged ions, and the flux of uncharged molecules across the skin is also enhanced in a process termed electroosmosis. This results from the combination of a reduced SC barrier and an induced solvent convective flow. Iontophoresis appears to drive molecules through discrete sites located in the SC, such as hair follicles and sweat glands. A quantitative comparison of the flux of ions through appendages and through the intercellular lipid domain is estimated to be between 50% and 95% during iontophoresis.247 After a clinically relevant exposure of 0.16 mA/cm2 for 1 hr, the subsequent permeability of human skin in vitro was reduced tenfold. The effect exists during and after application of the current and is fully reversible after approximately 24 hr. Enhancement of percutaneous absorption by iontophoresis has been studied for a wide variety of agents.247 At higher voltages (5–200 V/cm2 ) and short pulse exposure, electroporation of biological membranes may occure247 that also provides agent permeation into a tissue. It was shown recently that laser-generated stress waves (photomechanical waves) can also permeabilize the SC.341–344 The permeabilization of the SC was first demonstrated with δ-aminolevulenic acid (ALA) as a probe. The permeability of the SC depends on the peak stress. The onset of the permeability of the SC is observed at ∼380 bar and increase with increasing peak stress. The efficiency of ALA transport through the SC is nonlinear. A small increase of the peak

Chemical Agent Delivery

159

stress, from 440 bar to 500 bar (14% increase in peak pressure), caused the fluorescence intensity (protoporphyrin IX concentration induced by ALA application) to increase by ∼200%. The application of stress waves does not cause any pain and discomfort and does not appear to affect the structure and viability of the skin. The change in the permeability of the SC is transient and its barrier function recovers within a few minutes. The increased permeability allows macromolecules to diffuse through the SC to epidermis and dermis. The maximum size of particles that have been transported through the SC is 100 nm in diameter.341 Thus, lasergenerated stress waves can facilitate the transdermal delivery of large particles and molecules, such as novel probes (carbon, gold, melanin nanoparticles,137 quantum dots, encapsulated molecular probes), encapsulated drugs, or plasmid DNA. The combined action of laser-stress waves and anionic surfactant, such as sodium lauryl sulfate (2% of w/v), enhanced the delivery of nanoparticles through the SC.341 The application of sodium lauryl sulfate increases the size of the channels in the biomembrane and delays the recovery of the SC barrier function. The synergism of light and surfactant action was manifested as a significant reduction of time interval for providing the similar SC permeation: only 5 min of the application of sodium lauryl sulfate was enough at laser single-pulse action (∼7 J/cm2 ), providing the peak pressure of ∼600 bar and stress pulse duration of ∼250 ns instead of a few hours without laser pulse. The hydrostatic pressure has been reported to increase the permeability of the plasma membrane of the human RBC and peripheral blood mononuclear cells; the presence of strong stress gradients is essential for membrane permeabilization.343,344 The typical duration of an individual laser pulse is about 80 nsec, which induces a stress wave with a leading edge (0–100%) of about 10 nsec and peak stress of about 750 bar; a repetition rate of about 1 Hz and approximately 20 expositions of stress wave are needed.342,343 The mechanisms involved in the permeabilization and recovery of the plasma membrane are not well known, but it was found that the plasma membrane permeabilizes in less than 60 msec and recovers in 10–80 sec. The typical radius of the cell membrane acquaporins, 0.15–0.3 nm, is too small to accommodate bigger molecules, such as dextran with Stokes’ radius of the order of 2 nm (∼3 kDa). Therefore, the probe molecules of such size cannot move through the aquaporin in its native state. The action of the stress waves may be to denature the channel protein and form a temporary enlarged pore through the cell plasma membrane. Analysis of the kinetics of stress-wave-induced membrane permeability for several cells yields a diffusion constant of (2.2 ± 1.3) × 10−7 cm2 /s. The reduction in the diffusion constant depends on the size of the molecule and the average size of the pores in the fibrous structure. For small molecules with hydrodynamical radius of approximately 1 nm, the reduction in the diffusion constant ranges from 0.06 for 100 nm pores, to 0.5 for 10 nm pores. Therefore, with D = 30 × 10−7 cm2 /s in water, the diffusion constant for small molecules in fibrous structures with ∼100 nm (∼10 nm) pores is D ∼ = 28 × 10−7 cm2 /s (D ∼ = 15 × 10−7 cm2 /s). Experiments suggest that the average pore size of the cytoskelton is 30–40 nm, and in some regions as small as 20 nm.

160

Optical Clearing of Tissues and Blood

Thus, the synergism of chemicals and stress waves is that the stress waves induce a transient increase of the membrane permeability that results in the diffusion of molecules, present in the medium, into the cell. It should be noted that the drug has to be presented in the extracellular medium, during the application of the stress waves. Addition of the drug after the cells are exposed to stress waves does not increase the drug cytotoxicity (as a test for changeable permeability). In summary, it was shown that laser-generated stress waves increase the permeability of the cell plasma membrane and the SC. The increase of the permeability allows the introduction of macromolecules into the cytoplasm as well as into the epidermis and dermis. Thus, stress waves have the potential to deliver chemicals topically and noninvasively into the deep layers of the skin. Laser-induced shock waves making cell plasma membrane more permeable was used as a valuable tool for loading cells with large molecules or small beads.345 This technology is called laser-assisted optoporation. As an alternative, local heating of the plasma membrane using continuous laser irradiation together with an absorbing dye was applied to increase membrane permeability.147,346 Local heating is reversible and relatively harmless to the irradiated cells, being based on the phase transition of membrane lipids from the gel phase to the liquid crystalline phase.66 Such dye-assisted laser optoporation was recently provided using considerably reduced laser power and illumination spot size than it was done earlier.147 An argon laser with a 488 nm wavelength was used in the experiments—0.7 µm or 1 µm light spot diameters on the cell surface with the corresponding power 3.4 mW or 7.2 mW were provided; thus, power densities of 1.0 MW/cm2 were achieved in both cases. Phenol red was used as a light-absorbing dye. Small circular dark spots on the cell surface appeared immediately after laser irradiation (1 or 2.5 s) and disappeared within about 5 min. They were related to transient changes in the membrane properties, which could be visualized using the fluorescent marker laurdan, and were probably due to a local increase in temperature. Cell viability was maintained by using light doses up to 2.5 MJ/cm2 applied for 1 s. The efficiency of cell transfection was measured on the basis of single cells using a green fluorescence protein (GFP) coding plasmid as a reporter gene. The transfection rates after laser irradiation were around 30% for younger subcultures (Chinese hamster ovary cells) and less than 10% for aging cells. This may be due to age-dependent changes in the phase transition of membrane lipids. A method of enhancing skin permeability by creating a lattice of microzones (islets) of limited thermal damage in the SC by applying a few consequent optical pulses was recently proposed.196,347 The method has been used to create the lattice of islets of damage with size of ∼75 µm and lattice pitch of ∼450 µm, using a special applicator and pulse lamp with the fluence rate of 14–36 J/cm2 and 20 ms pulse duration. The tested permeation ability of several index matching agents, including glucose and glycerol, has shown a highly graded translucence of the skin in vitro. The possible mechanism of the enhanced permeability of the SC is the phase transition of SC intercellular lipids from the gel phase to the liquid crystalline phase due to local heating.66,348

Chemical Agent Delivery

161

Alternative techniques of clearing agents’ delivery based on injection of an agent into the skin with needle-free injection gun, laser skin surface ablation, and their combinations are also under development.349,350 A diode laser source with 980 nm wavelength in conjunction with an artificial absorber on the skin surface was used to facilitate enhanced penetration of the topically applied skin clearing agent glycerol into in vivo hamster and rat skin.350 Such a technique provides a sufficient skin surface heating, which leads to keratinocyte disruption and possibly skin surface ablation of less than 20 µm, with a treatment site of 16 mm2 at laser beam scanning. Results indicate an improvement of the ability to deliver NIR light of 1290 nm up to 36% deeper into in vivo rodent skin using a laser fluence of less than 96 J/cm2 . Higher fluences caused unwanted thermal denaturation of skin tissue. The SC ablation can be provided directly by application of pulsed erbium lasers with wavelengths of 2790–2940 nm, corresponding to the strong water absorption band.351,352 Laser ablation of 12.6% of the surface area of porcine SC produced a 2.8- and 2.1-fold increase in the permeability constant (Pa ) for 3 H-hydrocortisone and 125 I-γ-interferon, respectively (see Table 14).351 These studies demonstrate that a pulsed (250 µs pulse width) laser with a wavelength of 2790 nm and 1 J/cm2 of fluence density can reliably and precisely remove the SC at 10–14 laser pulses, facilitating penetration of large molecules such as 125 I-γ-interferon that cannot penetrate intact skin. Among such modalities as skin microdermabrasion, iontophoreses, electroporation, and Erb:YAG (λ = 2940 nm) ablation, tested by the authors of Ref. 352, laser ablation showed the greatest enhancement of ALA permeation through pig skin samples. The laser fluence was found to play an important role in controlling the drug flux, producing enhancement ratios from 4- to 246-fold relative to the control (see Table 15). The skin permeation of ALA across microdermabrasiontreated skin was approximately 5–15-fold higher than that across intact skin. The application of iontophoresis or electroporation alone also increased the ALA permeation by approximately 15-fold and 2-fold, respectively. The incorporation of iontophoresis or electroporation with the resurfacing techniques (laser ablation or microdermabrasion) caused a profound synergistic effect on ALA permeation. Another example of a biological structure of which enhanced permeation is desirable is tooth biofilm. Oral cavity microbial infections are caused by bacteria growing as a multispecies biofilm containing colonies of microorganisms.353 Biofilm is matrix-enclosed microbial aggregates associated with each other and a solid surface of a tooth. The matrix is penetrated by fluid channels that conduct the flow of nutrients, waste products, enzymes, metabolites, and oxygen. Bacteria within biofilms have an increased (up to 1.5 × 103 ) resistance to antimicrobial agents. Ultrasound, electrical field, and laser-induced photomechanical waves have been used to enhance the efficacy of antibiotics in killing biofilm bacteria.354 The action of ultrasound is primarily mediated by heat (up to 52◦ C for 5 min exposure) and cavitation. Electrical fields should be applied for 24–48 hr in order to enhance the efficacy of antiobiotics. Laser-induced photomechanical waves provide peak

162

Optical Clearing of Tissues and Blood

Table 14 In vitro percutaneous drug delivery through porcine skin at SC laser ablation by Erb:YSGG laser (λ = 2790 nm, τp = 250 µs, fluence 1 J/cm2 , beam diameter 2w = 2 mm).351 Coefficient of permeability is defined by Eq. (38).

Coefficient of permeability, P a × 105 (cm/h) Treatment Control Tape-stripped Laser ablation

Ablation (%) 0 100 3.1 6.3 9.4 12.6

3 H-hydrocortisone

1.46 5.32 1.85 2.59 2.95 4.17

125 I-γ-interferon

43.8 142 53.4 67.4 80.1 94.4

pressure at 600 bar and rise times of 50 ns and 110 ns in duration. A single laser pulse allows for 1–5 min enhancement of drug delivery into biofilm due to the possible increase in fluid movement caused by pressure-induced deformation of the bacterial microcolonies and the matrix.

7.5 Skin Reservoirs for Topically Applied Clearing Agents The SC is functioning not only as a barrier against a clearing agent penetration into skin, but also as a reservoir for topically applied substances.355,356 Skin appendages, in particular sebaceous glands, also serve as reservoirs for clearing agents.118–120,338,357 For the development of technologies for topical application of clearing agents, the knowledge of the reservoir function is of fundamental interest. The following definition of SC as a reservoir was set in Ref. 356: a reservoir is no temporal stable state, but means the storage of a substance on a certain interval with a continuous, concentration dependent release; a reservoir exists, if a concentration gradient is formed within the SC in direction to the viable epidermis; applied substances are penetrating out of this reservoir into deeper skin layers; with decreasing concentration within the reservoir the rate of discharge wanes, wherefore for small concentrations a reservoir exists for a long time. Based on this definition, the long-term functioning of a reservoir of the SC in human skin was investigated in vivo.355,356 Using laser confocal scanning microscopy and the tape stripping method, the long-term reservoir of the SC was determined both qualitatively and quantitatively depending on the polarity of the applied formulation. Formulations with different physicochemical properties were used. As hydrophilic formulations, the fluorescent dyes patent blue V or sodium fluorescein in an aqueous vehicle were used, and as a lipophilic formulation, the fluorescent dye curcumin in paraffin oil was used. Further, amphiphilic formulations with patent blue V or sodium fluorescein in propylene glycol, a penetration enhancer, were investigated. The fluorescent dyes curcumin and sodium fluorescein were added in a concentration of 0.5% to these vehicles. In such experiments,

Portable laser (τp = 300 µs, 2w = 3 mm) Microdermabrasion, 3 s Microdermabrasion, 5 s Microdermabrasion, 10 s

Ablation type Passive Laser, pulse duration τp = 250 µs, beam spot size 2w = 7 mm

– – –

Laser fluence rate (J/cm2 ) – 1.2 1.4 2.1 2.6 1.2 × 2 pulses 1.4 × 2 pulses 3.5 3.5

Normalized ALA flux (µg/cm2 h) – 2.72 ± 0.52 7.37 ± 1.66 114.86 ± 7.30 165.31 ± 32.97 204.45 ± 25.42 150.15 ± 23.56 83.32 ± 19.21 151.60 ± 40.46 3.44 ± 0.65 4.95 ± 1.14 10.15 ± 2.59

Original ALA flux (µg/cm2 h) 0.67 ± 0.13 1.43 ± 0.25 3.29 ± 1.61 56.34 ± 3.57 81.04 ± 16.15 100.21 ± 12.45 73.65 ± 11.54 40.87 ± 9.41 13.77 ± 3.64 3.31 ± 0.62 4.76 ± 1.09 9.75 ± 2.48

5.13 7.39 15.15

ER – 4.06 11.00 171.43 246.73 305.15 224.10 124.36 226.27

Table 15 Fluxes and enhancement ratio (ER) of 5-aminolaevulinic acid (ALA) across pig (1-week old) skin measured in vitro at treatment using Erb:YAG lasers (2940 nm) and microdermabrasion.352 The original flux was directly calculated from the flux across partially laser- or microdermabrasion-treated skin; the normalized flux was calculated by extrapolating the original flux data of the treated area to an area of 100% exposure and calibrated by the flux of the control group; ER is the normalized flux of the treated group/flux of control group; each value represents the mean ± SD (n = 4); for ALA flux measurements, the diffusion cell with the available area for skin sample of 0.785 cm2 was used; the receptor (pH 5 citrate-phosphate buffer) and donor (0.5% (w/v) ALA in pH 5 buffer) compartments were maintained at 37◦ C.

Chemical Agent Delivery 163

164

Optical Clearing of Tissues and Blood

paraffin oil and propylene glycol with refractive indices of 1.440 and 1.433, respectively, may be considered as optical immersion substances and fluorescence dyes as probes, showing immersion agents’ distribution in the skin. The quantitative formation of the SC long-term reservoir was crucially dependent on the polarity of the used vehicle.355,356 This vehicle influences both the amount of substances that penetrates into the skin and the distribution within the SC. The fluorescence signals of curcumin and sodium fluorescein were measured at a wavelength above 600 nm after excitation at 488 nm at 5, 24, 48, 72, and 120 hr after application. After treatment with sodium fluorescein in water, the fluorescence was detected both intra- and intercellular within the SC up to 120 hr after application. The fluorescence of the lipophilic curcumin applied in paraffin oil was measured between the superficial SC cells up to two days after application. The application of sodium fluorescein in propylene glycol led to an intra- and intercellular fluorescence (up to 5 days), as well as a fluorescence inside the follicular infundibulum (up to 5 days). These results indicate that the different localization of the reservoir, depending on the time after application, is determined by the physicochemical properties of both the dyes and the vehicles. The fluorescences of sodium fluorescein, applied in water, and curcumin, applied in paraffin oil, were only detected at the follicular orifices in the gaps between the follicular infundibulum and the hair shaft, and the SC, respectively.355,356 In contrast, a fluorescence of sodium fluorescein, applied in propylene glycol, was observed in the follicular infundibulum. It was detected up to 120 hr after topical application. Sodium fluorescein applied in water was mainly distributed intracellular, whereas, the application in propylene glycol resulted in both an intra- and intercellular localization of the dye. This difference in distribution was caused by the vehicle. Water is known to penetrate transcellularly into the SC, presumably causing the transport of the solved dye on this pathway even after 24 hr. In contrast, the intercellular permeation across the lipids is the main pathway of most substances. This pathway could additionally be opened for the hydrophilic sodium fluorescein by the vehicle propylene glycol, a known penetration enhancer disturbing the lipid layers. In contrast, the lipophilic curcumin applied in paraffin oil was only localized between the corneocytes on the skin surface and in the superficial SC.355,356 This result can be explained by the reduced amount of penetration of oils into SC, plus the good solubility of the dye and the oil within the intercellular lipids. In the study described in Refs. 355 and 356, a follicular long-term reservoir was only observed for the hydrophilic sodium fluorescein after application in propylene glycol. This result indicates that the follicular penetration also depends on the physicochemical properties of both the vehicle and the dye. Water and paraffin oil appear unsuited to produce the possible follicular penetration and storage. A follicular penetration of dyes was also reported for the application in emulsions,338 or solvents such as ethanol and glycerol,118–120 or 5 µm microspheres containing dye and suspended in silicon oil at w/w concentration of 4%.357 The penetration depth and reservoir properties of human skin in vivo for methylene blue

Chemical Agent Delivery

165

(MB) and indocyanine green (ICG) dissolved in ethanol/glycerol solvents were recently studied.118–120 These studies were applied for the improvement of sebaceous glands functioning at photodynamic acne treatment.118–120 In these studies, the most intensive sebaceous glands staining was just after the 15–20 min massage and heating procedure. It was also shown that highly porous nylon microspheres suspended in silicon oil provide penetration depth of MB into hairless rat skin in vivo up to 150 µm in 2 hr and 400 µm in 26 hr.357 The results of the study described in Refs. 118–120 and 355–357 led to the assumption that the better the penetration into the SC and the follicles, the longer the reservoir will remain there. Two pathways of release of the dyes from the reservoir are possible: (i) the desquamation of the SC and, the release of sebum, respectively, or (ii) the penetration into the viable tissue. In conclusion, it should be noted that for topically applied UV skin filters, their efficiency may be significantly reduced if an inappropriate cosmetic composition is used as a ground material for such a filter. Two main effects can be important: (i) reduction of light scattering in SC due to optical immersion and (ii) inhomogeneous distribution of the topically applied substances.340 Both effects lead to reduction of UV filters’ efficiency: first because of fewer interactions of migrating photons in skin with sunscreen material at less scattering, and second because of forming islands free of a sunscreen that are not blocking UV radiations. The second problem is analyzed in detail in Ref. 340.

8 Applications

8.1 Glucose Sensing 8.1.1 Introduction Approximately 17 million people in the USA (6% of the population) and 140 million people worldwide (this number is expected to rise to almost 300 million by the year 2025) suffer from diabetes mellitus.358,359 The development of a noninvasive test method would considerably improve the quality of life for diabetic patients, facilitate their compliance for glucose monitoring, and reduce complications and mortality associated with this disease. Noninvasive and continuous monitoring of glucose concentration in blood and tissues is one of the most challenging and exciting applications of optics in medicine. The major difficulty preventing development and clinical application of a noninvasive blood glucose sensor is associated with the very low signal produced by glucose molecules. This results in low sensitivity and specificity of glucose monitoring.256,360 8.1.2 Tissue and blood scattering spectroscopy The concept of noninvasive blood glucose sensing using the scattering properties of blood as an alternative to spectral absorption and polarization methods156,360 for the monitoring of physiological glucose concentrations in blood of diabetic patients has been under intensive discussion for the last decade.52−54,68−70,80,81,115,153, 154,156−159,361−363 Many of the considered early effects, such as RBC size, refractive index, packing, and aggregation changed under glucose variation are important for glucose monitoring in diabetic patients. Indeed, at physiological concentrations of glucose, ranging from 40 to 400 mg/dl, the role of some of the effects may be changed, and some other effects, such as glucose penetration inside the RBC and the following hemoglobin glycation may be important.68–70,115,358 Noninvasive determination of glucose was attempted using light scattering of skin tissue components measured by a spatially resolved diffuse reflectance54,362 and NIR frequency-domain reflectance techniques.52 Both approaches are based on change in glucose concentration, which affects the refractive index mismatch between the interstitial fluid and tissue fibers, and hence µ
s . A glucose clamp experiment (the concentrations of injected glucose and insulin are manipulated to result in a steady concentration of glucose ever a period of time156 ) showed that δµ
s at 650 nm qualitatively tracked changes in blood glucose concentration for the volunteer with diabetes studied (Fig. 79).54 The distances between the source 167

168

Optical Clearing of Tissues and Blood

and detector fibers were in the range: rsd = 1–10 mm, which corresponds to the approximate 0.5–5 mm in tissue depth upon which µ
s is determined. Drift in µ
s that was independent of glucose prevented statistical analysis and was attributed by the authors to other physiological processes contributing to δµ
s .54 Changes in µ
s did not exclusively result from changes in the refractive index of the interstitial fluid caused by increased glucose concentration. The spatially resolved reflectance measurements (at 800 nm and rsd = 0.8–10 mm) and oral glucose tolerance test were done to study five healthy volunteers and thirteen volunteers with type 2 diabetes using a probe continuously attached to the abdomen.362 For volunteers without diabetes, 80% of measurements showed tracking between δµ
s and blood glucose concentration, and the other 20% showed no correlation. For the volunteers without diabetes, 73% of the measurements resulted in a calibration model for µ
s versus blood glucose concentration. A poor correlation between the measured δµ
s and glucose concentration in these experiments may be connected with the sensitivity of the used probe (with a large probing depth) to the vascular effect of glucose (glucose-induced temporal variations of the blood flow in the skin and subcutaneous tissue).156

Figure 79 Reduced scattering coefficient at 650 nm (dots) and blood glucose concentration (solid curve) measured on a volunteer with insulin-dependent diabetes mellitus during a double clamping experiment.54 A multichannel CCD-based spatially resolved fiber-optic back-reflectance spectrometer and a neural network software were used to measure and extract the optical properties. Reflectance measurements were collected at 15 s intervals for ∼5 hr, and skin and room temperature were monitored throughout the course of the experiment; the volunteer remained as still as possible—food and drink were not permitted. Reprinted with permission.

The response of a nondiabetic male subject to a glucose load of 1.75 g/kg body weight, as a standard glucose tolerance test, was determined by continuously monitoring the product of nµ
s measured on muscle tissue of the subject’s thigh using a portable frequency-domain spectrometer (Fig. 80).52 The refractive index n of the interstitial fluid modified by glucose is defined by Eq. (57). As the subject’s blood glucose rose, the nµ
s decreased. Figure 80(b) shows the correlation plot obtained from the data of Fig. 80(a). The correlation plot fits well to a simple physical model based on the Rayleigh–Gans approximation and accounting for the refractive index matching concept. Key factors for the success of this approach are the

Applications

169

precision of the measurements of the reduced scattering coefficient and the separation of the scattering changes from absorption changes, as obtained with the NIR frequency-domain spectrometer.52 Evidently, other physiological effects related to glucose concentration could account for the observed variations of µ
s and, as it was mentioned earlier, the effect of glucose on the blood flow in the tissue may be one of the sources of the errors in µ
s measurements.

(a)

(b) Figure 80 Glucose tolerance test performed on a human subject with the portable frequency-domain (120 MHz) NIR (850 nm) spectrometer.52 (a) At time t = 45 min the subject ingested a glucose load of 160 g of table sugar (1.75 g/kg body weight); the solid curve is the continuous measurements of nµ
s on the thigh of the subject, n is defined by Eq. (57); the open circles indicate blood glucose concentration as determined by a home blood glucose monitor; the data acquired every 30 s were averaged in sets of five to produce the plot. (b) The corresponding to (a) correlation between the measured blood glucose and product nµ
s averaged over a time of 2.5 min centered on the time the finger was lanced for the measurement. The error in nµ
s is the standard deviations of the five measurements averaged to get a single point. The error in blood glucose concentration is estimated to be ± 2.5 mg/dl. The solid line is the theoretical result according to the Rayleigh–Gans model.

170

Optical Clearing of Tissues and Blood

The so-called occlusion spectroscopy approach is one that is based on light scattering from RBCs.153,154,156 This method suggests a controlled occlusion of finger blood vessels to slow blood flow in order to provide that the shear forces of blood flow be minimal and, thus, to allow RBCs to aggregate. Change in light scattering upon occlusion is measured. Occlusion will not affect the rest of the tissue components, since scattering properties of aggregated RBCs differ from those of the nonaggregated ones and from the rest of the tissue. As it was already discussed, change in glucose concentration affects the refractive index of blood plasma, and hence affects blood light scattering at occlusion due to refractive index match/mismatch between aggregates and plasma. Occlusion spectroscopy differs from that of spatially resolved reflectance and frequency-domain measurements in that it proposes measurements of glucose in blood rather than in the interstitial fluid. The occlusion spectroscopy method was tested in a human study using a hyperinsulinemic-hypoglycemic clamp.154 This technique offers the potential of directly measuring the change in the refractive index of blood plasma, but in clinical studies many other factors affecting scattering of RBCs and their aggregates should be accounted for: (i) the complexity of the RBC aggregation phenomenon and its dependence on glucose concentration and other pathological conditions and diseases;156,275,277 (ii) the effect of glucose on the shape and structure of RBCs. 8.1.3 OCT measurements An optical coherence tomography technique has recently been proposed for noninvasive assessment of glucose concentration in tissues.156–159,361,363,364 The high resolution of the OCT technique may allow high sensitivity, accuracy, and specificity of glucose concentration monitoring due to precise measurements of glucoseinduced changes in the tissue optical properties from the layer of interest (dermis). Unlike the diffuse reflectance method, OCT can provide depth-resolved qualitative and quantitative information about tissue optical properties of the three major layers of human skin: dead keratinized layer of squames (stratum corneum of epidermis); prickle cells layer (epidermis); and connective tissue of dermis. Dermis is the only layer of the skin containing a developed blood microvessel network. Since glucose concentration in the interstitial fluid is closely related to the blood glucose concentration, one can expect glucose-induced changes in an OCT signal detected from the dermis area of the skin. Two methods of OCT-based measurement and monitoring of tissue glucose concentration were proposed: (i) monitoring of the tissue scattering coefficient, µs , as a function of blood glucose concentration using standard OCT157–159 and (ii) measurement of glucose-induced changes in the refractive index, n, using a novel polarization-maintaining, fiber-based dualchannel phase-sensitive optical low-coherence reflectometer (PS-OLCR).361 The experiments were performed with a portable OCT system with the central wavelength of 1300 nm, power of 0.5 mW, and coherence length and lateral resolution of approximately 14 µm and 12 µm, respectively.157–159,364 The authors reported results obtained from phantom (aqueous suspension of polystyrene

Applications

171

microspheres and milk), animal (27 New Zealand rabbits and 13 hairless Yucatan micropigs), and human (20 healthy volunteers in 24 experiments) studies. OCT images were obtained from skin (ear of the rabbits, dorsal area of the micropigs, and arm of the volunteers). The slopes of the OCT signals were calculated at a depth of 150–900 µm. Glucose administration was performed using: (i) intravenous bolus injections for rapid increase of blood glucose concentration, (ii) an intravenous clamping technique for slow, controlled changes of the blood glucose concentration in animal studies, and (iii) standard oral glucose tolerance test (OGTT) in human studies. Blood samples were analyzed using OneTouch (Lifescan Inc., Milpitas, CA), HemoCue (Ryan Diagnostic, Inc., Naperville, IL), and Vitros 950 (OrthoClinical Diagnostics, Inc., Raritan, NJ) blood glucose analyzers. First, an OCT image of the layers of skin is taken, and the OCT signal as a function of depth is evaluated. The slope of the portion of the plot in the dermis layer is used to calculate µs . In an anesthetized animal skin experiment, OCT images demonstrate that glucose affects the refractive index mismatch in skin and decreases µs .364 The slope of the OCT signal versus depth line is determined and is correlated with the concentration of blood glucose (Fig. 81).

Figure 81 Representative OCT signals obtained from Yucatan micropig skin during a glucose clamping experiment at low and high blood glucose concentration (top) and a part of the OCT signal in the dermis area with the linear fit of the OCT signals in this layer (bottom).159

Typical results obtained in the clinical studies are shown in Fig. 82. OCT images and blood samples were taken from left and right forearm. Good correla-

172

Optical Clearing of Tissues and Blood

tion between the increase of the blood glucose concentration and decrease of the smoothed OCT signal slope has been observed at the depth of 200–600 µm during OGTT. Measurements performed in the layers of epidermis and upper dermis either did not show changes in the OCT signal slope at variations of blood glucose concentration or the changes were very weak. Most likely it is due to a gradient of glucose concentration from dermal blood microvessels to the SC. Thus, sensitivity and accuracy of the OCT measurements of blood glucose concentration would be maximal in the regions of developed blood microvessel network (that is, the dermis area).

(a)

(b) Figure 82 Slope of OCT signal and blood glucose concentration versus time. OCT images and blood samples were taken from human skin of (a) left and (b) right forearm during an OGTT.158 Copyright © 2002. American Diabetes Association. Reprinted with permission.

A comparison of results obtained from the skin of the rabbit ear during bolus glucose injection experiments and the skin of micropig during glucose clamping showed that the OCT signal slope changed approximately 8%/mM during the bolus glucose injection experiment and 2%/mM during glucose clamping. That suggests

Applications

173

the possibility of tissue physiological response to the sharp increase of analyte concentration in the interstitial fluid during bolus glucose injection experiments. The results obtained in phantom, animal, and clinical studies demonstrated the potential of the OCT technique to detect small glucose-induced changes in the scattering coefficient of the turbid media with high accuracy and sensitivity. However, additional studies should be performed on: (i) reduction of noise associated with speckles and tissue inhomogeneity; (ii) development of algorithms and methods for compensation of motion artifacts; and (iii) approbation of the system in clinical studies involving diabetic patients. Although, most likely, OCT-based glucose sensors may need calibration with invasive glucose sensors, they may dramatically reduce the number of invasive measurements and provide continuous monitoring of the blood glucose concentration. A question of specificity of the OCT technique to monitor blood glucose concentration in tissues has been addressed.159 The experimental and theoretical analysis of the influence of several physical and physiological parameters (such as altering the refractive index mismatch between the interstitial fluid and scattering centers, and structural modifications in tissue due to changes in glucose concentration) on the OCT signal slope was performed. Obtained results demonstrate that: (i) several body osmolytes may change the refractive index mismatch between the interstitial fluid and scattering centers in tissue—however, the effect of the glucose is approximately one to two orders of magnitude higher; (ii) an increase of the interstitial fluid glucose concentration in the physiological range (3–30 mM) may decrease the scattering coefficient by 0.22%/mM due to cell volume change; (iii) stability of the OCT signal slope is dependent on tissue heterogeneity and motion artifacts; and (iv) moderate skin temperature fluctuations (±1◦ C) do not decrease accuracy and specificity of the OCT-based glucose sensor—however, substantial skin heating or cooling (several ◦ C) significantly changes the OCT signal slope. These results suggest that the OCT technique may provide blood glucose concentration monitoring with sufficient specificity under normal physiological conditions. A new differential phase contrast OCT-based method (PS-OLCR) of monitoring glucose-induced changes in tissue optical properties has also been proposed.361 While conventional OCT uses the detection and analysis of the intensity of backscattered optical radiation, phase-sensitive OCT utilizes the phase information obtained by probing a sample simultaneously with two common-path lowcoherence beams. Variations in the sample refractive index will be exhibited in the phase difference, ϕ, between these two beams. The PS-OLCR technique is capable of measuring Å/nm scale path length change between the beams (associated with the phase difference as (λ/4π)ϕ) in clear and scattering media. The theoretical and experimental pilot studies on application of PS-OLCR for noninvasive, sensitive, and accurate monitoring of analyte concentration were reported by the authors of Ref. 361. They studied concentration-dependent changes of phase, dϕ/dC, and refractive index, dn/dC, in aqueous solutions of glucose, CaCl2 , MgCl2 , NaCl, KCl, KHCO3 , urea, bovine serum albumin (BSA), and globulin in clear and turbid tissue-like media (Table 16).361 Obtained results demonstrate: (i) good agreement between refractive indices measured with the PS-OLCR



lit

,

deg mM−1

dϕ dC

dϕ dC exp



,

deg mM−1




dn dC



mM−1




lit

× 105 ,

dn dC exp



mM−1




× 105 , Cphysiol , mM




 dn × 105 dC exp 59.23 1.20 0.68 9.61 1.66 4.42 4.94 0.019 0.022 × 10−5 , and mg dl−1 , Cphysiol ×

D-glucose (C6 H12 O6 ) 6.85 5.89 2.55 2.20 3–30 6.66 6.43 2.48 2.39 2.1–2.6 Calcium chloride (CaCl2 ) 6.04 6.08 2.25 2.26 0.7–1.0 Magnesium chloride (MgCl2 ) Sodium chloride (NaCl) 2.63 2.58 0.98 0.96 135–145 Potassium chloride (KCl) 2.58 2.97 0.96 1.11 3.5–5 2.79 2.97 1.04 1.11 22–26 Potassium bicarbonate (KHCO3 ) Urea (CH4 N2 O) 2.28 2.21 0.85 0.82 2.9–8.9 0.54 0.50 0.20 0.19 0.35–0.45 Bovine serum albumin (BSA)∗ 0.51 0.40 0.19 0.15 0.26–0.41 Bovine γ-globulin∗ * −dϕ/dC, dn/dC, and Cphysiol for bovine serum albumin (BSA) and globulin are shown in deg dl mg−1 , dl mg−1 respectively.

Analyte




Table 16 Concentration-dependent changes in the phase shift and the refractive index for major analytes in the body obtained from the literature (lit ) (589 nm) and measured using the PS-OLCR technique (exp) (1300 nm).361 The physiological range of the analytes concentration Cphysiol and its maximal possible effect on the refractive index change (last column) are also shown. The dϕ/dC values for the listed analytes were obtained in a 485 µm thick cell.

174 Optical Clearing of Tissues and Blood

Applications

175

technique and the conventional white-light refractometer (and previously reported in the literature for the visible spectral range); (ii) the effect of glucose on dn/dC is approximately one to four orders of magnitude greater than that of the other analytes at physiological concentrations; (iii) good agreement between results obtained in translucent and scattering media, suggesting that PS-OLCR could be applied for in vivo measurements; and (iv) high sub-mM sensitivity of PS-OLCR for measurement of glucose concentration. Like other scattering techniques, the detected phenomenon in OCT is the effect of glucose on the refractive index of the interstitial fluid. However, it does not allow for blood circulation and temperature changes. Unlike the spatially resolved back-reflectance and frequency-domain methods that use larger measuring volume and span multiple layers in tissues,52–54,156,362 OCT offers certain advantage, as it limits sampling depth to the upper dermis without unwanted signal from other layers. Precise sampling is very important in glucose monitoring in tissue because glucose uptake is different in different tissue layers, being lowest in connective tissue and smooth muscle, and highest in adipose tissue and skeletal muscle.365 Besides, when blood glucose changes rapidly, there is a time lag of 10–25 min, resulting in a transient difference between the blood and subcutaneous glucose concentrations, which should be accounted for.365

8.2 Speckle Technologies 8.2.1 Speckle topography and tomography Methods of speckle topography and tomography are rapidly progressing.7,14,15,23,24,29–31,55,85,86,143,366–369 Local statistical and correlation analysis offers much promise as a method for topographic mapping and structure monitoring of tissues. Local estimates of the intensity probability density distribution function p(I ) [see Eq. (64)], normalized statistical moments [the contrast VI and some other higher moments (for instance, the asymmetry coefficient Qa ) are usually employed, see Eqs. (63)], and correlation characteristics [see Eqs. (85)–(87)], are highly sensitive to structure parameters of an object, such as the correlation length Lc and the standard deviation σL of optical thicknesses (altitudes) of inhomogeneities or the relevant correlation length Lφ and the standard deviation σφ of phase fluctuations of the boundary optical field.23,24,31,55,82,83,85 In addition to the first-order speckle statistics [see Eqs. (63) and (64) and Fig. 29], statistics of the second order show how fast the intensity changes from point to point in the speckle pattern (see Fig. 14), i.e., characterize the size and the distribution of speckle sizes in the pattern. The statistics of the second order are usually described in terms of the autocorrelation function (AF) of intensity fluctuations,   g2 (ξ) = I (ξ + ξ)I (ξ) ,

(85)

176

Optical Clearing of Tissues and Blood

and its Fourier transform, representing the power spectrum of a random process; where ξ ≡ x or t is the spatial or temporal variable and ξ is the change in the variable. The angular brackets  in Eq. (85) stand for the averaging over an ensemble or time. To describe comparatively small intensity fluctuations, it is convenient to employ an autocorrelation function g˜ 2 of the fluctuating intensity component and the corresponding structure function DI ,    g˜ 2 (ξ) = I (ξ + ξ) − I  I (ξ) − I  ,  2  DI (ξ) = I (ξ + ξ) − I (ξ) ,   DI (ξ) = 2 g˜ 2 (0) − g˜ 2 (ξ) .

(86)

(87)

Analysis is usually performed in terms of normalized autocorrelation and structure functions. An autocorrelation function is preferable for the analysis of intensity fluctuations caused by comparatively large-size inhomogeneities in the scattering object. At the same time, the structure function is more sensitive to small-scale intensity oscillations. A tissue under study can be considered as an irregular system of lenslets with definite statistical characteristics that display intensity fluctuations similar to those observed when a focused laser beam is scanned over the surface of an object.31,369 Typically, intensity fluctuations include two components. The first component is a background with relatively small and comparatively smoothly varying amplitude. The second component is represented by infrequent high-intensity pulses related to matched inhomogeneities (see Figs. 29(a) and 29(b) for skin strippings; the distances between the plane of the waist of the incident laser beam and the object and between the object and the photodetector are matched with the effective focal length of the inhomogeneity, which ensures effective reimaging of the waist of the laser beam into the observation plane). Inhomogeneities can be classified by analyzing the contrast VI and the asymmetry coefficient Qa as functions of the distance between the waist plane of the laser beam and the surface of an object. Using this approach, statistical distributions of fluctuations of the refractive index of a medium can be reconstructed.24,85 The fact that VI and Qa abruptly increase when the ratio of the radius of the laser beam to the correlation length satisfies the condition w/Lφ ∼ 1 is a direct manifestation of the microfocusing effect in the far-field diffraction zone.31,369 The growth in the ratio w/Lφ is accompanied by a decrease in the quantities VI and Qa , which reach the values corresponding to completely developed speckles. These effects can occur in the case of weakly scattering objects, such as thin tissue layers or cellular monolayers, when Lc ∼ Lφ ∼ d (where d is the thickness of the sample), Lc  λ, and the standard deviation σφ related to phase fluctuations of the field is completely determined by fluctuations of the refractive index δn. Multiple scattering is characteristic of optically thick tissue layers. In this case, the spatial distribution of scattered light has a broad angular spectrum, and depo-

Applications

177

larization effects play an important role. The spatial distribution of the correlation properties of the scattered field, which is related to the structure of an object, can be considered in a manner similar to diffusion-wave spectroscopy55 with allowance for the fact that an object has a static structure, and a laser beam (or the object itself) is scanned over the surface of the object at a definite rate. Along with the chosen beam radius and the character of optical inhomogeneities of the medium, the rate of scanning determines fluctuations of the scattered field in the time domain. In such a situation, the normalized AF of intensity fluctuations is generally defined by Eq. (86) with ξ ≡ t and ξ ≡ τ. The behavior of the structure function [see Eq. (87)] near the zero value of its argument, which corresponds to the highest efficiency of high-frequency spatial intensity fluctuations, can be conveniently characterized in this case in terms of the so-called exponential factor νI :369,370   ln DI (τ2 )/DI (τ1 )

. (88) νI = ln |τ2 |/|τ1 | To analyze polarization properties of speckle fields, the following time-domain first- and second-order statistical characteristics of the intensity fluctuations of scattered light in the paraxial region can be employed. Measurements should be provided for two orthogonal linear polarizations (relative to the polarization of the probing beam)32 (1) The mean intensity of speckles, Isp  = I  + I⊥ .

(89)

(2) The cross-correlation function (correlation coefficient) for two polarization states,    (90) r⊥ (τ) = I (t) − I  I⊥ (t + τ) − I⊥  , where the indices (⊥, ) denote combinations of polarization states, and averaging is performed over the trajectory of scanning. The estimation of structure parameters of a tissue, such as the characteristic size of local inhomogeneities and spatial fluctuations of the refractive index, generally requires some assumption concerning the scattering model. One of the simplest models of scattering is the model of a random phase screen with Gaussian statistics of inhomogeneities.369–372 In many cases, statistical models for living structures are much more complicated. In particular, some of these models are nonlinear and may take into account multiple scattering. In spite of the lack of well-developed models for the structure of many tissues, any empirical information concerning statistical properties of scattered light is useful for the analysis of structure images of tissues. Such information may be also useful for the development of structure models themselves.

178

Optical Clearing of Tissues and Blood

The control of optical properties of tissues, in particular, the possibility of considerably decreasing the scattering coefficient, may become a key point in optical functional tomography of various human organs (e.g., eye, skin, cervix, and breast) in the process of searching for small tumors at early stages of their formation, and imaging of small living animals in studies of gene expression.373 At present, there are two thoroughly developed techniques for enhancement of optical translucence of a soft tissue based on the compression and OCA approaches. Both techniques employ the matching of refractive indices of inhomogeneities due to an enhanced diffusion of fluids in a tissue. In the former case, external mechanical pressure (compression) plays the role of an applied force. In the latter case, osmotic pressure of various solutions is important as well as OCA diffusion into a tissue. Specifically, mechanical pressure applied through the end of an optical fiber is successfully employed in transscleral laser surgery of eye tissues.35 An OCA-related decrease in the scattering coefficient was demonstrated above for the eye sclera, skin, dura mater, and gastric tissues both in vitro and in vivo. Experimental study of the spatial correlation of the coherent light transmitted through the human sclera allows one to understand the basic principles of speckle tomography improvement at OCA action. Such measurements were performed in vitro (samples of about 1 × 1 cm2 in area and 0.4–0.8 mm in thickness) for focused and unfocused laser beams, as well as for a low-coherent light source (superluminescent diode); see Figs. 14, 83–86. At scattering decrease, the regime of multiple scattering is switched to the regime of single scattering (or scattering with low multiplicity) when photons transmitted in the forward direction (ballistic photons) play a significant role and, consequently, coherent effects become predominant. Characteristic changes in speckle structures in the far-field zone were visually observed on a screen located in the plane of a photodetector and were recorded in reflected light by means of a CCD camera. Figure 14 presents two recorded fragments corresponding to the early and later stages of clearing at sharply focused beam diffraction. For native sclera, no speckle structures were observed within the first several minutes (usually, during 1–3 min) of the action of OCA at room temperature (the intensity of transmitted light was very low). Then, an isotropic speckle pattern with small speckles (multiple bulk scattering) was observed. Within 3–5 min, this pattern was replaced by a nonuniform speckle pattern with large speckles in the central area. Ultimately, within 7–10 min at room temperature, most of the transmitted light was concentrated around the direction of the laser beam incident on a sample. This regime is preferable for the analysis of structural large-scale inhomogeneitis (tumors) of a tissue, because laser beam penetration depth is increased from a few tenths of microns for normal tissue to 1 mm (scleral thickness) of cleared tissue, and statistical characteristics of the scattered field are more sensitive to tissue structure due to speckle transfer from the developed to the undeveloped mode. Speckle patterns for unfocused laser beams were recorded by using an Electrim digital video recorder. Figure 83 displays the changes of the recorded speckle patterns with the progress of optical clearing process.83,88 After the recording, data

Applications

179

Figure 83 Fragments of the images of speckle patterns induced by a single-mode He:Ne laser beam (633 nm) transmitted through a human scleral sample in the course of its impregnation by the OCA Trazograph-60.83,88 Time elapsed after OCA administration: (a) 1 min; (b) 2 min; (c) 4 min; (d) 5 min.

Figure 84 The corresponding to data of Fig. 83 evolution of the empirical probability density function of the speckle intensity fluctuations in the course of optical clearing of a human scleral sample.83,88 Time elapsed after Trazograph-60 administration: (a) 1 min; (b) 2 min; (c) 4 min; (d) 5 min.

image files were converted to 8-bit bitmap format for further processing, including calculation of the probability density function of speckle intensity. Evolution of form of the empirical probability density function of speckle intensity fluctua-

180

Optical Clearing of Tissues and Blood

tions with progress of sclera optical clearing is illustrated by Fig. 84.83,88 For the early stages of the optical clearing, speckle statistics are close to statistics of the so-called fully developed speckle fields with zero value of mean amplitude. Such a form of probability density function of intensity distributions is approximated by the negative exponential function (nonzero minimal value of intensity for presented histograms is caused by nonzero bias used for video recording). For the later stages of clearing, a qualitative change of intensity statistical distributions manifested in the appearing of bimodality is seen. This effect is typical for media with a single scattering causing only phase modulation of the probe beam and is characterized by relatively large values of standard deviation and correlation length of the phase fluctuations of the boundary field [see distributions in Figs. 84(c) and 84(d)]. The similar behavior can be explained in terms of the appearing of the specular component of the scattered field and the suppressing of the diffuse component. During optical clearing, far-zone speckles observed at the middle and later stages of the process demonstrate slow dynamic behavior, which can be interpreted as speckle “boiling” with very low frequencies. This dynamic is caused by the spatial-temporal inhomogeneities of the OCA concentration in the tissue volume. One of the most characteristic manifestations of such “boiling” is the presence of a ripple on the dependence of scleral transmittance on time elapsed after OCA application [Fig. 85(a)]. The transmission coefficient T was measured by a detector placed in the paraxial region of the far diffraction zone (at the axis of the incident beam, this is a “collimated” transmittance). For the off-axis detector position, the amplitude of random fluctuations of the detected signal caused by the speckle boiling can be adequately large [see, e.g., Fig. 85(b)]. Figure 85(c) shows the normalized power spectrum of the speckle intensity fluctuations calculated using the time series data presented in Fig. 85(a). It is clear that in the low-frequency region the spectral density decay obeys the power law. Very close behavior was observed for the focused laser and low-coherent radiation transmitted through the scleral sample (see Figs. 14 and 86).83,88 However, the contrast of the speckles from a light source with a low temporal coherence (spatial coherence is high) was essentially lower than for the coherent light scattering. This is due to the contribution of the incoherent part of the radiation in the speckle pattern formation. Figure 87 presents the optical scheme of a spatial speckle correlometer intended for topography or tomography of comparatively thin samples of tissues. This device employs a focused laser beam about 5 µm in diameter produced by a single-mode He:Ne laser. Structure patterns of an object were usually reconstructed through the two-dimensional scanning of the object and an appropriate analysis of statistical and correlation properties of scattered light. The scanning step in both coordinates was 5 µm. A photodetector was placed along the direction of the axis of the incident laser beam (detection of the scattered light exactly in the forward direction). The diameter of the entrance pinhole was about 25 µm, which is much less than the mean diameter of a speckle. The maximum rate of scanning was about 5 mm/s. The electronic units employed made it possible to obtain at least

Applications

181

(a)

(b)

(c) Figure 85 The slow dynamics of speckles induced by a single-mode He:Ne laser beam (633 nm) transmitted through a human scleral sample in the course of its impregnation by the OCA Trazograph-60.83,88 (a) The dependence of the “collimated” transmittance on time elapsed after OCA administration (on-axis detector position); (b) the time series of the normalized speckle intensity fluctuations for off-axis detector position; (c) the normalized power spectrum of the speckle intensity fluctuations.

20 equidistant counts per single step of scanning. An object was usually placed in the waist of the incident laser beam. The position of an object relative to the beam waist and the orientation of a polarization analyzer mounted in front of the photodetector were adjusted manually. The evolution of typical normalized AFs of intensity fluctuations in the process of scleral clearing measured with the use of a speckle correlometer (see Fig. 87) is shown in Fig. 88. Within small time intervals (1–2 min), the time evolution of the shape of the autocorrelation peak, which is associated with the transition of a tissue from one scattering regime to another, can be approximated by an exponential curve, whereas for large time intervals, this process can be approximately described by a Gaussian curve. The initial stage of the process is characterized by the existence of many scales of intensity fluctuations (the AF displays at least three distinguishable values of its slope). At later stages of clearing, the half-width of the AF peak tends to a value of (0.3–0.4) ×10−3 s, which is close to the ratio w/v of the waist radius of the incident beam to the scanning rate. This effect can be

182

Optical Clearing of Tissues and Blood

Figure 86 Fragments of the far-zone speckle patterns induced by a focused single-mode He:Ne laser beam (633 nm) ((a), (c)) and focused low-coherent light (superluminescent diode, 830 nm) ((b), (d)) transmitted through a human scleral sample in the course of its impregnation by the OCA Trazograph-60;83,88 ((a), (b)) early stage of tissue optical clearing and ((c), (d)) later stage.

Figure 87 Optical scheme of a scanning polarization-sensitive spatial speckle correlometer.31 1, single-mode He:Ne laser 633 nm; 2, telescope; 3, microobjective; 4, sample; 5, computer controlled scanning stage; 6, polarizer; 7, photomultiplier tube with a pinhole; 8, amplitude-digital convertor; 9, computer.

Applications

183

employed as a criterion of the completion of the transition from multiple scattering to single scattering.

Figure 88 Evolution of the normalized autocorrelation function of intensity fluctuations in the speckle field produced by light scattered from a sample of human sclera in the process of scleral enhanced translucence in Trazograph-60; the thickness of the sample is 0.6 mm; the measurements were performed with a sample processed in the solution during 120, 220, 420, and 820 s (from the down to the upper curve).31

Figure 89 Typical time dependencies of the exponential factor νI and the normalized mean intensity Is /I0 measured at the observation point for a scleral sample placed in Trazograph-60; λ = 633 nm.31

The time evolution of the exponential factor νI (see Eq. (88) and Fig. 89) characterizes the behavior of high-frequency components of intensity fluctuations. It should be noted that the exponential factor and the relative fluctuations of the mean intensity of transmitted light have substantially different rates of response to the action of OCA: νI is especially sensitive to the action of OCA at the early stages of clearing (within ∼1.5 min for Trazograph-60), whereas transmission reaches its maximum only by the tenth minute. This difference in response rates can be accounted for by the fact that the correlation properties of the field, characterized by transition from uniform to nonuniform speckle distributions, substantially change

184

Optical Clearing of Tissues and Blood

at the initial stages of matching of refractive indices. By contrast, maximum collimated transmission is achieved only when the refractive indices are completely matched. Straightforward modeling of the transport of initially collimated photons with a wavelength of 633 nm through a fibrous tissue consisting of collagen fibers with a mean diameter of 100 nm and a refractive index nc = 1.474 surrounded by a ground substance whose refractive index varies within the range n0 = 1.345–1.474 shows that, for a partial matching of refractive indices, n0 = 1.450 and nonscattered (67%) and singly scattered (25%) photons dominate in transmitted light (see Fig. 12). This prediction agrees well with the measured transmission and reflection spectra of sclera and the data on correlation measurements.32 Another specific feature of sclera clearing is the appearance of quasi-periodic oscillations of mean-intensity transmission (see Fig. 15), which are also manifested in the correlation characteristics (Fig. 89). These oscillations have small amplitude and can be clearly seen at later stages of translucence, when the main dynamic process associated with the directed diffusion of the OCA from the solution into the tissue and of water from the tissue to the solution is close to its completion. The characteristic oscillation time is 1.5–2.0 min. Apparently, a process with such a characteristic time can be attributed to the nonuniformity of diffusion of substances inside a tissue in space and time. This effect may be accounted for by the multistage character of the diffusion process. At the first stage, diffusion of OCA into a tissue and the flow of water out of the tissue partially equalize the refractive indices of hydrated collagen and the intercollagen substance. Under these conditions, optical transmission of a tissue grows until the dependence under study saturates. However, at the second stage, a relatively weak process of the interaction of the new ground substance with collagen is manifested. The ground substance somewhat lowers its refractive index through the dehydration of collagen, whereas the refractive index of collagen increases. The resulting mismatch of refractive indices slightly decreases optical transmission. At the next stage, a certain violation of the balance between the pressures of water and OCA in the solution and tissue gives rise to the diffusion of water from the tissue and OCA into the tissue, which leads to the more exact equalization of the refractive indices, and transmission grows again. Then, the process described above is repeated, and transmission oscillates in the time domain. Such oscillations are observed within the entire period of time when OCA acts on a tissue up to the fortieth to sixtieth minute. As mentioned earlier, transition from the regime of multiple scattering to the regime of single scattering should change the polarization properties of scattered radiation, which can be described in terms of the first- and second-order statistical characteristics [see Eqs. (89) and (90)]. At the early stage of sclera clearing, both polarization components of transmitted light have approximately equal intensities. However, in the process of clearing (transition of a tissue to the regime of low-step scattering), one of these components (namely, the component polarized along the polarization of the incident beam) begins to dominate over the other component (see Fig. 16).85 These experimental data demonstrate the reversibility of the clear-

Applications

185

ing process, which is important for living systems, and reveal a high sensitivity of polarization characteristics to structural changes in a tissue. The inhomogeneity of the OCA concentration within the tissue volume causes a macroscopic spatial inhomogeneity of the polarization structure of transmitted light; such polarization structure of the scleral tissue image can be easily observed using an optical scheme with a “white” light source and the tissue sample placed between two crossed polarizers. Figure 46 illustrates the evolution of such polarization images during scleral optical clearing. A tissue layer “works” as a phase plate (or number of phase plates175,255 ), whose linear birefringence is spatially and temporally dependent. The large-scale spatial inhomogeneities of images may be due to spatial variations of the sample thickness and structure, both of which may influence the efficiency of OCA impregnation and corresponding phase shift between the orthogonal optical field components.34 It is also seen that translucent sclera features small-scale spatial inhomogeneities of scattering and polarization properties in the form of a domain structure, with domain areas on the order of 0.1–1 mm2 . Such a structure is associated with a spatially nonuniform distribution of the diffusion rate of substances and is clearly manifested in both time-domain realizations of intensity fluctuations for separate polarization components and the behavior of the cross-correlation function for these components (see Fig. 90) measured with a laser beam scanned over a sample.31 It is obvious that, similar to the correlation characteristics of speckle fields, polarization and cross-correlation characteristics can be employed for imaging the structure of tissues, as well as for tissue topography and tomography.

Figure 90 Evolution of the cross-correlation coefficient for two orthogonal polarizations of intensity fluctuations in the speckle field produced in the observation plane by a single-mode He:Ne laser beam (633 nm) scanned over a scleral sample impregnated in Trazograph-60 during (1) 200 s and (2) 400 s. The scanning rate was 5 mm/s.31

186

Optical Clearing of Tissues and Blood

8.2.2 Diffusion-wave spectroscopy and functional imaging Experimental implementation of diffusion-wave spectroscopy is very simple: a measuring system should irradiate the scattering object under investigation with a light beam produced by a continuous-wave laser and measure intensity fluctuations in scattered radiation within a single speckle with the use of a photomultiplier and an electronic correlator (Fig. 91). The single-mode fiber collects light emerging from a certain area of the object and transmits it to the photomultiplier. Fibers are designed in such a manner as to pick up radiation from any area at the surface of the sample. The output signal of the photomultiplier is fed to a digital autocorrelator that reconstructs the time-domain autocorrelation function (AF) of the intensity fluctuations. For Gaussian statistics, this AF is related to the time-domain autocorrelation function of the field by the Siegert formula:7   g2 (τ) = A 1 + βsb |g1 (τ)|2 , (91) where τ is the delay time; A = i2 is the square of the mean value of the photocurrent, or the base line of the autocorrelation function; βsb is the parameter of self-beating efficiency, βsb ≈ 1; and   E ∗ (t + τ)E(t)   (92) g1 (τ) = |E(t)|2 is the normalized autocorrelation function of the optical field. Different arrangements of optical fibers connected to the source of radiation and the detector allow one to get images of a dynamic inhomogeneity within a tissue, because AF decays faster when the source of radiation and the detector are located close to the dynamic inhomogeneity, which gives rise to fluctuations in the time domain. It is in this area that most of the detected photons pass through the dynamic volume. Such a behavior of AFs allows one to employ the variation in their slopes (decay rates) as a parameter for the imaging of dynamic inhomogeneities in a medium. This model corresponds to a situation where the microcirculation rate of blood locally increases near, e.g., a growing tumor. Using a similar approach, one can also model a directed blood flow. This is the basis of one of the methods of tissue functional imaging. The correlation time τc of intensity fluctuations depends on two main processes: mean velocity of blood microcirculation and multiplicity of scattering events. If the mean velocity of blood and the scattering coefficient of surrounding tissue go down, the correlation time increases. So, the correlation time is an informative parameter in blood microcirculation imaging but its dependence on the light scattering of the surrounding tissue may distort information about blood flow. Thus, optical clearing technology may have good perspectives in the improvement of diffusion-wave spectroscopy and imaging as well. On the other hand, the correlation time τc , with the assumption that the microcirculation is not changed, can be

Applications

187

Figure 91 The experimental setup for measurement of the autocorrelation function of intensity fluctuations of the backscattered light from living tissues (designed at Pennsylvania University in the group of Prof. A. Yodh).86 It uses a tunable (770–800 nm) single-mode diode laser with a very narrow linewidth (∼5 MHz) and output power of 400 mW and polarization-maintaining optical fibers that transfer light through the camera objective to the sample and back to photomultipliers (PMT HC120, Hamamatsu) connected with a correlator (BIC-9000).

a good measure of the reduction of the tissue scattering coefficient due to refractive indices matching using OCAs. Evidently, some OCAs other than the reduction of the scattering coefficient can increase or decrease the mean velocity of blood microcirculation. This should be taken into account in the interpretation of experimental results and when used in the complex study of pathological tissues. Some results on in vivo studies of OCAs’ effect on correlation time τc (ms) of intensity fluctuations of backscattered NIR laser light (785 nm) from human skin with topically applied OCAs are summarized in Table 17. These measurements show that aloe vera (skin gel) influences at least up to a depth of about 0.75 mm in the skin (approximately a half of the maximal displacement in the experiments) during first 5–8 min. Because the random blood flow, which defines the width of the autocorrelation function, may be supposed to be constant during the OCA application, the changes of the decay time should reflect the reduction of the scattering multiplicity due to tissue immersion. For the small source-detector separation, this dependence falls very slowly; and for larger one, it grows. Due to the opposite response on reduction of the scattering coefficient of the upper tissue layer of the strongly backscattered photons (more multiple scat-

188

Optical Clearing of Tissues and Blood

Table 17 Temporal evolution of the autocorrelation function decay time τc measured in vivo for human skin with a topically applied OCA.86 Channel A—source-detector separation is 0.8 mm and channel B—source-detector separation is 1.5 mm. Each measurement of τc was done during 1 min (time of averaging).

OCA Aloe vera (skin gel)

40% Glucose

Time interval of OCA application, min

τc (ms) for channel A (0.8 mm)

τc (ms) for channel C (1.5 mm)

0 3 10 22 34 0 12

0.96 0.91 0.90 0.83 0.88 1.07 1.04

0.47 0.52 0.73 0.85 0.70 0.76 0.89

tered photons coming from tissue depth to detector) and laterally traveling photons (fewer multiple scattered photons coming to detector), the decay time also has opposite dependencies on reduction of the scattering coefficient (time interval of OCA application) for the small (0.8 mm) and relatively large (1.5 mm) sourcedetector separations. Besides aloe vera gel and glucose, a few other OCAs were tested, including polyethylene glycol and glycerol. Observed changes in the decay correlation time have the same tendency with maximum change of 10–20%.

8.3 Precision Tissue Photodisruption and Tattoo Removal Femtosecond laser pulses can generate high-precision subsurface photodisruption in transparent tissues such as cornea.36 The strong optical intensities required for photodisruption can be achieved at the focus of a high-peak-power laser beam. The location of optical breakdown can be controlled to occur only at the focus of the beam where the intensity exceeds the threshold level of breakdown. If the laser is focused beneath the surface of a tissue, subsurface breakdown occurs only at the focus. No damage takes place in the tissue layers that the beam was focused through. In contrast to transparent tissues, turbid tissues scatter light, spreading the pulse in both space and time, making it difficult to maintain the tight focus and short pulse duration needed for well-confined photodisruption.36 The same problems are characteristic for nonlinear spectroscopy including multiphoton fluorescence microscopy and SHG imaging.101,127,191,263–266,272 The ability to focus light through turbid tissue is limited, especially at wavelengths less than 1300 nm. Tissue optical clearing technology using an appropriate immersion agent can be applied for a temporal reduction of scattering needed for providing an effective nonlinear study or underlying tissue photodisruption. In particular, a femtosecond laser technology was used to demonstrate early proof of this concept for high-precision subsurface photodisruption in translucent human sclera.36 Approximately 5-mJ femtosecond pulses from two laser sources, 1060 nm

Applications

189

(500 fs) and 775 nm (150 fs) with a repetition rate of 1 kHz, were used to make a subsurface incision in sclera in vitro. The beam was focused to a 1.5 (775 nm) or 5 µm (1060 nm) spot size and scanned below the tissue surface at various depths to produce incision patterns. Tissue samples were impregnated by Hypaque-76 (x-ray contrast) to make them transparent, usually within 15 min. The measured axial transmission spectra of normal scleral tissue and scleral tissue treated by Hypaque-76 (the light forward scattered in a small cone angle around the incident beam was detected), as well as by saline and Hypaque-76 spectra, are presented in Fig. 92. As is seen, transmission of normal sclera is never greater than 10%, with a broad maximum at 1600–1800 nm. For sclera treated by Hypaque-76, the transmittance greatly increases across the entire spectrum, especially for wavelengths in the NIR from 800 to 1350 nm. A transparent window is also created between two strong water absorption bands, from 1500 to 1800 nm. Transmission at 775 and 1060 nm is above 60%. The difference in transmission is not expected to affect the results of photodisruption itself since photodisruption depends on the intensity of the pulse and not the linear absorption. Thus, spectral windows with less scattering and absorption allow for a focused beam to penetrate into a tissue with less attenuation and distortion, and do not influence photodisruption efficiency.

Figure 92 Transmission spectra of normal and treated by Hypaque-76 human scleral samples along with saline and Hypaque-76.36 All samples were approximately 0.5 mm thick. Reprinted with permission.

Figure 93 shows the minimum spots obtained for 775 nm and 1060 nm beams after focusing through a full thickness (0.5 mm) of sample of tissue. The white square in the top right corner is 10 µm on edge. In the bottom left corner is the beam

190

Optical Clearing of Tissues and Blood

after passing through an equivalent thickness of saline. The unscattered beams are less than 10 µm in diameter, but somewhat bigger than the initial ones due to the nonideal optical system and aberrations at focusing through the material. As expected at 775 nm, the size of the intensity distribution emerging from the sclera did not change with the position of the focusing lens since scattering is very strong. The emergent beam is many times larger than the unscattered spot and strongly speckle modulated (see also Fig. 35).85 At a longer wavelength (1060 nm), due to less scattering, the distribution is smaller but still heavily scattered. These images clearly show that focusing to the back surface of the tissue is not possible using these wavelengths; thus, for the normal scleral tissue, breakdown was only possible at the front surface at first.36

(a)

(b) Figure 93 Spot sizes of 775 nm and 1060 nm focused beams through (a) normal and (b) made transluscent by Hypaque-76 human sclera samples (0.5 mm).36 The bottom left corner shows the smallest spot obtained by focusing through an equivalent thickness 0.5 mm slab of water. A square 10 µm on edge is shown in the top right corner. The bar beneath each picture is the two color map: left (black) indicates not light, right (white) is maximum light intensity. Reprinted with permission.

After treating the tissue with Hypaque-76, the spot size decreases to almost that of the unscattered beam for both wavelengths [see Fig. 93(b)]; thus, the strong focusing of the beam should permit controlled back-surface photodisruption. Several types of intrascleral incisions were experimentally demonstrated using optical clearing technology such as: partial-thickness channel creation, which could be used to perform a transscleral procedure analogous to deep sclerotomy, where a block of inner-surface sclera is removed with minimal disruption to overlying

Applications

191

layers, or in altering the mechanical properties of the sclera for the treatment of presbyopia; full-thickness channel creation (see Fig. 94), which may be useful for draining aqueous fluid for the treatment of glaucoma; creation of a grid of tissue pores that may be useful in changing the bulk properties of the tissue, including the tissue’s hydraulic conductivity; and scleral pocket creation, which may be useful for implant insertion to treat presbyopia.36

Figure 94 Femtosecond-laser-created transscleral channel in the sample of human sclera.36 (a) A channel with a minimum width was demonstrated (the electron micrograph). The channel width is estimated at 10 µm or less from this cross section. Laser parameters: 4.2 µJ, 150 fs, 775 nm, 1 kHz, 2 µm FWHM spot, 1.5 µm spot separation. (b) Potential application of creating a drainage passageway. (c) Scan pattern of the laser spot. Reprinted with permission.

Data from a Monte Carlo simulation presented in Fig. 25 allow one to estimate the fractions of 700 nm light absorbed by different tissue layers, such as conjunctiva, sclera, ciliary muscle, and ciliary pigment epithelium, of the rabbit eye in the course of optical clearing at tissue impregnation by 40% glucose.173 It is well seen that at optical clearing more light is absorbed by ciliary muscle and ciliary pigment epithelium due to more effective light-beam penetration through the low-scattering sclera. Recently, optical clearing technology was successfully applied to assist laser tattoo removal.374,375 Six subjects with multicolored decorative tattoos were treated by a Q-switched alexandrite and Nd:YAG lasers (755, 532, and 1064 nm). Treatment parameters varied depending upon the color of the tattoo ink and skin phototype. Minimal or no epidermal damage was observed in the areas treated with the laser in combination with optical clearing. For these skin areas, significantly greater tattoo lightening was achieved as compared to the test sites treated with laser alone. Optical clearing of the upper skin layers provides smoother in-depth distribution of laser light and, therefore, an increased threshold for injury of the epidermis, as well as more light transportation in tissue depth to induce lightening of the deep tissue layers.

9 Other Methods of Tissue Optical Properties Control

9.1 Tissue Compression and Stretching As it was already mentioned in the Introduction, squeezing (compressing) or stretching of a soft tissue produces a significant increase in its optical transmission.73 The major reasons for this are the following: (i) increased optical tissue homogeneity due to removal of blood and interstitial fluid from the compressed site [see Eq. (6)]; (ii) closer packing of tissue components causes less scattering due to cooperative (interference) effects;37,46 and (iii) less tissue thickness. Mechanisms underlying the effects of optical clearing and changing of light reflection by soft tissues at compression and stretching were proposed in a number of theoretical and experimental studies.21,22,35,37,46,56,57,112,135,145,198,376,377 It should be emphasized, however, that squeezing-induced effects in tissues containing little blood, such as sclera, are characterized by a marked inertia (a few minutes) because of the relatively slow diffusion of water from the compressed region.35 It was suggested that compression of sclera may displace water from the interspace of collagen fibrils, increasing the protein and mucopolysaccharide concentrations. Since these proteins and sugars have refractive indexes closer to that of the collagen fibrils, a more index-matched environment can be created. On the other hand, compression reduces specimen thickness d, which might increase the effective scatterer concentration inside the tissue.57 Therefore, compression may also give rise to an increase in tissue scattering coefficient µs . However, the total effect on the change in optical properties, which is proportional to the product of µs d, is characterized by less scattering. Sometimes the increase in scatterer concentration is likely to be more dominant than the reduction in index mismatch.57 In addition, reduction of tissue thickness causes an increase in local chromophore concentration (for bloodless tissue, or tissue specimens having aggregated and/or coagulated blood), i.e., the absorption coefficient increases. The authors of Ref. 57 observed that compression caused leaking around the specimen. Some of the extracellular fluids along the edge of the tissue sample were forced out upon compression. Unless sufficient pressure was applied to rupture the cell walls, the intracellular fluids would be retained by the cells in the bulk of the sample. When compressed, tissue thickness was reduced so that the volumetric water concentration was increased. This may explain the increase of the absorption coefficient at the wavelengths of water bands with compression. The authors of Ref. 57 have studied in vitro optical proper193

194

Optical Clearing of Tissues and Blood

ties of human skin, bovine aorta, bovine sclera, and porcine sclera in the spectral range from 400 to 1800 nm using the integrating sphere (IS) technique and inverse adding-doubling (IAD) method for convolution of absorption and reduced scattering coefficients. The diffuse reflectance and transmittance of these tissue samples of about 2 × 2 cm2 were measured at no pressure and at pressures of 0.1, 1, and 2 kg/cm2 uniformly distributed over the sample surface. They generally observed a decrease in reflectance, while transmittance, absorption, and scattering coefficients increased owing to compression. Some of these data for the human skin are presented in Table 18.378–386 As was explained earlier, the amount of scattering depends on refractive index mismatch, as well as on scatterer concentration and spacing. Along the load direction, the spacing between tissue components is reduced and, due to water escape from the compression site, refractive index matching should occur; both effects decrease the averaged light scattering (transmittance increases and reflectance decreases). On the other hand, compression reduces specimen thickness, which might increase the effective scatter and chromophore concentration inside the tissue. In experiments with a uniformly distributed compression, the scatter concentration increase was likely more dominant than the reduction in index mismatch and scatterer packing affect. It should be noted that the relative balance of the contributions of the listed mechanisms is expected to be changed if a pointwise compression is applied.35,41,42,376 To understand the optical properties of compressed tissue, time-resolved studies in the range of minutes are important.35 Spectral properties of skin can be effectively controlled by applying an external localized pressure in in vivo experiments when UV-induced erythema (skin redness) is developed.56,121,369 Figure 95 shows the apparent optical density (OD) evaluated from the in vivo measured back-reflectance spectra of erythematous human skin for different values of external localized mechanical pressure. To the third day after UV irradiation, the developed erythema is seen as increased absorbance (OD) in the 520–580 nm spectral range due to increased blood volume in the skin. Blood hemoglobin blocks the backscattered intensity from the deep skin layers. For the longer wavelengths from 600 to 700 nm, increased blood volume causes increase of light scattering from tissue, which is seen as the increase of apparent OD. At pressures of (8.4–14) ×105 Pa, blood is leaving the compressed area of the skin; thus, OD spectral dependence becomes smoother due to less absorption in the range 520–580 nm and there is less scattering in the range 600–700 nm. The intensity of skin autofluorescence is also well controlled with external localized pressure applied to the skin site. As it follows from Fig. 96, the external localized pressure in the range from 0 to 1.4 × 105 Pa considerably changes the fluorescence output with the wavelength 460 nm at induced erythema. Due to more effective flurescence light attenuation by blood hemoglobin at more intensive erythema (14 days after UV irradiation), skin compression more effectively controls (increases) fluorescence output. As it was also shown, application of a pressure cuff to the upper arm of healthy volunteers at levels of 0, 20, 30, 40, and 60 mm Hg did not significantly alter

White matter coagulated

White matter

White matter coagulated (n = 7)

White matter (n = 7)

Gray matter coagulated (n = 7)

800 1064

360 640 1060 360 740 1100 360 640 860 1060 360 860 1060 800 1064

1064 1064

Normal Coagulated

Brain: Gray matter (n = 7)

308 308

λ, nm

Tissue In vitro measurements Aorta: Normal Coagulated 239(45) 293(73)

– –

µs , cm−1

0.9(0.18) 0.5(0.1)

170(17) 130(13)

3.33(2.19) 141.3(42.6) 0.17(0.26) 90.1(32.5) 0.56(0.7) 56.8(18.0) 9.39(1.70) 426(122) 0.45(0.27) – 1.0(0.45) 179.8(32.6) 2.53(0.55) 402.0(91.8) 0.8(0.2) 408.2(88.5) 0.97(0.4) 353.1(68.1) 1.08(0.51) 299.5(70.1) 8.3(3.65) 604.2(131.5) 1.7(1.3) 417.0(272.5) 2.15(1.34) 363.3(226.8) 0.8(0.16) 140(14) 0.4(0.08) 110(11)

0.53(0.09) 0.46(0.18)

33 44

µa , cm−1

10.2 9.1

– – – – –

– – – – – – – – –

23.9 29.3

77 270

µ
s , cm−1

0.94(0.02) 0.93(0.02)

0.818(0.093) 0.89(0.04) 0.90(0.05) 0.868(0.031) – 0.954(0.001) 0.702(0.093) 0.84(0.05) 0.871(0.028) 0.889(0.010) 0.800(0.089) 0.922(0.025) 0.930(0.015) 0.95(0.02) 0.95(0.02)

0.9 0.9

– –

g

DIS, IMC; samples 0.5–3 hrs post mortem, fast frozen and homogenized; coagulation in a bath at 75◦ C381

Double integrating sphere (DIS), inverse Monte Carlo (IMC)379,380 Coagulation: saline bath, 80◦ C, 2 hr

Post mortem (6 h), excised, in 4◦ C saline, slab, water bath (85◦ C), data from Ref. 378 Post mortem, slab, 70◦ C water bath, 10 min, data from Ref. 378

Remarks

Table 18 Optical properties of normal, coagulated, and compressed human tissues measured in vitro and in vivo (rms values are given in parentheses)

Other Methods of Tissue Optical Properties Control 195

Cerebellum (n = 7)

coagulated

White matter coagulated (n = 7) Cerebellum (n = 7)

Tissue White matter (n = 7)

λ, nm 450 510 630 670 850 1064 850 1064 400 500 600 700 800 900 1000 1100 400 500 600 700 800 900 1000 1100

µa , cm−1 1.4 1.0 0.8 0.7 1.0 1.0 0.9 0.1 4.7(0.8) 1.4(0.2) 0.8(0.2) 0.6(0.1) 0.6(0.1) 0.7(0.1) 0.8(0.1) 0.7(0.1) 19.3(7.7) 5.1(1.7) 2.9(1.4) 1.7(0.4) 1.1(0.2) 1.1(0.3) 1.0(0.4) 1.1(0.5)

µs , cm−1 420 426 409 401 342 296 300 270 276.7(19.1) 277.5(32.6) 272.1(12.3) 266.8(12.1) 250.3(17.2) 229.6(15.8) 215.4(14.7) 202.1(13.9) 560.0(25.5) 512.2(47.8) 458.2(65.6) 489.9(70.1) 458.2(54.0) 458.2(65.6) 419.1(49.4) 428.5(40.0)

µ
s , cm−1 92.4 80.94 65.44 60.15 41 32.56 36.0 29.7 – – – – – – – – – – – – – – – –

Table 18 (Continued).

g 0.78 0.81 0.84 0.85 0.88 0.89 0.88 0.89 0.80(0.03) 0.85(0.02) 0.87(0.02) 0.89(0.01) 0.90(0.01) 0.90(0.01) 0.90(0.01) 0.90(0.01) 0.61(0.01) 0.77(0.02) 0.78(0.01) 0.85(0.01) 0.87(0.02) 0.89(0.02) 0.90(0.03) 0.91(0.03)

Remarks Integrating sphere (IS), IMC, quasi-Newton inverse algorithm, HGPF; hemoglobin-free cryosections (<48 h post mortem): gray matter 100–200 µm; white matter 80–150 µm; coagulation: saline bath 80◦ C, 2 hr;382 data from tables of Ref. 382 IS, IMC, quasi-Newton inverse algorithm, HGPF; hemoglobin-free cryosections of normal tissues (<48 hr post mortem): cerebellum, gray matter, pons, and thalamus 100–200 µm; white matter 80– 150 µm; and tumors excised from patients of ≈300 µm in thickness; coagulation: saline bath 80◦ C, 2 hr;382 data from graphs of Ref. 382, taken from Ref. 383 with corrections

196 Optical Clearing of Tissues and Blood

Pons (n = 7)

Gray matter coagulated (n = 7)

Tissue Gray matter (n = 7)

λ, nm 400 500 600 700 800 900 1000 1100 400 500 600 700 800 900 1000 1100 400 500 600 700 800 900 1000 1100

µa , cm−1 2.6(0.6) 0.5(0.2) 0.3(0.1) 0.2(0.1) 0.2(0.1) 0.3(0.2) 0.6(0.3) 0.5(0.3) 7.5(0.4) 1.8(0.2) 0.7(0.1) 0.7(0.1) 0.8(0.1) 0.9(0.1) 1.4(0.2) 1.5(0.2) 3.1(0.7) 0.9(0.3) 0.6(0.2) 0.5(0.2) 0.6(0.3) 0.7(0.3) 1.0(0.4) 0.9(0.4)

µs , cm−1 128.5(18.4) 109.9(13.0) 94.1(13.5) 84.1(12.0) 77.0(11.0) 67.3(9.6) 61.6(5.7) 55.1(6.5) 258.6(18.8) 326.5(7.7) 319.0(15.2) 319.0(7.5) 252.7(18.3) 214.6(10.3) 191.0(18.7) 186.6(13.5) 163.5(15.3) 133.7(19.2) 109.4(18.5) 93.5(20.9) 83.6(21.0) 74.8(18.7) 69.9(17.5) 64.0(17.8)

µ
s , cm−1 – – – – – – – – – – – – – – – – – – – – – – – –

Table 18 (Continued).

g 0.87(0.02) 0.88(0.01) 0.89(0.02) 0.90(0.02) 0.90(0.02) 0.90(0.02) 0.90(0.02) 0.90(0.02) 0.78(0.04) 0.85(0.03) 0.87(0.03) 0.88(0.03) 0.87(0.02) 0.87(0.02) 0.88(0.03) 0.88(0.03) 0.89(0.02) 0.91(0.01) 0.91(0.01) 0.91(0.01) 0.91(0.01) 0.92(0.01) 0.91(0.01) 0.92(0.01)

Remarks

Other Methods of Tissue Optical Properties Control 197

Thalamus coagulated (n = 7)

Thalamus (n = 7)

Tissue Pons coagulated (n = 7)

λ, nm µa , cm−1 410 17.2(1.6) 510 8.5(0.8) 610 7.7(0.5) 710 6.9(0.6) 810 6.5(0.6) 910 5.9(1.0) 1010 5.7(1.0) 1100 6.5(0.9) 410 3.2(1.0) 510 0.9(0.3) 610 0.6(0.2) 710 0.5(0.3) 810 0.7(0.3) 910 0.7(0.3) 1010 0.8(0.3) 1100 0.8(0.3) 400 15.0(3.3) 500 4.2(0.9) 600 1.6(0.6) 700 1.4(0.3) 800 1.1(0.3) 900 1.1(0.3) 1000 1.4(0.4) 1100 1.5(0.4)

µs , cm−1 685.7(63.7) 627.5(73.6) 510.5(70.5) 402.5(67.7) 329.7(55.4) 276.0(46.4) 241.6(34.4) 221.1(31.5) 146.7(49.4) 188.7(31.9) 176.3(34.5) 169.0(28.7) 158.5(35.3) 155.4(22.3) 139.3(34.9) 146.0(36.6) 391.1(56.1) 399.9(67.7) 365.7(43.2) 327.0(30.6) 286.0(33.8) 267.4(31.6) 233.8(39.7) 223.6(32.1)

µ
s , cm−1 – – – – – – – – – – – – – – – – – – – – – – – –

Table 18 (Continued).

g 0.85(0.02) 0.89(0.01) 0.89(0.01) 0.89(0.01) 0.89(0.01) 0.88(0.01) 0.88(0.01) 0.88(0.01) 0.86(0.03) 0.87(0.03) 0.88(0.02) 0.89(0.03) 0.89(0.02) 0.90(0.02) 0.90(0.02) 0.91(0.02) 0.83(0.04) 0.90(0.01) 0.92(0.01) 0.92(0.01) 0.93(0.01) 0.93(0.01) 0.93(0.01) 0.94(0.01)

Remarks

198 Optical Clearing of Tissues and Blood

Coagulated

Prostate: Normal

White matter coagulated (n = 7)

Tissue White matter (n = 7)

850 980 1064 850 980 1064

λ, nm 400 500 600 700 800 900 1000 1100 410 510 610 710 810 910 1010 1100 0.6(0.2) 0.4(0.2) 0.3(0.2) 7.0(0.2) 5.0(0.2) 4.0(0.2)

µa , cm−1 3.1(0.2) 0.9(0.1) 0.8(0.1) 0.8(0.1) 0.9(0.1) 1.0(0.1) 1.2(0.2) 1.0(0.2) 8.7(1.7) 2.9(0.6) 1.7(0.4) 1.4(0.5) 1.5(0.5) 1.7(0.6) 1.9(0.6) 2.4(0.5) 100(20) 90(20) 80(20) 230(30) 190(30) 180(30)

µs , cm−1 413.5(21.4) 413.5(43.9) 413.5(21.4) 393.1(30.9) 364.5(28.6) 329.5(35.0) 305.4(15.9) 283.2(22.2) 568.7(111.9) 513.2(116.9) 500.2(129.9) 475.2(108.3) 440.0(114.3) 407.4(92.8) 367.7(95.5) 358.4(81.6) – – – – – –

µ
s , cm−1 – – – – – – – – – – – – – – – –

Table 18 (Continued).

0.94(0.02) 0.95(0.02) 0.95(0.02) 0.94(0.02) 0.95(0.02) 0.95(0.02)

g 0.75(0.03) 0.80(0.02) 0.83(0.02) 0.85(0.02) 0.87(0.01) 0.88(0.01) 0.88(0.01) 0.88(0.01) 0.83(0.03) 0.87(0.02) 0.90(0.02) 0.91(0.01) 0.92(0.01) 0.93(0.01) 0.93(0.01) 0.93(0.01)

Shock frozen sections of 60– 500 µm, 0.5–3 hr post mortem, Ref. 303 Sections of 60–500 µm, 0.5–3 hr post mortem, water bath (75◦ C, 10 min), Ref. 303

Remarks

Other Methods of Tissue Optical Properties Control 199

Hispanic male skin (n = 3), external pressure 0.1 kg/cm2 Hispanic male skin (n = 3), external pressure 1 kg/cm2

Caucasian female skin (n = 3), external pressure 0.1 kg/cm2 Caucasian female skin (n = 3), external pressure 1 kg/cm2 Hispanic male skin (n = 3)

Caucasian male skin (n = 3), external pressure 0.1 kg/cm2 Caucasian male skin (n = 3), external pressure 1 kg/cm2 Caucasian female skin (n = 3)

Skin: Caucasian male skin (n = 3)

Tissue Normal Coagulated

500 810 500 810 500 810 500 810 500 810 500 810 500 810 500 810

500 810

λ, nm 1064 1064

15.3 0.63 13.6 0.57 5.2 0.97 7.4 1.4 10.0 1.7 3.8 0.87 5.1 0.93 6.2 0.87

5.1 0.26

µa , 1.5(0.2) 0.8(0.2)

cm−1

– – – – – – – – – – – – – – – –

– –

µs , cm−1 47(13) 80(12)

167.4 52.7 156.7 53.4 23.9 8.2 31.5 11.3 40.2 13.1 24.2 7.5 37.6 11.4 40.4 10.2

50 15.8

µ
s , cm−1 0.64 1.12

Table 18 (Continued).

0.862 0.861

– – – – – – – – – – – – – – – –

– –

g

IS, inverse adding-doubling (IAD) method; sample thickness: 0.40, 0.23, 0.25 mm57 IS, IAD, sample thickness: 0.15, 0.05, 0.13 mm57 IS, IAD; sample thickness: 0.12, 0.05, 0.13 mm57 IS, IAD; sample thickness: 0.42, 0.50, 0.50 mm57 IS, IAD; sample thickness: 0.30, 0.30, 0.34 mm57 IS, IAD; sample thickness: 0.27, 0.20, 0.23 mm57 IS, IAD; sample thickness: 0.70, 0.78, 0.63 mm57 IS, IAD; sample thickness: 0.35, 0.62, 0.48 mm57 IS, IAD; sample thickness: 0.28, 0.48, 0.33 mm57

Remarks Freshly excised, slab, water bath (70◦ C, 10 min), Ref. 378

200 Optical Clearing of Tissues and Blood

Volar side of lower arm (epidermis and dermis) Normal Treated Abdominal skin: chosen direction perpendicular direction (along collagen fibers)

skin temperature 38◦ C

In vivo measurements Forearm (light skin, n = 7): skin temperature 22◦ C

Subdermis (primarily globular fat cells) (n = 12)

Negroid dermis (n = 5)

Tissue Ex vivo measurements Skin: Caucasian dermis (n = 12)

810 810

1300 1300 0.014 0.07

– –

– –

140 80

– – – – – –

– – – – – – – –

–

µs , cm−1

20 10

– –

9.191(0.931) 7.340(0.901) 6.067(0.847) 9.613(0.894) 7.649(0.971) 6.234(0.928)

23.2(4.1) 16.3(2.5) 32.1(20.4) 26.8(14.1) 18.1(0.4) 12.6(3.4) 12.1(3.2) 10.8(2.7)

27.3(5.4)

µ
s , cm−1

Table 18 (Continued).

2.372(0.282) 0.966(0.110) 0.981(0.073) 2.869(0.289) 1.157(0.106) 1.135(0.123)

0.19(0.06) 0.13(0.07) 2.41(1.53) 1.49(0.88) 0.45(0.18) 0.13(0.05) 0.09(0.03) 0.12(0.04)

700 900 633 700 900 633 700 900

590 750 950 590 750 950

0.33(0.09)

µa ,

633

λ, nm

cm−1

– –

– – – – – –

0.9 0.9 0.9 0.9 0.9 0.9 0.9 0.9

0.9

g

SRR; CCD detector, rsd ≤ 10 mm; diffusion approximation384

OCT312 ; depth up to 350 µm; skin treated with a detergent solution (2% of anionic tensides in water)

Spatially resolved reflectance (SRR); MC-generated grid60

A single integrating sphere “comparison” method, IMC; samples from abdominal and breast tissue obtained from plastic surgery, g = 0.9;385,386 for dermis, the relative change in µ
s with temperature (25– 40◦ C) is (4.7 ± 0.5) × 10−3 ◦ C−1 and for subdermis (–1.4 ± 0.28) × 10−3 ◦ C −1

Remarks

Other Methods of Tissue Optical Properties Control 201

202

Optical Clearing of Tissues and Blood

Figure 95 The apparent optical density (OD) spectral distributions of erythematous human skin (three days after UV irradiation) for different values of external mechanical pressure: (1) without pressure; (2) 5.6 × 104 Pa; (3) 8.4 × 104 Pa; (4) 1.4 × 105 Pa.56,121

Figure 96 Autofluorescence (AF) intensity of erythematous human skin (λf = 460 nm) with dependence on external mechanical pressure: (1) seven days after UV rradiation (less developed erythema); (") fourteen days after UV irradiation (more developed erythema). AF intensity of human skin without erythema and compression (dotted line) is marked as a reference value.56,377

the levels of oxy-hemoglobin or melanin of preliminary UV-irradiated volar forearm skin (induced erythema and melanin pigmentation).135 In contrast, deoxyhemoglobin significantly contributes to the skin color appearance. Blood pooling, expressed as increased deoxy-hemoglobin, can contribute to what is visually perceived as pigmentation. Oxy-hemoglobin values increased to a maximum on the first day after UV irradiation, correlating well with the clinical evaluation of erythema, and then decreased exponentially to the baseline value. Melanin contents showed a significant increase on the seventh day and remained relatively constant for the next three weeks, correlating well with the clinical evaluation of pigmentation (tanning). Deoxy-hemoglobin increased slightly on the first day and remained

Other Methods of Tissue Optical Properties Control

203

elevated for the next two weeks. Thus, it correlated moderately with the clinical erythema scoring on the first day only, while it contributes significantly to what is clinically perceived as skin tanning on days seven and fourteen. Application of pressure below the diastolic level increased deoxy-hemoglobin concentration as measured by diffuse reflectance spectroscopy. This increase corresponded to a decrease of a “pigmentation” parameter in a similar fashion that has been documented for increases in melanin concentration. Topical application of H2 O2 reduced deoxy-hemoglobin levels as measured by reflectance spectroscopy. This reduction coincided kinetically with a visible skin blanching. Steady-state diffuse reflectometry with a variable source-detector separation rsd allows one to understand the characteristic properties of light propagation in a tissue with mechanical tension.384 The corresponding experimental (in vivo) and theoretical polar diagrams for the normalized back reflectance from stretched human skin are shown in Fig. 97. Measurements and calculations were done for a few rsd . To examine the effect of skin tension on the optical properties, the skin was maintained in a stretched position by pulling it mechanically in a defined direction and then fixing it to an aluminum ring with double-sided adhesive tape. For medium and far distances of the detector from the light source (2 and 7 mm), it was found that anisotropy of light propagation followed the stretching direction as shown in Fig. 97(a). The direction of stretching complies with that of the maximal reflectance signal for a given distance greater than about 2 mm. Thus, the scattering coefficient is minimal when measured along the direction of stretching, because intensity is maximal for the distant detector. Figure 97(b) presents results of Monte Carlo (MC) modeling that account for the anisotropy of the scattering coefficient, caused by different photon interactions with the medium when it travels along tissue fibers and across them,
  µs = µs0 1 + f 0.5 − | cos ψ| ,

(93)

where µs0 is a base scattering coefficient, f is the fraction of scatterers oriented in the preferential direction, and ψ is the angle between the current photon direction and the cylinder axis, with the fraction of scatterers (cylinders) oriented in the preferential direction, f = 0.35. As it follows from Eq. (93), the scattering cross section varies with the direction cosine of the incident photon with respect to the axis of the cylinder. It is maximal for perpendicular incidence and minimal for parallel incidence. Figure 88(b) demonstrates that for short source-detector separations, signals are much higher (scattering coefficients are also higher) perpendicular to the preferential direction of the collagen fibers, although the ellipses are much sharper than for the in vivo data. For farther detectors, the signal is higher in the parallel direction (scattering coefficient is lower). Such correlated or anticorrelated dependencies between the back-reflectance and the scattering coefficient, respectively, for short and long source-detector separations are also followed from data in Fig. 53.

204

Optical Clearing of Tissues and Blood

(a)

(b) Figure 97 Light propagation in the skin at mechanical tension.384 (a) Skin at the abdomen of the human volunteer was stretched in two directions—parallel (triangles) and perpendicular (squares) to the body axis (0–180◦ ). For both directions, full-rotation experiments were conducted. The diagram shows the signals measured by two detectors at medium and far distances (rsd = 2 mm and 7 mm) from the light source. The reflectance is maximal in the direction of maximal tissue tension. (b) Monte Carlo simulation including collagen fiber model orientation. Normalized reflectance at five distances from the source, generated using the MC simulation with 35% preferentially oriented collagen fibers in the skin. The 0◦ direction is parallel to the aligned collagen fibers and 90◦ is perpendicular.

Other Methods of Tissue Optical Properties Control

205

From this analysis, it follows that with skin stretching, the scattering coefficient and corresponding back-reflectance and transmittance can be effectively controlled. On the other hand, intact skin has its own anisotropy, which is believed to be caused by the preferential orientation of collagen fibers in the dermis, as described by Langer’s skin tension lines.384 Thus, the human skin’s reduced scattering coefficient varies by up a factor of two between different directions of light propagation at the same position (see Table 18). With external forced tension, more significant damping of scattering along the direction of mechanical stress is expected. The measurements of the deformations and applied loads and estimating the biomechanical properties of tissue are critical to many areas of the health sciences, including monitoring of the tension in wound closures, skin flaps, and tissue expanders.145 Such measurements, which can be provided by polarized light reflectivity detection, will allow surgeons to treat wounds more successfully by minimizing scar tissue and maximizing the speed of treatment, by letting them know how much the skin can be stretched at each treatment step. In vivo human experiments showed that the specular reflection from skin changes with stretch.145 For small values of stretch, the specular reflectivity measured for a He:Ne laser (λ = 633 nm) beam with the 45 deg angle of incidence increases linearly with strain (Fig. 97). The linear relationship between applied stretch and polarized reflectivity can be understood if the skin surface is approximated by a sinusoidal profile in the resting stage. Stretching reduces amplitude and increases spatial scale of the skin profile, thereby making it smoother and flatter, resulting in a corresponding increase in reflectivity. For ten tested subjects with various skin complexions, the slope is in the range from 0.0074 to 0.0391, with the linear correlation coefficient R 2 being from 0.88 to 0.99. For larger stretches (for strains above 8.8% (5 mm stretch) for the human subject tested, Fig. 98) the dependence is saturated and even goes down. The stretches in two perpendicular directions (parallel and perpendicular to the long axis of the forearm) yield good correlation between stretch and reflected light intensity and show that skin has anisotropic properties that can be detected by light reflection. For example, the slope measured in the direction parallel to the long axis of the forearm was 0.0095 ± 0.0002, and in the perpendicular direction it was 0.0065 ± 0.0008.145

9.2 Temperature Effects and Tissue Coagulation A reproducible effect of temperature between 25 and 40◦ C on the reduced scattering coefficient of human dermis and subdermis was found in an ex vivo study in the NIR.385,386 For dermis, the relative change in the reduced scattering coefficient showed an increase [(4.7 ± 0.5) × 10−3 ◦ C−1 ]) and for subdermis a decrease [(−1.4 ± 0.28) × 10−3 ◦ C−1 ]. The absolute values of the coefficients are presented in Table 18. The observed positive and negative temperature coefficients of scattering for dermis and subdermis, respectively, were explained in the following way.385,386 The main scattering components of subdermis were assumed to

206

Optical Clearing of Tissues and Blood

Figure 98 Measurement of the specular reflectivity of polarized laser light (633 nm) from the forearm of the volunteer.145 For small values of stretch, the change in reflectivity is linear (R 2 is the liner correlation coefficient). As stretch continues beyond about 8.8% strain (5 mm stretch), the reflectivity approaches a limit.

be lipids in membranes and vacuoles. It is known that lipids undergo phase transitions at certain temperatures, which alter their orientation, mobility, and packing order.66,247,248 Glycolipids found in human cells’ membranes undergo phase transitions in the temperature range from 25 to 45◦ C, namely, a transition from a gel phase through a stable crystalline phase to a liquid-crystalline phase with increasing temperature. The decrease in the scattering coefficient seen experimentally with increasing temperature is therefore consistent with an increase in fluidity known to occur in lipids with increasing temperature. Modifications of the collagen fiber structure of dermis caused by increasing temperature, possibly through changes in hydration, is the most plausible explanation of the increased scattering properties.385,386 As claimed by the authors of Refs. 385 and 386, the finding that a tissue that is largely protein has a positive temperature coefficient and a tissue that is largely lipid has a negative temperature coefficient leads to interesting possibilities in tissues where the protein/lipid ratio is intermediate, such as brain tissue. The absorption, µa , and the reduced scattering, µ
s , coefficients of human forearm skin have been also determined in the course of in vivo studies for two skin surface temperatures, 22◦ C and 38◦ C60 (see data in Table 18). A rather high increase of 16–21% of µa and much smaller increase of 2.7–4.6% of µ
s coefficient values were found with temperature change from 22◦ C to 38◦ C for the wavelengths 950–590 nm.60 It is important to note that blood glucose concentrations alter thermally modulated optical signals from skin.126,156 This is due to some physiologic and physical effects induced in temperature-modulated skin—the temperature modulation affecting mostly cutaneous vascular circulation (physiological effect), and change of glucose concentration affecting mostly cutaneous light scattering (physical effect).

Other Methods of Tissue Optical Properties Control

207

A device based on the thermo-optical response of human skin was used to collect signals from the forearms of volunteers.60,126 Glucose concentrations were correlated with temperature-modulated localized reflectance signals at wavelengths between 590 and 935 nm. There are no known NIR glucose absorption bands in this range; thus, µa is mostly defined by blood absorption reflecting hemodynamic changes in cutaneous tissue. Evidently, µ
s is a measure of the refractive index mismatch between the interstitial fluid (ISF) and tissue connective fibers. Localized reflectance data were collected continuously over a 90 min period of probe–skin contact as temperature was repetitively stepped between 22◦ C and 38◦ C for 15 temperature modulation cycles.126,156 Each cycle comprised the following steps: skin was equilibrated for 2 min at a probe temperature of 22◦ C, and the temperature was raised to 38◦ C over the course of 1 min, maintained for 2 min, and then lowered to 22◦ C over a 1-min period. At each temperature limit (during the 2 min window), four optical data packets were collected and values of µa and µ
s were determined. Temperature modulation between 38◦ C and 22◦ C caused a periodic set of cutaneous refractive index and vascular changes, leading to periodic changes in skin reflectance.126,156 A four-term linear least-squares fitting of glucose to the reflectance data was used: [Glucose] = a0 + i ai × Ri
(r, λ, T ).

(94)

The reflectance parameter R
(r, λ, T ), as defined by Eq. (94), equals loge (measured localized reflectance). Thirty-two sequences of R
(at T22◦ C ) = loge R(r, λ, T22◦ C ) and R
(at T38◦ C ) = loge R(r, λ, T38◦ C ) were used in the linear least-squares correlation. For each meal tolerance test (MTT) over the 2 hr period, the temperature sequences encompassed 20 data points.126,156 The correlation between glucose values and optical signals in this spectral range was attributed to the effect of glucose on the refractive index and on the cutaneous hemodynamic response; the correlation coefficient was in the range from 0.69 to 0.94 for two volunteers tested for six days each.126,156 The thermo-optical response method offers certain compartmentalization advantages over localized reflectance measurements that use large source-detector separations,52–54,350 as it limits sampling depth to the dermis by virtue of the probe design (short source-detector separations) and the use of temperature control. Low-intensity laser radiation, when used for spectroscopy (diagnostics) or therapy, may heat tissue and therefore distort results of tissue optical properties measurements or may induce uncontrolled change in a photobiological response of a tissue caused by a local heating. Human skin temperature increase at continuous wave (CW) laser irradiation can be estimated on the basis of experimental and theoretical studies presented in Refs. 125, 387, and 388. It was found experimentally that for near-infrared laser radiation (789 nm) guided to the skin surface of the forearm of an awake human volunteer by a 1 mm optical fiber, the temperature increased linearly with power level as (0.101±0.001)◦ C/mW at the depth of 0.5 mm, as (0.038 ± 0.001)◦ C/mW at the depth of 1 mm, and as (0.029 ± 0.0005)◦ C/mW at the depth of 1.5 mm in the range of illuminating power up to 10 mW.125

208

Optical Clearing of Tissues and Blood

The combination of the Monte Carlo technique to calculate the fluence rate distribution of light and the adaptive finite element method to solve the heat transfer equation was applied to investigate the process of hyperthermia induced by trans-skin irradiation with a He:Ne laser (633 nm).387,388 It was shown that overheated tissue volume, heating depth, and temperature can be effectively controlled by changing of the free convection boundary conditions on the tissue surface and varying power, radius, and shape of the incident laser beam. The four-layer model of human skin (epidermis, upper dermis, blood plexus, and lower dermis) with optical and thermal properties of the tissue layers taken from the literature was used for modeling. The modeling was done for Gaussian and rectangular incident light beams at noncoagulating intensities. By variation of the value of the heat transfer coefficient A (corresponding to free convection at an initial skin surface temperature equal to 34◦ C, A = 0.009 W/cm2 K; to weak isolation, A = 0.004 W/cm2 K; and to a strong one, A = 0.0005 W/cm2 K), the skin surface thermal insulation on tissue temperature distributions was studied. For a 25 mW Gaussian incident beam of 1 mm in diameter, the subsurface temperature maxima at the depths of 0.20 mm, 0.18 mm, and 0.10 mm and equal to 36.7◦ C, 41.3◦ C, and 42.8◦ C, as isolation increased were found. At the depth of 1 mm, the calculated temperature was 37.7◦ C, 38.1◦ C, and 38.5◦ C as isolation increased. It should be noted that the above mentioned experimental data are well fit to the modeled ones for free convection boundary conditions on the skin surface. From experiments, it follows that for light power of 25 mW, temperature increase at the depth of 1 mm is expected as (0.038◦ C/mW) × 25 mW = 0.95◦ C. In its turn, calculations showed that the temperature increased from the initial value without radiation of 36.4◦ C to 37.7◦ C at laser action, i.e., 1.1◦ C. The somewhat higher temperature increase in theory than expected from the experimental estimation may be explained by a shorter wavelength of the light used, which is more effectively absorbed by tissue chromophores. The loss of water by tissue due to temperature effects (freezing in a refrigerator or noncoagulating heating) seriously influences its optical properties. For instance, in an in vitro study of human aorta, the absorption coefficient increased by 20–50%, especially in the visible range, when an average 46.4% of total tissue weight was lost as a result of dehydration at tissue sample prolonged freezing in a refregirator.43,44 The weight loss was accompanied by an average shrinkage in thickness of 19.5%. Primarily because of shrinkage (more dense packing of tissue components), the absorption coefficient was increased in the spectral range 400– 1300 nm. There was only a slight increase of 2–15% in the reduced scattering coefficient in the visible range, again due to closer packing of tissue components. The slope of the wavelength dependence of the reduced scattering coefficient µ
s , which is proportional to λ−h [see Eq. (65)] is a good test for the alteration of tissue morphology at its heating or freezing. Data summarized in Table 19 demonstrate experimental values of parameter h for normal, dehydrated, and coagulated human aorta received at an in vitro study in the spectral range 400–1300 nm.43,44 Tissue dehydration (by its slow freezing) increases the slope (h) from 1.15 (control) to 1.22, which reflects denser scatterer packing at tissue shrinkage. Sample

Other Methods of Tissue Optical Properties Control

209

heating during 5 min in a saline bath at temperatures in the vicinity of the tissue coagulation threshold may increase h [for instance, for 60◦ C from 1.21 (control) to 1.28 (possibly due to a local protein coagulation)], or decrease h [for instance, for 70◦ C from 1.30 (control) to 1.10 (possibly due to more extensive protein coagulation)]. This result reflects the fact that collagen denaturation starts dominating tissue behavior between 55◦ C and 70◦ C.43 Due to aorta tissue heterogeneity, its components may have reached different end points at the end of the 5 min heating period. At 100◦ C heating of the samples in a saline bath, h was reduced from 1.38 for the normal tissue samples to 1.06 for the heated ones, and for the samples preliminarily wrapped in aluminium foil, h was reduced from 1.26 to 1.03. Table 19 Power relationship between the wavelength and the reduced scattering coefficient, and the significance of h (see Eq. (65)) values for control and experimental reduced scattering spectra (400–1300 nm) for human aorta as obtained from a t-test (rms values in parentheses).43,44

Description Dehydration Heating at 60◦ C Heating at 70◦ C Heating at 100◦ C (direct heating) Heating at 100◦ C (wrapped heating)

hcontrol

hexper

Significance, %

1.15(0.10) 1.21(0.12) 1.30(0.01)

1.22(0.13) 1.28(0.04) 1.10(0.10)

∼15 ∼25 <5

1.38(0.11)

1.06(0.07)

<5

1.26(0.08)

1.03(0.05)

<5

With tissue heating in a bath, absorption coefficients may increase up to 28% (60◦ C) or decrease up to 22% (70◦ C) on selected wavelengths. At the same time, values of reduced scattering coefficient were increased for all used heating temperatures in a wide range from 1.1 to 76% depending on the wavelength in the range from 350 to 1320 nm and heating temperature. At temperature of 60◦ C, the increase was rather smooth, 16–19% in this wavelength range. At 70◦ C, the increase of µ
s was not so smooth, changing from 1.1 to 24.8%. Heating up to 100◦ C gave an increase in the reduced scattering coefficient of 22–76%, but tissue wrapping decreased these values to 15–54%. Such complex behavior of scattering and absorption properties of precoagulated and coagulated tissues reflects tissue heterogeneity, the specificity of the protein denaturation process leading to the appearance of coarse and small thermally coagulated granular cellular proteins and some tissue chromophores damage, as well as interactions of heated saline with tissue resulting in saline and chromophores diffusion.43 Following Ref. 43,* it may be verified that tissue progresses from normal to denaturated states between 60◦ C and 70◦ C, and that at ∼60◦ C some changes in the optical properties caused by thermal damage are still reversible, even though * The errata of the paper is presented in Ref. 3, p. 379.

210

Optical Clearing of Tissues and Blood

the thermal threshold for protein coagulation is exceeded. In general, the complex behavior of tissue optical properties at heating can be explained by particular changes in tissue morphology. To model optical properties of heated tissues, the modified morphology should be expressed in such terms as scatterer size distribution, particle and interstitial fluid refractive index mismatch, particle packing, and chromophore concentrations. Long-pulsed laser heating may induce reversible and irreversible changes in the optical properties of tissue.44,127 The total transmittance decreases and the diffuse reflectance increased in both fresh and precoagulated human skin and canine aorta samples when they were irradiated by a 0.2 ms pulsed Nd:YAG laser emitting at 1064 nm with a repetition rate of 10 Hz (20 pulses of 0.9 J/pulse) and a 1.5 mm light spot.44 The existence of nonlinear behavior in the optics of biological media was indicated.44 Possible mechanisms responsible for this nonlinear optical response are listed in Table 20. The in vitro skin-equivalent raft tissue irradiated with one pulse from a perovskite (λ = 1341 nm) laser with fluence of 20 J/cm2 and pulse duration of 20 ms showed thermal injury areas characterized by less scattering as seen on the OCT images and loosing of the SHG signal.127 Such behavior was interpreted by authors of Ref. 127 as collagen fibers’ disintegration at thermal tissue protein denaturation, which can be rejuvenated at tissue healing. Table 20 Possible mechanisms responsible for inducing reversible changes in tissue optical response on laser long-pulsed irradiation.44

Mechanism

Description

Optical response

Thermal lensing, n(T ) = n(273 K) + T (r, z, t)(dn/dT )

Gradient in the index of refraction caused by nonuniform heating Change in the size and/or shape of scatterers due to temperature rise Temporary local dehydration during laser heating Decrease in tissue density and increase in tissue thickness caused by thermal expansion of tissue

Decrease in Tt and increase in Rd

Temperature dependence of the reduced scattering coefficient: µ
s (T ) = µs (T )[1 − g(T )] Water transport Thermal expansion

Increase in Tt and crease in Rd (as µ
s creases) Increase in Tt and crease in Rd Decrease in Tt and crease in Rd

dededein-

Laser ablation or coagulation is usually accompanied by a change in the normal optical properties of the tissue.44,194 For example, ablation of aortal tissue using an excimer laser (308 nm) results in a 2.3–3.7-fold increase in its optical density compared with the untreated material.194 Published results on optical properties of coagulated tissues are presented in Table 14. In spite of some variations in dependence on tissue type, wavelength studied, and sample preparation technique, the

Other Methods of Tissue Optical Properties Control

211

general tendency is the increase both of the absorption and scattering coefficients from a few dozen to 200–300% at tissue coagulation. Low temperatures (+12◦ C) sometime result in the so-called cold cataract, i.e., a sharp rise in the scattering coefficient due to protein aggregation.148,149 This process is reversed with an increase in temperature. Cryogenic temperatures used in cryosurgery may also change scattering properties of tissues due to local variations in refractive index, such as at the boundary between liquid and frozen water in tissue.134 The corresponding subsurface morphological changes were evident during freezing (−80◦ C) of in vivo hamster skin.

9.3 Tissue Whitening Sometimes, in order to provide higher contrast images of intracellular components of the epithelial tissues, instead of optical clearing the usage of induced tissue turbidity (whitening) is more preferable.128,133,136,146 For example, a fundamental part of the colposcopic exam is the use of acetic acid, which when applied to the cervix induces transient whitening changes in the epithelial tissues.133 The spatial and temporal changes of aceto-whitening are the major visual diagnostic indicators in the determination of the location of the most severe dysplastic regions. The acetowhitening effect causes a differential brightening of dysplastic tissue relative to normal tissue, and besides cervical disorders is used for skin and other epithelial disease screening. The brightening of nuclei enhances the contrast and significantly improves detectability of nuclear morphology in basal cell cancers.134 Under normal conditions, the nucleus contains a diffuse network of thin chromatin filaments that are typically 30–100 nm in diameter and occupy a small volume within. Because of the small dimensions and small difference between the refractive index of the chromatin, which can be estimated as 1.39, and surrounding tissue components (cell cytoplasm and interstitial fluid) of about 1.35, its backscattering is low. The acetic acid causes the assembling of the chromatin into thick fibers that are 1–5 µm in diameter; the compacted chromatin fills a large fraction of the intranuclear volume134 and some increase in its refractive index is also expected;133 thus, the backscattering signal from the nuclei is increased and they appear bright. After washing ex vivo samples of human epidermis with 5% acetic acid for three minutes, the epidermal cell nuclei appear bright when viewed in the confocal reflectance images.134 The temporal kinetics of the acetowhitening process as measured by the reflected light maximizes during the first 1–2 min and then decays over several minutes (5–10 min) thereafter, allowing one to clearly distinguish high-grade cervical intraepithelial neoplasia (CIN 2/3) from normal cervical epithelium when the ratio of green to red light intensities of the backscattered light is analyzed.133 Normalizing the green by the red light preserves kinetics in the reflected signal, indicating that the reflected light has a spectral change. Roughly, acetowhitening in CIN 2/3 cervical tissue in vivo causes a 20–80% rise of the original reflectance. In contrast, mature squamous epithelium appears to increase in reflectance by only

212

Optical Clearing of Tissues and Blood

approximately 5%; moreover, reflectance is constant in time after acetic acid application. All of the contrast agents, such as acetic acid, hyperosmotic saline, toluidine blue, and Lugol’s iodine solutions, when added to normal and malignant human cervical cancer cells had an effect on the backscattering properties of cells, as viewed under reflectance confocal microscopy at the wavelength 808 nm.146 The source of changes in the appearance of the confocal images is likely related to changes in the refractive index profile of the cells. The tested agents increase the relative difference between gray-scale values for representative regions of the confocal images of the nucleus and cytoplasm of cells, with the highest difference for cancer cells. Among the studied agents, acetic acid provides the highest difference.

Conclusion

This book shows that optical immersion technology allows one to effectively control the optical properties of tissues and blood. Such control leads to the essential reduction of scattering and therefore causes much higher transmittance (optical clearing) and the appearance of a large amount of least-scattered (snake) and ballistic photons, allowing for successful application of coherent-domain and polarization imaging techniques. The dynamics of tissue optical clearing, defined by the dynamics of refractive index matching, is characterized by a time response of about 5–30 min, which in its turn depends on the diffusivity of the immersion agent in a tissue layer, water diffusion rate, and tissue layer thickness. The tissue and cell swelling or shrinkage may play an important role in the tissue clearing process at application of hyperosmotic agents. In vivo reflectance spectrophotometry and frequency-domain measurements for immersed tissues show that the refractive index matching technique provided by the appropriate chemical agent can be successfully used in tissue spectroscopy and imaging when radical reduction of scattering properties is needed. Hyperdermal injection of glucose causes the essential clearing of human skin. For such tissues as sclera or cornea, some drops of glucose is sufficient to make up very high and rather prolonged tissue clearing. In in vivo experiments, the impregnation of a tissue by an agent is more effective than in in vitro studies due to the higher diffusivity of an agent at physiological temperature and by involvement of blood and lymph microvessels into the process of agent distribution. Optical clearing may play a significant role in the designing of optical technologies for imaging of small living animals in studies of gene expression. Dynamic optical characteristics can be used for the determination of the diffusion coefficient and concentration of endogenous (metabolic) and exogenous (chemical agent) fluids in human sclera, skin, and other tissues. The immersion technique has great potential in noninvasive medical diagnostics using OCT due to the rather small thickness of tissue layers usually examined by OCT, which allows for fast impregnation of a target tissue during a topical application of an immersion liquid. It has been demonstrated that the body’s interior tissues such as the blood vessel wall, esophagus, stomach, cervix, and colon can usually be imaged at a depth of about 1–2 mm. For more effective diagnosis using OCT, a higher penetration depth can be provided by applying immersion substances. The method of clearing tissue is convenient, inexpensive, and simple for diagnostic purposes; in particular, it can be applied for in vivo monitoring of microcirculation. It may be useful for the study of structure and function of blood 213

214

Optical Clearing of Tissues and Blood

microvessels—diameters of arterioles and venules, capillary density, bifurcation angles, etc. Optical clearing might be a fruitful technique for various methods of tissue spectroscopy, microscopy, and imaging (Raman, fluorescence, confocal, laser scanning, near-field, multiphoton, SHG, etc.), where scattering is a serious limitation.

(a)

(b) Figure 99 Optical clearing of the cranial bone samples.181 (a) The reflectance spectra of porcine temple bone measured concurrently with administration of propylene glycol at different time intervals (a fiber-optical spectrometer with a source-detector separation of about 200 µm). (b) The human bone samples: the reduced scattering coefficient wavelength dependence before (upper) and after 1 hr administration of the glycerol (lower); integrating sphere spectrometer, inverse adding-doubling method.

Conclusion

215

Encouraging results were recently received in the enhancement of a single-photon fluorescent signal, as well as in improvement of image depth and contrast of twophoton fluorescence and SHG imaging at tissue immersion. It is important to note that the reduction of light scattering may help in the differentiation of various fluorophores in the depth of a tissue—for instance, of skin. The concept of index matching improving the optical penetration depth of whole blood has been proved experimentally in in vitro studies. It should be accounted for that blood optical clearing is defined not only by the refractive index matching effect, but also by changes of red blood cells’ size and their aggregation ability when chemicals are added. Immersion optical clearing technology is applicable not only to soft tissues and blood, but also to hard tissues. At present, tendon (see Figs. 49 and 50),33,34,101 cranial bones (see Fig. 99),181 and tooth182 have been tested. As it follows from Fig. 99(a), bone optical reflectance is very sensitive to impregnation of the superficial bone layers by clearing agents. This is a viable method for the study of bone demineralization (bone porosity). Figure 99(b) also demonstrates a high sensitivity of the scattering coefficient (reduction) to an immersion agent action, especially for the longer wavelengths, where water dispersion is essential. Many of the tested agents and methods of their delivery have some advantages and disadvantages. The main disadvantages are the osmotic stress, which occurs at high concentration of osmotically active agent applied, and low permeability of tissue cell structures for the clearing agents. Therefore, the finding of new agents and delivery techniques is important. For example, it was recently proved that the mixture of prepolymers PPG/PEG has excellent clearing properties; the diffusivity in human skin is on the order of 0.2 mm/hr and it has no detectable influence on the tissue blood vascular system.102,274,325 A few methods of enhancing skin permeability based on limited thermal damage in the SC was also recently proposed.196,347,349–352 Tissue optical clearing technology using an appropriate immersion agent can be applied for the temporal reduction of scattering needed for providing a precise laser photodisruption of the underlying tissue in ophthalmology36 or skin tattoo removal.374,375,389 Evidently, reversible tissue optical clearing technology has valuable features to be applied not only to tissue spectroscopy and diagnostics, but also to a variety of laser and photothermal therapies and surgeries.

References

1. G. Müller, B. Chance, and R. Alfano, et al. (eds.), Medical Optical Tomography: Functional Imaging and Monitoring, vol. IS11, SPIE Press, Bellingham, WA, 1993. 2. G. Müller and A. Roggan (eds.), Laser-Induced Interstitial Thermotherapy, vol. PM25, SPIE Press, Bellingham, WA, 1995. 3. V. V. Tuchin (ed.), Selected Papers on Tissue Optics: Applications in Medical Diagnostics and Therapy, vol. MS102, SPIE Press, Bellingham, WA, 1994. 4. B. Chance, M. Cope, E. Gratton, N. Ramanujam, and B. Tromberg, “Phase Measurement of Light Absorption and Scatter in Human Tissue,” Rev. Sci. Instrum., vol. 69, no. 10, 1998, pp. 3457–3481. 5. A. V. Priezzhev, V. V. Tuchin, and L. P. Shubochkin, Laser Diagnostics in Biology and Medicine, Nauka, Moscow, 1989. 6. V. V. Tuchin, Lasers and Fiber Optics in Biomedical Science, Saratov Univ. Press, Saratov, 1998. 7. V. V. Tuchin, Tissue Optics: Light Scattering Methods and Instruments for Medical Diagnosis, SPIE Tutorial Texts in Optical Engineering, vol. TT38, SPIE Press, Bellingham, WA, 2000. 8. V. V. Tuchin, “Optical Clearing of Tissue and Blood Using Immersion Method,” J. Phys. D: Appl. Phys., vol. 38, 2005, pp. 2497–2518. 9. V. V. Tuchin, “Lasers and Fiber Optics in Biomedicine,” Laser Phys., vol. 3, 1993, pp. 767–820; 925–950. 10. V. V. Tuchin, “Lasers Light Scattering in Biomedical Diagnostics and Therapy,” J. Laser Appl., vol. 5, nos. 2, 3, 1993, pp. 43–60. 11. V. V. Tuchin, “Light Scattering Study of Tissues,” Phys.—Uspekhi, vol. 40, no. 5, 1997, pp. 495–515. 12. S. Bown and G. Buonaccorsi (eds.), “Special Issue on VI Biennial Meeting of the International Photodynamic Association,” Lasers Med. Sci., vol. 12, no. 3, 1997, pp. 180–284. 13. G. J. Mueller and D. H. Sliney (eds.), Dosimetry of Laser Radiation in Medicine and Biology, vol. IS5, SPIE Press, Bellingham, WA, 1989. 14. V. V. Tuchin (ed.), Coherent-Domain Optical Methods: Biomedical Diagnostics, Environmental and Material Science, Kluwer Academic Publishers, Boston, vols. 1 and 2, 2004. 15. V. V. Tuchin, L. V. Wang, and D. A. Zimnyakov, Optical Polarization in Biomedical Applications, Springer-Verlag, New York, 2005.

218

Optical Clearing of Tissues and Blood

16. A. Roggan, M. Friebel, K. Dorschel, A. Hahn, and G. Mueller, “Optical Properties of Circulating Human Blood in the Wavelength Range 400–2500 nm,” J. Biomed. Opt., vol. 4, no. 1, 1999, pp. 36–46. 17. A. M. K. Nilsson, G. W. Lucassen, W. Verkruysse, S. Andersson-Engels, and M. J. C. van Gemert, “Changes in Optical Properties of Human Whole Blood In Vitro Due to Slow Heating,” Photochem. Photobiol., vol. 65, no. 2, 1997, pp. 366–373. 18. H. J. S. M. Sterenborg and J. C. Van der Leun, “Change in Epidermal Transmission Due to UV-Induced Hyperplasia in Hairless Mice: A First Approximation of the Action Spectrum,” Photodermatology, vol. 5, 1988, pp. 71–82. 19. R. R. Anderson and J. A. Parrish, “Optical Properties of Human Skin,” in The Science of Photomedicine, J. D. Regan and J. A. Parrish (eds.), Plenum Press, New York, 1982, pp. 147–194. 20. A. Dunn, C. Smithpeter, A. J. Welch, and R. Richards-Kortum, “FiniteDifference Time-Domain Simulation of Light Scattering from Single Cells,” J. Biomed. Opt., vol. 2, no. 3, 1997, pp. 262–266. 21. V. V. Bakutkin, I. L. Maksimova, P. I. Saprykin, V. V. Tuchin, and L. P. Shubochkin, “Light Scattering by the Human Eye Sclera,” J. Appl. Spectrosc. (USSR), vol. 46, no. 1, 1987, pp. 104–107. 22. I. L. Maksimova, V. V. Tuchin, and L. P. Shubochkin, “Light Propagation in Anisotropic Biological Objects,” in Laser Beams, Khabarovsk Technical Inst. Press, Khabarovsk, USSR, 1985, pp. 91–95. 23. V. V. Tuchin (ed.), SPIE CIS Selected Papers: Coherence-Domain Methods in Biomedical Optics, vol. 2732, SPIE Press, Bellingham, WA, 1996. 24. V. V. Tuchin, “Coherence-Domain Methods in Tissue and Cell Optics,” Laser Phys., vol. 8, no. 4, 1998, pp. 807–849. 25. B. W. Henderson and T. J. Dougherty (eds.), Photodynamic Therapy: Basic Principles and Clinical Applications, Marcel-Dekker, New York, 1992. 26. H.-P. Berlien and G. J. Mueller (eds.), Applied Laser Medicine, SpringerVerlag, Berlin, 2003. 27. J. R. Mourant, J. P. Freyer, A. H. Hielscher, A. A. Eick, D. Shen, and T. M. Johnson, “Mechanisms of Light Scattering from Biological Cells Relevant to Noninvasive Optical-Tissue Diagnostics,” Appl. Opt., vol. 37, 1998, pp. 3586–3593. 28. R. Drezek, A. Dunn, and R. Richards-Kortum, “Light Scattering from Cells: Finite-Difference Time-Domain Simulations and Goniometric Measurements,” Appl. Opt., vol. 38, no. 16, 1999, pp. 3651–3661. 29. V. V. Tuchin (ed.), Handbook of Optical Biomedical Diagnostics, vol. PM107, SPIE Press, Bellingham, WA, 2002. 30. D. A. Zimnyakov and V. V. Tuchin, “Optical Tomography of Tissues: Overview,” Quantum Electron., vol. 32, no. 10, 2002, pp. 849–867. 31. D. A. Zimnyakov, V. V. Tuchin, and A. A. Mishin, “Spatial Speckle Correlometry in Applications to Tissue Structure Monitoring,” Appl. Opt., vol. 36, 1997, pp. 5594–5607.

References

219

32. V. V. Tuchin, I. L. Maksimova, D. A. Zimnyakov, I. L. Kon, A. H. Mavlutov, and A. A. Mishin, “Light Propagation in Tissues with Controlled Optical Properties,” J. Biomed. Opt., vol. 2, 1997, pp. 401–417. 33. G. V. Simonenko, T. P. Denisova, N. A. Lakodina, and V. V. Tuchin, “Measurement of an Optical Anisotropy of Biotissues,” Proc. SPIE 3915, 2000, pp. 152–157. 34. G. V. Simonenko, V. V. Tuchin, and N. A. Lakodina, “Measurement of the Optical Anisotropy of Biological Tissues with the Use of a Nematic Liquid Crystal Cell,” J. Opt. Technol., vol. 67, no. 6, 2000, pp. 559–562. 35. P. O. Rol, Optics for Transscleral Laser Applications: Dissertation No. 9655 for Doctor of Natural Sciences, Swiss Federal Institute of Technology, Zurich, Switzerland, 1992. 36. Z. S. Sacks, R. M. Kurtz, T. Juhasz, and G. A. Mourau, “High Precision Subsurface Photodisruption in Human Sclera,” J. Biomed. Opt., vol. 7, no. 3, 2002, pp. 442–450. 37. N. G. Khlebtsov, I. L. Maksimova, V. V. Tuchin, and L. Wang, “Introduction to Light Scattering by Biological Objects,” Chapter 1 in Handbook of Optical Biomedical Diagnostics, vol. PM107, V. V. Tuchin (ed.), SPIE Press, Bellingham, WA, 2002, pp. 31–167. 38. Z. S. Sacks, D. L. Craig, R. M. Kurtz, T. Juhasz, and G. Mourou, “Spatially Resolved Transmission of Highly Focused Beams Through Cornea and Sclera Between 1400 and 1800 nm,” Proc. SPIE 3726, 1999, pp. 522–527. 39. I. L. Maksimova, D. A. Zimnyakov, and V. V. Tuchin, “Controlling of Tissue Optical Properties I. Spectral Characteristics of Eye Sclera,” Opt. Spectrosc., vol. 89, 2000, pp. 78–86. 40. D. A. Zimnyakov, I. L. Maksimova, and V. V. Tuchin, “Controlling of Tissue Optical Properties II. Coherent Methods of Tissue Structure Study,” Opt. Spectrosc., vol. 88, 2000, pp. 1026–1034. 41. B. Nemati, H. G. Rylander III, and A. J. Welch, “Optical Properties of Conjunctiva, Sclera, and the Ciliary’s Body and Their Consequences for Transscleral Cyclophotocoagulation,” Appl. Opt., vol. 35, no. 19, 1966, pp. 3321– 3327. 42. B. Nemati, A. Dunn, A. J. Welch, and H. G. Rylander III, “Optical Model for Light Distribution during Transscleral Cyclophotocoagulation,” Appl. Opt., vol. 37, no. 4, 1998, pp. 764–771. 43. I. F. Cilesiz and A. J. Welch, “Light Dosimetry: Effects of Dehydration and Thermal Damage on the Optical Properties of the Human Aorta,” Appl. Opt., vol. 32, 1993, pp. 477–487. 44. W.-C. Lin, M. Motamedi, and A. J. Welch, “Dynamics of Tissue Optics During Laser Heating of Turbid Media,” Appl. Opt., vol. 35, no. 19, 1996, pp. 3413–3420. 45. G. Vargas, E. K. Chan, J. K. Barton, H. G. Rylander III, and A. J. Welch, “Use of an Agent to Reduce Scattering in Skin,” Laser. Surg. Med., vol. 24, 1999, pp. 133–141.

220

Optical Clearing of Tissues and Blood

46. A. P. Ivanov, S. A. Makarevich, and A. Ya. Khairulina, “Propagation of Radiation in Tissues and Liquids with Densily Packed Scatterers,” J. Appl. Spectrosc. (USSR), vol. 47, no. 4, 1988, pp. 662–668. 47. V. V. Tuchin, X. Xu, and R. K. Wang, “Dynamic Optical Coherence Tomography in Optical Clearing, Sedimentation and Aggregation Study of Immersed Blood,” Appl. Opt. – OT, vol. 41, no. 1, 2002, pp. 258–271. 48. R. Barer and S. Joseph, “Refractometry of Living Cells,” Q. J. Microsc. Sci., vol. 95, 1954, pp. 399–406. 49. D. M. Zhestkov, A. N. Bashkatov, E. A. Genina, and V. V. Tuchin, “Influence of Clearing Solutions Osmolarity on the Optical Properties of RBC,” Proc. SPIE 5474, 2004, pp. 321–330. 50. X. Wu, L. Stinger, and G. W. Faris, “Determination of Tissue Properties by Immersion in a Matched Scattering Fluid,” Proc. SPIE 2979, 1997, pp. 300– 306. 51. V. V. Tuchin, “Fundamentals of Low-Intensity Laser Radiation Interaction with Biotissues: Dosimetry and Diagnostical Aspects,” Bull. Russian Acad. Sci., Phys. Ser., vol. 59, no. 6, 1995, pp. 120–143. 52. J. S. Maier, S. A. Walker, S. Fantini, M. A. Franceschini, and E. Gratton, “Possible Correlation between Blood Glucose Concentration and the Reduced Scattering Coefficient of Tissues in the Near Infrared,” Opt. Lett., vol. 19, 1994, pp. 2062–2064. 53. M. Kohl, M. Cope, M. Essenpreis, and D. Böcker, “Influence of Glucose Concentration on Light Scattering in Tissue-Simulating Phantoms,” Opt. Lett., vol. 19, 1994, pp. 2170–2172. 54. J. T. Bruulsema, J. E. Hayward, T. J. Farrell, M. S. Patterson, L. Heinemann, M. Berger, T. Koschinsky, J. Sandahal-Christiansen, H. Orskov, M. Essenpreis, G. Schmelzeisen-Redeker, and D. Böcker, “Correlation between Blood Glucose Concentration in Diabetics and Noninvasively Measured Tissue Optical Scattering Coefficient,” Opt. Lett., vol. 22, no. 3, 1997, pp. 190–192. 55. D. A. Zimnyakov, V. V. Tuchin, and A. G. Yodh, “Characteristic Scales of Optical Field Depolarization and Decorrelation for Multiple Scattering Media and Tissues,” J. Biomed. Opt., vol. 4, 1999, pp. 157–163. 56. Yu. P. Sinichkin, N. Kollias, G. Zonios, S. R. Utz, and V. V. Tuchin, “Reflectance and Fluorescence Spectroscopy of Human Skin In Vivo,” Chapter 13 in Optical Biomedical Diagnostics, V. V. Tuchin (ed.), SPIE Press, Bellingham, WA, 2002, pp. 725–785. 57. E. K. Chan, B. Sorg, D. Protsenko, M. O’Neil, M. Motamedi, and A. J. Welch, “Effects of Compression on Soft Tissue Optical Properties,” IEEE J. Select. Tops Quant. Electron., vol. 2, no. 4, 1996, pp. 943–950. 58. E. Chan, T. Menovsky, and A. J. Welch, “Effect of Cryogenic Grinding of Soft-Tissue Optical Properties,” Appl. Opt., vol. 35, no. 22, 1996, pp. 4526– 4532. 59. M. Kohl, M. Essenpreis, and M. Cope, “The Influence of Glucose Concentration upon the Transport of Light in Tissue-Simulating Phantoms,” Phys. Med. Biol., vol. 40, 1995, pp. 1267–1287.

References

221

60. O. Khalil, S.-J. Yeh, M. G. Lowery, X. Wu, C. F. Hanna, S. Kantor, T.-W. Jeng, J. S. Kanger, R. A. Bolt, and F. F. de Mul, “Temperature Modulation of the Visible and Near Infrared Absorption and Scattering Coefficients of Human Skin,” J. Biomed. Opt., vol. 8(2), 2003, pp. 191–205. 61. A. N. Bashkatov, “Controlling of Optical Properties of Tissues at Action by Osmotically Active Immersion Liquids,” Cand. Science Thesis, Saratov State Univ., Saratov, 2002. 62. E. A. Genina, A. N. Bashkatov, V. I. Kochubey, and V. V. Tuchin, “Optical Clearing of Human Dura Mater,” Opt. Spectrosc., vol. 98, no. 3, 2005, pp. 470–476. 63. L. S. Dolin, F. I. Feldchtein, G. V. Gelikonov, V. M. Gelikonov, N. D. Gladkova, R. R. Iksanov, V. A. Kamensky, R. V. Kuranov, A. M. Sergeev, N. M. Shakhova, and I. V. Turchin, “Fundamentals of OCT and Clinical Applications of Endoscopic OCT,” Chapter 17, in Coherent-Domain Optical Methods: Biomedical Diagnostics, Environmental and Material Science, vol. 2, V. V. Tuchin (ed.), Kluwer Academic Publishers, Boston, 2004, pp. 211– 270. 64. V. V. Tuchin, D. M. Zhestkov, A. N. Bashkatov, and E. A. Genina, “Theoretical Study of Immersion Optical Clearing of Blood in Vessels at Local Hemolysis,” Optics Express, vol. 12, 2004, pp. 2966–2971. 65. R. Barer, K. F. A. Ross, and S. Tkaczyk, “Refractometry of Living Cells,” Nature, vol. 171(4356), 1953, pp. 720–724. 66. M. Haruna, K. Yoden, M. Ohmi, and A. Seiyama, “Detection of Phase Transition of a Biological Membrane by Precise Refractive-Index Measurement Based on the Low Coherence Interferometry,” Proc. SPIE 3915, 2000, pp. 188–193. 67. D. J. Faber, M. C. G. Aalders, E. G. Mik, B. A. Hooper, M. J. C. van Gemert, and T. G. van Leeuwen, “Oxygen Saturation-Dependent Absorption and Scattering of Blood,” Phys. Rev. Lett., vol. 93, 2004, 028102. 68. V. V. Tuchin, R. K. Wang, E. I. Galanzha, N. A. Lakodina, and A. V. Solovieva, “Monitoring of Glycated Hemoglobin in a Whole Blood by Refractive Index Measurement with OCT,” Conference Program CLEO/QELS (2003) Baltimore, June 1–6, 2003, pp. 120–122. 69. V. V. Tuchin, R. K. Wang, E. I. Galanzha, J. B. Elder, and D. M. Zhestkov, “Monitoring of Glycated Hemoglobin by OCT Measurement of Refractive Index,” Proc. SPIE 5316, 2004, pp. 66–77. 70. G. Mazarevica, T. Freivalds, and A. Jurka, “Properties of Erythrocyte Light Refraction in Diabetic Patients,” J. Biomed. Opt., vol. 7, 2002, pp. 244–247. 71. A. N. Bashkatov, E. A. Genina, V. I. Kochubey, Yu. P. Sinichkin, A. A. Korobov, N. A. Lakodina, and V. V. Tuchin, “In Vitro Study of Control of Human Dura Mater Optical Properties by Acting of Osmotical Liquids,” Proc. SPIE 4162, 2000, pp. 182–188. 72. A. N. Bashkatov, E. A. Genina, V. I. Kochubey, and V. V. Tuchin, “Estimation of Wavelength Dependence of Refractive Index of Collagen Fibers of Scleral Tissue,” Proc. SPIE 4162, 2000, pp. 265–268.

222

Optical Clearing of Tissues and Blood

73. G. A. Askar’yan, “The Increasing of Laser and Other Radiation Transport through Soft Turbid Physical and Biological Media,” Sov. J. Quant. Electron., vol. 9, no. 7, 1982, pp. 1379–1383. 74. F. Veretout, M. Delaye, and A. Tardieu, “Molecular Basis of Lens Transparency. Osmotic Pressure and X-ray Analysis of α-Crystalline Solutions,” J. Mol. Biol., vol. 205, 1989, pp. 713–728. 75. R. Barer and S. Joseph, “Refractometry of Living Cells,” Q. J. Microsc. Sci., vol. 95, 1954, pp. 399–406. 76. B. A. Fikhman, Microbiological Refractometry, Medicine, Moscow, 1967. 77. E. Eppich, J. Beuthan, C. Dressler, and G. Müller, “Optical Phase Measurements on Biological Cells,” Laser Physics, vol. 10, 2000, pp. 467–477. 78. B. Beauvoit, T. Kitai, and B. Chance, “Contribution of the Mitochondrial Compartment to the Optical Properties of the Rat Liver: A Theoretical and Practical Approach,” Biophys. J., vol. 67, 1994, pp. 2501–2510. 79. B. Chance, H. Liu, T. Kitai, and Y. Zhang, “Effects of Solutes on Optical Properties of Biological Materials: Models, Cells, and Tissues,” Anal. Biochem., vol. 227, 1995, pp. 351–362. 80. H. Liu, B. Beauvoit, M. Kimura, and B. Chance, “Dependence of Tissue Optical Properties on Solute-Induced Changes in Refractive Index and Osmolarity,” J. Biomed. Opt., vol. 1, 1996, pp. 200–211. 81. J. Qu and B. C. Wilson, “Monte Carlo Modeling Studies of the Effect of Physiological Factors and Other Analyses on the Determination of Glucose Concentration In Vivo by Near Infrared Optical Absorption and Scattering Measurements,” J. Biomed. Opt., vol. 2, no. 3, 1997, pp. 319–325. 82. D. A. Zimnyakov, V. V. Tuchin, A. A. Mishin, I. L. Kon, and A. N. Serov, “In Vitro Human Sclera Structure Analysis Using Tissue Optical Immersion Effect,” Proc. SPIE 2673, 1996, pp. 233–243. 83. V. V. Tuchin, D. A. Zimnykov, I. L. Maksimova, G. G. Akchurin, ´ A. A. Mishin, S. R. Utz, and I. S. Peretochkin, “The Coherent, Low-Soherent and Polarized Light Interaction with Tissues Undergo the Refractive Indices Matching Control,” Proc. SPIE 3251, 1998, pp. 12–21. 84. A. N. Bashkatov, V. V. Tuchin, E. A. Genina, Yu. P. Sinichkin, N. A. Lakodina, and V. I. Kochubey, “The Human Sclera Dynamic Spectra: In Vitro and In Vivo Measurements,” Proc. SPIE 3591, 1999, pp. 311–319. 85. V. V. Tuchin, “Coherent Optical Techniques for the Analysis of Tissue Structure and Dynamics,” J. Biomed. Opt., vol. 4, no. 1, 1999, pp. 106–124. 86. V. V. Tuchin, J. Culver, C. Cheung, S. A. Tatarkova, M. A. DellaVecchia, D. Zimnyakov, A. Chaussky, A. G. Yodh, and B. Chance, “Refractive Index Matching of Tissue Components as a New Technology for Correlation and Diffusing-Photon Spectroscopy and Imaging,” Proc. SPIE 3598, 1999, pp. 111–120. 87. V. V. Tuchin (ed.), Controlling of Tissue Optical Properties: Applications in Clinical Study, Proc. SPIE 4162, 2000.

References

223

88. V. V. Tuchin, “Controlling of Tissue Optical Properties,” Proc. SPIE 4001, 2000, pp. 30–53. 89. V. V. Tuchin, “Advances in Immersion Control of Optical Properties of Tissues and Blood,” Proc. SPIE 5254, 2003, pp. 1–13. 90. V. V. Tuchin, A. N. Bashkatov, E. A. Genina, Yu. P. Sinichkin, and N. A. Lakodina, “In Vivo Investigation of the Immersion-Liquid-Induced Human Skin Clearing Dynamics,” Tech. Phys. Lett., vol. 27, no. 6, 2001, pp. 489–490. 91. R. K. Wang, X. Xu, V. V. Tuchin, and J. B. Elder, “Concurrent Enhancement of Imaging Depth and Contrast for Optical Coherence Tomography by Hyperosmotic Agents,” J. Opt. Soc. Am. B, vol. 18, 2001, pp. 948–953. 92. R. K. Wang and V. V. Tuchin, “Enhance Light Penetration in Tissue for High Resolution Optical Imaging Techniques by the Use of Biocompatible Chemical Agents,” J. X-Ray Sci. Technol., vol. 10, 2002, pp. 167–176. 93. R. K. Wang and J. B. Elder, “Propylene Glycol as a Contrast Agent for Optical Coherence Tomography to Image Gastrointestinal Tissue,” Lasers Surg. Med., vol. 30, 2002, pp. 201–208. 94. A. N. Bashkatov, E. A. Genina, and V. V. Tuchin, “Optical Immersion as a Tool for Tissue Scattering Properties Control,” in Perspectives in Engineering Optics, K. Singh and V. K. Rastogi (eds.), Anita Publications, New Delhi, 2003, pp. 313–334. 95. V. V. Tuchin, “Optical Spectroscopy of Tissue,” in Encyclopedia of Optical Engineering, R. G. Driggers (ed.), Marcel-Dekker, New York, 2003, pp. 1814–1829. www.dekker.com/servlet/product/DOI/ 101081EEOE120009712 96. V. V. Tuchin, “Light-Tissue Interactions,” in Biomedical Photonics Handbook, T. Vo-Dinh (ed.), CRC Press, Boca Raton, 2003, pp. 3-1–3-26. 97. R. K. Wang, X. Xu, Y. He, and J. B. Elder, “Investigation of Optical Clearing of Gastric Tissue Immersed with Hyperosmotic Agents,” IEEE J. Select. Tops. Quant. Electron., vol. 9, 2003, pp. 234–242. 98. X. Xu and R. K. Wang, “Synergetic Effect of Hyperosmotic Agents of Dimethyl Sulfoxide and Glycerol on Optical Clearing of Gastric Tissue Studied with Near Infrared Spectroscopy,” Phys. Med. Biol., vol. 49, 2004, pp. 457–468. 99. I. V. Meglinsky, A. N. Bashkatov, E. A. Genina, D. Yu. Churmakov, and V. V. Tuchin, “The Enhancement of Confocal Images of Tissues at Bulk Optical Immersion,” Quantum Electron., vol. 32, no. 10, 2002, pp. 875–882. 100. I. V. Meglinsky, A. N. Bashkatov, E. A. Genina, D. Yu. Churmakov, and V. V. Tuchin, “Study of the Possibility of Increasing the Probing Depth by the Method of Reflection Confocal Microscopy upon Immersion Clearing of Near-Surface Human Skin Layers,” Laser Phys., vol. 13, no. 1, 2003, pp. 65– 69. 101. A. T. Yeh, B. Choi, J. S. Nelson, and B. J. Tromberg, “Reversible Dissociation of Collagen in Tissues,” J. Invest. Dermatol., vol. 121, 2003, pp. 1332–1335.

224

Optical Clearing of Tissues and Blood

102. M. H. Khan, B. Choi, S. Chess, K. M. Kelly, J. McCullough, and J. S. Nelson, “Optical Clearing of In Vivo Human Skin: Implications for Light-Based Diagnostic Imaging and Therapeutics,” Lasers Surg. Med., vol. 34, 2004, pp. 83– 85. 103. F. Zhou and R. K. Wang, “Theoretical Model on Optical Clearing of Biological Tissue with Semipermeable Chemical Agents,” Proc. SPIE 5330, 2004, pp. 215–221. 104. G. Vargas, K. F. Chan, S. L. Thomsen, and A. J. Welch, “Use of Osmotically Active Agents to Alter Optical Properties of Tissue: Effects on the Detected Fluorescence Signal Measured through Skin,” Lasers Surg. Med., vol. 29, 2001, pp. 213–220. 105. V. V. Tuchin and A. B. Pravdin, “Dynamics of Skin Diffuse Reflectance and Autofluorescence at Tissue Optical Immersion,” in CD-Edition of the Materials on European Workshop “BioPhotonics 2002,” October 18–20, 2002, Heraklion, Crete, Greece. 106. D. Y. Churmakov, I. V. Meglinski, and D. A. Greenhalgh, “Amending of Fluorescence Sensor Signal Localization in Human Skin by Matching of the Refractive Index,” J. Biomed. Opt., vol. 9, 2004, pp. 339–346. 107. Y. He, R. K. Wang, and D. Xing, “Enhanced Sensitivity and Spatial Resolution for In Vivo Imaging with Low-Level Light-Emitting Probes by Use of Biocompatible Chemical Agents,” Opt. Lett., vol. 28, no. 21, 2003, pp. 2076– 2078. 108. E. I. Galanzha, V. V. Tuchin, Q. Luo, H. Cheng, and A. V. Solov’eva, “The Action of Osmotically Active Drugs on Optical Properties of Skin and State of Microcirculation in Experiments,” Asian J. Phys., vol. 10, no. 4, 2001, pp. 503–511. 109. G. Vargas, A. Readinger, S. S. Dosier, and A. J. Welch, “Morphological Changes in Blood Vessels Produced by Hyperosmotic Agents and Measured by Optical Coherence Tomography,” Photochem. Photobiol., vol. 77, no. 5, 2003, pp. 541–549. 110. E. I. Galanzha, V. V. Tuchin, A. V. Solovieva, T. V. Stepanova, Q. Luo, and H. Cheng, “Skin Backreflectance and Microvascular System Functioning at the Action of Osmotic Agents,” J. Phys. D: Appl. Phys., vol. 36, 2003, pp. 1739–1746. 111. M. Brezinski, K. Saunders, C. Jesser, X. Li, and J. Fujimoto, “Index Matching to Improve OCT Imaging through Blood,” Circulation, vol. 103, 2001, pp. 1999–2003. 112. X. Xu, R. K. Wang, J. B. Elder, and V. V. Tuchin, “Effect of Dextran-Induced Changes in Refractive Index and Aggregation on Optical Properties of Whole Blood,” Phys. Med. Biol., vol. 48, 2003, pp. 1205–1221. 113. B. Grzegorzewski and E. Kowaliáska, “Optical Properties of Human Blood Sediment,” Acta Physica Polonica A, vol. 101, no. 1, 2002, pp. 201–209. 114. B. Grzegorzewski, E. Kowaliáska, A. Gãrnicki, and A. Gutsze, “Diffraction Measurement of Erythrocyte Sedimentation Rate,” Optica Applicata, vol. 32, no. 1, 2002, pp. 15–21.

References

225

115. A. K. Amerov, J. Chen, G. W. Small, and M. A. Arnold, “The Influence of Glucose upon the Transport of Light through Whole Blood,” Proc. SPIE 5330, 2004, pp. 101–111. 116. A. N. Bashkatov, E. A. Genina, Yu. P. Sinichkin, V. I. Kochubey, N. A. Lakodina, and V. V. Tuchin, “Estimation of the Glucose Diffusion Coefficient in Human Eye Sclera,” Biophys., vol. 48, no. 2, 2003, pp. 292–296. 117. A. N. Bashkatov, E. A. Genina, Yu. P. Sinichkin, V. I. Kochubey, N. A. Lakodina, and V. V. Tuchin, “Glucose and Mannitol Diffusion in Human Dura Mater,” Biophy. J., vol. 85, no. 5, 2003, pp. 3310–3318. 118. E. A. Genina, A. N. Bashkatov, Yu. P. Sinichkin, V. I. Kochubey, N. A. Lakodina, G. B. Altshuler, and V. V. Tuchin, “In Vitro and In Vivo Study of Dye Diffusion into the Human Skin and Hair Follicles,” J. Biomed. Opt., vol. 7, 2002, pp. 471–477. 119. V. V. Tuchin, E. A. Genina, A. N. Bashkatov, G. V. Simonenko, O. D. Odoevskaya, and G. B. Altshuler, “A Pilot Study of ICG Laser Therapy of Acne Vulgaris: Photodynamic and Photothermolysis Treatment,” Lasers Surg. Med., vol. 33, no. 5, 2003, pp. 296–310. 120. E. A. Genina, A. N. Bashkatov, G. V. Simonenko, O. D. Odoevskaya, V. V. Tuchin, and G. B. Altshuler, “Low-Intensity ICG-Laser Phototherapy of Acne Vulgaris: A Pilot Study,” J. Biomed. Opt., vol. 9, no. 4, 2004, pp. 828– 834. 121. Yu. P. Sinichkin, S. R. Utz, and H. A. Pilipenko, “In Vivo Human Skin Spectroscopy: I Remittance Spectra,” Opt. Spectrosc., vol. 80, no. 2, 1996, pp. 228–234. 122. Yu. P. Sinichkin, S. R. Utz, L. E. Dolotov, H. A. Pilipenko, and V. V. Tuchin, “Technique and Device for Evaluation of the Human Skin Erythema and Pigmentation,” Radioeng., no. 4, 1997, pp. 77–81. 123. L. E. Dolotov, Yu. P. Sinichkin, V. V. Tuchin, S. R. Utz, G. B. Altshuler, and I. V. Yaroslavsky, “Design and Evaluation of a Novel Portable ErythemaMelanin-Meter,” Lasers Surg. Med., vol. 34, no. 2, 2004, pp. 127–135. 124. B. C. Wilson, M. S. Patterson, and L. Lilge, “Implicit and Explicit Dosimetry in Photodynamic Therapy: A New Paradigm,” Lasers Med. Sci., vol. 12, 1997, pp. 182–199. 125. Y. Ito, R. P. Kennan, E. Watanabe, and H. Koizumi, “Assessment of Heating Effects in Skin During Continuous Wave Near Infrared Spectroscopy,” J. Biomed. Opt., vol. 5, 2000, pp. 383–390. 126. S.-J. Yeh, O. Khalil, Ch. F. Hanna, and S. Kantor, “Near-Infrared ThermoOptical Response of the Localized Reflectance of Intact Diabetic and Nondiabetic Human Skin,” J. Biomed. Opt., vol. 8, 2003, pp. 534–544. 127. A. T. Yeh, B. Kao, W. G. Jung, Z. Chen, J. S. Nelson, and B. J. Tromberg, “Imaging Wound Healing Using Optical Coherence Tomography and Multiphoton Microscopy in an In Vitro Skin-Equivalent Tissue Model,” J. Biomed. Opt., vol. 9, no. 2, 2004, pp. 248–253.

226

Optical Clearing of Tissues and Blood

128. T. Sakuma, T. Hasegawa, F. Tsutsui, and S. Kurihara, “Quantitative Analysis of the Whiteness of Atypical Cervical Transformation Zone,” J. Reprod. Med., vol. 30, 1985, pp. 773–776. 129. Ya. Holoubek, “Note on Light Attenuation by Scattering: Comparison of Coherent and Incoherent (Diffusion) Approximations,” Opt. Eng., vol. 37, 1998, pp. 705–709. 130. J. S. Balas, G. C. Themelis, E. P. Prokopakis, I. Orfanudaki, E. Koumantakis, and E. S. Helidonis, “In Vivo Detection and Staging of Epithelial Dysplasias and Malignancies Based on the Quantitative Assessment of Acetic Acid-Tissue,” J. Photochem. Photobiol., vol. 53, 1999, pp. 153–157. 131. A. Agrawal, U. Utzinger, C. Brookner, C. Pitris, M. F. Mitchell, and R. Richards-Kortum, “Fluorescence Spectroscopy of the Cervix: Influence of Acetic Acid, Cervical Mucus, and Vaginal Medications,” Lasers Surg. Med., vol. 25, 1999, pp. 237–249. 132. R. A. Drezek, T. Collier, C. K. Brookner, A. Malpica, R. Lotan, R. RichardsKortum, and M. Follen, “Laser Scanning Confocal Microscopy of Cervical Tissue before and after Application of Acetic Acid,” Am. J. Obstet. Gynecol., vol. 182, 2000, pp. 1135–1139. 133. B. W. Pogue, H. B. Kaufman, A. Zelenchuk, W. Harper, G. C. Burke, E. E. Burke, and D. M. Harper, “Analysis of Acetic Acid-Induced Whitening of High-Grade Squamous Intraepithelial Lesions,” J. Biomed. Opt., vol. 6, 2001, pp. 397–403. 134. B. Choi, T. E. Milner, J. Kim, J. N. Goodman, G. Vargas, G. Aguilar, and J. S. Nelson, “Use of Optical Coherence Tomography to Monitor Biological Tissue Freezing During Cryosurgery,” J. Biomed. Opt., vol. 9, 2003, pp. 282– 286. 135. G. N. Stamatas and N. Kollias, “Blood Stasis Contributions to the Perception of Skin Pigmentation,” J. Biomed. Opt., vol. 9, 2004, pp. 315–322. 136. M. Rajadhyaksha, S. Gonzalez, and J. M. Zavislan, “Detectability of Contrast Agents for Confocal Reflectance Imaging of Skin and Microcirculation,” J. Biomed. Opt., vol. 9, 2003, pp. 323–331. 137. J. K. Barton, N. J. Halas, J. L. West, and R. A. Drezek, “Nanoshells as an Optical Coherence Tomography Contrast Agent,” Proc. SPIE 5316, 2004, pp. 99–106. 138. R. K. Wang and V. V. Tuchin, “Enhance Light Penetration in Tissue for High Resolution Optical Techniques by the Use of Biocompatible Chemical Agents,” Proc. SPIE 4956, 2003, pp. 314–319. 139. R. K. Wang, Y. He, and V. V. Tuchin, “Effect of Dehydration on Optical Clearing and OCT Imaging Contrast after Impregnation of Biological Tissue with Biochemical Agents,” Proc. SPIE 5316, 2004, pp. 119–127. 140. M. Lazebnik, D. L. Marks, K. Potgieter, R. Gillette, and S. A. Boppart, “Functional Optical Coherence Tomography of Stimulated and Spontaneous Scattering Changes in Neural Tissue,” Proc. SPIE 5316, 2004, pp. 107–112.

References

227

141. X. Xu, R. Wang, and J. B. Elder, “Optical Clearing Effect on Gastric Tissues Immersed with Biocompatible Chemical Agents Investigated by Near Infrared Reflectance Spectroscopy,” J. Phys. D: Appl. Phys., vol. 36, 2003, pp. 1707–1713. 142. R. K. Wang and V. V. Tuchin, “Optical Coherence Tomography: Light Scattering and Imaging Enhancement,” Chapter 13 in Coherent-Domain Optical Methods: Biomedical Diagnostics, Environmental and Material Science, V. V. Tuchin (ed.), Kluwer Academic Publishers, Boston, vol. 2, 2004, pp. 3–60. 143. Q. Luo, H. Cheng, Z. Wang, and V. V. Tuchin, “Laser Speckle Imaging of Cerebral Blood Flow,” Chapter 5 in Coherent-Domain Optical Methods: Biomedical Diagnostics, Environmental and Material Science, V. V. Tuchin (ed.), Kluwer Academic Publishers, Boston, vol. 1, 2004, pp. 165–195. 144. Y. He and R. K. Wang, “Dynamic Optical Clearing Effect of Tissue Impregnated with Hyperosmotic Agents and Studied with Optical Coherence Tomography,” J. Biomed. Opt., vol. 9, 2004, pp. 200–206. 145. N. Guzelsu, J. F. Federici, H. C. Lim, H. R. Chauhdry, A. B. Ritter, and T. Findley, “Measurement of Skin Strech via Light Reflection,” J. Biomed. Opt., vol. 8, 2003, pp. 80–86. 146. A. F. Zuluaga, R. Drezek, T. Collier, R. Lotan, M. Follen, and R. RichardsKortum, “Contrast Agents for Confocal Microscopy: How Simple Chemicals Affect Confocal Images of Normal and Cancer Cells in Suspension,” J. Biomed. Opt., vol. 7, 2002, pp. 398–403. 147. H. Schneckenburger, A. Hendinger, R. Sailer, W. S. L. Strauss, and M. Schmitt, “Laser-Assisted Optoporation of Single Cells,” J. Biomed. Opt., vol. 7, 2002, pp. 410–416. 148. H. S. Dhadwal, R. R. Ansari, and M. A. DellaVecchia, “Coherent Fiber Optic Sensor for Early Detection of Cataractogenesis in the Human Eye Lens,” Opt. Eng., vol. 32, 1993, pp. 233–238. 149. B. Grzegorzewski and S. Yermolenko, “Speckle in Far–Field Produced by Fluctuations Associated with Phase Separation,” Proc. SPIE 2647, 1995, pp. 343–349. 150. G. C. Beck, N. Akgun, A. Rück, and R. Steiner, “Developing Optimized Tissue Phantom Systems for Optical Biopsies,” Proc. SPIE 3197, 1997, pp. 76– 85. 151. G. C. Beck, N. Akgun, A. Rück, and R. Steiner, “Design and Characterization of a Tissue Phantom System for Optical Diagnostics,” Lasers Med. Sci., vol. 13, 1998, pp. 160–171. 152. C.-L. Tsai and J. M. Fouke, “Noninvasive Detection of Water and Blood Content in Soft Tissue from the Optical Reflectance Spectrum,” Proc. SPIE 1888, 1993, pp. 479–486. 153. L. D. Shvartsman and I. Fine, “Optical Transmission of Blood: Effect of Erythrocyte Aggregation,” IEEE Trans. Biomed. Eng., vol. 50, 2003, pp. 1026– 1033.

228

Optical Clearing of Tissues and Blood

154. O. Cohen, I. Fine, E. Monashkin, and A. Karasik, “Glucose Correlation with Light Scattering Patterns—A Novel Method for Non-Invasive Glucose Measurements,” Diabetes Technol. Ther., vol. 5, 2003, pp. 11–17. 155. A. N. Yaroslavskaya, I. V. Yaroslavsky, C. Otto, G. J. Puppels, H. Duindam, G. F. J. M. Vrensen, J. Greve, and V. V. Tuchin, “Water Exchange in Human Eye Lens Monitored by Confocal Raman Microspectroscopy,” Biophys., vol. 43, no. 1, 1998 pp. 109–114. 156. O. Khalil, “Non-Invasive Glucose Measurement Technologies: An Update from 1999 to the Dawn of the New Millenium,” Diabetes Technol. Ther., vol. 6, no. 5, 2004, pp. 660–697. 157. R. O. Esenaliev, K. V. Larin, I. V. Larina, and M. Motamedi, “Noninvasive Monitoring of Glucose Concentration with Optical Coherent Tomography,” Opt. Lett., vol. 26, no. 13, 2001, pp. 992–994. 158. K. V. Larin, M. S. Eledrisi, M. Motamedi, and R. O. Esenaliev, “Noninvasive Blood Glucose Monitoring with Optical Coherence Tomography: A Pilot Study in Human Subjects,” Diabetes Care, vol. 25, no. 12, 2002, pp. 2263– 2267. 159. K. V. Larin, M. Motamedi, T. V. Ashitkov, and R. O. Esenaliev, “Specificity of Noninvasive Blood Glucose Sensing Using Optical Coherence Tomography Technique: A Pilot Study,” Phys. Med. Biol., vol. 48, 2003, pp. 1371–1390. 160. C. Dunsby and P. M. W. French, “Techniques for Depth-Resolved Imaging Through Turbid Media Including Coherence-Gated Imaging,” J. Phys. D: Appl. Phys., vol. 36, no. 14, 2003, R207–R227. 161. V. Tuchin, I. Maksimova, D. Zimnyakov, I. Kon, A. Mavlutov, and A. Mishin, “Light Propagation in Tissues with Controlled Optical Properties,” Proc. SPIE 2925, 1996, pp. 118–140. 162. A. N. Bashkatov, E. A. Genina, Yu. P. Sinichkin, N. A. Lakodina, V. I. Kochubey, and V. V. Tuchin “Estimation of Glucose Diffusion Coefficient in Scleral Tissue,” Proc. SPIE 4001, 2000, pp. 345–355. 163. E. A. Genina, A. N. Bashkatov, N. A. Lakodina, S. A. Murikhina, Yu. P. Sinichkin, and V. V. Tuchin, “Diffusion of Glucose Solution through Fibrous Tissues: In Vitro Optical and Weight Measurements,” Proc. SPIE 4001, 2000, pp. 255–261. 164. B. O. Hedbys and S. Mishima, “The Thickness-Hydration Relationship of the Cornea,” Exp. Eye Res., vol. 5, 1966, pp. 221–228. 165. A. I. Kholodnykh, K. Hosseini, I. Y. Petrova, R. O. Esenaliev, and M. Motamedi, “In vivo OCT Assessment of Rabbit Corneal Hydration and Dehydration,” Proc. SPIE 4956, 2003, pp. 295–298. 166. X. Xu and R. K. Wang, “The Role of Water Desorption on Optical Clearing of Biotissue: Studied with Near Infrared Reflectance Spectroscopy,” Med. Phys., vol. 30, 2003, pp. 1246–1253. 167. J. Jiang and R. K. Wang, “Comparing the Synergetic Effects of Oleic Acid and Dimethyl Sulfoxide as Vehicles for Optical Clearing of Skin Tissue In Vitro,” Phys. Med. Biol., vol. 49, 2004, pp. 5283–5294.

References

229

168. E. Lankenau, J. Welzel, R. Birngruber, and R. Engelhardt, “In Vivo Tissue Measurements with Optical Low Coherence Tomography,” Proc. SPIE 2981, 1997, pp. 78–84. 169. V. V. Tuchin, T. G. Anishchenko, A. A. Mishin, and O. V. Soboleva, “Control of Bovine Sclera Optical Characteristics with Various Osmolytes,” Proc. SPIE 2982, 1997, pp. 284–290. 170. I. L. Kon, V. V. Bakutkin, N. V. Bogomolova, S. V. Tuchin, D. A. Zimnyakov, and V. V. Tuchin, “Trazograph Influence on Osmotic Pressure and Tissue Structures of Human Sclera,” Proc. SPIE 2971, 1997, pp. 198–206. 171. D. A. Zimnyakov, V. V. Tuchin, and K. V. Larin, “Speckle Patterns Polarization Analysis as an Approach to Turbid Tissue Structure Monitoring,” Proc. SPIE 2981, 1997, pp. 172–180. 172. A. N. Bashkatov, E. A. Genina, Yu. P. Sinichkin, and V. V. Tuchin, “The Influence of Glycerol on the Transport of Light in the Skin,” Proc. SPIE 4623, 2002, pp. 144–152. 173. V. V. Tuchin, I. L. Maksimova, A. N. Bashkatov, Yu. P. Sinichkin, G. V. Simonenko, E. A. Genina, and N. A. Lakodina, “Eye Tissues Study—Scattering and Polarization Effects,” OSA TOPS, 1999, pp. 120–125. 174. V. V. Tuchin, A. N. Bashkatov, E. A. Genina, and Yu. P. Sinichkin, “Scleral Tissue Clearing Effects,” Proc. SPIE 4611, 2002, pp. 54–58. 175. A. V. Papaev, G. V. Simonenko, and V. V. Tuchin, “A Simple Model for Calculation of Polarized Light Transmission Spectrum of Biological Tissue Sample,” J. Opt. Technol., vol. 71, no. 5, 2004, pp. 3–6. 176. S. Yu. Shchyogolev, “Inverse Problems of Spectroturbidimetry of Biological Disperse Systems: An Overview,” J. Biomed. Opt., vol. 4, 1999, pp. 490–503. 177. V. V. Tuchin, X. Xu, and R. K. Wang, “Sedimentation of Immersed Blood Studied by OCT,” Proc. SPIE 4241, 2001, pp. 357–369. 178. V. V. Tuchin, X. Xu, R. K. Wang, and J. B. Elder, “Whole Blood and RBC Sedimentation and Aggregation Study using OCT,” Proc. SPIE 4263, 2001, pp. 143–149. 179. A. N. Bashkatov, E. A. Genina, I. V. Korovina, Yu. P. Sinichkin, O. V. Novikova, and V. V. Tuchin, “In Vivo and In Vitro Study of Control of Rat Skin Optical Properties by Action of 40%-Glucose Solution,” Proc. SPIE 4241, 2001, pp. 223–230. 180. A. N. Bashkatov, E. A. Genina, I. V. Korovina, V. I. Kochubey, Yu. P. Sinichkin, and V. V. Tuchin, “In Vivo and In Vitro Study of Control of Rat Skin Optical Properties by Acting of Osmotical Liquid,” Proc. SPIE 4224, 2000, pp. 300–311. 181. A. N. Bashkatov, E. A. Genina, V. I. Kochubey, N. A. Lakodina, and V. V. Tuchin, “Optical Clearing of Human Cranial Bones by Administration of Immersion Agents,” Biophotonics: From Fundamental Principles to Health, Environment, Security and Defense Applications, Ottawa, Ontario, Canada, Sept. 29–Oct. 9, 2004.

230

Optical Clearing of Tissues and Blood

182. E. V. Cruz, K. Kota, J. Huque, M. Iwaku, and E. Hoshino, “Penetration of Propylene Glycol into Dentine,” Int. Endodontic J., vol. 35, 2002, pp. 330– 336. 183. H. Y. Cheng, Q. M. Luo, S. Q. Zeng, S. B. Chen, J. Cen, and H. Gong, “A Modified Laser Speckle Imaging Method with Improved Spatial Resolution,” J. Biomed. Opt., vol. 8, no. 3, 2003, pp. 559–564. 184. A. N. Bashkatov, D. M. Zhestkov, E. A. Genina, and V. V. Tuchin, “Immersion Optical Clearing of Human Blood in the Visible and Near Infrared Spectral Range,” Opt. Spectrosc., vol. 98, no. 4, 2005, pp. 638–646. 185. D. M. Zhestkov, A. N. Bashkatov, E. A. Genina, and V. V. Tuchin, “Influence of Clearing Solutions Osmolarity on the Optical Properties of RBC,” Proc. SPIE 5474, 2004, pp. 321–330. 186. S. P. Chernova, N. V. Kuznetsova, A. B. Pravdin, and V. V. Tuchin, “Dynamics of Optical Clearing of Human Skin In Vivo,” Proc. SPIE 4162, 2000, pp. 227–235. 187. P. L. Walling and J. M. Dabney, “Moisture in Skin by Near-Infrared Reflectance Spectroscopy,” J. Soc. Cosmet. Chem., vol. 40, 1989, pp. 151–171. 188. K. A. Martin, “Direct Measurement of Moisture in Skin by NIR Spectroscopy,” J. Soc. Cosmet. Chem., vol. 44, 1993, pp. 249–261. 189. K. Wichrowski, G. Sore, and A. Khaiat, “Use of Infrared Spectroscopy for In Vivo Measurement of the Stratum Corneum Moisturization after Application of Cosmetic Preparations,” Int. J. Cosmet. Sci., vol. 17, 1995, pp. 1–11. 190. J. M. Schmitt, J. Hua, and J. Qu, “Imaging Water Absorption with OCT,” Proc. SPIE 3598, 1999, pp. 36–46. 191. T. Yasui, Y. Tohno, and T. Araki, “Characterization of Collagen Orientation in Human Dermis by Two-Dimensional Second-Harmonic-Generation Polarimetry,” J. Biomed. Opt., vol. 9, no. 2, 2004, pp. 259–264. 192. V. V. Tuchin, “Tissue and Blood Optical Properties Control,” in Biophotonics—Principles and Applications, NATO Advanced Study Institute, September 29–October 9, 2004, Ottawa, Canada, 2005, 49 pp. Advances in Biophotonics, Brian W. Wilson, Valery V. Tuchin, and Stoyan Tanev (eds.), IOS Press, Amsterdam, 2005. 193. K. F. Kolmel, B. Sennhenn, and K. Giese, “Investigation of Skin by Ultraviolet Remittance Spectroscopy,” British J. Dermatol., vol. 122, no. 2, 1990, pp. 209–216. 194. H.-J. Schwarzmaier, M. P. Heintzen, and W. Müller, et al., “Optical Density of Vascular Tissue before and after 308-nm Excimer Laser Irradiation,” Opt. Eng., vol. 31, 1992, pp. 1436–1440. 195. R. Splinter, R. H. Svenson, and L. Littman, et al., “Computer Simulated Light Distributions in Myocardial Tissues at the Nd-YAG Wavelength of 1064 nm,” Lasers Med. Sci., vol. 8, 1993, pp. 15–21. 196. V. V. Tuchin, D. Tabatadze, J. Childs, I. Yaroslavsky, A. B. Pravdin, A. A. Gavrilova, and G. Altshuler, “Optical Clearing of Skin Using Flashlamp-Induced Permeability Enhancement for Accelerated Delivery of

References

197. 198.

199. 200. 201.

202.

203.

204. 205.

206. 207.

208.

209.

210. 211.

231

Index-Matching Agents,” 25th Annual Meeting of the American Society for Laser Medicine and Surgery, March 30–April 3, 2005, Lake Buena Vista, Florida, Abstracts, Lasers Surg. Med., Suppl. 17, March 2005, p. 5. V. V. Tuchin, “Optical Immersion as a New Tool to Control Optical Properties of Tissues and Blood,” Laser Phys., vol. 15, 2005, pp. 1109–1136. P. Rol, P. Niederer, U. Dürr, P.-D. Henchoz, and F. Fankhauser, “Experimental Investigation on the Light Scattering Properties of the Human Sclera,” Laser Light Ophthalmol., vol. 3, 1990, pp. 201–212. J. M. Schmitt and G. Kumar, “Optical Scattering Properties of Soft Tissue: A Discrete Particle Model,” Appl. Opt., vol. 37, no. 13, 1998, pp. 2788–2797. I. L. Maksimova, V. V. Tuchin, and L. P. Shubochkin, “Polarization Features of Eye’s Cornea,” Opt. Spectrosc. (USSR), vol. 60, no. 4, 1986, pp. 801–807. I. L. Maksimova, V. V. Tuchin, and L. P. Shubochkin, “Light Scattering Matrix of Crystalline Lens,” Opt. Spectrosc. (USSR), vol. 65, no. 3, 1988, pp. 615–619. D. E. Freund, R. L. McCally, and R. A. Farrell, “Effects of Fibril Orientations on Light Scattering in the Cornea,” J. Opt. Soc. Am. A., vol. 3, 1986, pp. 1970–1982. R. A. Farrell, D. E. Freund, and R. L. McCally, “Hierarchical Structure and Light Scattering in the Cornea,” Mat. Res. Soc. Symp. Proc., vol. 255, 1992, pp. 233–246. R. A. Farrell, D. E. Freund, and R. L. McCally, “Research on Corneal Structure,” Johns Hopkins APL Technol. Digest., vol. 11, 1990, pp. 191–199. J. R. Mourant, I. J. Bigio, and J. Boyer, et al., “Elastic Scattering Spectroscopy as a Diagnostic Tool for Differentiating Pathologies in the Gastrointestinal Tract: Preliminary Testing,” J. Biomed. Opt., vol.1, no. 2, 1996, pp. 192–199. F. A. Duck, Physical Properties of Tissue: a Comprehensive Reference Book, Academic Press, London, 1990. J. R. Mourant, J. Boyer, A. H. Hielscher, and I. J. Bigio, “Influence of the Scattering Phase Function on Light Transport Measurements in Turbid Media Performed with Small Source-Detector Separations,” Opt. Lett., vol. 21, no. 7, 1996, pp. 546–548. J. R. Mourant, M. Canpolat, C. Brocker, O. Esponda-Ramos, T. M. Johnson, A. Matanock, K. Stetter, and J. P. Freyer, “Light Scattering from Cell: the Contribution of the Nucleus and the Effects of Proliferative Status,” J. Biomed. Opt., vol. 5, no. 2, 2000, pp. 131–137. K. Sokolov, R. Drezek, K. Gossagee, and R. Richards-Kortum, “Reflectance Spectroscopy with Polarized Light: is it Sensitive to Cellular and Nuclear Morphology,” Opt. Express, vol. 5, 1999, pp. 302–317. A. Brunsting and P. F. Mullaney, “Differential Light Scattering from Spherical Mammalian Cells,” Biophys. J., vol. 10, 1974, 439–453. J. Beuthan, O. Minet, J. Helfmann, M. Herring, and G. Mueller, “The Spatial Variation of the Refractive Index in Biological Cells,” Phys. Med. Biol., vol. 41, no. 3, 1996, pp. 369–382.

232

Optical Clearing of Tissues and Blood

212. K. W. Keohane and W. K. Metcalf, “The Cytoplasmic Refractive Index of Lymphocytes, Its Significance and Its Changes during Active Immunization,” Quart. J. Exper. Physiol. Cognate Med. Sci., vol. 44, 1959, pp. 343–346. 213. Tuan Vo-Dinh (ed.), Biomedical Photonics Handbook, CRC Press, Boca Raton, 2003. 214. F. P. Bolin, L. E. Preuss, R. C. Taylor, and R. J. Ference, “Refractive Index of Some Mammalian Tissues using a Fiber Optic Cladding Method,” Appl. Opt., vol. 28, 1989, pp. 2297–2303. 215. R. Graaff, J. G. Aarnoudse, J. R. Zijp, P. M. A. Sloot, F. F. M. de Mul, J. Greve, and M. H. Koelink, “Reduced Light Scattering Properties for Mixtures of Spherical Particles: A Simple Approximation Derived from Mie Calculations,” Appl. Opt., vol. 31, 1992, pp. 1370–1376. 216. L. T. Perelman, V. Backman, M. Wallace, G. Zonios, R. Manoharan, A. Nusrat, S. Shields, M. Seiler, C. Lima, T. Hamano, I. Itzkan, J. Van Dam, J. M. Crawford, and M. S. Feld, “Observation of Periodic Fine Structure in Reflectance from Biological Tissue: A New Technique for Measuring Nuclear Size Distribution,” Phys. Rev. Lett., vol. 80, 1998, pp. 627–630. 217. H. C. van de Hulst, Light Scattering by Small Particles, Wiley, New York, 1957 [reprint, Dover, New York, 1981]. 218. F. A. Marks, “Optical Determination of the Hemoglobin Oxygenation State of Breast Biopsies and Human Breast Cancer Xenografts in Unde Mice,” Proc. SPIE 1641, 1992, pp. 227–237. 219. C. F. Bohren and D. R. Huffman, Absorption and Scattering of Light by Small Particles, Wiley, New York, 1983. 220. M. I. Mishchenko, J. W. Hovenier, and L. D. Travis (eds.), Light Scattering by Nonspherical Particles, Academic Press, San Diego, 2000. 221. M. I. Mishchenko, L. D. Travis, and A. A. Lacis, Scattering, Absorption, and Emission of Light by Small Particles, Cambridge Univ. Press, Cambridge, 2002. 222. K. S. Shifrin, Physical Optics of Ocean Water, American Institute of Physics, New York, 1988. 223. J. M. Schmitt and G. Kumar, “Turbulent Nature of Refractive-Index Variations in Biological Tissue,” Opt. Lett., vol. 21, 1996, pp. 1310–1312. 224. G. Kumar and J. M. Schmitt, “Micro-Optical Properties of Tissue,” Proc. SPIE 2679, 1996, pp. 106–116. 225. J. R. Mourant, T. M. Johnson, S. Carpenter, A. Guerra, T. Aida, and J. P. Freyer, “Polarized Angular Dependent Spectroscopy of Epithelial Cells and Epithelial Cell Nuclei to Determine the Size Scale of Scattering Structures,” J. Biomed. Opt., vol. 7, no. 3, 2002, pp. 378–387. 226. R. K. Wang, “Modeling Optical Properties of Soft Tissue by Fractal Distribution of Scatterers,” J. Modern Opt., vol. 47, no. 1, 2000, pp. 103–120. 227. A. Kienle, L. Lilge, M. S. Patterson, R. Hibst, R. Steiner, and B. C. Wilson, “Spatially Resolved Absolute Diffuse Reflectance Measurements for Noninvasive Determination of the Optical Scattering and Absorption Coefficients of Biological Tissue,” Appl. Opt., vol. 35, no. 13, 1996, pp. 2304–2314.

References

233

228. A. Kienle, M. S. Patterson, N. Dognitz, R. Bays, G. Wagnieres, and H. van de Bergh, “Noninvasive Determination of the Optical Properties of Two-Layered Turbid Medium,” Appl. Opt., vol. 37, 1998, pp. 779–791. 229. R. C. Haskell, L. O. Svaasand, T.-T. Tsay, T. C. Feng, M. N. McAdams, and B. J. Tromberg, “Boundary Conditions for the Diffusion Equation in Radiative Transfer,” J. Opt. Soc. Am. A, vol. 11, no. 10, 1994, pp. 2727–2741. 230. A. Kienle and M. S. Patterson, “Improved Solutions of the Steady-State and Time-Resolved Diffusion Equations for Reflectance from a Semi-Infinite Turbid Media,” J. Opt. Soc. Am. A, vol. 14, 1997, pp. 246–254. 231. T. J. Farrell and M. S. Patterson, “Experimental Verification of the Effect of Refractive Index Mismatch on the Light Fluence in a Turbid Medium,” J. Biomed. Opt., vol. 6, no. 4, 2001, pp. 468–473. 232. Y. Kamai and T. Ushiki, “The Three-Dimensional Organization of Collagen Fibrils in the Human Cornea and Sclera,” Invest. Ophthalmol. Visual Sci., vol. 32, 1991, pp. 2244–2258. 233. A. Kotyk and K. Janacek, Membrane Transport: An Interdisciplinary Approach, Plenum Press, New York, 1977. 234. A. Pirie and van R. Heyningen, Biochemistry of the Eye, Blackwell Scientific Publications, Oxford, 1966. 235. Y. Huang and K. M. Meek, “Swelling Studies on the Cornea and Sclera: The Effect of pH and Ionic Strength,” Biophys. J., vol. 77, 1999, pp. 1655–1665. 236. Y. L. Kim, J. T. Walsh Jr., T. K. Goldstick, and M. R. Glucksberg, “Variation of Corneal Refractive Index with Hydration,” Phys. Med. Biol., vol. 49, 2004, pp. 859–868. 237. V. V. Tuchin, “Optics of the Human Sclera: Photon Migration, Imaging and Spectroscopy,” OSA TOPS, vol. 21, 1998, pp. 99–104. 238. R. W. Hart and R. A. Farrell, “Light Scattering in the Cornea,” J. Opt. Soc. Am., vol. 59, no. 6, 1969, pp. 766–774. 239. S. A. Prahl, “The Inverse Adding-Doubling Program,” htttp://omlc.ogi.edu/ software/iad/index.html 240. I. S. Grigor’eva and E. Z. Meilikhova (eds.), Handbook of Physical Constants, Energoatomizdat, Moscow, 1991. 241. I. K. Kikoin (ed.), Handbook of Physical Constants, Atomizdat, Moscow, 1976. 242. A. J. Gordon and R. D. Ford, The Chemist’s Companion: A Handbook of Practical Data, Techniques, and References, Wiley, New York, 1972. 243. I. L. Knunyants (ed.), Big Encyclopedic Dictionary: Chemistry, Big Russian Encyclopedia Publishers, Moscow, 1998. 244. http://www.dmso.org 245. A. I. Korenkov, J. Pahnke, K. Frei, R. Warzok, H. W. S. Schroeder, R. Frick, L. Muljana, J. Piek, Y. Yonekawa, and M. R. Gaab, “Treatment with Nimodipine or Mannitol Reduces Programmed Cell Death and Infarct Size Following Focal Cerebral Ischemia,” Neurosurg. Rev., vol. 23, 2000, pp. 145– 150.

234

Optical Clearing of Tissues and Blood

246. P. L. Gourley, “Biocavity Laser for High-Speed Cell and Tumor Biology,” J. Phys. D: Appl. Phys., vol. 36, 2003, pp. R228–R239. 247. H. Schaefer and T. E. Redelmeier, Skin Barier: Principles of Percutaneous Absorption, Karger, Basel, 1996. 248. F. Pirot, Y. N. Kalia, A. L. Stinchcomb, G. Keating, A. Bunge, and R. H. Guy, “Characterization of the Permeable Barrier of Human Skin In Vivo,” Proc. Natl. Acad. Sci., vol. 94, 1997, pp. 1562–1567. 249. I. H. Blank, J. Moloney, and A. G. Emslie, et al., “The Diffusion of Water Across the Stratum Corneum as a Function of its Water Content,” J. Invest. Dermatol., vol. 82, 1984, pp. 188–194. 250. T. von Zglinicki, M. Lindberg, G. H. Roomans, and B. Forslind, “Water and Ion Distribution Profiles in Human Skin,” Acta Derm. Venerol. (Stockh), vol. 73, 1993, pp. 340–343. 251. V. V. Tuchin, S. R. Utz, and I. V. Yaroslavsky, “Tissue Optics, Light Distribution, and Spectroscopy,” Opt. Eng., vol. 33, 1994, pp. 3178–3188. 252. A. N. Yaroslavsky, S. R. Utz, S. N. Tatarintsev, and V. V. Tuchin, “Angular Scattering Properties of Human Epidermal Layers,” Proc. SPIE 2100, 1994, pp. 38–41. 253. G. W. Lucassen, G. N. A. van Veen, and J. A. J. Jansen, “Band Analysis of Hydrated Human Skin Stratum Corneum Attenuated Total Reflectance Fourier Transform Infrared Spectra In Vivo,” J. Biomed. Opt., vol. 3, 1998, pp. 267–280. 254. B. E. Bouma and G. J. Tearney (eds.), Handbook of Optical Coherence Tomography, Marcel-Dekker, New York, 2002. 255. A. G. Ushenko and V. P. Pishak, “Laser Polarimetry of Biological Tissues: Principles and Applications” in Coherent-Domain Optical Methods: Biomedical Diagnostics, Environmental and Material Science, vol. 1, V. V. Tuchin (ed.), Kluwer Academic Publishers, Boston, 2004, pp. 94–138. 256. L. V. Wang, G. L. Coté, and S. L. Jacques (eds.), “Special Section on Tissue Polarimetry,” J. Biomed. Opt., vol. 7, no. 3, 2002, pp. 278–397. 257. V. Shankaran, J. T. Walsh, Jr., and D. J. Maitland, “Comparative Study of Polarized Light Propagation in Biological Tissues,” J. Biomed. Opt., vol. 7, no. 3, 2002, pp. 300–306. 258. S. Jacques, “Strengths and Weaknesses of Various Optical Imaging Techniques,” Saratov Fall Meeting’01, Internet Plenary Lecture, Saratov, Russia, 2001, http://optics.sgu.ru/SFM. 259. S. L. Jacques, R. J. Roman, and K. Lee, “Imaging Superficial Tissues with Polarized Light,” Lasers Surg. Med., vol. 26, 2000, pp. 119–129. 260. L. Jacques, J. C. Ramella-Roman, and K. Lee, “Imaging Skin Pathology with Polarized Light,” J. Biomed. Opt., vol. 7, no. 3, 2002, pp. 329–340. 261. D. A. Zimnyakov and Yu. P. Sinichkin, “A study of Polarization Decay as Applied to Improved Imaging in Scattering Media,” J. Opt. A: Pure Appl. Opt., vol. 2, 2000, pp. 200–208.

References

235

262. R. G. Kessel, Basic Medical Histology: The Biology of Cells, Tissues, and Organs, Oxford Univ. Press, New York, 1998. 263. W. Denk, “Two-Photon Excitation in Functional Biological Imaging,” J. Biomed. Opt., vol. 1, no. 3, 1996, pp. 296–304. 264. P. T. C. So, C. Y. Dong, and B. R. Masters, “Two-Photon Excitation Fluorescence Microscopy,” in Biomedical Photonics Handbook, T. Vo-Dinh (ed.), CRC Press, Boca Rotan, 2003, pp. 11-1–11-17. 265. B. R. Masters and T. P. C. So, Handbook of Multiphoton Excitation Microscopy and Other Nonlinear Microscopies, Oxford University Press, New York, 2004. 266. A. K. Dunn, V. P. Wallace, M. Coleno, M. W. Berns, and B. J. Tromberg, “Influence of Optical Properties on Two-Photon Fluorescence Imaging in Turbid Samples,” Appl. Opt., vol. 39, no. 7, 2000, pp. 1194–1201. 267. M. Gu, X. Gan, A. Kisteman, and M. G. Xu, “Comparison of Penetration Depth Between Two-Photon Excitation and Single-Photon Excitation in Imaging Through Turbid Tissue Media,” Appl. Phys. Lett., vol. 77, 2000, pp. 1551–1553. 268. E. Beaurepaire, M. Oheim, and J. Mertz, “Ultra-Deep Two-Photon Fluorescence Excitation in Turbid Media,” Opt. Commun., vol. 188, 2001, pp. 25–29. 269. R. Cicchi, F. S. Pavone, D. Massi, and D. D. Sampson, “Contrast and Depth Enhancement in Two-Photon Microscopy of Human Skin Ex Vivo by Use of Optical Clearing Agents,” Opt. Express, vol. 13, 2005, pp. 2337–2344. 270. B. R. Masters and P. T. C. So, “Confocal Microscopy and Multi-Photon Excitation Microscopy of Human Skin In Vivo,” Opt. Express, vol. 8, 2001, pp. 2–10. 271. K. König and I. Riemann, “High-Resolution Multiphoton Tomography of Human Skin with Subcellular Spatial Resolution and Picosecond Time Resolution,” J. Biomed. Opt., vol. 8, 2003, pp. 432–439. 272. P. Stoller, K. M. Reiser, P. M. Celliers, and A. M. Rubenchik, “PolarizationModulated Second-Harmonic Generation in Collagen,” Biophys. J., vol. 82, no. 2, 2002, pp. 3330–3342. 273. M. Han, L. Zickler, G. Giese, M. Walter, F. H. Loesel, and J. F. Bille, “Second-Harmonic Imaging of Cornea after Intrastromal Femtosecond Laser Ablation,” J. Biomed. Opt., vol. 9, no. 4, 2004, pp. 760–766. 274. B. Choi, J. Lotfi, J. C. Ramirez, and J. S. Nelson, “Laser Speckle Imaging of Blood Flow Velocity Dynamics During Optical Clearing of Rodent,” 25th Annual Meeting of the American Society for Laser Medicine and Surgery, March 30–April 3, 2005, Lake Buena Vista, Florida, Abstracts, Lasers Surg. Med., Suppl. 17, March 2005, p. 1. 275. A. V. Priezzhev, O. M. Ryaboshapka, N. N. Firsov, and I. V. Sirko, “Aggregation and Disaggregation of Erythrocytes in Whole Blood: Study by Backscattering Technique,” J. Biomed. Opt., vol. 4, no. 1, 1999, pp. 76–84. 276. A. H. Gandjbakhche, P. Mills and P. Snabre, “Light-Scattering Technique for the Study of Orientation and Deformation of Red Blood Cells in a Concentrated Suspension,” Appl. Opt., vol. 33, 1994, pp. 1070–1078.

236

Optical Clearing of Tissues and Blood

277. A. V. Priezzhev, N. N. Firsov, and J. Lademann, “Light Backscattering Diagnostics of Red Blood Cells Aggregation in Whole Blood Samples,” Chapter 11 in Handbook of Optical Biomedical Diagnostics, vol. PM107, V. V. Tuchin (ed.), SPIE Press, Bellingham, WA, 2002, pp. 651–674. 278. A. Podoleanu, “En-Face OCT Imaging,” Ch. 16 in Coherent-Domain Optical Methods: Biomedical Diagnostics, Environmental and Material Science, vol. 2, V. V. Tuchin (ed.), Kluwer Academic Publishers, Boston, 2004, pp. 163–209. 279. A. F. Fercher, W. Drexler, C. K. Hitzenberger, and T. Lasser, “Optical Coherence Tomography—Principles and Applications,” Rep. Progr. Phys., vol. 66, 2003, pp. 239–303. 280. U. Utzinger and R. Richards-Kortum, “Fiber-Optic Probes for Biomedical Optical Spectroscopy,” J. Biomed. Opt., vol. 8, 2003, pp. 121–147. 281. R. R. Alfano and B. R. Masters (eds.), Biomedical Optical Biopsy and Optical Imaging: Classic Reprints on CD-ROM Series, Optical Society of America, Washington, DC, 2004. 282. W. R. Chen, V. V. Tuchin, Q. Luo, and S. L. Jacques, “Special Issue on Biophotonics,” J. X-Ray Sci. and Technol., vol. 10, nos. 3-4, 2002, pp. 139–243. 283. P. French and A. I. Ferguson (eds.), “Special Issue on Biophotonics,” J. Phys. D: Appl. Phys., vol. 36, no. 14, 2003, R207–R258; 1655–1757. 284. E. D. Hanlon, R. Manoharan, T.-W. Koo, K. E. Shafer, J. T. Motz, M. Fitzmaurice, J. R. Kramer, I. Itzkan, R. R. Dasari, and M. S. Feld, “Prospects for In Vivo Raman Spectroscopy,” Phys. Med. Biol., vol. 45, 2000, pp. R1–R59. 285. R. R. Ansari and J. Sebag (eds.), “Ophthalmic Diagnostics,” J. Biomed. Opt., vol. 9, no. 1, 2004, pp. 8–179. 286. A. N. Yaroslavsky, A. V. Priezzhev, J. Rodriguez, I. V. Yaroslavsky, and H. Battarbee, “Optics of Blood,” in Handbook of Optical Biomedical Diagnostics, Chapter 2, vol. PM107, V. V. Tuchin (ed.), SPIE Press, Bellingham, WA, 2002, pp. 169–216. 287. S. M. Bertoluzzo, A. Bollini, M. Rsia, and A. Raynal, “Kinetic Model for Erythrocyte Aggregation,” Blood Cells, Molecules, Diseases, vol. 25, no. 22, 1999, pp. 339–349. 288. S. Chien, “Physiological and Pathophysiological Significance of Hemorheology,“ in Clinical Hemorheology, S. Chien, J. Dormandy, E. Ernst, and A. Matarai (eds.), Martinus Nijhoff, Dordrecht, 1987, pp. 125–132. 289. D. H. Tycko, M. H. Metz, E. A. Epstein, and A. Grinbaum, “Flow-Cytometric Light Scattering Measurement of Red Blood Cell Volume and Hemoglobin Concentration,” Appl. Opt., vol. 24, 1985, pp. 1355–1365. 290. M. Hammer, D. Schweitzer, B. Michel, E. Thamm, and A. Kolb, “Single Scattering by Red Blood Cells,” Appl. Opt., vol. 37, no. 31, 1998, pp. 7410– 7418. 291. M. Yu. Kirillin and A. V. Priezzhev, “Monte Carlo Simulation of Laser Beam Propagation in a Plane Layer of the Erythrocytes Suspension: Comparison of Contributions from Different Scattering Orders to the Angular Distribution of Light Intensity,” Quant. Electron., vol. 32, no. 10, 2002, pp. 883–887.

References

237

292. A. Ishimaru, Wave Propagation and Scattering in Random Media, IEEE Press, New York, 1997. 293. S. F. Shumilina, “Dispersion of Real and Imaginary Part of the Complex Refractive Index of Hemoglobin in the Range 450 to 820 nm,” Bullet. Belarus. USSR Acad. Sci., Phys.-Math. Ser., no. 1, 1984, pp. 79–84. 294. A. G. Borovoi, E. I. Naats, and U. G. Oppel, “Scattering of Light by a Red Blood Cell,” J. Biomed. Opt., vol. 3, no. 3, 1998, pp. 364–372. 295. N. G. Khlebtsov and S. Yu. Shchyogolev, “Account for Particle Nonsphericity at Determination of Parameters of Disperse Systems by a Turbidity Spectrum Method, 1: Characteristic Functions of Light Scattering by Nonspherical Particle Systems in Rayleigh-Gans Approximation,” Opt. Spectrosc., vol. 42, 1977, pp. 956–962. 296. V. Twersky, “Absorption and Multiple Scattering by Biological Suspensions,” J. Opt. Soc. Am., vol. 60, 1970, pp. 1084–1093. 297. J. M. Steinke and A. P. Shephard, “Diffusion Model of the Optical Absorbance of Whole Blood,” J. Opt. Soc. Am. A, vol. 5, 1988, pp. 813–822. 298. A. H. Hielsher, J. R. Mourant, and I. J. Bigio, “Influence of Particle Size and Concentration on the Diffuse Backscattering of Polarized Light from Tissue Phantoms and Biological Cell Suspensions,” Appl. Opt., vol. 36, 1997, pp. 125–135. 299. L. O. Svaasand and Ch. J. Gomer, “Optics of Tissue,” in Dosimetry of Laser Radiation in Medicine and Biology, SPIE Press, Bellingham, vol. IS5, 1989, pp. 114–132. 300. A. Ya. Khairullina and S. F. Shumilina, “Determination of Size Distribution Function of the Erythrocytes According to Size by the Spectral Transparency Method,” J. Appl. Spectrosc., vol. 19, 1973, pp. 1078–1083. 301. J. Beuthan, O. Minet, M. Herring, G. Mueller, and C. Dressler, “Biological Cells as Optical Phase-Filters—A Contribution to Medical Functional Imaging,” Minimal Invasive Medizin, vol. 5, no. 2, 1994, pp. 75–78. 302. S. Tanev, W. Sun, N. Loeb, and V. Tuchin, “The Finite-Difference TimeDomain Approach and its Application to the Modeling of Light Scattering by Biological Cells in Absorbing and Controlled Extra-cellular Media,” in Biophotonics—Principles and Applications, NATO Advanced Study Institute, Sept. 29–Oct. 9, 2004, Ottawa, Canada, 2005, 34 pp. Advances in Biophotonics, Brian W. Wilson, Valery V. Tuchin, and Stoyan Tanev (eds.), IOS Press, Amsterdam, 2005. 303. A. Roggan, K. Dörschel, O. Minet, D. Wolff, and G. Müller, “The Optical Properties of Biological Tissue in the Near Infrared Wavelength Range— Review and Measurements,” in Laser-Induced Interstitial Thermotherapy, vol. PM25, G. Müller and A. Roggan (eds.), SPIE Press, Bellingham, WA, 1995, pp. 10–44. 304. B. B. Das, F. Liu, and R. R. Alfano, “Time-Resolved Fluorescence and Photon Migration Studies in Biomedical and Random Media,” Rep. Prog. Phys., vol. 60, 1997, pp. 227–292.

238

Optical Clearing of Tissues and Blood

305. S. Cheng, H. Y. Shen, G. Zhang, C. H. Huang, X. J. Huang, “Measurement of the Refractive Index of Biotissue at Four Laser Wavelengths,” Proc. SPIE 4916, 2002, pp. 172–176. 306. H. Liu and S. Xie, “Measurement Method of the Refractive Index of Biotissue by Total Internal Reflection,” Appl. Opt., vol. 35, 1996, pp. 1793–1795. 307. D. Arifler, M. Guillaud, A. Carraro, A. Malpica, M. Follen, R. RichardsKortum, “Light Scattering from Normal and Dysplastic Cervical Cells at Different Epithelial Depths: Finite-Difference Time-Domain Modeling with a Perfectly Matched Layer Boundary Condition,” J. Biomed. Opt., vol. 8, 2003, pp. 484–494. 308. G. J. Tearney, M. E. Brezinski, J. F. Southern, B. E. Bouma, M. R. Hee, and J. G. Fujimoto, “Determination of the Refractive Index of Highly Scattering Human Tissue by Optical Coherence Tomography,” Opt. Lett., vol. 20, 1995, pp. 2258–2260. 309. A. Knüttel and M. Boehlau-Godau, “Spatially Confined and Temporally Resolved Refractive Index and Scattering Evaluation in Human Skin Performed with Optical Coherence Tomography,” J. Biomed. Opt., vol. 5, 2000, pp. 83– 92. 310. S. A. Alexandrov, A. V. Zvyagin, K. K. M. B. D. Silva, and D. D. Sampson, “Bifocal Optical Coherence Refractometry of Turbid Media,” Opt. Lett., vol. 28, 2003, pp. 117–119. 311. A. V. Zvyagin, K. K. M. B. D. Silva, S. A. Alexandrov, T. R. Hillman, J. J. Armstrong, T. Tsuzuki, and D. D. Sampson, “Refractive Index Tomography of Turbid Media by Bifocal Optical Coherence Refractometry,” Optics Express, vol. 11, 2003, pp. 3503–3517. 312. A. Knüttel, S. Bonev, and W. Knaak, “New Method for Evaluation of In Vivo Scattering and Refractive Index Properties Obtained with Optical Coherence Tomography,” J. Biomed. Opt., vol. 9, 2004, pp. 265–273. 313. D. W. Leonard and K. M. Meek, “Refractive Indices of the Collagen Fibrils and Extrafibrillar Material of the Corneal Stroma,” Biophys. J., vol. 72, 1997, pp. 1382–1387. 314. K. M. Meek, S. Dennis, and S. Khan, “Changes in the Refractive Index of the Stroma and its Extrafibrillar Matrix When the Cornea Swells,” Biophys. J., vol. 85, 2003, pp. 2205–2212. 315. R. A. Farrell and R. L. McCally, “Corneal Transparency” in Principles and Practice of Ophthalmology, D. A. Albert and F. A. Jakobiec (eds.), W. B. Saunders, Philadelphia, 2000, pp. 629–643. 316. Y. L. Kim, J. T. Walsh Jr., T. K. Goldstick, and M. R. Glucksberg, “Variation of Corneal Refractive Index with Hydration,” Phys. Med. Biol., vol. 49, 2004, pp. 859–868. 317. R. C. Lin, M. A. Shure, A. M. Rollins, J. A. Izatt, and D. Huang, “Group Index of the Human Cornea at 1.3-µm Wavelength Obtained In Vitro by Optical Coherence Domain Reflectometry,” Opt. Lett., vol. 29, 2004, pp. 83–85.

References

239

318. W. Drexler, C. K. Hitzenberger, A. Baumgartner, O. Findl, H. Sattmann, and A. F. Fercher, “Investigation of Dispersion Effects in Ocular Media by Multiple Wavelength Partial Coherence Interferometry,” Exp. Eye Res., vol. 66, 1998, pp. 25–33. 319. V. N. Grisimov, “Refractive Index of the Ground Material of Dentin,” Opt. Spectrosc., vol. 77, 1994, pp. 272–273. 320. X. J. Wang, T. E. Milner, J. F. de Boer, Y. Zhang, D. H. Pashley, and J. S. Nelson, “Characterization of Dentin and Enamel by Use of Optical Coherence Tomography,” Appl. Opt., vol. 38, 1999, pp. 2092–2096. 321. M. Ohmi, Y. Ohnishi, K. Yoden, and M. Haruna, “In Vitro Simultaneous Measurement of Refractive Index and Thickness of Biological Tissue by the Low Coherence Interferometry,” IEEE Trans. Biomed. Eng., vol. 47, 2000, pp. 1266–1270. 322. X. J. Wang, T. E. Milner, R. P. Dhond, W. V. Sorin, S. A. Newton, and J. S. Nelson, “Characterization of Human Scalp Hairs by Optical LowCoherence Reflectometry,” Opt. Lett., vol. 20, 1995, pp. 524–526. 323. D. K. Sardar and L. B. Levy, “Optical Properties of Whole Blood,” Lasers Med. Sci., vol. 13, 1998, pp. 106–111. 324. X. Wang, T. E. Milner, M. C. Chang, J. S. Nelson, “Group Refractive Index Measurement of Dry and Hydrated Type I Collagen Films Using Optical Low-Coherence Reflectometry,” J. Biomed. Opt., vol. 1, no. 2, 1996, pp. 212– 216. 325. B. Choi, L. Tsu, E. Chen, T. S. Ishak, S. M. Iskandar, S. Chess, and J. S. Nelson, “Determination of Chemical Agent Optical Clearing Potential Using In Vitro Human Skin,” Lasers Surg. Med., vol. 36, 2005, pp. 72–75. 326. P. Michailova, Medical Cosmetics, Medicine, Moscow, 1984. 327. M. Kirjavainen, A. Urtti, and I. Jaaskelainen, et al., “Interaction of Liposomes with Human Skin In Vitro—The Influence of Lipid Composition and Structure,” Biochem. Biophys. Acta, vol. 1304, 1996, pp. 179–189. 328. S. S. Olmsted, J. L. Padgett, A. I. Yudin, K. J. Whaley, T. R. Moench, and R. A. Cone, “Diffusion of Macromolecules and Virus-Like Particles in Human Cervical Mucus,” Biophys. J., vol. 81, 2001, pp. 1930–1937. 329. S. Papadopoulos, K. D. Jürgens, and G. Gros, “Protein Diffusion in Living Skeletal Muscle Fibers: Dependence on Protein Size, Fiber Type, and Contraction,” Biophys. J., vol. 79, 2000, pp. 2084–2094. 330. K. D. Peck, A.-H. Ghanem, and W. I. Higuchi, “Hindered Diffusion of Polar Molecules Through and Effective Pore Radii Estimates of Intact and Ethanol Treated Human Epidermal Membrane,” Pharmaceutical Res., vol. 11, 1994, pp. 1306–1314. 331. T. Inamori, A.-H. Ghanem, W. I. Higuchi, and V. Srinivasan, “Macromolecule Transport in and Effective Pore Size of Ethanol Pretreated Human Epidermal Membrane,” Int. J. Pharmaceutics, vol. 105, 1994, pp. 113–123. 332. M. Sznitowska, “The influence of Ethanol on Permeation Behavior of the Porous Pathway in the Stratum Corneum,” Int. J. Pharmacol., vol. 137, 1996, pp. 137–140.

240

Optical Clearing of Tissues and Blood

333. A. K. Levang, K. Zhao, and J. Singh, “Effect of Ethanol/Propylene Glycol on the In Vitro Percutaneous Absorption of Aspirin, Biophysical Changes and Macroscopic Barrier Properties of the Skin,” Int. J. Pharmacol., vol. 181, 1999, pp. 255–263. 334. D. Bommannan, R. O. Potts, and R. H. Guy, “Examination of the Effect of Ethanol on Human Stratum Corneum In Vivo Using Infrared Spectroscopy,” J. Control Release, vol. 16, 1991, pp. 299–304. 335. C. A. Squier, M. J. Kremer, and P. W. Wertz, “Effect of Ethanol on Lipid Metabolism and Epidermal Permeability Barrier of Skin and Oral Mucosa in the Rat,” J. Oral Pathol. Med., vol. 32, 2003, pp. 595–599. 336. U. Jacobi, J. Bartoll, W. Sterry, and J. Lademann, “Orally Administered Ethanol: Transepidermal Pathways and Effects on the Human Skin Barrier,” Arch. Dermatol. Res., vol. 296, 2005, pp. 332–338. 337. http://www.dmso.org 338. J. Lademann, N. Otberg, H. Richter, H.-J. Weigmann, U. Lindemann, H. Schaefer, and W. Sterry, “Investigation of Follicular Penetration of Topically Applied Substances,” Skin Pharmacol. Appl. Skin Physiol., vol. 14, 2001, pp. 17–22. 339. J. Lademann, U. Jacobi, H. Richter, N. Otberg, H.-J. Weigmann, H. Meffert, H. Schaefer, U. Blume-Peytavi, and W. Sterry, “In Vivo Determination of UVPhotons Entering into Human Skin,” Laser Phys., vol. 14, 2004, pp. 1–4. 340. J. Lademann, A. Rudolph, U. Jacobi, H.-J. Weigmann, H. Schaefer, W. Sterry, and M. Meinke, “Influence of Nonhomogeneous Distribution of Topically Applied UV Filters on Sun Protection Factors,” J. Biomed. Opt., vol. 9, 2004, pp. 1358–1362. 341. S. Lee, D. J. McAuliffe, N. Kollias, T. J. Flotte, and A. G. Doukas “Photomechanical Delivery of 100-nm Microspheres Through the Stratum Corneum: Implications for Transdermal Drug Delivery,” Laser Surg. Med., vol. 31, 2002, pp. 207–210. 342. S. Lee, T. Anderson, H. Zhang, T. J. Flotte, and A. G. Doukas, “Alteration of Cell Membrane by Stress Waves In Vitro,” Ultrasound Med. Biol., vol. 22, 1996, pp. 1285–1293. 343. S. Lee, D. J. McAuliffe, H. Zhang, Z. Xu, J. Taitelbaum, T. J. Flotte, and A. G. Doukas, “Stress-Waves-Induced Membrane Permeation of Red Blood Cells is Facilitated by Aquaporins,” Ultrasound Med. Biol., vol. 23, 1997, pp. 1089–1094. 344. D. J. McAuliffe, S. Lee, H. Zhang, T. J. Flotte, and A. G. Doukas, “StressWaves-Assisted Transport Through the Plasma Membrane In Vitro,” Laser Surg. Med., vol. 20, 1997, pp. 216–222. 345. J. S. Soughayer, T. Krasieva, S. C. Jacobson, J. M. Ramsey, B. J. Tromberg, and N. L. Allbritton, “Characterization of Cellular Optoporation with Distance,” Anal. Chem., vol. 15, 2000, pp. 1342–1347. 346. G. Palumbo, M. Caruso, E. Crescenzi, M. F. Tecce, G. Roberti, and A. Colasanti, “Targeted Gene Transfer in Eukaryotic Cells by Dye-Assisted Laser Optoporation,” J. Photochem. Photobiol. B, vol. 36, 1996, pp. 41–46.

References

241

347. A. A. Gavrilova, D. Tabatadze, J. Childs, I. Yaroslavsky, G. Altshuler, A. B. Pravdin, and V. V. Tuchin, “In Vitro Optical Clearing of Rat Skin Using Lattice of Photoinduced Islands for Enhancement of Transdermal Permeability,” Proc. SPIE 5771, 2005, pp. 344–348. 348. C. L. Gay, R. H. Guy, G. M. Golden, V. H. W. Mak, and M. L. Francoeur, “Characterization of Low-Temperature (i.e., <65◦ C) Lipid Transitions in Human Stratum Corneum,” J. Invest. Dermatol., vol. 103, 1994, pp. 233–239. 349. O. Stumpp and A. J. Welch, “Injection of Glycerol into Porcine Skin for Optical Skin Clearing with Needle-Free Injection Gun and Determination of Agent Distribution Using OCT and Fluorescence Microscopy,” Proc. SPIE 4949, 2003, pp. 44–50. 350. O. Stumpp, A. J. Welch, and J. Neev, “Enhancement of Transdermal Skin Clearing Agent Delivery Using a 980 nm Diode Laser,” Lasers Surg. Med., vol. 33, 2005 (in press). 351. J. S. Nelson, J. L. McCullough, T. C. Glenn, W. H. Wright, L.-H. L. Liaw, and S. L. Jacques, “Mid-Infrared Laser Ablation of Stratum Corneum Enhances In Vitro Percutaneous Transport of Drugs,” J. Invest. Dermatol., vol. 97, 1991, pp. 874–879. 352. J.-Y. Fang, W.-R. Lee, S.-C. Shen, Y.-P. Fang, and C.-H. Hu, “Enhancement of Topical 5-Aminolaevulinic Acid Delivery by Erbium:YAG Laser and Microdermabrasion: A Comparison with Iontophoresis and Electroporation,” British J. Dermatol., vol. 151, 2004, pp. 132–140. 353. Study of the Composition of Dental Plaque, last updated: May 23, 2004, www.ClinicalTrials.gov 354. N. S. Soukas, A. S. Socransky, S. E. Mulholland, S. Lee, and A. G. Doukas, “Photomechanical Drug Delivery into Bacterial Biofilms,” Pharm. Res., vol. 17, no. 4, 2000, pp. 405–409. 355. U. Jacobi, E. Waibler, W. Sterry, and J. Lademann, “In Vivo Determination of the Long-Term Reservoir of the Horny Layer Using Laser Scanning Microscopy,” Laser Phys., vol. 15, no. 4, 2005, pp. 565–569. 356. E. Waibler, “Investigation of the Long-Term Reservoir of the Stratum Corneum-Quantification, Localization and Residence Time,” Summary of Doctoral Thesis, Jan. 2005, Charité-Universitätsmedizin, Berlin. 357. S. Mordon, C. Sumian, and J. M. Devoisselle, “Site-Specific Methylene Blue Delivery to Pilosebaceous Structures Using Highly Porous Nylon Microspheres: An Experimental Evaluation,” Lasers Surg. Med., vol. 33, 2003, pp. 119–125. 358. F. O. Nuttall, M. C. Gannon, W. R. Swaim, and M. J. Adams, “Stability Over Time of Glycohemoglobin, Glucose, and Red Blood Cells Survival in Hemologically Stable People Without Diabetes,” Metabolism, vol. 53, no. 11, 2004, pp. 1399–1404. 359. National Institute of Diabetes and Digestive and Kidney Diseases: Diabetes Overview. NIH, Bethesda, MD, 1994 (NIH publ. no. 94-3235).

242

Optical Clearing of Tissues and Blood

360. R. J. McNichols and G. L. Coté, “Optical Glucose Sensing in Biological Fluids: An Overview,” J. Biomed. Opt., vol. 5, 2000, pp. 5–16. 361. K. V. Larin, T. Akkin, M. Motamedi, R. O. Esenaliev, and T. E. Millner, “Phase-Sensitive Optical Low-Coherence Reflectometry for Detection of Analyte Concentration,” Appl. Opt., vol. 43, 2004, pp. 3408–3414. 362. L. Heinemann, U. Kramer, H. M. Klotzer, M. Hein, D. Volz, M. Hermann, T. Heise, and K. Rave, “Non-Invasive Task Force: Noninvasive Glucose Measurement by Monitoring of Scattering Coefficient During Oral Glucose Tolerance Tests,” Diabetes Technol. Ther., vol. 2, 2000, pp. 211–220. 363. M. Essenpreis, A. Knüttel, and D. Boecker, inventors; Boehringer Mannheim GmbH, assignee: “Method and Apparatus for Determining Glucose Concentration in a Biological Sample,” U.S. Patent 5,710,630, Jan. 20, 1998. 364. K. V. Larin, I. V. Larina, M. Motamedi, V. Gelikonov, R. Kuranov, and R. O. Esenaliev, “Potential Application of Optical Coherent Tomography for Non-Invasive Monitoring of Glucose Concentration,” Proc. SPIE 4263, 2001, pp. 83–90. 365. D. W. Schmidtke, A. C. Freeland, A. Heller, and R. T. Bonnecaze, “Measurements and Modeling of the Transient Difference Between Blood and Subcutaneous Glucose Concentrations in the Rat After Injection of Insulin,” Proc. Natl. Acad. Sci., vol. 95, 1998, pp. 294–299. 366. D. A. Zimnyakov, J. D. Briers, and V. V. Tuchin, “Speckle Technologies for Monitoring and Imaging of Tissues and Tissuelike Phantoms,” in Handbook of Optical Biomedical Diagnostics, Chapter 2, vol. PM107, V. V. Tuchin (ed.), SPIE Press, Bellingham, WA, 2002, pp. 987–1036. 367. S. J. Kirkpatrick and D. D. Duncan, “Optical Assessment of Tissue Mechanics,” in Handbook of Optical Biomedical Diagnostics, Chapter 2, vol. PM107, V. V. Tuchin (ed.), SPIE Press, Bellingham, WA, 2002, pp. 1037–1084. 368. D. A. Zimnyakov and V. V. Tuchin, “Speckle Correlometry” in Biomedical Photonics Handbook, Tuan Vo-Dinh (ed.), CRC Press, Boca Raton, 2003, pp. 14–1–14–23. 369. S. S. Ul’yanov, D. A. Zimnyakov, and V. V. Tuchin, “Fundamentals and Applications of Dynamic Speckles Induced by Focused Laser Beam Scattering,” Opt. Eng., vol. 33, 1994, pp. 3189–3201. 370. D. A. Zimnyakov and V. V. Tuchin, “Fractality of Speckle Intensity Fluctuations,” Appl. Opt., vol. 35, 1996, pp. 4325–4333. 371. J. C. Dainty (ed.), Laser Speckle and Related Phenomena, 2nd Ed., SpringerVerlag, New York, 1984. 372. J. W. Goodman, Statistical Optics, Wiley, New York, 1985. 373. J. Huisken, J. Swoger, F. Del Bene, J. Wittbrodt, and E. H. K Stelzer, “Optical Sectioning Deep Inside Live Embryos by Selective Plane Illumination Microscope,” Sci., vol. 305, 2004, pp. 1007–1009. 374. M. H. Khan, S. Chess, B. Choi, K. M. Kelly, and J. S. Nelson, “Can Topically Applied Optical Clearing Agents Increase the Epidermal Damage Threshold and Enhance Therapeutic Efficacy,” Lasers Surg. Med., vol. 35, 2004, pp. 93– 95.

References

243

375. M. H. Khan, C. Przeklasa, B. Choi, K. M. Kelly, and J. S. Nelson, “Laser Assisted Tattoo Removal in Combination with Topically Applied Optical Clearing Agents,” 25th Annual Meeting of the American Society for Laser Medicine and Surgery, March 30–April 3, 2005, Lake Buena Vista, Florida, Abstracts, Lasers Surg. Med., Suppl. 17, March 2005, p. 85. 376. P. Rol, P. Niederer, U. Dürr, P.-D. Henchoz, and F. Fankhauser, “Experimental Investigation on the Light Scattering Properties of the Human Sclera,” Laser Light Ophthalmol., vol. 3, 1990, pp. 201–212. 377. Yu. P. Sinichkin, S. R. Utz, A. H. Mavlutov, and H. A. Pilipenko, “In Vivo Fluorescence Spectroscopy of the Human Skin: Experiments and Models,” J. Biomed. Opt., vol. 3, 1998, pp. 201–211. 378. W.-F. Cheong, S. A. Prahl, and A. J. Welch, “A Review of the Optical Properties of Biological Tissues,” IEEE J. Quantum Electron., vol. 26, no. 12, 1990, pp. 2166–2185. Updated by W.-F. Cheong, further additions by L. Wang and S. L. Jacques, Aug. 6, 1993. 379. A. N. Yaroslavsky, I. V. Yaroslavsky, T. Goldbach, and H.-J. Schwarzmaier, “Inverse Hybrid Technique for Determining the Optical Properties of Turbid Media from Integrating-Sphere Measurements,” Appl. Opt., vol. 35, no. 34, 1996, pp. 6797–6809. 380. H.-J. Schwarzmaier, A. N. Yaroslavsky, I. V. Yaroslavsky, et al., “Optical Properties of Native and Coagulated Human Brain Structures,” Proc. SPIE 2970, 1997, pp. 492–499. 381. A. Roggan, O. Minet, C. Schröder, and G. Müller, “The Determination of Optical Tissue Properties with Double Integrating Sphere Technique and Monte Carlo Simulations,” Proc. SPIE 2100, 1994, pp. 42–56. 382. A. N. Yaroslavsky, P. C. Schulze, I. V. Yaroslavsky, R. Schober, F. Ulrich, and H.-J. Schwarzmaier, “Optical Properties of Selected Native and Coagulated Human Brain Tissues In Vitro in the Visible and Near Infrared Spectral Range,” Phys. Med. Biol., vol. 47, 2002, pp. 2059–2073. 383. J. Mobley and T. Vo-Dinh, “Optical Properties of Tissues,” in Biomedical Photonics Handbook, Tuan Vo-Dinh (ed.), CRC Press, Boca Raton, 2003, pp. 2-1–2-75. 384. S. Nickell, M. Hermann, M. Essenpreis, T. J. Farrell, U. Krämer, and M. S. Patterson, “Anisotropy of Light Propagation in Human Skin,” Phys. Med. Biol., vol. 45, 2000, pp. 2873–2886. 385. C. R. Simpson, M. Kohl, M. Essenpreis, and M. Cope, “Near-Infrared Optical Properties of Ex Vivo Human Skin and Subcutaneos Tissues Measured Using the Monte Carlo Inversion Technique,” Phys. Med. Biol., vol. 43, 1998, pp. 2465–2478. 386. J. Laufer, C. R. Simpson, M. Kohl, M. Essenpreis, and M. Cope, “Effect of Temperature on the Optical Properties of Ex Vivo Human Dermis and Subdermis,” Phys. Med. Biol., vol. 43, 1998, pp. 2479–2489. 387. V. V. Tuchin, Yu. N. Scherbakov, A. N. Yakunin, and I. V. Yaroslavsky, “Numerical Technique for Modeling of Laser-Induced Hyperthermia,” in Laser–

244

Optical Clearing of Tissues and Blood

Induced Interstitial Thermotherapy, G. Müller and A. Roggan (eds.), SPIE Press, Bellingham, WA, 1995, pp. 100–113. 388. Yu. N. Scherbakov, A. N. Yakunin, I. V. Yaroslavsky, and V. V. Tuchin, “Thermal Processes Modeling During Uncoagulating Laser Radiation Interaction with Multi-Layer Biotissue, 1. Theory and Calculating Models, 2. Numerical Results,” Opt. Spectrosc., vol. 76, no. 5, 1994, pp. 754–765. 389. R. J. McNichols, M. A. Fox, A. Gowda, S. Tuya, B. Bell, and M. Motamedi, “Temporary Dermal Scatter Reduction: Quantitative Assessment and Implications for Improved Laser Tattoo Removal,” Laser Surg. Med., vol. 36, 2005, pp. 289–296.

Index

A aberrations, 190 ablation, 210 abrasive creams, 151 absolute contrast, 100 absorption, 32 absorption bands, 70 absorption coefficient, 11, 112, 132, 193, 209 acanthosis, 79 acetic acid, xi, 212 aceto-whitening, 211 aceto-whitening effect, xi actual index profile, 1 adding-doubling (IAD) method, 194 adhesive tape, 203 adipose tissue, 175 adventitia, 88 aequorin, 154 agent diffusion coefficient, 14 agent flux, 148 aggregation-disaggregation effects, 123 aggregation processes, xii albumin, xi ambient humidity, 70 amide I, 70 amide II, 70 δ-aminolevulenic acid (ALA), 158 amphiphilic drugs, 156 Amphotericin B, 157 amplitude scattering matrix, 5 angle of observation, 63 angular-scattering functions, 8 anhydrous glycerol, 65 anisotropic scattering, 10 anisotropic scattering of blood, 112 anterior eye layers, 56 anti-inflammatory properties, 157 antimicrobial agents, 161 aorta, 91 apparent absorbance, 68, 83 apparent absorbance spectra, 83 apparent optical density (OD), 194 appearance, 61 appliqué with a pattern of absorbing centers, 69 arterioles, 107 aspirin, 156

asymmetry coefficient, 175 attenuated total reflectance Fourier transform infrared (ATR FTIR) spectroscopy, 70 autocorrelation analysis, 98 autocorrelation function, 175 autocorrelation peak, 181 average contrast, 100 axial, 78

B back-scattered photons, 26 background index, 1 bacteria, 133 ballistic photons, 26, 98, 178 barrier functions, 151 barrier functions of the skin, 61, 71 basal cell cancer, 211 beam, 208 beam defocusing, 99 biaxial birefringent structures, 87 bifurcation angles, 111 binary fluctuations, 3 biocompatible agents, 112 biofilm bacteria, 161 biomechanical properties of tissue, 205 biphasic effect, 123 birefringence of form, 87 blood, 33, 62 blood circulation, 86 blood flow, 109, 169, 186 blood glucose, 168 blood glucose analyzer, 171 blood glucose sensing, 167 blood hemolysis, 131 blood microvessels, 73, 107, 170 blood mononuclear cells, 159 blood optical clearing, 123 blood perfusion, 55 blood plasma, 112 blood plasma index, 117 blood plasma osmolarity, 112 blood pooling, 202 blood smear microscopy, 122 blood stases, 108 blood vessels, 96, 109 blood volume, 194

246 bone, 87 bone demineralization, 215 Bouguer–Beer–Lambert law, 11 boundary condition, 11, 14 boundary optical field, 175 bovine aorta, 194 bovine cornea, 22 bovine sclera, 194 bovine serum albumin, 134 brain tissue, 108, 206 breast, 178 brittleness, 61 bulk, 70 bulk absorption coefficient, 9 1,3-butanediol, 147 1,4-butanediol, 147 butylene glycol, 70

C cancer, 111 cancer cells, 212 capillaries, 107 capillary density, 111 carbohydrates, 157 carbon, 159 carbonyl absorption, 70 cartilage, x, 87 catalase, 155 cationized horse ferritin, 152 cavitation, 158, 161 CCD camera, 178 cells, 133 cell cornification, 133 cell damage, 133 cell death, 133, 143 cell membrane, x, 147 cell motility, 133 cell permeability, 133 cell shrinkage, 134 cell transfection, 160 cellular monolayer, 176 cerebral blood flow, 108 cerebral membrane, x cervical intraepithelial neoplasia, 211 cervical mucus, 152 cervical tissue, xi cervix, 178 chemical agent delivery, 147 chemical enhancers, 156 chemical potential, 46 chemical potential of water, 18 chemiluminescence, 97 chicken skin, 79, 104 Chinese hamster ovary cells, 160 chirality, 134 choroid, 48 chromaticity coordinates, 48

Optical Clearing of Tissues and Blood chromatin filaments, 211 ciliary body, 32, 48, 56 ciliary muscle, 55, 191 ciliary pigmented epithelium, 55, 191 circulating, 112 coagulated tissue, 209 coagulation, xi coefficient of diffusion, 13 coefficient of permeability, 16 coherence length, 170 coherent beam, 63 cold cataract, 211 collagen, x, 3, 21, 62, 98, 100 collagen denaturation, 209 collagen dissociation, xi collagen fibers, 12, 203 collagen fibers’ reversible dissociation, 104 collagen fibrils, 73, 193 collimated transmittance, 18, 32 colloid isoelectric point, 19 colonic mucosa, 84 color triangle, 48 colposcopic exam, 211 complex amplitude of the scattered light, 64 complex refractive indices, 4 complexly shaped scatterers, 7 compressing, 193 confocal microscopy, 94, 151 confocal probe, 94 confocal reflectance images, 211 confocal scanning microscopy, 162 conjunctiva, 48, 54 connective tissue, 1, 83 contrast, 63, 175 contrasting of tissue abnormalities, 74 cooling, 173 cooperative (interference) effects, xi, 193 cornea, 21, 87 corneal stroma, x corneal thickness, 22 corneocytes, 155 cornified layer, 61 correlated scattering, 6 correlation analysis, 175 correlation length, 63, 175 correlation time, 187 correlation-corrected distribution, 7 cosmetic lotions, x, 62, 148 cranial bones, 215 cranium, 59 creams, 155 cross-correlation function, 177 cross-linking of collagen, 104 crossed polarizers, 93 curcumin, 162 cytochrome c, 154 cytochrome oxidase, x

Index cytoplasma, x cytoplasmic organelles, 1, 2 cytoskelton, 159

D D periodicity, 20 decylmethylsulfoxide, 157 degree of permeability, 18 dehydration, 93, 103, 208 dehydration effect, 65 densely packed disperse system, 23 densely packed particles, 6 deoxyhemoglobin, 202 depolarization depth, 131 depolarization effects, 177 depolarization efficiency, 89 depolarization length, 89 derivatives of amino acids, 151 dermal vasculature, 82 dermis, 62, 103, 170 desquamation of the SC, 165 destructive interference, 64 deterioration, 71 developed speckles, 64, 176 dextrans, 112 diabetes mellitus, 167 dielectric spheres, 3 differentiation of vital cells, 143 diffuse reflectance, xii, 32, 83 diffuse reflectance spectra, 68 diffusion coefficient, 51, 54, 150 diffusion coefficient of glucose in dermis, 73 diffusion coefficient of water, 51, 68 diffusion equation, 12 diffusion model, 101 diffusion process, 184 diffusion through a biological membrane, 51 diffusion time, 15 diffusion-wave spectroscopy, 177, 186 diffusional pathlength, 156 diffusional resistance, 148 diffusivity of proteins, 152 diffusivity of water, 61 digital autocorrelator, 186 dimethylsulfoxide (DMSO), 62, 68, 157 diode-pumped solid-state laser, 99 dipropylene glycol, 70 directed blood flow, 186 disaggregation processes, xii dissolved molecule, 18 distribution of speckle sizes, 175 distribution size function, 6 DOPE (dioleylphosphatidylethanolamine), 151 dorsal skin, 65, 95 dried tissue, 21 drug cytotoxicity, 160

247 dry volume, 20 dura mater, 19, 87, 108 dye laser, 95 dynamics of tissue optical clearing, 47 dysplastic nuclei, 3 dysplastic tissue, 211

E edema, xii effective pore radii, 156 elastic scattering, 3 elasticity, 61 elastin, x, 21 elastin fibers, 3, 62, 88, 100 electrical resistance, 61 electron micrographs, 23 electronic correlator, 186 electroosmosis, 158 electrophoresis, 70 electrophoretic applicator, 71 electroporation, 158 empirical probability density function, 63 emulsions, 150 encapsulated drugs, 159 encapsulated molecular probes, 159 endogenous osmic liquid, xi engineered tissue model (raft), 103 enhanced agent delivery, 82 enhanced molecule delivery, 148 enhancement of transmittance, 79 enzymes, 161 epidermal cell nuclei, 211 epidermal glue stripping, 71 epidermis, 61 epithelial cells, 7 epithelial disease, 211 epithelium, 83, 85 erbium laser, 161 erythema, xii, 71 erythematous human skin, 194 erythrocyte cytoplasm, 112 erythrocyte size, 125 erythrocytes aggregation, 112 erythrocytes’ volume, 112 esophagus, 84 ethanol, 156 ethanol/glycerol solvents, 165 ethanol/propylene glycol composition, 156 eukaryotic ribosomes, 155 excimer laser, 210 exogenous osmotic liquid, xi exponential factor, 177 extracellular media, 87 extrafibrillar material, 21 eye, 178

248

F far-field diffraction zone, 176 fat cellular layer, 62 fat-free dry weight, 3 femtosecond laser pulses, 188 fiber probe, 32 fiber-based dual-channel phase-sensitive optical low-coherence reflectometer (PS-OLCR), 170 fibroblasts, 3, 88 fibrous tissue, x, 19–20 Fick’s law, 16 fine porous medium, 150 finite-difference time-domain (FDTD) approach, 144 first-order speckle statistics, 175 fixed tissue samples, 55 flashlamp system, 68 flow of water, 150 flow velocity, 107 fluence rate, 11, 12 fluid diffusion, 46 fluid transport, 43 fluorescence, 95 fluorescence probe, 97 fluorescence recovery after photobleaching (FRAP), 152 fluorescence scanning microscopy, 98 fluorescent source, 97 fluorophores, 97 flux of matter, 16 focal area, 98 focused laser beam, 63 follicular infundibulum, 164 follicular orifices, 164 formalin fixing, 104 forward-scattered photons, 26 Fourier transform, 176 fractal object, 7 Franz diffusion chamber, 147 free convection, 208 free diffusion, 12 free water, 70 freezing, xi, 208 frequency-domain technique, 52 Fresnel reflection coefficient, 11 fungi, 133

G β-galactosidase, 155 gamma distribution, 7 gastrointestinal mucus, 151 gastrointestinal tissues, 83–84 gating technique, 103 Gaussian, 208 Gaussian statistics, 64, 177 gels, x, 62, 70

Optical Clearing of Tissues and Blood gel fiber, 152 gene, 160 gene expression, 178 glass substrate-glue technology, 63 glaucoma, 191 globular proteins, 151 glucose, x, 32, 58, 62, 100, 107, 112, 129 glucose clamp experiment, 167 glucose diffusion in water, 73 glucose gel, 71 glucose uptake, 175 glucose-water solution, 32 glycation, 112 glycerine-aldehyde-3-P-dehydrogenase, 155 glycerol, x, 58, 62, 70, 100, 107, 109, 112, 147 glycerol gel, 71 glycerol injection, 71 glycerol-water solutions, 62 glycerol-water-urea solution, 64 glycolipids, 206 gold, 159 granular layer, 61 green fluorescence protein (GFP), 160 ground matter, xi, 62 growth cycle, 133 gyaluronic acid, 151

H hair follicles, 82, 151, 158 hair shaft, 164 hamster skin, 71 hard tissues, 215 heartbeats, 76 heating, 70 helix of collagen secondary structure, 103 hematocrit, 112, 129 hematous tissue, x hemoglobin, 112 hemoglobin absorption, 58, 74 hemoglobin glycation, 167 hemoglobin oxygenation, 112 hemoglobin saturation, x hemolysis, 107 hemolyzed blood, 125 hemorheological status of patients, 124 hexagonal arrangement, 21 hindered diffusion, 16, 156 homogeneous slab, 15 human aorta, 208 human cervical mucus, 147 human dermis, 205 human dura mater, 58 human epidermal membrane, 147, 156 human fat tissue, 107 human immunoglobulin molecules, 152 human sclera, 23, 52

Index human serum albumin, 134 human skin, 147, 194 human skin in vivo, 99 humidity, 61 hydration, 20 hydration of spores, 143 hydration of the dermis, 61 hydrodynamic diameter, 151 hydrophilic component, 82 hydrophilic compounds, 156 hydrophilic polyethylene glycol (PEG)-based prepolymers, 82 hydrostatic pressure, 19, 48, 159 α-hydroxy acids, 151 hydroxyapatite crystals, 87 hydroxyethyl cellulose, 70 Hypaque-76, 189 hyperinsulinemic-hypoglycemic clamp, 170 hyperosmotic agents, 10, 100, 112 hyperosmotic saline, 212 hypodermic fatty layer, 65

I ideal partially permeable membrane, 19 image contrast, 91, 94, 99 imaging, 103 imaging contrast, 79 immersion liquid, 23 immersion technique, 143 impermeable wrap, 155 implant, 191 in vivo frequency-domain measurements, 75 in vivo measurements, 54 indocyanine green (ICG), 74, 165 induced erythema, 202 injection, 52 injection of distillate water, 71 insulin, 158 integrating sphere, 32 integrating sphere (IS) technique, 194 integrating sphere spectrophotometer, 33 intensity modulation, 52 intensity probability density distribution function, 63, 175 interacting fluxes, 43 intercellular lipids, 155, 164 intercellular motility, 98 interference, 63 interfibrillar space, 14, 17, 20, 47 interfibrillar substance, 12 intermolecular spacing, 20 interparticle correlation effects, 7, 112 interstitial liquid, xi intracellular diffusion, 155 intracellular fluids, 193 intracellular organelles, 134 intradermal injection, 71, 98, 107

249 intravascular structures, 112 intravenous bolus injection, 171 intravenous clamping technique, 171 intravessel injection, 129 intrinsic birefringence, 89 intrinsic muscle proteins, 154 inverse adding-doubling method, 32 inverse Monte Carlo method, 32 ionic strength, 20 iontophoresis, 158 irreversible changes, 46 irreversible thermodynamics, 18 isoosmotic condition, 134 isotonic condition, 134 isotropic scattering, 10 isthmus, 85

K keratinization, 61 keratinocytes, xi, 62

L lamina propia, 85 laminar blood flow, 113 Langer’s skin tension lines, 205 large-size inhomogeneities, 176 laser, 210 laser heating, xii laser skin surface ablation, 161 laser speckle imaging, 109 laser-assisted optoporation, 160 laser-generated stress waves, 158 lateral resolution, 170 lattice of microzones (islets), 68, 160 law of Gladstone and Dale, 1 lenslets, 176 leukocytes, 112 limited thermal damage, 68, 160 linear anisotropy, 87 linear birefringence, x, 134 linear optical properties, 104 linear polarization, 47 lipids, 3, 150 lipid bridges, 61 lipid lamellae, 158 lipophilic, 156 lipophilic polypropylene glycol-based prepolymers (PPG), 82 liposomes, 70, 151 living scattering etalon, 48 local heating, 48 local heating of the plasma membrane, 160 long-pulsed laser heating, 210 loss of water, 208 lotions, 155 low-frequency oscillations, 43 low-step scattering, 12

250 Lugol’s iodine solution, 212 lymph microvessels, 213 lymphocytes, 3 lysozyme, 152

M malignant melanoma, 80 mannitol, 58, 156 mass fractal, 8 mass loss, 68 mass measurements, 68 mean diameter of a speckle, 180 mean free path length, 113 mean free path of a photon, 26 mean index variation, 2 melanin granules, 1 melanin nanoparticles, 159 melanin pigmentation, 202 melanin synthesis, xii membrane acquaporins, 159 membrane brightness, 144 membrane permeabilization, 159 membrane porosity, 156 mesh effective spacing, 152 metabolic activity, 55 metabolites, 161 method of T-matrices, 7 methylene blue (MB), 165 microbial infections, 161 microcirculation, 107 microfocusing effect, 176 microinjection, 152 microoptical model, 6 microorganisms, 61, 133 microspectrophotometry, 152 mitochondria, 2, 88 mixture of prepolymers PPG/PEG, 215 mode-locked dye laser, 99 modulation depth, 52 moisture factor, 70 moisturizers, 70 molecular displacement processes, 18 monodisperse tissue model, 3 Monte Carlo modeling, 23, 55 Monte Carlo simulation, 94, 191 motion artifacts, 173 mucins, 151 mucopolysaccharide, 151, 193 mucosa, 83 multichannel fiber-optic switcher, 52 multifiber detection system, 52 multilayered tissue structure, 97 multiphoton fluorescence scanning microscopy, 98 multiple scattering, 12, 26, 176 multiple-single scattering transition, 47 muscle, x–xi, 87 muscle tissue, 85, 168

Optical Clearing of Tissues and Blood myelin, 87 myocardium, 88 myofibrils, 88 myofilamental lattice, 155 myofilaments, 88 myoglobin, 154

N natural moisturizing vegetal compounds, 70 Nd:YAG laser, 99, 191 needle-free injection gun, 161 negative uniaxial crystal, 89 nerve, 87 New Zealand rabbits, 171 NIR frequency-domain reflectance techniques, 167 noncoagulating heating, 208 noninteracting thin cylinders, 17 nonlinear spectroscopy, 103 nonlinear susceptibility, 103 normalized statistical moment, 175 nourishing (nutritive) creams, 150 nuclear scatter, xi nucleic acid, 157 nucleus, 211

O OCA, 187 occlusion, 155 occlusion spectroscopy, 170 oils, 150 ointments, 155 oleic acid, 62, 68, 157 optical activity, x optical breakdown, 188 optical clearing, 111, 178 optical clearing potentials (OCP), 147 optical coherence tomography (OCT), 78 OCT images, 107 OCT signal slope, 173 OCT technique, 170 OCT-based glucose sensor, 173 optical depth, 113 optical functional tomography, 178 optical immersion of cells, 133 optical immersion technique, 12, 100, 105 optical properties of blood, 112 optical sectioning, 99 optical transmission, 193 optically clarified blood, 112 oral glucose tolerance test, 168 oscillations, 46 osmolarity, 129, 147 osmotic dehydration of erythrocytes, 112 osmotic pressure, 19, 48, 134 osmotic strength, 51

Index osmotic stress, x oxygen, 161

P packing factor, 6 paraffin oil, 162, 164 paraxial region, 177 particle separation distance, 25 passive transport, 156 patent blue V, 162 peeling creams, 151 PEG, 32 penetration depth, 86, 99 percutaneous absorption, 155 permeability, 66 permeable membrane, 16 permeation efficiency, 147 permeation pathways, 156 perovskite, 210 petrolatum, 155 pharmaceutical products, x, 62 phase contrast, 145 phase contrast microscopy, xi phase delay, 88 phase difference, 88 phase fluctuations, 175 phase plate, 89, 185 phase refractometry, 133 phase retardation, 88 phase shift, 52 phase transition of membrane lipids, 160 phenol red, 160 phonophoresis, 158 phospholipids, 151 photodisruption, 188 photodynamic acne treatment, 165 photomechanical waves, 158 photomultiplier, 186 photon-density wave, 52 photon-diffusion theory, 6 photon diffusion coefficient, 12 photon migration paths, 55 photon statistics, 23 phototherapy, xii photothermal therapy, 111 physiological hydration, 20, 21 physiological pH, 20, 157 physiological solution, 32 physiological temperature, 42 pigmentation, 33 plasma membrane, 159 plasmid DNA, 159 point-spread function, 94 pointwise compression, 194 pointwise mechanical stress, 47 polar molecules, 156 polarization, 177

251 polarization anisotropy, x polarization components, 88 polarization images, 185 polarization immersion effect, 93 polarization-speckle microscope, 93 polarized light microscope, 93 polarized reflectivity, 205 polidispersity, 129 pollutions, 61 polydispersion, 6, 112 polyglycerylmethacrylate, 70 polymer gel, 152 polymer mixture, 82 porcine sclera, 194 porcine skin, 68, 97 porcine stomach, 86 porosity coefficient, 15 porous material, 15 porous nylon microspheres, 165 positive uniaxial birefringent medium, 88 posterior interface, 56 powders, 155 power law, 180 power spectrum, 176 precancerous condition, xi precoagulated tissue, 209 presbyopia, 191 prolated ellipsoidal particles, 7 propylene glycol, x, 62, 70, 100, 107, 162 protective function, 61 protein, 3, 33, 151, 193 protein aggregation, 211 protein-bound water, 70 protein denaturation, 209 protein solutions, 142 protein/lipid ratio, 206 protein/lipid/water matrix, 87 proteoglycans, 20, 22 “pseudohexagonal” arrangement, 22 psoriatic erythrodermia, 79 pulse shape, 99 pulsed Nd:YAG laser, 210 pyrrolidonecarboxylic acid, 151

Q Q-switched alexandrite laser, 191 quality of images, 59 quantum dots, 159 quasi-periodic oscillations, 46, 184

R rabbit cornea, 22 rabbit dura mater, 58 radial distribution function, 23 raffinose, 156 raft tissue, 210 random phase screen, 177

252 random phases, 63 random process, 176 random relief, 63 rat fibroblast cell clone, 143 rat mesentery, 107 rat prostate carcinoma, 7 rate of contrast, 103 Rayleigh limit, 88 Rayleigh–Gans approximation, 168 RBC packing function, 129 RBC sedimentation, 120 RBC shrinking (acanthocytes), 124, 129 RBC swelling (spherocytes), xii, 124 RBC-plasma interface, 120 red blood cells, 112 reduced scattering coefficient, 11, 32, 208 reflectance, x, 48 reflection spectra, 23 refractive index, 32–33, 147, 170 refractive index matching, 26, 129, 146 refractive index mismatch, 112, 167, 207 refractive index of blood plasma, 127 refractive index of erythrocyte cytoplasm, 117 refractive index of water, 32 refractive indices of bacteria, 143 rehydrated sample, 65 rehydration of lyophilized bacterial cells, 143 rehydration process, 65, 104 relative contrast, 100 release of sebum, 165 removal of blood, 193 reproductive tract, 151 respiratory tract, 151 resurfacing techniques, 161 reticular dermis, 94 reticular fibers, 88 reticulin, x, 21 retina, 32, 48, 87 reversible dissociation of collagen, 93 rhodamine fluorescent film, 95 ribosomes, 2 rod-shaped particles, 6 rodent tail tendon, 93

S saline bath, 209 sampling depth, 175 sarcoplasmic protein, 155 saturation of contrast, 101 saturation of the optical transmittance, 66 scanning rate, 181 scattered photons, 98 scattering angle, 5 scattering anisotropy factor, 129 scattering coefficient, 112, 129, 170 scattering phase function, 4

Optical Clearing of Tissues and Blood sclera, 32, 87, 191 sclera color, 48 sclera inflammation, 54 sclera transmission, 23 scleral optical clearing, 185 scleral stroma, x sclerotomy, 190 sea collagen, 151 sebaceous glands, 162, 165 second-harmonic generation (SHG), 103 SHG photon recycling, 104 SHG polarimetry, 103 SHG radar graph, 104 sedimentation process, 113 sedimenting blood, 112 self-beating efficiency, 186 self-diffusion coefficient, 150 semipermeable membrane, 18 shape of eryhrocytes, 112 shape of the aggregates, 124 shear forces, 170 sheet-like particles, 6 shrinkage, xii, 19, 43, 208 shunts, 108 Siegert formula, 186 silicon oil, 164 silicon wax, 150 single scattering model, 79 single-mode fiber, 186 single-mode He:Ne laser, 180 single-scattering, 12 size distribution, 6 size of a speckle, 63 skeletal muscle fibers, 147, 154, 175 skewed log-normal distribution, 6 skin, 61, 151, 178 skin appendages, 162 skin autofluorescence, 194 skin barrier function, 156 skin color appearance, 202 skin compression, 194 skin dermis, x skin flap window, 96, 107, 109 skin flaps, 205 skin heating, 173 skin irritation, 157 skin microdermabrasion, 161 skin permeation, 148 skin phototype, 191 skin reservoirs, 162 skin stripping, 63, 158, 176 skin surface temperature, 206 skin temperature fluctuations, 173 slow diffusion, 193 small living animals, 178 small-scale intensity oscillations, 176 smooth muscle, 175 sodium fluorescein, 162

Index sodium lactate, 70, 151 sodium lauryl sulfate, 159 soft tissue, 1 solvent convective flow, 158 sonophoresis, 70, 158 source term, 11 source-detector separation, 11 spatial correlation, 25, 178 spatial fluctuations, 63 spatial resolution, 94 spatial-temporal fluctuations, 43 spatially resolved diffuse reflectance, 18, 167 spatially resolved steady-state reflectance, 11 specific refractive increment, 33 speckle, 63, 173 speckle “boiling”, 180 speckle correlometer, 180 speckle fields, 177 speckle intensity fluctuations, 63 speckle patterns, 63 speckle topography, 175 spectral measurements, 63 specular component, 180 spores, 133 squeezing, 193 standard deviation, 175 statistical analysis, 175 statistical approach, 1 statistical mean index profile, 1 statistics of the second order, 175 stomach, 84 stratum corneum (SC), 61 SC ablation, 161 SC hydration, 70 stretched human skin, 203 stretching, xi, 193 stromal edema, 22 structure function, 176 subcutaneous glucose concentration, 175 subdermis, 205 subepidermal blisters, 80 sucrose, 156 sugars, 151 sunscreen creams, x, 62 superluminescence diode, 78 surfactants, 157 sweat glands, 61, 156 swelling, 19, 22, 134 swelling area, 73 swelling of the corneocytes, 156 swelling white ring, 73 synergetic effect, 83

T tape stripping method, 162 tattoo ink, 191 tattoo lightening, 191

253 tattoo removal, 191 technical lecithin, 150 temperature coefficients, 205 temperature-modulated skin, 206 temporal intensity fluctuations, 63 tendon, x, 87, 103 Tenon’s capsule, 55 tensile strength, 61 tension in wound closures, 205 testis, 87 thermal conductivity, 61 thermal damage, 209 thermal denaturation, 161 thermal insulation, 208 thermo-optical response, 207 thermodynamic activity, 157 thin membrane, 16 threshold for injury, 191 thrombocytes, 112 time-dependent transmittance, 48 tissue, x, xi, 19, 20, 22 tissue birefringence, 87 tissue dehydration, xi, 86 tissue expanders, 205 tissue fixation, 104 tissue hydration, 20 tissue matrix, 1 tissue optical clearing, 188 tissue proliferation, xii tissue swelling, 43 tissue whitening, 211 tissue-discrete particle model, 1 tissue-enhanced translucence, 26 titanium sapphire laser, 99 toluidine blue, 212 tomography, 175 tooth, 87, 161, 215 tooth biofilm, 161 topical application, 95 topographic mapping, 175 total attenuation coefficient, 78, 112 total transmittance, xii, 32 transdermal delivery, 148, 159 transdermal delivery of immersion agents, 61 transdermal skin resistance measurements, 68 transepidermal water lost (TEWL), 61, 150 transfection rate, 160 transmembrane permeation, 147 transmittance spectra, x transparent window, 107 transport free path length, 131 transport mean free path, 11 transport scattering coefficient, 129 transscleral laser surgery, 178 transscleral selective photodestruction, 58 transverse resolution, 78 trisaccharides, 148 tumor, 56, 178

254 two log-normal size distributions, 7 two-photon microscopy, 98

U ultrasound, 158 uniaxial birefringent medium, 87 uniaxial second-order nonlinear susceptibility, 103 unsaturated fatty acids, 157 urea, 70, 156 UV filter substance, 157 UV irradiation, xii UV skin filters, 165 UV-induced erythema, 194

V vacuoles, 206 Van de Hulst approximation, 3 vascular disease, 111, 112 vascular tissues, 94 vein, 91 velocity images, 108 venules, 107 vessel dilation, 108 virtual transparent window, 73 viscoelastic gel, 151 viscoelasticity, 61 viscosity, 142 viscous flow, 61, 150 volume density, 3

Optical Clearing of Tissues and Blood volume fraction, 1 volume fraction of RBCs, 128 volume-averaged concentration, 15 volume-averaged phase function, 8

W waist of the laser beam, 176 water absorption bands, xi water balance, 18 water channels, 155 water content, 61 water desorption, 83 water flux, 68 water loss, 83 water loss measurements, 68 water pools, 155 water transport, 18 water-glycerol gelatinous gel, 134 wavelets, 63 weakly scattering objects, 176 whole eyeball, 52

X x-ray contrast, 112, 189 x-ray contrast agent, x, 32

Y yeasts, 133 Yucatan micropigs, 171

Valery V. Tuchin was born February 4, 1944. He received his degrees in radiophysics and electronics (1966), Candidate of Sciences in Optics (Ph.D., 1973), and Doctor of Science in Quantum Radiophysics (1982) from Saratov State University, Saratov, Russia. He is a Professor and holds the optics and biomedical physics chair, and he is a director of the research-educational Institute of Optics and Biophotonics at Saratov State University. Prof. Tuchin also heads the Laboratory of Laser Diagnostics of Technical and Living Systems at the Insitute of Precision Mechanics and Control, Russian Academy of Sciences. He was dean of the Faculty of Physics of Saratov University from 1982 to 1989. His research interests include biomedical optics and laser medicine, nonlinear dynamics of laser and biophysical systems, and physics of optical and laser measurements. He has authored more than 250 peer-reviewed papers and books, including Tissue Optics. Light Scattering Methods and Instrumentation for Medical Diagnosis (SPIE Tutorial Texts in Optical Engineering, Vol. TT38, 2000), Handbook of Optical Biomedical Diagnostics (SPIE Press, Vol. PM107, 2002), CoherentDomain Optical Methods for Biomedical Diagnostics, Environmental and Material Science (Kluwer Academic Publishers, Boston, USA, Vol. 1 & 2, 2004), and Optical Polarization in Biomedical Applications, Springer-Verlag, 2005 (Lihong Wang and Dmitry Zimnyakov—co-authors). He is a holder of more than 20 patents. Prof. Tuchin currently teaches courses on optics, tissue optics, laser and fiber optics in biomedicine, laser dynamics, biophysics and medical physics for undergraduate and postgraduate students. Since 1992, Prof. Tuchin has been the instructor of SPIE and OSA short courses on biomedical optics for an international audience of engineers, students, and medical doctors; and he is an editorial board member of J. of Biomedical Optics (SPIE), Lasers in the Life Sciences, and J. of X-Ray Science and Technology, and the Russian journals Izvestiya VUZ, Applied Nonlinear Dynamics and Laser Medicine. He is a member of the Russian Academy of Natural Sciences and the International Academy of Information, a member of the board of SPIE/RUS and a fellow of SPIE (since 2004). He has been awarded the title and scholarship “Soros Professor” (1997–1999), and the Russian Federation Scholarship for the outstanding scientists (1994–2003); and Honored Science Worker of the Russian Federation (since 2000).