UNIVERSITY of CALIFORNIA, SAN DIEGO Leveraging the Photothermal Effect for Cosmetic Applications a Dissertation Submitted in Pa

Home , Photothermal therapy

UNIVERSITY OF CALIFORNIA, SAN DIEGO

Leveraging the Photothermal Effect for Cosmetic Applications

A dissertation submitted in partial satisfaction of the

requirements for the degree Doctor of Philosophy

Materials Science and Engineering

Wangzhong Sheng

Committee in charge:

Professor Adah Almutairi, Chair Professor Shaochen Chen Professor Yi Chen Professor Jesse Jokerst Professor Robert Sah

2017

Wangzhong Sheng, 2017

The Dissertation of Wangzhong Sheng is approved, and it is acceptable in quality and

form for publication on microfilm and electronically:

Chair

University of California, San Diego

2017

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DEDICATION

I dedicate this thesis to my parents, Zhouying Zhao and Yongchen Sheng, for their sacrifice and unwavering support over the years, but especially to my mother, a fellow doctorate in Materials Science and Engineering, for pushing me, inspiring me, and supporting me through this incredible journey.

TABLE OF CONTENTS

Signature Page………………………………………………………………………. iii

Dedication…………………………………………………………………………... iv

Table of Contents…………………………………………………………………… v

List of Abbreviations……………………………………………………………….. ix

List of Figures………………………………………………………………………. xi

List of Tables……………………………………………………………………..... xv

Acknowledgements…………………………………………………………………. xvi

Vita………………………………………………………………………………….. xix

Abstract of the Dissertation…………………………………………………………. xx

Introduction: A Review of the Progress towards Achieving Heat Confinement: the Holy

Grail of Photothermal Therapy ...... 1

I. ABSTRACT ...... 1

II. INTRODUCTION ...... 2

III. LIGHT TISSUE INTERACTIONS ...... 6

IV. PHOTOTHERMAL MATERIALS PLATFORMS ...... 14

V. NANOMEDICINE ENABLED SYNERGISTIC PHOTOTHERMAL

THERAPY PLATFORMS ...... 26

VI. PROGRESS AND FUTURE PERSPECTIVES ...... 34

VII. CONCLUDING REMARKS ...... 36

VIII. ACKNOWLEDGEMENTS ...... 38

IX. REFERENCES ...... 38

Chapter I: Near Infrared-Induced Heating of Confined Water in Polymeric Particles for

Efficient Payload Release ...... 50

1.0 CHAPTER ONE PREAMBLE ...... 50

1.1 ABSTRACT ...... 51

1.2 INTRODUCTION ...... 52

1.3 RESULTS AND DISCUSSION ...... 55

1.4 CONCLUSION ...... 72

1.5 METHODS ...... 74

1.6 ACKNOWLEDGMENTS ...... 80

1.7 REFERENCES ...... 81

1.8 SUPPORTING INFORMATION ...... 87

Chapter II: Selective Photothermolysis Theory and Gold-Nanoparticle Laser-Induced

Photothermal-Assisted Liposuction (Nanolipo) in a Porcine Model ...... 98

2.0 CHAPTER TWO PREAMBLE ...... 98

2.1 ENGINEERING ANALYSIS OF SELECTIVE PHOTOTHERMOLYSIS

AND MODELING ...... 99

2.2 ABSTRACT ...... 109

2.3 INTRODUCTION ...... 110

2.4 MATERIALS & METHODS ...... 111

2.5 RESULTS ...... 115

2.6 DISCUSSION ...... 117

2.7 CONCLUSIONS ...... 119

2.8 ACKNOWLEDGEMENTS ...... 119

2.9 REFERENCES ...... 120

2.10 FIGURE LEGENDS AND TABLES ...... 123

2.11 BIODISTRIBUTION AND CLEARANCE OF GOLD NANOPARTICLES

FROM PIGS ...... 128

Chapter III: A Single-Blind Study Evaluating the Efficacy of NanoLipo in an ex vivo

Human Tissue Model ...... 132

3.0 PREAMBLE: FUNCTION AND PHYSIOLOGY OF HUMAN ADIPOSE

TISSUE ...... 132

3.1 ABSTRACT ...... 138

3.2 INTRODUCTION ...... 139

3.3 MATERIALS AND METHODS ...... 141

3.4 RESULTS ...... 146

3.5 DISCUSSION ...... 150

3.6 CONCLUSION ...... 155

3.7 CONFLICT OF INTEREST ...... 156

3.8 ACKNOWLEDGEMENTS ...... 156

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3.9 REFERENCES ...... 158

3.10 FIGURES ...... 161

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LIST OF ABBREVIATIONS

Introduction

PTT, photothermal therapy; OD, optical density; LITT, laser interstitial thermal therapy; NIR, near infra-red; TR, thermal relaxation; SPT, selective photothermolysis; LSPR, localized surface plasmon resonance; AuNR, gold nanorods;

AuNS, gold nanoshells; ICG, indocyanine green; NCs, nanocrystals; HGNS, hollow gold nanoshells; FDA, food and drug administration; PEG, poly(ethylene) glycol;

Hsp, heat shock protein; DOX, doxorubicin; PDT, photodynamic therapy; PAc, photoacoustic imaging; NP, nanoparticles.

Chapter 1

PLGA, poly(lactic-co-glycolic) acid; Tg, glass transition temperature; NIR, near infra-red; UCNP, upconverting nanoparticles; FLIM, fluorescence lifetime imaging microscopy; GPC, gel-permeation chromatography; CW, continuous wave;

D2O, deuterated water; DSC, differential scanning calorimetry; Cp, heat capacity;

FDA, fluorescein diacetate; CHCl3, chloroform; DMF, dimethylformamide; TFH, tetrahydrofuran.

Chapter 2

LCST, lower critical solution temperature; TRT, thermal relaxation time; pNIPAM, poly n-isopropylacrylamide; SAL, suction-assisted lipectomy; NIR, near

infra-red; AuNR, gold nanorod; PEG, poly(ethylene) glycol; OD, optical density; ICP-

MS, inductively coupled plasmas-mass spectroscopy.

Chapter 3

AT, adipose tissue: WAT, BAT, VAT, SAT, sSAT, dSAT, white, brown, visceral, subcutaneous, superficial subcutaneous, deep subcutaneous; ECM, extracellular matrix; GNP, gold nanoparticles; PBS, phosphate buffered saline; PEG, poly(ethylene) glycol; NBTC, Nitroblue tetrazolium chloride; SPT, selective photothermolysis.

LIST OF FIGURES

Figure 0.1. Schematic highlighting factors crucial to achieving material and heat confinement in photothermal therapy …………………….…. 6

Figure 0.2. Wavelengths between 700-900 nm are minimally absorbed by endogenous species and are able to penetrate deepest into tissue…….…………………………………….……… 10

Figure 0.3. Examples of photosensitizers applied for photothermal therapy..……………………………………………………………. 17

Figure 0.4. A triple-combination local delivery platform………...…………… 25

Figure 0.5. Photothermal theranostic systems ………………………………… 33

Figure 1.1. Schematic representation of NIR-induced release mechanism ………………………………………………………… 55

Figure 1.2. NIR-induced release occurs only at water-resonant wavelengths ………………………………………………………. 60

Figure 1.3. Metal-free photothermal release requires efficient vibrational absorption by confined water …………………………………….. 63

Figure 1.4. Fluorescence lifetime imaging (FLIM) reveals an increase in particles’ internal temperature.…………………………...…….. 70

Figure 1.5. NIR light-triggered intracellular FDA release.……………………. 72

Figure 1.S1. Electrosprayed PLGA particles are uniform………………………. 88

Figure 1.S2. Electrosprayed PLGA particles efficiently encapsulate the dye molecules …………………………………………………………. 89

Figure 1.S3. Digital photographs of a solution of fluorescein-loaded PLGA particles …………………………………………………………… 89

Figure 1.S4. SEM images of fluorescein-loaded PLGA particles…..…………... 90

Figure 1.S5. GPC chromatograms of the PLGA ……………………………….. 90

Figure 1.S6. Photographs displaying laser-induced particles agglomeration over time ……………………………………………………………….. 91

Figure 1.S7. NIR irradiation at a water-resonant wavelength induces release of both hydrophobic and hydrophilic compounds …………………. 92

Figure 1.S8. Fluorescence microscopy images and size distribution histogram of Nile Red-loaded polyester 1 particles…………...……………… 93

Figure 1.S9. NIR irradiation at a water-resonant wavelength can induce release from capsules made of different polymeric materials …………….. 94

Figure 1.S10. Changes in average lifetime of fluorescein encapsulated in PLGA particles embedded in a polyacrylamide gel upon alternating cycles of irradiation ………………………………………………. 95

Figure 1.S11. Light microscopy images and size distribution of electrosprayed FDA-doped polyester 1 particles …………………………………. 95

Figure 1.S12. Cytotoxicity study on the effect of NIR-irradiation ……………… 96

Figure 2.1.1. Schematic showing various achievable temperature profiles during selective photothermolysis ……………………………………….. 104

Figure 2.1.2. Continuous light exposure induces heating of the photothermal layer ………………………………………………………………. 107

Figure 2.1.3. Discontinuous light exposure with sufficient cool-down times in between exposure events enables selective photothermal heating …………………………………………….. 108

Figure 2.2.1. Schematic of NanoLipo procedure ……………………………….. 123

Figure 2.2.2 Tunable absorption of AuNR solution used in NanoLipo ………… 124

Figure 2.2.3 Histological effects of NanoLipo in ex vivo porcine skin and subcutaneous tissue ……………………………………………….. 124

Figure 2.2.4 NanoLipo allows removal of more tissue and more fat than standard suction-assisted lipectomy ……………………………… 125

Figure 2.2.5 Ultrasound imaging analysis reveals that NanoLipo enhances reductions in tissue thickness …………………………………….. 126

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Figure 2.2.S1 Treatment regions were marked with non-absorbable sutures…………………………………………………………...... 127

Figure 2.2.S2 Details of ex vivo experiments…………………………………….. 127

Figure 2.3.1 Clearance profile of AuNRs from pig blood ……………………… 129

Figure 2.3.2 Biodistribution of gold in major organs of pigs after one month ………………………………………………………… 130

Figure 2.3.3 Biodistribution of gold in major organs of pigs after six month …………………………………………………………. 131

Figure 3.1. NanoLipo procedure results in the removal of visibly more fat……………………………………………………………. 161

Figure 3.2. Heating profiles of tissue specimens for NanoLipo and Control pre-treatments ……………………………………………………... 162

Figure 3.3. (Screen Shot) Liberated triglyceride in lipoaspirated tissues are a hallmark of the NanoLipo procedure…..………………………….. 163

Figure 3.4. Smoothness within the treatment zone was used as a key marker to distinguish the NanoLipo treated specimen from Control specimens under blinded conditions ……………………………… 164

Figure 3.5. Micro computer-aided tomography images of the distribution of infused GNP solution within adipose tissue compared to non- injected Control …………………………………………………… 165

Figure 3.6. Absence of nitroblue tetrazolium staining (arrows) shows selective thermal damage of connective tissue membranes at the fat pearl level ………………………………………………………………. 166

Figure 3.S1. An ex vivo human tissue model was developed to test the ability of NanoLipo pretreatment to aid liposuction ………………………… 167

Figure 3.S2. Thin tissue specimens with skin removed after treatment ………………………………………………………….. 168

Figure 3.S3. Thicker tissue specimens exhibited large indentations in the skin after treatment …………………………………………………….. 168

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Figure 3.S4. A set of high Fitzpatrick skin-type specimen treated generated consistent results ………………………………………………….. 169

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LIST OF TABLES

Table 0.1. Significant ‘milestones’ in the development of photothermal therapy…………………………………………………………….. 20

Table 1.S1. Fluorescence lifetime data for fluorescein molecules in a polyacrylamide gel heated at various temperatures …...... 96

Table 1.S2. FLIM data of fluorescein within particles embedded in a polyacrylamide gel after different irradiation times at 980 nm ………………………………………. 97

Table 1.S3. FLIM data of fluorescein free in a polyacrylamide gel after different irradiated time at 980 nm ………………………………. 97

Table 2.1.1. Calculation of Parameters for Achieving Selective Photothermolysis…………………………………………..……… 105

Table 2.2.1. Device specifications for NanoLipo solution…………………….. 123

Table 3.1. Physical characteristics of human tissue specimens tested……… 170

ACKNOWLEDGEMENTS

First and foremost, I would like to thank the chair of my committee, Professor

Adah Almutairi, for her guidance and support, both academically and professionally, as well as her friendship over the years. I would also like to thank the other members of my committee: Professors Shaochen Chen, Yi Chen, Jesse Jokerst, Robert Sah, and previously Nathan Gianneschi, for their help and support. Other faculty members who have supported my academic endeavors include Professors Karen Christman, Adam

Engler, Joanna McKittrick, and others.

I would like to thank all of the members of the Almutairi lab, both past and present, for their help, peer-mentorship, and collaboration during my doctoral work. In particular I would like to thank Mathieu Lessard-Viger for ‘showing me the ropes’ in the lab during my first year, Ali H.Alhasan and Arnie Garcia for the hours we labored away over pig studies, Amy Moore for her support when she was lab manager, as well as various other lab members, including Alex Stubelius, Carina Arboleda, Sangeun

Lee, Vincent Tran, and others for their collaboration and the times we spent attending seminars or grabbing coffee. I would like to especially thank Viet Anh Nguyen Huu and Noah Johnson for their close friendship and the hours we spent pondering the mysteries of science and life over beer and spicy roast beef sandwiches, as well as Sha

He for the many intellectual and philosophical discussions we’ve had.

Looking back, one of the most positive aspects/benefits of being in Adah’s lab was the opportunity to make connections/network with various coaches/mentors outside of academia. Among these people, I would like to thank Khalid Almutairi,

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Stephen Keane, Richard Kuntz, Eboneé Williams, William J. Seare, Daniel Masters,

Rosibel Ochoa, and Phil Yeung, for all they’ve taught me.

Also, my work would not have been possible without the support of the wonderful faculty and staff at UCSD and UCSD affiliate institutions who maintain the facilities/equipment that were instrumental for my research. Among these people, I would like to thank Jennifer Santini from the CMME microscopy core, Professors

Paterno Castillo and James Day (and his post-doctoral fellow, Magali Porrachia), who maintain the ICP-AES and ICP-MS machines at the Scripps Institute of

Oceanography, Esther Cory of Professor Bob Sah’s lab for her help with micro computed tomography, and the UCSD Animal Care Program staff for their support during the various animals studies I conducted as part this dissertation work. I will also take the opportunity to thank my graduate program coordinators Charlotte Lauve

(who has since retired), Patrick Mallon, and especially Katherine Hamilton for the support regarding various administrative matters.

Finally, I would like to thank all of my friends, including Ryan O’neil, Eric

Hahn, Andre Pissarenko, Keyur Karandikar, Meghan O’Donell, Alex Phan, Joey Yim,

Chen Chen, to name a few, who have made these past five years a lot more bearable, as well as my wonderful, loving girlfriend, Yuling Luo, for her unending support.

Introduction, in full, has been accepted and recommended for publication with minor revisions in the Journal of Biomedical Optics. Wangzhong Sheng, Sha He,

William J. Seare, and Adah Almutairi; SPIE 2017. The dissertation author was the primary author of this work.

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Chapter 1, with the exception of the preamble, is in full, a reprint of the material as it appears in ACS Nano 2014. Mathieu L. Viger, Wangzhong Sheng, Kim

Doré, Ali H. Alhasan, Carl-Johan Carling, Jacques Lux, Caroline de Gracia Lux,

Madeleine Grossman, Roberto Malinow, and Adah Almutairi; ACS Press 2014. The dissertation author and M.L. Viger were the primary co-investigators and co-authors of this manuscript and contributed equally.

Chapter 2, is in part, a reprint of the material as it appears in Plastic and

Reconstructive Surgery - Global Open. Wangzhong Sheng, Ali H. Alhasan, Gabriella

DiBernardo, Khalid M. Almutairi, J. Peter Rubin, Barry E. DiBernardo, Adah PhD

Almutairi; LLW Journals 2015. The dissertation author was the primary investigator and author of this work.

Chapter 3, in part, is currently being prepared for submission for the American

Society for Plastic Surgeon’s Aesthetic Surgery Journal. Wangzhong Sheng, William

J. Seare, Ali H. Alhasan, Esther Cory, Paul Chasan, Robert Sah, Khalid Almutairi,

Barry DiBernardo, and Adah Almutairi; Oxford Academic 2017. The dissertation author was the primary investigator and author of this work.

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VITA

2012 Bachelor of Science, Cornell University

2013 Master of Science, University of California, San Diego

2017 Teaching Assistant, Department of Mechanical and Aerospace

Engineering, University of California, San Diego

2017 Doctor of Philosophy, University of California, San Diego

PUBLICATIONS

“Gold Nanoparticle-assisted Selective Photothermolysis of Adipose Tissue (NanoLipo)”. Wangzhong Sheng, Ali H. Alhasan, Gabriella DiBernardo, Khalid M. Almutairi, J. Peter Rubin, Barry E. DiBernardo, Adah Almutairi; Plastic and Reconstructive Surgery: Global Open 2014, 2:12, 283.

“Near-infrared-induced heating of confined water in polymeric particles for efficient payload release”. Mathieu L Viger*, Wangzhong Sheng*, Kim Doré, Ali H Alhasan, Carl-Johan Carling, Jacques Lux, Caroline de Gracia Lux, Madeleine Grossman, Roberto Malinow, Adah Almutairi; ACS Nano 2014, 8:5, 4815-4826. *Contributed Equally

“Application of time-resolved fluorescence for direct and continuous probing of release from polymeric delivery vehicles” Mathieu L Viger, Wangzhong Sheng, Cathryn L McFearin, Mikhail Y Berezin, Adah Almutairi; Journal of Controlled Release 2013, 171:3, 308-314.

FIELDS OF STUDY

Major Field: Materials Science and Engineering Studies in Chemical Engineering and Pharmaceutical Sciences Professor Adah Almutairi

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ABSTRACT OF THE DISSERTATION

Leveraging the Photothermal Effect for Cosmetic Applications

Wangzhong Sheng

Doctor of Philosophy in Materials Science and Engineering

University of California, San Diego, 2017

Professor Adah Almutairi, Chair

The photothermal effect is a phenomenon in which matter absorbs light energy and converts it into heat. This process can be leveraged to enable the application of light for biological applications, especially for otherwise benign light with wavelengths in the “optical window” regime of the electromagnetic spectrum between

700-900 nm that can penetrate deeply into tissue. The objective of this dissertation work was to apply and adapt the photothermal effect in the context of cosmetic

surgery in a safe and effective manner to aid in fat removal during liposuction. This thesis begins with a discussion of an area in which the photothermal effect has been applied extensively in biology, specifically, for the eradication of cancer, called photothermal therapy. Materials used for this application, known as photothermal agents, are briefly introduced. Light-matter interactions, including a variation of the photothermal effect called selective photothermolysis, are discussed in detail. Chapter

One introduces an example where the photothermal effect was observed to be the key underlying mechanism enabling the release of encapsulated payloads from inherently non-light responsive polymer particles in the context of drug delivery. In Chapter

Two, starting from basic engineering principles, specifically the heat equation, the photothermal effect is derived and mathematically modeled. Subsequently, the ability to achieve selective and confined photothermal heating in a target pigmented region was tested in a physical model involving a thermoresponsive polymer as an indicator to allow for simple visualization. After successful demonstration of these concepts, parameters such as photothermal agent concentration and manner of light exposure were optimized, and subsequent studies were performed in an ex vivo porcine tissue, in an in vivo porcine model, and finally, in an ex vivo human adipose tissue liposuction model. This work culminated in a single-blind study in human adipose tissue comparing photothermally assisted liposuction (dubbed NanoLipo), to a control involving saline injection and near infrared light laser-assisted liposuction absent gold nanoparticles. The results achieved in this work demonstrate the significant potential of this promising technology in making a clinical impact.

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Introduction:

A Review of the Progress towards Achieving Heat Confinement: the Holy

Grail of Photothermal Therapy

I. ABSTRACT

Photothermal therapy (PTT) involves the application of normally benign light wavelengths in combination with efficient photothermal agents that convert the absorbed light to heat to ablate selected cancers. The major challenge in PTT is the ability to confine heating and thus direct cellular death to precisely where photothermal agents are located. The dominant strategy in the field has been to create large libraries of photothermal agents with increased absorption capabilities, and to enhance their delivery and accumulation to achieve sufficiently high concentrations in the tissue targets of interest. While the challenge of material confinement is important for achieving “heat and lethality confinement”, this review article suggests another key prospective strategy to make this goal a reality. In this approach, equal emphasis is placed on selecting parameters of light exposure, including wavelength, duration, power density, and total power supplied, based on the intrinsic properties and geometry of tissue targets that influence heat dissipation, to truly achieve heat confinement. This review highlights significant milestones researchers have achieved, as well as examples that suggest future research directions, in this promising

2 technique, as it becomes more relevant in clinical cancer therapy and other non-cancer applications.

II. INTRODUCTION

Cancer, characterized by uncontrolled cellular growth, is one of the leading causes of death in the United States.1 Nearly 1.7 million newly diagnosed cases of cancer were reported in 2015, of which about 35% of affected patients died from the disease.2 Since cancer incidence rates track closely with our growing aging population, these numbers are expected to continue to increase in the future. Cancer is not just one disease; it is extremely complex and can be found in many forms. Through continuous and enormous effort, we have gained a better understanding of cancer, although there is still much we do not know. What is clear is that cancer treatments almost universally require a multi-pronged approach to improve overall disease prognosis.3-5

Currently, the primary prong of cancer therapy is surgery, which is often complemented by the sub-prongs of chemo6 and radiation therapy7 to bolster effectiveness. Nevertheless, resistance to treatment often leads to cancer recurrence, especially in the cases of certain malignant cancers.8 Additionally, the highly invasive nature, limited indication and success, substantial cost, and pain of surgery, as well as potential complications with expected and unexpected side-effects of adjuvant therapies, all lead to reduced quality of life and reduced life expectancy for surviving patients.9-11 Photothermal therapy (PTT) has recently emerged as a potential area of promise for this multi-pronged approach to cancer therapy.12-16 In PTT, the core concept is to apply otherwise benign near infrared (NIR) light that can reach several

3 centimeters beneath the skin surface17 in combination with highly efficient photothermal (PT) agents placed non-invasively within tissues to create hypothermia to damage cancers only where the agents are placed, offering unprecedented safety, specificity, and selectivity, in eradicating cancer.

Historically, hyperthermia through heating has proven to be extremely effective in inducing damage, especially to biological tissues.18 In general, there is a time-temperature dependent relationship such that the higher the temperature, the more rapidly the tissues become damaged, with the likelihood of damage exponentially doubling above 43 °C.19 In tumors, the capacity for heat dissipation is significantly reduced due to the tortuous blood vessel network the cancers form, making them even more susceptible to thermal treatments.20 Despite this difference, care must be taken to precisely define where the heating occurs so that the surrounding healthy tissue is not accidentally damaged in the heating process. This concept is called heat confinement, and is often not achievable with traditional heating methods.21 Light-induced heating with the aid of PT agents has the advantage of affording more control over where the selected tissue heating occurs. However, the general drawback of poor tissue penetration-depth in light-based biological applications still must be overcome. Thus, the overall problem of selectively delivering heat to, and confining it within tissues, is one of the major challenges of photothermal therapy.

In order to achieve light-induced heat confinement, highly efficient PT agents must first be sequestered inside the tumor at sufficient concentrations to absorb the deep-penetrating NIR light. Early incarnations of PTT involved the direct injection of

4 concentrated mixtures of PT agents into tumor xenografts followed by continuous and long-term exposure to high powered lasers.13, 15, 22 These initial outcomes often caused non-selective and uncontrolled heating to extreme temperatures, not unlike that of direct heating methods. These treatments resulted in immediate cellular death of both cancerous and non-cancerous tissues in the exposed regions. The mechanism of damage was that of well-known tissue hyperthermia, dominated by necrosis accompanied by irreparable denaturation of proteins and other biological macromolecules.23 While some of these proof-of-concept models demonstrated the ability of PTT to reduce tumor volume or prolong the survival of treated animals,24 clinically, they are generally not feasible or practical. Additionally, as is usually the case with introducing any foreign substance into biological systems, it is also important to mitigate any potential toxicity associated with these exogenous agents.25,

26 This will be no easy feat, and will likely require careful consideration of the multi- factorial interactions between diverse and complex disease states found in cancer biology with the various engineered nanomedicine platforms delivering the cancer therapies. Thus, the delivery of PTT agents to the desired targets of interest is a second major challenge.

This review focuses on the current state of development of this promising technology over many years, its recent advances, as well as obstacles that must still be overcome. We begin by introducing the theory behind the photothermal conversion process and examining examples where the photothermal effect was successfully applied towards heating of naturally existing pigments. In this context, the principle of selective photothermolysis is presented, where specific parameters of light exposure

5 are chosen based on the size and geometry of biological targets of interest to achieve heat confinement in pigmented tissues. We propose this principle as a strategy to potentially achieve more selective light-induced heat confinement in PTT. Some of the leading, highly efficient, man-made photothermal agents that enable this technique are then summarized, although we opted to take a succinct approach due to the vast amount of research published on this topic. We also briefly comment on the potential sources of toxicity of popular photothermal agents as well as strategies that have been employed to address these concerns. In the subsequent section, we expand on the topic of drug delivery and provide an overview of cases in which the PT effect was leveraged synergistically with diagnostic or other therapeutic agents, to garner theranostic or multi-modal therapeutic nanomedicine platforms, respectively, for cancer therapy. Finally, we speculate on critical avenues for further research and development that could make this technique more clinically relevant.

Figure 0.1. Schematic highlighting factors crucial to achieving material and heat confinement in photothermal therapy; Current research directions involve alternative parameters of light exposure to modulate tissue heating, or alternative methods of delivery, such as intravenous or local sustained-delivery, as opposed to bolus intratumoral injections that are typically considered the norm in photothermal therapy. Abbreviations: (Application of Laser Light) λ = wavelength, W/cm2 = power density, t = exposure time; (Prospective Photothermal Agents) ε = extinction coefficient, c = concentration, η = photothermal conversion efficiency.

III. LIGHT TISSUE INTERACTIONS

The interaction of light with matter is extremely complex.27-31 In general, when light impinges on matter, it interacts with the electromagnetic field of the atoms in the material and as a result, many complex phenomena occur. The interaction depends on important factors such as the wavelength and power density of light, but just as importantly, it depends on the properties of the interacting species itself. Different species can reflect, diffract, polarize, absorb and disperse light. Additionally, many complex nonlinear optical effects may occur.30 Nevertheless, in describing the PT effect, especially in diffuse anisotropic media such as biological tissue, many of these

7 observed phenomena can be categorized as absorption or scattering.32, 33

Simplistically, absorption describes the amount of incident energy that is captured, whereas scattering describes any energy that is not.33-35 Absorption can be further distinguished into absorption by the target species or by unwanted species. Light which is competitively absorbed or scattered is said to be attenuated, or reduced in intensity. Additionally, depending on the PT absorber, not all of the electromagnetic energy may be converted to heat. For example, instead it could be used to drive chemical reactions in applications such as photodynamic therapy.36, 37 Absorbed light may also be re-emitted as another wavelength of light, e.g. fluorescence.38, 39 Good PT agents will not only capture a large percentage of the incident light, but also effectively transform it into heat, and are said to have high PT conversion efficiencies.40-42 The most commonly used model to describe absorption is the Beer-

Lambert law43:

퐴휆 = 휀 푐 푙

where Aλ is the absorption (optical density) of the material at wavelength λ, ε is the molar extinction coefficient, c is the molar concentration, and l is the path length that the light travels through the absorbing medium. The Beer-Lambert law is quite accurate in describing light-matter phenomenon in cuvette settings,44 although actual observed behavior starts to deviate from this law if any major assumptions are not met.

Unfortunately, biological tissue is one such medium. It contains several non- independently absorbing pigmented species that may competitively absorb incident light in the highly inhomogeneous and turbid tissue environment.45, 46 Therefore, the

8 expected trends defined by this equation should only serve as a general reference when assessing absorbers for photothermal applications.

The Biological Optical Window

One of the biggest limitations in light-based biological applications is penetration depth.47-49 The strongest naturally existing species contributing to light attenuation by surface tissues are melanin, hemoglobin, and water.50 Melanin is a pigmented protein produced by melanocytes concentrated in the basal layer of the epidermis. Melanin has a local absorption maximum around 350 nm that decreases gradually progressing into the visible wavelengths of the electromagnetic spectra and beyond.51 As a result, longer wavelengths of light can penetrate deeper into tissues.

The amount of melanin in skin varies across different populations of people, and increases with sun exposure.52 Melanin content in skin is typically classified by the

Fitzpatrick Scale of skin colors, with level I being very pale through level VI being very dark.53, 54 In all skin types, some of the penetrating light energy is absorbed by melanin. In darker skin types, melanin is present in significantly higher concentrations so even longer light wavelengths will have difficulty penetrating beyond the epidermis.43, 44

Like melanin, hemoglobin also exhibits large absorptions at longer wavelengths, with a sharp drop at approximately 600 nm for both the oxygenated and deoxygenated forms.55 Dragan et al. measured the absorption of whole human blood sandwiched between two glass slides and found that at 800 nm, the absorption of a 1 mm thick sample, which could be representative of a small blood vessel, was 0.5

9 optical density (OD).56 Described in terms of the Beer-Lambert law, this optical density means that nearly 70% of incident light is absorbed. Fortunately, hemoglobin’s contribution to light attenuation is more variable than melanin’s, since its absorption is dictated by the size and number of blood vessels, which, generally, are distributed non-homogenously in any given tissue. Thus, it appears that longer wavelengths of light beyond 600 nm would be preferred for transcutaneous biological applications.

However, there exists an upper limit of around 1,000 nm, where another strongly absorbing “pigment”, water,57 can significantly limit light penetration into tissue. Since human epidermis is on average 64.6% water,58, 59 in addition to being the primary harbinger of melanin, it is the largest barrier to entry for light in the infrared regime of the EM spectrum. Because of this, techniques that have shown promise in achieving heat confinement, such as laser interstitial thermal therapy (LITT), a minimally invasive, contrast-agent-free, laser photothermal technique that relies on absorption of water-resonant wavelengths of light by the water molecules in tissue to directly ablate benign and malignant lesions, requires the percutaneous insertion of a fiber-optic catheter directly into tissues to bypass the skin barrier to achieve therapeutic efficacy.60 It would be desirable to be able to use wavelengths of light that can more readily bypass the optical skin barrier. Incidentally, when the absorption profiles of melanin, hemoglobin, and water are overlaid, a region of minimal absorption between 600 nms and 1000 nms emerges. This range of wavelengths is commonly referred to as the biological optical window (Fig. 0.2).61, 62 Light in the middle of this near-infrared (NIR) region (around 800 nm) can penetrate through skin

10 and up to several centimeters into tissue,17 making NIR absorbing species especially attractive for transcutaneous photothermal applications.39, 63

Figure 0.2. A. Wavelengths between 700-900 nm are minimally absorbed by endogenous species and are able to penetrate deepest into tissue.49 Reproduced from Avci P et al. Low-level laser (light) therapy (LLLT) in skin: stimulating, healing, restoring. Seminars in cutaneous medicine and surgery. 2013;32(1):41-52. © 2013 B. Schematic demonstrating underlying principles behind laser hair removal technologies. Reproduced from Dr. Gayne Dolyan Descornet, “How does laser hair removal work”, Woman Health and Life. C. Thermal relaxation times (TR) for spherical structures, e.g. tumors, of various sizes, according to selective photothermolysis theory. Treatment times in PTT are typically well above this regime.

Selective Photothermolysis to Overcome Attenuation by Endogenous Species

Although endogenous species are responsible for the limited penetration depth of light, they have also been successful targets of photothermal applications in the tangential field of cosmetics. Photothermal targeting of melanin, in the form of laser hair removal,64, 65 and photothermal targeting of hemoglobin, in the treatment of port- wine stains66 and varicose veins,67, 68 and have been widely and successfully applied in the clinic for decades. These applications take advantage of the fact that in biological structures such as hair follicles and blood vessels, there is a locally higher concentration of melanin and hemoglobin respectively, than is present within surrounding tissues. If the targeted structure absorbs the defined wavelength at least twice as much as its immediate surroundings, it is possible to selectively heat the target structure to damage it by optimizing the way in which light energy is applied to achieve heat confinement. Cooling of the skin surface can help to counteract some of the heating of the melanin, especially in darker skin types, to allow for the preservation of the sensitive dividing epidermal cells in the basal skin layer, while the deeper hair shafts are heated much more, relatively. In general, for transcutaneous applications, the higher the difference in absorption between the target structures and the surrounding tissues, including the skin, the easier it is to achieve heat confinement.

This process, first proposed by R.R. Anderson,69, 70 is known as selective photothermolysis (SPT).

In SPT, light is applied in a controlled pulsatile manner over a large region.

Here, the strongly absorbing target structure will heat up much more than surrounding less-absorbing structures. The time it takes for the target structure to exponentially

12 decay from the maximum temperature reached is described by the thermal relaxation time, τR, which is dictated primarily by the size, geometry, and local density of the target structures. A core concept in SPT is to apply short, repeated bursts of light energy absorbed preferentially by the target structures before heat can dissipate or diffuse to adjacent tissues (relaxation), resulting in the continual build-up of heat, thereby creating a selective-heating effect. The overall exposure time should be less than the thermal relaxation time of the target structure to ensure that the surrounding areas experience only mild temperature elevations. Another caveat is that pulses must not be too short in order to avoid any cavitation effects which may be induced by the high powers achieved as a result of the short pulses.71

Heat transfer within tissues is governed by intrinsic properties of tissue such as thermal diffusivity, which is on the order of 10-3 cm2/s. Approximating this term to be the same for both the target and its surroundings, Anderson et al. calculated the τR for spherical targets with a wide range of sizes. For example, a spherical structure 1 mm

69 in diameter has a τR of 0.3 seconds . By taking into consideration other intrinsic properties of tissue such as density and specific heat, appropriate powers and durations of light exposure could be calculated for SPT of target structures of various shapes and sizes (Fig. 0.2).70 Altshuler et al. later expanded on this theory and proposed that due to the inexactness of the estimates, in practice, a buffer zone slightly larger than the absorbing structure should be defined with a threshold temperature set at this boundary that should not be exceeded. These engineering principles formed the foundation, and continue to serve as the guiding principles for the development of laser–based cosmetic technologies.72

Although SPT applied towards endogenous species does not overcome the problem of penetration depth, the unique ability to confine heat to specific targetable pigmented structures, while sparing the surrounding tissue, could prove extremely powerful in the context of practical photothermal therapy. In the case of photothermal heating of exogenous species, the major advantage, and challenge, is that one could create an absorbing structure by virtue of where the photothermal absorber is introduced. If we examine the optical aspects and apply the principles of SPT theory towards PTT, we find that in the majority of demonstrations of photothermal therapy, heating is rarely confined. Researchers typically continuously expose target regions to high laser powers, resulting in uncontrolled damage leading to necrosis of the treated region and beyond.73 Instead, we might propose that the thermal relaxation time could be estimated for different sized target tumors or metastases to determine the optimum parameters that should be used to limit the extent of photothermal heating and improve the specificity of PTT. Extending Anderson’s calculations further, spherical structures

2 cm and 3 cm in diameter would have τR of approximately 2 and 5 minutes, respectively (Fig. 0.2). SPT for structures of any larger sizes may not be as applicable since rarely do tumors studied in a research setting exceed this size, and tumors of this size are likely to be resected clinically74. We have found examples in the literature where researchers have carefully considered how light is best applied based on the geometry of the target structures in order to maximize the specificity of the light-based application.5, 16, 72, 75-77 Furthermore, our own research lab is attempting to leverage this phenomenon in the context of liposuction to modulate heating of adipose tissues to aid in their removal by plastic surgeons.72

IV. PHOTOTHERMAL MATERIALS PLATFORMS

Ultimately, the selective photothermolysis of tumors is enabled by the large differential light absorption between the target and the surrounding tissues. To that end, in this section we briefly summarize some of the leading man-made photothermal agents that have been applied for PTT.

Exogenous Agents for Photothermal Therapy

Many candidates have been explored for the photothermal killing of cancer both in vitro and in vivo (Fig. 0.3), with varying degrees of success. Arguably, the most widely applied class of photothermal agents is gold-based nanostructures, whose synthesis, characterization, standardization, and diverse properties have been extensively studied.78-84 The mechanism of light-to-heat conversion is a localized form of surface plasmon resonance (LSPR).85 This process is extremely efficient, and as a result, these materials exhibit one of the highest extinction coefficients reported,86, 87 on the order of 109 M-1cm-1. Since the absorbing LSPR peak and PT conversion efficiency are strongly dependent upon the geometry of the nanostructure,88 researchers have engineered gold nanoparticles of many shapes and sizes to absorb preferentially within the biological window regime for photothermal therapy.

The simplest of all the geometries is probably the solid gold spherical nanoparticle. Around the early 2000s, the El-Sayed group synthesized and studied the plasmonic properties of solid gold nanospheres and found that although they were able

15 to control the particle’s LSPR peak by tuning its size, the peak was still less than 600 nm,89 making it not practical as a candidate for PTT. To overcome this limitation, the group investigated the plasmonic properties of gold nanorods (AuNRs) and found that the anisotropic geometry of these nanostructures resulted in the establishment of a second LSPR peak corresponding to the longitudinal direction of the elongated rod- like structure.12 Simulations suggested that by varying the aspect ratio of the nanorods, the LSPR peak could be precisely tuned to absorb a wide range of wavelengths in the

NIR regime.90 It took several years for a reproducible technique to synthesize gold nanorods to be developed,91 and several more years before AuNRs were applied for photothermal therapy in an in vivo setting.13 In this study, Dickerson et al. demonstrated the ability of plasmonic photothermal therapy to treat cancer in a squamous cell carcinoma xenograft model. The authors injected AuNRs either intratumorally (abbrev: IT; 15 μL, ODλ=800 = 40) or intravenously (abbrev: IV; 100 μL,

ODλ=800 = 120) followed by 10 min of laser exposure, and compared the outcome to a negative saline control. Both procedures outperformed the control with the authors observing a >57 % tumor resorption for the IT-injected group versus 25% for the IV- injected group. However, the IV injected group, which likely relied on passive accumulation via the enhanced permeation and retention (EPR) effect92 for delivery, required twenty times more AuNRs in combination with a laser power density almost double that of the intratumorally injected group, yet resulted in less tumor destruction

(25% vs 57%, respectively).13 This is not surprising, given that a recent analysis of the literature found that in less than 1% of the time did injected nanoparticles successfully penetrate solid tumors.93

Around the same time, the Halas and West groups utilized a slightly different approach to engineer gold nanoparticles which could absorb in the NIR regime in the form of gold nanoshells (AuNS). AuNS are created by depositing a thin layer of gold on a dielectric nanoparticle substrate, most commonly, silica.94, 95 Similar to nanorods, the LSPR peak of nanoshells could be controlled by tuning the thickness of the shell, with the absorption maxima of thinner shells more red-shifted than thicker shells.96

Hirsch et al.. applied these 100 nm gold-silica nanoshells in an in vitro and in vivo setting for plasmonic photothermal therapy.15 In the in vitro study, the authors incubated human breast epithelial carcinoma cells in a serum-free medium containing

4.4 x 109 PEGylated nanoshells per milliliter. After one hour, nanoshells were rinsed away and a coherent NIR light source at 820 nm, 35 W/cm2 was used to irradiate the cells for seven minutes to induce photothermal damage. Compared to the area treated with the laser in the absence of nanoshells, the photothermally treated regions saw significantly more cell death. For the in vivo demonstration of PTT, 20-50 μL of gold- silica nanoshells were intratumorally injected 5 mm into the tumor volume and subsequently exposed to 820 nm light at 4 W/cm2 at a 5 mm spot diameter for up to six minutes. The authors observed temperature elevations of 37.4 ± 6.6° C within 4-6 minutes of light exposure which induced irreversible damage to the tumor. However, examination of the thermal profiles as a function of depth revealed that the maximum temperature elevation was found at just 1 mm below the skin surface, even though the nanoshells were injected 5 mm deep into the tumor.15 It is likely that the AuNSs were distributed throughout the injection site, so photothermal ablation of the surface

17 tissues with the high powers applied may have created a carbonaceous layer which would have exacerbated light attenuation as the photothermal treatment proceeded.22

Figure 0.3. Examples of photosensitizers applied for photothermal therapy; A. Graphene,97 B. Indocyanine green-loaded polymer nanoparticles,98 C. Melanin-human serum albumin nanoparticles,99 D. Gold nanorods,100 E. Multi-walled carbon nanotubes,76 F. Gold nanocages,100 G. Gold nanostars,101 H. Gold nanoshells,94 and I. MoS2 nanoplates.102 A., B., D., E., F., I. are reprinted with permission from the American Chemical Society, © 2010, 2013, 2013, 2009, 2013, 2014, respectively. C. is reprinted from permission from John Wiley & Sons, Inc. © 2013. G. is not subject to U.S. Copyright. H. is reprinted with permission from The Royal Society of Chemistry. © 2015.

Many more photothermal agents with large effective absorptions have been developed, albeit they faced many of the same challenges described above. In terms of plasmonic photothermal agents, additional nanostructures with exotic geometries have

18 been synthesized and studied, including gold nanocubes, nanotetrahedron, and nanoicosahedrons,103 nanocages,104 nanostars,44, 101, 105, 106 and nanohexatripods.100

None-gold based plasmonic nanostructures such as quantum dots,107 porous palladium nanoparticles,108 and silver nanoparticles109 have also been developed for plasmonic photothermal therapy. In addition to plasmonic photothermal nanomaterials, carbon- based agents such as single-walled110 and multi-walled76, 111 carbon nanotubes,112 and graphene sheets,97, 113-115 which rely on the vibration of excited, delocalized sp2 electrons in the lattice of the carbon structure for heat generation,76 have also been developed. Various small-molecule organic dyes absorbing in the NIR window such as indocyanine green (ICG),22 cypate,116, 117 and porphyrins,118, 119 have also been applied for PTT. In the case of these agents, various engineered nanostructures such as micelles14, 120-125 and biopolymer-based nanoparticles38, 126, 127 need to be employed to prolong the residence time of the agents to make them applicable for PTT, because they are prone to photo-bleaching128 and are rapidly cleared from the body.129

Sometimes, the polymer particles themselves can be applied directly as photothermal agents, as is the case with the conducting polymer, polypyrrole.85, 130, 131 Other creative researchers have even engineered melanin-loaded polymeric nanoparticles for use in

PTT.99, 132

The plethora of existing agents that have been reported can attest to the incredible amount of research effort dedicated to this topic. Unfortunately, there is inherent variability between different photothermal platforms relating to their intrinsic properties such as their extinction coefficients and PT conversion efficiencies, which are useful for determining the absorption cross-section of the agent used.133 However,

19 additional factors such as the wide range of laser parameters and the diverse cancer models used, must be considered just as carefully. As a result, a comparison between different photothermal agent platforms, or even across the same photothermal agent platform82 is difficult, and does not yield meaningful conclusions (Tab. 1). For this reason, we opted to abstain from summarizing these studies in detail. Instead, we chose to highlight a few key milestone examples, as well as some trends in the development of this technique in Tab. 1, some of which we have already discussed above.

Systemic Toxicity and the Challenge of Delivery

Although novel agents are still being reported today, one issue that has persisted is the delivery of PT agents. PTT involving intratumoral routes of administration are still consistently outperforming intravenous ones. The necessity for direct injection into tumors can be problematic because of size, location, and the need to visualize the tumor and its borders with costly equipment, often in an operating or procedural room setting. Unfortunately, intravenous injection of PT agents, in addition to suffering from the disadvantage of poor accumulation in target tissues, also poses the potential problem of systemic toxicity. Recently, del Rosal et al. reported on the

134 application of NdVO4 nanocrystals (NCs) for photothermal therapy. The authors observed an absorption cross-section of 2.0 x 10-19 cm2, which was eight orders of magnitude lower than gold’s reported absorption cross-section86 of 2.93 x 10-11 cm2.

As a result, much more NCs were needed to achieve the desired absorption. The authors injected the NCs intratumorally at a concentration nearly 2,000 times the maximum amount which they showed to be safe in vitro in order to achieve minimal photothermal heating that was only slightly more (~8 °C) than the laser-alone group after 4 minutes of laser exposure. If the IT injected amount was diluted by the whole blood volume of the animal, which would be the case if the agent had been injected

IV, the concentrations reached would still be an order of magnitude above the documented safety threshold shown by the authors. We then have to consider whether the photothermal agent itself could induce acute toxicity, not to mention the challenge associated with administering such a large amount of material. Ultimately, the general determining factor for toxicity of foreign agents is whether the concentrations injected

22 are above the known acceptable thresholds for the therapeutic agents in question.26

This means that PT agents with high absorptions but a low toxicity threshold could exhibit concerns, as well.

In general one of the biggest concerns of nanoparticles is the fate of the nanoparticles after it is introduced into the body.25, 26, 135 Ideally, injected NPs, especially those of an inorganic nature, should be renally cleared, which means that their hydrodynamic radii need to be less than 5.5 nm.136-139 For this reason, the majority of IV injected inorganic nanoparticles are not approved by the FDA for systemic use clinically.137 Unfortunately, almost all of the useful inorganic nanoparticle-based photothermal agents studied are above this accepted size range.

When Sonavane et al. investigated the biodistribution of gold nanoparticles of a broad size range (15 nm up to 200 nm), they found that 15 and 50 nm NPs were able to penetrate the blood brain barrier in mice.140 This result is alarming since the majority of gold nanoparticles used for PTT are in this size regime.141 Nevertheless, some studies have shown that gold nanoparticles induced no acute toxicity when uptaken by human cells.142

The geometry of nanoparticles may also influence their biodistribution. Black et al. investigated biodistribution of gold nanostructures with varying shapes and found that gold nanorods and cages were distributed throughout tumors while nanospheres and nanodisks were only around the tumor periphery.143 In a similar study, Wang et al. found that nanohexapods exhibited even higher cellular uptake, and more importantly, lower cytotoxicity compared to nanorods and nanocages of similar sizes,100 suggesting that geometry may also be implicated in the potential toxicity of

23 nanoparticles. It is critical to ensure that injected nanoparticles are stable and structurally sound, especially because they can be easily heated to extreme temperatures due to their high effective absorptions. Recently, Goodman et al. reported on the surprising in vivo instability of hollow gold nanoshells (HGNS).144

Although their HGNS were observed to be completely stable under laser irradiation in solution, biodistribution studies suggested that the nanoshells were heavily fragmented in vivo, with the authors attributing the cause of the instability to residual silver ions remaining after synthesis of the HGNS. The findings are in line with what other researchers have observed for gold nanorods, which are known to fragment or shorten upon exposure to high laser power densities, although smaller particles were found to be more resistant to fragmentation due to coupling to the environment.108, 145 Overall, these results are not unexpected, however, since systems often do not behave the same in aqueous dispersions as they do in vivo.44 From these examples, we see that inorganic nanostructures still face many challenges before they can be translated into useful clinical therapies.146

Organic photothermal agents have not fared much better, mostly for the same reasons of biodistribution, stability, and toxicity related to processing.25 For example, the high aspect ratio of carbon nanotubes, reminiscent of the geometry of carcinogens such as asbestos, led to long-term residence in the lungs of mice studied.147 Residual metallic impurities from nanotube synthesis and undesirable aggregation are also commonly attributed to toxicity of carbon nanostructures.135 Various strategies have been employed to functionalize photothermal agents with passivating polymers. Most commonly, polyethylene glycol (PEG) is employed to impart a stealth-like character

24 to these agents to reduce their toxicity.148, 149 The attachment of targeting moieties to photothermal nanoparticle platforms for functional groups overexpressed in certain cancers may help improve the local accumulation of photothermal agents within the tumor to increase treatment specificity and hopefully their effectiveness, while also reducing their systemic toxicity (Tab. 0.1).150

Finally, the challenges of systemic delivery could be avoided altogether, if local routes of administration were developed and perfected. Conde et al. recently reported on a well-characterized triple-combination photothermal, chemo, and gene therapy applied at the tumor location in the form of a hydrogel patch with or without tumor resection (Fig. 0.4).77 In their prophylaxis study without pre-resection, a hydrogel patch loaded with gold nanoparticles and nanorods was implanted 15 days after tumor induction, with multiple photothermal treatment events starting on day 18.

The remarkable part of this study was in its methodology. Instead of choosing traditional PTT methods involving a single exposure at a high power density, the authors opted for multiple treatment events of short durations (120 s) applied once per day over the span of several days. Although the authors did not discuss their reasoning for the short exposure times, these times corresponded well, and likely not by coincidence, to the thermal relaxation times we calculated in Sec. 2 for the corresponding tumor sizes in this study. This prophylactic triple combination treatment, which is especially relevant for non-resectable tumors, was able to achieve complete tumor remission whereas the control group with tumor resection only saw

40% tumor recurrence. Furthermore, treatment following intratumoral or systemic injection of the gold nanomaterials resulted in only 60% and 40% tumor reduction,

25 respectively. This result is consistent with most of the existing studies comparing IT and IV modes of administration.151, 152 Strategies such as this could help reduce the chance for systemic toxicity, although it would be important to ensure that the photothermal agents did not extravasate elsewhere. However, choosing known low- toxicity agents with well-documented excretion pathways can obviate some of these concerns.

Figure 0.4. A triple-combination local delivery platform developed by Conde et al. in which chemo, gene, and photothermal therapy are applied in conjunction to eradicate cancer. For the photothermal therapy component of the treatment, the authors performed four sessions of laser exposure for120 seconds each (1W, 808 nm) on days 18, 19, 20, 25 to reach temperatures as high as 60 °C to damage colorectal cancer in a xenograft tumor model in mice.77 Reprinted with permission from Macmillan Publishers Ltd: Nature Materials; J. Conde, N. Oliva, Y. Zhang and N. Artzi, Nat. Mat. 15, 1128–1138 (2016), © 2016.

In this section, we saw that a significant amount of effort has been dedicated towards finding novel agents with high effective absorptions for photothermal therapy.

However, the potential systemic toxicity of photothermal agents and their route of administration are factors which must also be carefully considered regardless of the effective absorption of the agent. In Sec. 2, we suggested one possible way the parameters of light exposure could be optimized is to limit the heating to tumors, while sparing surrounding tissues. In this section, we saw a recent example of researchers already moving forward in that direction. In the next sections, we discuss in more detail, how nanomedicine153-156 has helped, and may continue to help mitigate attenuation effects resulting from improperly introduced photothermal agents and address toxicity concerns157, 158.

V. NANOMEDICINE ENABLED SYNERGISTIC PHOTOTHERMAL

THERAPY PLATFORMS

Photothermal therapy, as we have presented so far, involves light-induced heating to high temperatures with the assistance of strongly absorbing photothermal agents to ablate cancerous tissues. Apart from being used to directly ablate tumors cells, light induced heating is useful in many other processes related to cancer therapy, as well. In this section, we present some examples in which photothermal agents are integrated with other cancer therapeutic or imaging modalities to garner multifunctional theranostic nanoparticle delivery platforms to enhance the outcome of cancer treatment.

Benefits of Mild Hyperthermia

Hyperthermia involves heating of biological tissues to none physiological temperatures. While temperatures above 43 °C are employed to ablate tissue, heating to milder temperatures could also prove to be beneficial in the context of cancer therapy.159-161 One of the first responses by the body to local hyperthermia is enhanced blood flow to the region to dissipate heat. As mentioned in Sec. 1, the tortuous nature of the improperly formed blood vessels found in tumors creates a selective mechanism for heat build-up. This property, combined with the typically lower threshold of tumor cells to tolerate heat, naturally make tumors more susceptible to thermal damage.20

Increased blood flow also contributes to amplifying the EPR effect to enhance the amount of material accumulating in the tumor,162 such as circulating photothermal agents or chemotherapy drugs, perhaps all loaded into and delivered by a single nanoconstruct. Mild temperature elevations may also trigger subcellular events to render cells susceptible to damage through apoptosis.23, 111, 161, 163 In addition, it may drive a shift towards anaerobic metabolism in cells, resulting in higher oxygen availability to the tissue environment to overcome the hypoxic conditions typically found in tumors that have limited the efficacy of oxygen-dependent cancer therapies.131, 164 Unfortunately, one major undesirable heating-associated effect is the activation of immunological responses, including the upregulation of heat shock proteins (Hsp)165 that lends itself to protecting cancer cells from being damaged.166, 167

In fact, studies have shown that the ability of cancers to upregulate Hsp production is well-correlated with the aggressiveness of the cancer,165 resulting in various gene therapy-based strategies to block Hsp production to improve cancer treatment

28 outcomes.117, 168 Nevertheless, by integrating photothermal capabilities into theranostic nanoparticle platforms, researchers could use the mild heating regime to initially amplify the efficacy of other treatment modalities, followed by controlled photothermal ablation at higher laser power densities, creating a two-pronged attack, to help to ensure tumor destruction.

Photothermally Triggered Payload Release

Controlled drug release is important in cancer chemotherapy because it enables potent chemotherapeutic drugs to kill the cancers while minimizing systematic toxicity. Many physical, chemical and physiological signals have been explored to act as triggers for controlled release of chemotherapy drugs. As discussed previously, light is a unique trigger but is still limited by the availability of effective photochemical processes. In addition to synergistic enhancement of other therapies, heating, generated by photothermal agents with the irradiation of light, can expand the utility of light-controlled release to non-photoresponsive systems. In general, various drug carriers such as liposomes,169 polyelectrolyte capsules, polymeric nanoparticles,75 and even blood-cell-mimicking capsules170 all become less stable and more permeable upon heating, leading to the therapeutic cargo eluting from their interior. Moreover, the burst nature of heating-induced release can effectively enhance the efficacy of treatments such as chemotherapy in the appropriate environment. Following this concept, various studies using laser-triggered controlled release systems involving photothermal transducers and thermoresponsive systems or phase-changing materials as release substrate have been demonstrated.98, 171, 172 In one example, Kim et al.

29 reported an elegant system utilizing gold nanoshells for the photothermally-triggered cargo release. They first load the cargo into liposomes and encapsulate the liposomes with a thin layer of the nanoshells. The integrity of the AuNSs make the liposome leakage-free under ambient conditions. Upon the exposure to NIR light, the gold nanoshells generated heating via the photothermal effect and deform in situ, triggering the release of cargo from the interior.173 In a similar example, Zheng et al. designed a synergistic chemo/photothermal therapy platform containing the common chemotherapy agent, doxorubicin (DOX), and ICG loaded into a lipid-polymer nanoparticle system (DINP).98 The authors showed that the DINPs were stable in the absence of light, but with a single, eight-minute treatment of 808 nm laser at 1 W/cm2, they were able to trigger the release of DOX and also heat the tumor to more than 50

°C to ablate it. Compared to the all of the controls, the DINPs were the most effective and seemed to be the only system able to suspend tumor growth for up to 17 days.

Although not applied to cancer, our own research group recently demonstrated the exciting potential for using selective photothermal heating of water nanodomains confined inside non-light responsive polymer microparticles, as a means to induce payload release.75

Photothermally Enhanced Cellular Uptake

Photothermal heating could also be leveraged to increase cellular uptake, as well.174 It is known that many nano-sized objects are uptaken by cells via receptor- mediated endocytosis, macropinocytosis, and other membrane-related processes. A number of studies have shown that the permeability of cell membrane can be

30 significantly enhanced when the environmental temperature is increased to 43 oC, resulting in the uptake of more therapeutic agents in the targeted regime.175 Provided that most therapeutic agents require a minimum concentration to be effective, enhancing localized uptake to cells may potentially increase the utilization percentage of delivered therapeutic agents, while lowering the overall dose that needs to be administrated.44 Following this concept, Sherlock et al. demonstrated that ultra-small iron-cobalt (FeCo)-core-graphitic-carbon-shell nanoparticles loaded with the DOX exhibited accelerated cellular uptake upon NIR laser irradiation.176 Particularly, the

FeCo-graphitic-shell nanoparticles with a large absorption band can heat the targeted cells from 37oC to 43oC at laser power densities as low as 0.3 W/cm2. The 6 oC incremental temperature increase in the cell media induced a two-fold increase in the cellular uptake of the FeCo-graphitic-shell nanoparticles loaded with DOX. Similarly,

Tian et al. showed that the uptake of nano-graphene, which has proven to be an effective carrier for gene transfection without any other aids, can be enhanced three- fold with NIR exposure.113 The Lapotko group also demonstrated an intriguing pathway of applying the photothermal effect for enhanced drug delivery. They found that plasmonic noble metal nanoparticles exposed to short laser pulses could generate transient vapor bubbles which could adhere to cell membranes and disrupt them upon the burst of these bubbles. The authors demonstrated that the burst event can be either used as trigger for drug release from the carrier or to enhance endosome escape of the engulfed drugs.177 These strategies to enhance cell membrane permeability are promising, and are likely to become more popular in translational nanomedicine

31 research because many of the drug delivery vehicles could be engineered to have these capabilities.

Combination Photo-therapies

Combinatorial therapies which integrate several therapeutic approaches together, such as the PTT/Chemotherapy system described above,176 have generally proven to be more effective than stand-alone techniques.178 Photodynamic therapy

(PDT) is an oxygen-dependent light-based cancer therapy very similar to PTT in which photosensitizer molecules transfer the energy of the absorbed photon to the surrounding molecular oxygen to locally generate toxic singlet oxygen molecules, instead of heat, resulting in local toxicity to effectively damage the structure and function of cancer cells. Like PTT, PDT has many of the same advantages of high specificity and low systemic toxicity compared to conventional chemotherapy.

However, an inevitable drawback that has hindered the efficacy of PDT in tumor treatment is its requirement for oxygen. The hypoxic conditions found inside tumors are exacerbated by the depletion of the already limited oxygen supply from the oxygen-consuming reactions of PDT. It is also important to isolate PDT agents inside drug delivery vehicles because these highly sensitive species could become toxic with exposure to even ambient levels of light.179 Nevertheless, due to the high level of similarity between PDT and PTT, it made sense to design systems with both modes of therapy integrated. Jang et al. demonstrated an interesting example by using gold nanorods as both carrier and quencher for the PDT agent, aluminum phthalocyanine tetrasulfonate (AlPcS4). AlPcS4 was assembled onto the surface of gold nanorods and

32 quenched by the surface plasmon resonance. Therefore, they were non-phototoxic even when exposed to light of proper wavelengths. Only when the gold nanorods were used to generate heat under laser irradiation were the photosensitizers released from the surface, thus becoming phototoxic. In a similar example, gold nanostars were surface functionalized with another PDT agent, chlorin e-6, to perform dual

PDT/PTT.105 In concept, PDT initially performs according to its design parameters due to the presence of oxygen, but as oxygen becomes depleted, PTT takes over as the predominant mode of tumor damage, and in combination, conspires to doubly damage the tumor.

Photothermal Therapy/Photoacoustic Imaging

Adding diagnostic functions to nanoparticle drug delivery systems can be useful for tracking their location, as well as help provide information and feedback regarding the therapy. Photoacoustic (PAc) imaging, a natural “derivative” of the photothermal effect, is likely the easiest to implement along with PTT because it utilizes the heat generated from laser irradiation to produce acoustic waves that can convey both the anatomic and functional information about the tissues probed.180, 181

Huang et al. reported a biodegradable gold nanovesicle composed of a simple design comprised of PEG-b-PCL tethered to AuNPs for photoacoustic imaging and photothermal therapy, simultaneously, while under the irradiation of the same laser.182

183 More recently, Yong et al. and Song et al. reported using WS2 quantum dots and

184 two-dimensional Co9Se8 nanosheets, respectively, as agents for both PTT and PAc imaging. These materials belong to a group of novel PT agents classified as metal

33 chalcogenides,102, 185-187 which exhibit broad absorption spectra, high extinction coefficients, and photothermal conversion efficiencies, as well as low cytotoxicity.188

Moreover, these nanostructures can be easily functionalized to enable other popular imaging modalities, such as fluorescence imaging, magnetic resonance imaging,189 and computed tomography,183 without affecting the photothermal conversion process, thereby making them even more powerful as platforms for photothermal applications.

Figure 0.5. Photothermal theranostic systems such as indocyanine-human serum albumin nanoparticles are capable of multi-modal imaging, as well as multi-modal therapy to allow for the tracking of the photothermal agents in conjunction with combination therapy to improve cancer treatment outcomes.16 Adapted with permission from the American Chemical Society © 2016.

VI. PROGRESS AND FUTURE PERSPECTIVES

Case Study: ICG Then and Now

Through the examples presented thus far, it is evident that photothermal therapy has progressed significantly since the idea of light-induced heating to kill cancer was first conceptualized. To provide some perspective, we present, to our knowledge, one of the earliest examples of photothermal therapy, or “chromophore- enhanced tumor destruction”, as it was referred to in 1995 (Tab. 0.1).22 After having characterized their system in vitro,190 Chen et al. subcutaneously injected 0.5 to 1% indocyanine green in free dye form, into mammary fat-pad tumors of rats, exposed them to various laser power densities, and monitored the animals for up to 9 days.

Several important conclusions were drawn from this study which reiterates many of the challenges for PTT presented thus far. First, the authors found that laser powers in the 3-5 W range worked unequivocally better than higher powers in the 10-15 W range, since the higher powers rapidly denatured and burned tissue, creating a carbonized tissue layer which effectively shielded NIR light from penetrating into tissue, suggesting that the parameters of light exposure must be carefully considered.

Second, the authors conceded that the model showed efficacy, likely because intratumoral administration of the dye into the capsule-like tumor allowed the chromophore to be sequestered long enough to allow for photothermal treatment, suggesting that encapsulation, delivery, and accumulation of photothermal agents, especially small molecule dyes like ICG, via alternate routes of administration, might be critical to the success of PTT. Lastly, the authors found that although tumor

35 destruction was apparent in their period of observation, a sufficient amount of tumor cells remained viable and was able to perpetuate the tumor growth in all of groups. A study extending out to 30 days revealed that at time of death, the laser-ICG treated tumors were still about half the size of tumors treated by lasers only191. This result suggested that PTT alone, at least as a single treatment, may not be sufficient in eradicating a tumor. Cancer has proven to be a complex and formidable challenge for medical sciences. Multi-pronged strategies involving therapies such as PTT at specified times and parameters, following surgical interventions such as reduction of tumor bulk combined with other modalities such as chemo, radiation, immunotherapy, and stem cell targeting, are likely needed to improve treatment outcomes.

In this review we have highlighted significant advances in nanomedicine; over the past two decade, researchers in many fields have helped us understand and overcome many challenges. As an example, a simple yet elegant system, comprising

ICG as the primary photosensitizer, was recently reported by Sheng et al. (Fig. 0.5).16

Indocyanine green, in addition to being applied for PTT, and by extension, photoacoustic imaging, can also serve as a photosensitizer in PDT when exposed to low-power short-term laser exposures.192, 193 The authors attached ICG to human serum albumin (HSA) using thiol chemistry, and the system subsequently self- assembled into a theranostic nanoparticle platform (HSA-ICG NP).16 By formulating

ICG into NPs, they were able to enhance its uptake into cancer cells in addition to improving the stability of the dye in terms of its rapid clearance and photo-bleaching, while also enabling fluorescence imaging. The authors intravenously injected the

HAS-ICG-NPs and in terms of imaging, were able to track the NPs through dual

PAc/fluorescence imaging. In terms of therapy, they were able to deliver the ICG-NPs, as well as achieve dual PDT/PTT, by applying an 808 nm laser continuously for 5 min at 0.8 W/cm2. In the xenograft model, this combined therapy was able to completely destroy the tumor and ensure the survival of all the animals in the study group, out to

50 days, while only stunted tumor growth was observed in all of the control groups

(particles alone, laser alone, or interval exposure), with no animals surviving beyond

25 days. Interestingly, when the same study was carried out in an orthotopic tumor model with the same type of cancer (4T1 breast carcinoma), interval exposure was able to prolong the survival of the animals an additional 10 days, out to 35 days, while other control groups of particles alone or laser alone performed about the same as the xenograft study, demonstrating at least some degree of effectiveness akin to selective photothermolysis with interval exposure. The reduced effectiveness for the interval group compared to the continuous exposure group could be attributed to the long rest times between exposure events (1 min exposures with 5 minutes rest intervals) in the interval exposure group. Overall, compared to the work by Chen and colleagues, this work involved only a single photosensitizing agent, and it was still able to accomplish tumor destruction through dual-therapy, and also provide diagnostic information by tracking the injected nanoparticles.

VII. CONCLUDING REMARKS

The development of various engineered NIR-light-absorbing nanoscale agents will continue to attract attention in the field of cancer therapy and beyond. In this

37 review, we presented many examples of how photothermal therapy has incrementally evolved over the years. The idea of applying the photothermal effect to treat cancer arguably arose from the desire to expand upon the therapeutic benefits of mild hyperthermia therapy. As vast libraries of powerful photothermal agents were developed, PTT research deviated from mild hyperthermia therapy and evolved into ablation therapy, where high laser powers were used to heat up these highly efficient agents to uncontrollably damage tissue. However, as researchers began to understand the implications of uncontrolled heating, attention began to shift towards carefully controlling the light-exposure to achieve more defined thermal damage. There was also a shift back towards applying the photothermal effect to achieve mild hyperthermia to synergistically bolster other treatment modalities in cancer therapy.

Although our review is far from being exhaustive, present efforts support this claim, and indicate continued development and refinement of some of the more complex theranostic photothermal platforms incorporating photothermal agents of all types for use in multiple imaging and therapeutic applications. The addition of targeting capabilities to improve accumulation of nanoparticle therapeutic platforms within tissues of interest, as well as selective heat confinement through controlled light exposure, are strategies vital to expanding the ability of this platform to assist preclinical and clinical research. The field of nanomedicine has helped tremendously on the material-confinement front. Looking forward, careful manipulation of the parameters of light exposure, perhaps through concepts borrowed from selective photothermolysis theory, could push a shift in paradigm in this field that may one day lead to truly selective heating of tumors, which if achieved, would bring photothermal

38 therapy one step closer to becoming a vital prong in the clinical cancer therapy approach.

VIII. ACKNOWLEDGEMENTS

The authors would like to thank Dr. Paul Chasan and Dr. Amy Moore for their comments and suggestions during editing of this work. The authors would also like to thank NIH R01EY024134 and Air Force office of Scientific Research (AFOSR)

FA9550-15-1-0273 for funding. The Introduction chapter, in full, has been accepted and recommended for publication with minor revisions in the Journal of Biomedical

Optics. Wangzhong Sheng, Sha He, William J. Seare, and Adah Almutairi. SPIE 2017.

The dissertation author was the primary author of this work.

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Chapter I:

Near Infrared-Induced Heating of Confined Water in Polymeric Particles

for Efficient Payload Release

1.0 CHAPTER ONE PREAMBLE

In this chapter, I describe my findings in a project involving light-activated release of various payloads from inherently none light-responsive polymer micro and nanoparticles. In a previous study investigating the use of light-sensitive polymers for the release of payloads from particles formulated by electro-spray, it was discovered that the control group involving microparticles comprised of poly-lactic-co-glycolic acid (PLGA), a none light-responsive polymer, released significant amounts of encapsulated payload in response to exposure to water-resonant wavelengths of light at 980 nm. Along with my collaborator, M.L. Viger, we explored this phenomenon further and discovered that nano-domains of water were formed as a result of the hydration process when microparticles were placed into an aqueous environment.

These nanodomains of water not only plasticized the polymer and lowered its glass transition temperature (Tg), but were also preferentially heated upon exposure to light, inducing a rise in temperature to above the plasticized Tg of the polymer, leading to payload release. We found this mechanism to be highly universal, and to be applicable

51 for the release of payloads with a wide range of hydrophobic characteristics, as well as from a wide range of none light-responsive polymers.

As is often the case, scientific discoveries such as this result from chance or perceived failures in preceding experiments. This chapter introduces the chance encounter that led to my first exposure to the photothermal effect, specifically, to a variation of selective photothermolysis that would form the underlying basis of this dissertation. The remainder of Chapter 1, below, is in full, a reprint of the material as it appears in ACS Nano 2014. Mathieu L. Viger, Wangzhong Sheng, Kim Doré, Ali

H. Alhasan, Carl-Johan Carling, Jacques Lux, Caroline de Gracia Lux, Madeleine

Grossman, Roberto Malinow, and Adah Almutairi, ACS Press 2014. The dissertation author and M.L. Viger were the primary co-investigators and co-authors of this manuscript and contributed equally.

1.1 ABSTRACT

Near-infrared (NIR) light-triggered release from polymeric capsules could make a major impact on biological research by enabling remote and spatio-temporal control over the release of encapsulated cargo. The few existing mechanisms for NIR- triggered release have not been widely applied because they require custom synthesis of designer polymers, high-powered lasers to drive inefficient two-photon processes, and/or co-encapsulation of bulky inorganic particles. In search of a simpler mechanism, we found that exposure to laser light resonant with the vibrational absorption of water (980 nm) in the NIR region can induce release of payloads encapsulated in particles made from inherently non-photo-responsive polymers. We

52 hypothesize that confined water pockets present in hydrated polymer particles absorb electromagnetic energy and transfer it to the polymer matrix, inducing a thermal phase change. In this study, we show that this simple and highly universal strategy enables instantaneous and controlled release of payloads in aqueous environments as well as in living cells using both pulsed and continuous wavelength lasers without significant heating of the surrounding aqueous solution.

1.2 INTRODUCTION

Light-triggered release from polymeric particles is a key tool for delivering encapsulated molecules with a high degree of spatial and temporal control.194-199

Given the deep penetration depth and low attenuation of near infrared (NIR) radiation in biological tissues, systems that respond to these wavelengths (750-1300 nm) are particularly attractive.200 Though many technologies facilitating release of encapsulated cargo from polymeric carriers using NIR radiation have been developed, all of them originate from only a handful of photochemical and photo-physical strategies, which present drawbacks that limit the extent of their application. For example, gold or other metal nanostructures are often incorporated to generate heat upon absorption of NIR light, loosening the surrounding polymer matrix to release therapeutics.201, 202 This approach is limited by the controversial biocompatibility of metal nanostructures and the considerable amount of heat generated that may be detrimental to surrounding tissues.203-205 Alternatively, carriers can also be formulated from polymers containing photo-responsive modalities that respond through chemical changes such as isomerization, oxidation, dimerization, and bond cleavage.206, 207 In

53 order to make use of the inherently low energy of NIR light, these mechanisms all require simultaneous two-photon absorption, necessitating the use of high-powered and focused pulsed NIR lasers, with energies sometimes exceeding the damage threshold of biological tissue.208, 209 Recently, the limitations of two-photon photochemistry have been overcome by coupling photosensitive polymer carriers to upconverting nanoparticles (UCNPs) that sequentially absorb multiple NIR photons and convert them into higher-energy photons in the UV region.210, 211 Since excitation photons do not have to be absorbed simultaneously, UCNPs could enable triggered release in response to more biologically benign power densities. However, UCNPs contain heavy rare-earth elements that may prove to be toxic in vitro and in vivo.212

These significant barriers to widespread application encourage further work to discover metal-free, universal strategies for remote-controlled photo-release using NIR light.

In search of such a mechanism, we found that exposure to NIR laser light resonant with the vibrational overtone of water at 980 nm (Figure 1.1A) could induce release of an encapsulated payload from particles made of inherently non-light- sensitive polymers, e.g. poly(lactic-co-glycolic acid) (PLGA). We hypothesize that this universal release mechanism arises from the interaction between the electromagnetic radiation and confined nano-domains of water formed upon hydration of the particles (Figure 1.1B).213, 214 Upon exposure to 980 nm NIR light, water molecules absorb optical energy through vibrational transitions and the excitation energy is rapidly converted into heat.215-217 In the case of bulk water, the heat is rapidly dissipated through the whole aqueous environment via convective heat

54 transfer.218 However, for small water droplets trapped inside polymer microstructures and partially shielded from the external bulk water, the absorption of resonant electromagnetic energy increases the internal droplets’ temperature. To equilibrate the resulting temperature gradients, part of the generated heat is hypothesized to be lost by conductive heat transfer to the surrounding constituting matrix (Figure 1.1C), thus locally increasing the temperature of the polymer. Localized heating of the polymer matrix above the polymer’s glass transition temperature would induces a phase change from a rigid, glassy state to a compliant, rubbery state, activating diffusion of encapsulated payload out of the particles.

Using real-time monitoring of steady-state fluorescence, we found that this strategy can be readily applied to precisely control the release of encapsulated contents of varying hydrophobicities from different polymer particles in aqueous environments as well as in living cells. We chose PLGA as the main polymer matrix because it is

FDA-approved, widely used for a variety of biomaterials applications, and extensively studied as a drug carrier.214, 219 In the NIR region, water has vibrational absorption bands around 1950, 1450, 1200, and 980 nm.220 While the absorption coefficient increases at longer wavelengths, we chose 980 nm light for this study because wavelengths above 1100 nm were not readily available to us. However, the other overtone vibrational absorption bands should be capable of inducing release in a similar fashion. With fluorescence lifetime imaging microscopy (FLIM), we measured the temperature inside polymeric particles with high spatial resolution, which revealed increases in temperature upon irradiation at 980 nm. This simple and generalizable release mechanism represents a major innovation in the field of remote-controlled

55 release, as it can be applied to a large number of systems with no inherent light sensitivity.

Figure 1.1. Schematic representation of NIR-induced release mechanism: (A) Absorption spectrum of water in the NIR region; (B) Formation of isolated nanodomains of water in the polymeric structure; (C) Release of encapsulated molecules following photothermal heating of water droplets inside the polymer particles.

1.3 RESULTS AND DISCUSSION

NIR Irradiation at a Water-Resonant Wavelength Induces Payload Release from

Polymeric Capsules

PLGA polymer capsules incorporating model compounds of varying hydrophobicities (fluorescein, Nile blue, Nile red, IR780) were produced by electrospray. Representative dark field images and optical epifluorescence of the produced dye-loaded PLGA particles are shown in Figure 1.S1 and Figure 1.S2, respectively. The wavelength selective response was first explored by monitoring the change in fluorescence intensity of a suspension of fluorescein-loaded particles (size:

0.5 ± 0.1 µm, dye loading: ~10% w/w) before and after 15 min irradiation at 50 nm increments of wavelengths, ranging from 780 nm to 1030 nm (Figure 1.2A).

Fluorescein emission increases upon release because it fluoresces more intensely in a polar environment, and like most organic dyes, self-quenches when its relative concentration exceeds ~10-3 M, such as when it is entrapped within particles (0.2 mol of fluorescein per kg of PLGA).221 A strong increase in fluorescence intensity was obtained only when the irradiation wavelength coincided with the maximum water absorption band, i.e., 980 nm (Figure 1.2A, open circles); profiles following irradiation at non-resonant wavelengths were very similar to those of non-irradiated control samples (Figure 1.2A, solid circles). The rate of light-induced release of fluorescein was also monitored by fluorescence spectroscopy (see Methods section for experimental details). When excited at 980 nm (1 W), the amount of released material rapidly increased in the first 30 min, and saturated at around 60 min (Figure 1.2B, open circles). The quantity of released material was found to be proportional to the amount of energy (number of photons) provided to the system, where a cumulative 47,

76, and 89% of fluorescein, corresponding to 1.51 ± 0.06 µg, 2.44 ± 0.01 µg and 2.84

± 0.05 µg, was released from the PLGA particles after 10, 30 and 60 min of

57 irradiation, respectively. Fluorescein release could be monitored visually as the suspension changed from colorless to bright green (Figure 1.S3A). Epifluorescence microscopic images (Figure 1.S3B) further confirmed NIR-triggered release: Before

NIR irradiation, fluorescein was clearly loaded inside PLGA particles; after irradiation, fluorescence appeared as irregularly shaped dried patterns, rather than well-defined domains. In contrast, particles not exposed to light (off-state) or irradiated at non-resonant wavelengths (i.e., 800 nm and 900 nm) released negligible amounts of dye over the course of several minutes (Figure 1.2B). This wavelength selective response demonstrates the necessity of resonant photon interactions with water. The observed low background leakage of non-irradiated controls suggests that the fluorescein molecules remain well encapsulated inside the PLGA particles. We calculated a 25-fold increase in the rate of fluorescein release upon exposure to 980 nm NIR light (on/off ratio).

SEM images of irradiated particles (Figure 1.S4) showed no obvious changes in particle morphology (i.e., size, shape, and surface texture); even after 90 minutes of

980 nm pulsed irradiation at 1 W, no degradation was observed. To further confirm that 980 nm irradiation does not degrade the polymer, PLGA particles were run through gel permeation chromatography (GPC) before and after irradiation at 980 nm for 1 hr (pulsed laser, 1 W). The GPC data did not show any change in retention time of the polymer and/or the appearance of small molecule peaks at longer retention times corresponding to products of degradation (Figure 1.S5). As this mechanism induced release of the encapsulated payload without disruption of the polymer matrix, possible mechanisms such as optical cavitation or thermally-induced ester degradation

58 could be ruled out. More specifically, optical cavitation, or the formation of bubbles of vaporized gas in response to light, results from aqueous absorption of short laser pulses which generate a dielectric breakdown at the focal point, creating a plasma that expands and produces an acoustic shockwave.222 Since laser-induced cavitation requires intense laser light (> 200 nJ/pulse) focused at a spot size of 0.5 – 5 µm,223 and laser beams with pulse energy ≤ 60 nJ and spot size diameters ≥ 2 mm were used in this release study, it is unlikely that the observed photo-release resulted from optical cavitation. Although no clear evidence of particle deformation was observed by SEM, irradiation did appear to reduce particle stability, as reflected by collapse and agglomeration (Figure 1.S6). Aggregation could result from the particles being heated

224, 225 above their glass transition (Tg) and reaching a more rubbery state. Nonetheless, the temperature increase of the bulk solution was less than a few degrees above ambient. We found that temperature elevations of the bulk solution from ambient temperature were dependent on laser power and sample volume. Irradiation at 980 nm

(1 W, 15 min) of 0.25, 1, and 2 mL aliquots of a particle suspension resulted in maximum temperature elevations of approximately 10, 5, and 2 °C, respectively. In order to avoid unnecessary variability in the release experiments due to bulk heating, we can minimize temperature elevation of the bulk by working with large sample volumes.

As mentioned previously, the photo-physical process involved in this release scheme occurs through the excitation of an overtone vibration absorption of water, and so, is not limited to delivery of 980 nm light as short, focused light pulses, in contrast to simultaneous two-photon absorption processes. Thus, efficient light-triggered

59 release can also be achieved at lower excitation powers using a more economical and biologically relevant continuous wave (CW) laser set-up (Figure 1.2C). By varying the excitation power from 170 mW to 350 mW, we found that the rate of release depends on the average photon energy used to excite PLGA particles. As expected, the release of fluorescein is faster at higher CW power, with an on/off ratio of 18 when irradiating at 350 mW compared to 14 and 5 when irradiating at 260 and 170 mW, respectively.

Since robust control over release is desired, we assessed the activatable nature of this release mechanism. A permanent effect would cause continuous release after irradiation ends, while an activatable effect would involve slowed release profiles following termination of laser irradiation.226, 227 A fluorescein-doped PLGA particle suspension was irradiated repeatedly using NIR light (pulsed laser, 980 nm, 1 W) for 5 min, followed by 15-min intervals with the laser turned off. A rapid increase in fluorescein release from the PLGA particles was observed upon NIR irradiation, but the release rate practically decreased to its initial rate when the NIR irradiation was switched off (Figure 1.2D). A similar on/off release ratio was observed over multiple consecutive exposures. This stepwise triggered release indicates that the PLGA particles retain their integrity upon NIR irradiation. The small payload release observed in the off-state may be attributed in part to water molecules diffusing into the space created by the loss of fluorescein from the particles, which in turn encourages outward diffusion.213

Figure 1.2. NIR-induced release occurs only at water-resonant wavelengths. (A) Action spectrum of the fluorescence intensity of a suspension of fluorescein-doped PLGA particles after 15 min irradiation (open circles) or 15 min at room temperature (solid circles) superimposed on the water absorption spectrum (black trace) (N = 3); (B) Cumulative fluorescein release after irradiation of a suspension of fluorescein- doped PLGA particles with pulsed laser light (1 W) at 980 nm (open circles), 900 nm (solid pentagon), and 800 nm (open triangles). Control (no irradiation, solid squares); (C) Cumulative fluorescein release after irradiation of a suspension of fluorescein- doped PLGA particles with 980 nm CW laser light at varying powers; (D) Stepwise triggered release from fluorescein-loaded PLGA particles using six cycles of non- irradiation (15 min) and five cycles of irradiation (5 min, 980 nm, pulsed laser, 1 W). Fluorescein acquisition parameters: λex = 480 nm, λem = 520 nm.

Since the PLGA particles are capable of encapsulating both hydrophobic and hydrophilic compounds at high efficiency, the ability of this photothermal process to release small payloads of varying hydrophobicities was investigated. IR780, Nile red, and Nile blue (in order of increasing polarity) were loaded by electrospray into spherical PLGA capsules (~10% w/w) of 0.7 ± 0.1 µm, 1.4 ± 0.3 µm, and 1.2 ± 0.4

µm in size, respectively (dark field microscopy: Figure 1.S1, fluorescence microscopy: Figure 1.S2). For all aqueous suspensions of dye-loaded PLGA particles,

NIR exposure (980 nm, pulsed laser, 1 W) resulted in release of the dyes over time

(Figure 1.S7). These findings revealed that this photothermal process is capable of sustaining release of both hydrophilic and hydrophobic compounds. A noticeable trend in the kinetics of release related to differences in dye polarity could be observed, i.e., more polar dyes were released faster. This difference could result from variations in water absorption: Hydrophilic content facilitates water penetration in polymeric carriers and leads to the creation of more porous and swelled polymer networks,195 while more hydrophobic compounds hinder diffusion of water into the structure.213, 228

Variations in release rate could also relate to the energy needed to induce diffusion out of the matrix, which should be lower for hydrophilic than hydrophobic compounds, since hydrophobic compounds would have higher affinity for the hydrophobic carrier.

Also, as we have shown release with cargos that cannot be excited at 980 nm through two-photon absorption (Nile blue, IR780), the possibility that this light-assisted mechanism of release derives from thermal relaxation of two-photon excited states can be discarded.

Role of Water in Inducing Payload Release

In the proposed mechanism, entrapped water plays a crucial role in payload release by absorbing the NIR light to induce localized heating inside the particles. To test this hypothesis and rule out the possibility of direct absorption of optical energy by the polymer, we compared the release profiles of fluorescein from PLGA particles in deuterated water (D2O) to those in deionized water (H2O). A comparison of the absorption spectra of H2O and D2O shows that D2O does not absorb significantly at

980 nm (Figure 1.3A). All of its vibrational transitions are shifted to lower energy by the increase in isotope mass; the vibrational band of H2O at 980 nm is shifted to 1300

229 nm in D2O. Therefore, light-induced release in D2O, if observed, could only be the result of photonic absorption by the polymer capsules. Fluorescein-loaded particles suspended in D2O did not release cargo upon irradiation; the kinetic profiles matched those of a non-irradiated sample (Figure 1.3B). This is consistent with previous results showing that the laser photonic densities (≤ 1 W) used in this study do not result in polymer degradation (Figure 1.S5) and/or changes in PLGA particle morphology (Figure 1.S4), and supports the conclusion that release requires strong absorbance of NIR light by water (Figure 1.2).

Figure 1.3. Metal-free photothermal release requires efficient vibrational absorption by confined water. (A) Absorption spectrum of deionized (H2O) and deuterated water (D2O); (B) Cumulative fluorescein release after irradiation of a suspension of fluorescein-doped PLGA particles dispersed in H2O (solid circles) or D2O (open pentagons) with pulsed laser light at 980 nm. Control (no irradiation, solid squares). Laser power: 0.8 W.

In the first stage of hydrolytic degradation, particles made of biodegradable polymers such as PLGA swell and take up water213, 214 to form domains of water in the hydrophobic nanostructure (Figure 1.1B).230 While in this state, polymers’ thermal behavior is affected, as the confined water acts as a plasticizer inside the polymer matrix by increasing the free volume of the polymer chains and consequently depressing the glass transition temperature (Tg); this process is called hydroplasticization.231 Using differential scanning calorimetry (DSC), we measured the Tg of the PLGA particles before (dry Tg) and after dispersion in water (wet Tg).

Based on the DSC measurements, the presence of a saturated water environment induced a depression of the Tg of PLGA particles from 42.6 °C to 27.9 °C, thus confirming the presence of water inside the capsules. The measured wet Tg value was

64 in good agreement with the value predicted by a model proposed by Tsavalas et al.,231 i.e., 27.1 °C, which we used to calculate a 5% water content weight fraction in the particles. PLGA carriers are complex 3D systems made of interconnecting pores and channels of different size and tortuosity distributed throughout the entire volume of the spheres in which water can penetrate rapidly.232 Pre-existing pores (in contrast to dynamically formed pores created upon degradation) found in PLGA particles of similar composition and size present diameters between 3 – 20 nm;232 water likely fills these pores.233

Since this mechanism of release relies on excitation of confined water, it should work on any polyester system without inherent light sensitivity at 980 nm. This versatility could have significant impact because linear polyesters (e.g., PLGA, poly(lactic acid) and polycaprolactone) constitute one of the most important classes of synthetic biodegradable polymers234, 235 due to their biodegradability and biocompatibility, as well as their easily tailored physicochemical and mechanical properties.236, 237 We examined the applicability of this mechanism of release to other polyesters using Nile red-loaded polymer capsules (see inset in Figure 1.S9 for molecular structures).238 Fluorescence microscopy images of the resulting electrosprayed particles (polyester 1: 0.9 ± 0.1 µm, polyester 2: 2.3 ± 0.5 µm) can be found in Figure 1.S8. For both aqueous suspensions of dye-loaded particles, NIR exposure (980 nm, pulsed laser, 0.9 W) resulted in release of the Nile red molecules over time (Figure 1.S9). Both sets of particles presented similar stability in the absence of irradiation, suggesting that the confined water-assisted mechanism of release applies to multiple polyesters.239

Another important class of delivery vehicles are particles made from derivatized polysaccharides (e.g. methacrylated hyaluronic acid,240 chitosan,241 and modified dextran242, 243), which have attracted considerable attention for their tunable degradation rates, high processibility, and biocompatible byproducts. To study the applicability of this photo-physical mechanism to induce release from sugar-based polymers, we formulated Nile red-doped acetalated dextran particles using electrospray (size: 1.6 ± 0.5 µm, see Figure 1.S8 for fluorescence microscopy images and size distribution histogram of the resulting particles) and tested the responsiveness of these particles to irradiation at 980 nm (pulsed laser, 0.9 W). Similarly to polyester particles, non-irradiated particles showed low background leakage (Figure 1.S9C, solid squares), whereas irradiation at 980 nm induced a strong increase in payload release (Figure 1.S9C, open circles). These results support the inference that remote- controlled photo-release using a water-resonant wavelength is applicable to a large number of systems with no inherent light sensitivity.

Evaluation of the Internal Temperature of the Particles upon 980 nm Irradiation

Fluorescent molecular thermometer represents a promising tool for intra- particle thermometry, as it functions at the molecular level and, thus, would be effective in monitoring temperature within micron-sized domains. The strong effect of temperature on the fluorescence of molecular probes has led to various sensing strategies based on changes in emission wavelength and/or fluorescence intensity.244,

245 In addition, excited-state lifetimes of fluorescent molecules are intensely affected by temperature, showing in most cases shorter lifetimes at higher temperatures.246 This

66 temperature dependence results from the increasing importance of non-radiative decay rates at higher temperatures. In this study, FLIM was used to extract intra-particle temperatures by comparing fluorescein lifetime to a standard curve generated by direct heating of the dye. This thermometric methodology allowed temperature measurements inside polymeric particles with high spatial and temperature resolution.

To image individual particles for an extended period of time, particles were stabilized in polyacrylamide hydrogels (see Methods). We first acquired a calibration curve relating free fluorescein lifetime and temperature by manually heating a polyacrylamide gel containing fluorescein from 20 °C (room temperature) to 70 °C and acquiring lifetime decay curves using a Time-Correlated Single Photon Counting

(TCSPC) system on a spectrofluorometer. The average excited- av.) were then extracted (see Table 1.S1) av. av.

(X°C) – av. (20°C)) as a function of temperature (Figure 1.4A). The excited-state lifetime deceased with increasing temperature following a negative linear relationship with a calculated sensitivity of -8 picoseconds (ps)/°C. For the FLIM measurements, gels containing fluorescein-doped PLGA particles were placed in a reading chamber containing water. Single particles or small particle aggregates were irradiated at 980 nm (pulsed laser, 10 mW) in a raster scan motion to ensure complete irradiation of the entire particle/aggregate. FLIM images were acquired every 5 min and an average lifetime was extracted from each image by integration of all pixels. Fluorescein’s absorption/emission wavelengths and fluorescence lifetime are relatively insensitive to the environment’s polarity, therefore changes in lifetime recorded for the free dye in the hydrogel (Figure 1.5A) should correlate well to those recorded in the particles.

Since fluorescein can be excited at 980 nm through two-photon absorption, the decay curves could be acquired while particles were continuously irradiated. Measurement of particle temperature via lifetime of encapsulated fluorescein would not otherwise be possible, as the particles’ internal temperature equilibrates with that of the surrounding solution rapidly after irradiation. Lifetime measurements reveal increases in the lifetime of fluorescein within particles following cessation of irradiation (Figure

1.S10).

The average lifetime of fluorescein within particles, as extracted from FLIM images, clearly decreases with increasing irradiation time (Figure 1.4B, open circles, see also Table 1.S2 for lifetime data av. = -267 ± 31 ps, tirr. = 25 min) following an exponential trend. According to the change in lifetime versus temperature calibration curve in Figure 1.4A, the average internal temperature of PLGA particles reached 34, 45 and 54 °C following 5, 15 and 25 min irradiation, respectively, well above the wet Tg of the dye-doped PLGA particles. On the other hand, the lifetime of free fluorescein embedded in polyacrylamide gels did not decrease upon irradiation

(Figure 1.4B, solid circles, see also Table 1.S3 for lifetime data). This data suggests that the absorbed optical energy can be efficiently dissipated throughout the bulk hydrogel, preventing excessive localized heating. The observed selective heating of polymeric particles versus the bulk aqueous environment can be attributed to the enormous volume difference between the two; the heat capacity of water within nanodomains is lower than that of bulk water. Also, it has been shown that water molecules confined in hydrophobic micro/nanostructures possess a much higher thermal

68 conductivity than bulk water.247 The space constraint causes an increase in collision frequency, which enhances energy and heat transfer.

Example FLIM images of small aggregates of particles within irradiated hydrogels show the consistency of decreasing fluorescein lifetime with increasing irradiation at 980 nm (Figure 1.4C). Though dye-doped particles presented different average lifetime values before irradiation (likely due to variations in particle size and dye loading), the change in lifetime of each particle as a function of irradiation time was comparable regardless of its initial value. These images also clearly show that thermal changes happened exclusively inside particles; no noticeable changes in the surrounding hydrogel were observed. Furthermore, because of the high spatial resolution of the FLIM system, regions of intense changes in lifetime were discernible throughout the whole polymer matrix, which suggested that water-rich areas capable of generating substantial heat were present within the PLGA particles. Using the relation found in Figure 1.4A, the corresponding average internal temperature was added to each FLIM image. These temperature changes were more modest than in other thermally induced release mechanisms, especially compared to those involving gold nanostructures, which are reported to be well above 250 °C.203, 204 Compared to gold, for example, water has a much higher heat capacity (Cp, water = 4.186 J/g·°C;

Cp, gold = 0.129 J/g·°C),248 which means that it requires more optical energy to raise its temperature by the same amount. Ultimately, because the Tg of most biodegradable polymers is below 60 °C,249 the resulting localized heating is enough to soften polymer matrices and induce release of the encapsulated payloads. An attractive option to increase the photon-to-heat conversion efficiency, and hence conserve the

69 same release efficacy at lower laser powers, would be to exploit water absorption bands with greater absorptivity (i.e., 1200, 1450, and 1950 nm). This could prove especially useful for release experiments in highly scattering environments, as the laser beam would quickly broaden and lose its high peak power while travelling through the sample. Also, the polymeric capsules’ sensitivity towards 980 nm excitation should be related to the amount of confined water in the polymeric capsules, which is greatly influenced by polymer composition (e.g. proportion of hydrophilic functionalities, potential to form intra-polymer hydrogen bonds, elastic modulus).230,

231 Thus, studying the effect of polymer composition (e.g. introducing hydrophilic functionalities) on particles’ sensitivity to NIR light represents an exciting future direction and logical continuation of this work.

Figure 1.4. Fluorescence lifetime imaging (FLIM) reveals an increase in particles’ internal temperature. (A) Change in average lifetime (picoseconds, ps) of free fluorescein in a polyacrylamide gel caused by heating to various temperatures; (B) Changes in average lifetime of free fluorescein in a polyacrylamide gel (solid circles) (N = 4) or encapsulated in PLGA particles and embedded in a polyacrylamide gel (open circles) (N = 6) upon irradiation at 980 nm (pulsed laser, 10 mW) for varying periods; (C) two examples of FLIM images of fluorescein-doped PLGA particles embedded in a polyacrylamide gel after varying periods of irradiation at 980 nm (pulsed laser, 10 mW). The internal temperature was extracted from the calibration curve in (A).

Light-activated Intracellular Release

To examine whether this mechanism allows release of cargo from particles within cells, we first formulated polyester 2.1 particles (size: 0.7 ± 0.1 µm, see Figure

1.S11 for dark field microscopy and size distribution histogram) encapsulating the model release compound fluorescein diacetate (FDA). FDA is non-fluorescent in its native form but becomes highly fluorescent upon uptake by cells, as intracellular esterases hydrolyze the diacetate groups to produce fluorescein. Polyester 2.1 was chosen for these experiments because it possesses a glass transition in aqueous environments close to 37 °C (calculated wet Tg of 40 °C), which translates to less temperature-dependent release than PLGA (measured wet Tg of 28 °C). Particles were incubated with macrophages for 3 h, and cells were analyzed by flow cytometry following separation from free particles by centrifugation. Irradiation of cells incubated with FDA-loaded particles at 980 nm for 15 and 30 min resulted in 28- and

71-fold greater cell-associated fluorescence, respectively, than cells alone, while fluorescence increased by only 14-fold in non-irradiated cells (Figure 1.5). Not all cells took up FDA-loaded particles, as indicated by the left shoulder of the curves representing cells incubated with particles. The increase in released FDA with irradiation time, as indicated by fluorescence of its cleaved product, indicates tunable control over the dose of cargo released. We believe that light-triggered FDA release is the result of heating of the confined water in the polyester particles because the increase in temperature of the bulk environment was minimal, from 37 ± 0.1 °C to

38.8 ± 0.1 °C after 30 min of irradiation (980 nm, pulsed laser, 0.5 W).

As irradiation at 980 nm could be harmful to cells, its effects on cell count and viability were measured. Upon exposure to 980 nm light (pulsed laser, 0.5 W, 30 min), no change in cell viability was observed (non-irradiated cells: 98.5 ± 0.6% viable; irradiated cells: 97.9 ± 0.6% viable; see Figure 1.S12A) and only a slight decrease in

72 cell density (19%) was detected (non-irradiated cells: 350 ± 10 cells/µL; irradiated cells: 290 ± 20 cells/µL; Figure 1.S12B), which indicates that irradiation at this laser power causes minimal damage to living cells. The cytoxicity of the particles themselves has previously been reported as minimal up to a particle concentration of

100 µg/mL. These results demonstrate that 980 nm irradiation of polyester particles is a suitable means of triggering cargo release in living cells.

Figure 1.5. NIR light-triggered intracellular FDA release. Background, cells incubated without particles (background); remaining traces, cells incubated with FDA-loaded particles and irradiated for varying periods. Irradiation: 980 nm, pulsed laser, 0.5 W.

1.4 CONCLUSION

In the present study, we have demonstrated the feasibility of exploiting the unusual behavior of water confined within biodegradable polymeric particles to

73 thermally induce a phase change upon NIR irradiation in polymer carriers with no inherent light sensitivity and achieve controlled release of an encapsulated payload in aqueous environment and in living cells. This metal-free photo-induced release strategy involves: (1) water diffusion into the polymer particles, (2) light-induced heating of confined water by targeting the overtone vibrational absorption band of water at 980 nm, (3) conductive heat transfer from the excited water droplets to the polymer matrix, (4) thermal phase change of the polymer particles to a more rubbery state, and (5) increase of the particles diffusivity and release of their encapsulated content. NIR light induced significant release of both hydrophilic and hydrophobic small molecules. The on-demand rate of release was found to depend on the average

NIR photon energy administered to the system. Multiple consecutive NIR exposures can be used to obtain multiple release doses without irreversible rupture of the carriers, and, given the high encapsulation efficiency of the electrospraying technique, allows a large number of release cycles. This new release mechanism provides additional benefits: wavelength selectivity and high sensitivity, which allows the use of low CW laser power and avoids excessive heating. With more research, we foresee multiple applications such as light-activated self-healing capsules, extracellular scaffolds (nanofibers, hydrogels) providing on-demand delivery of cues for cell proliferation, differentiation, or migration, activatable fluorescent particles based on thermochromic dyes, and light-triggered drug delivery systems.194, 207, 250 Given the successful demonstration of 980 nm laser-driven in vivo diagnosis251-253 and therapy254-256 at depths of several millimeters without excessive heat generation or tissue damage, we are hopeful that NIR-induced heating of particle-confined water

74 may be widely adopted for site-specific photorelease in animal models, especially where light has direct access, e.g. in the eye or optically transparent organisms such as

C. elegans or zebrafish embryos.

1.5 METHODS

Materials

Poly(lactic-co-glycolic acid) (PLGA, ratio: 50:50; Mw: 7-17 kDa; alkyl ester terminated), polystyrene (Mw: 35 kDa), fluorescein (acid free, 95%), Nile blue chloride (85%), Nile red (technical grade), IR780 (98%), fluorescein diacetate (FDA), dextran (9-11 kDa), acrylamide (molecular biology grade), and N,N’-methylene bisacrylamide (99%) were purchased from Sigma Aldrich. Pluronic F127 (13 kDa) was purchased from O-BASF. Lithium acylphosphinate salt was synthesized according to a procedure published by Fairbanks et al. 257 Dextran was acetalated according to a procedure described by Suarez et al.258 The synthesis of polyester 1 and

238 2 can be found elsewhere. Chloroform (CHCl3, 99.8%, EMD), dimethylformamide

(DMF, 99%, Aldrich), tetrahydrofuran (THF, 99%, Fischer Scientific), and deuterated water (D 99.9%, Cambridge Isotope Inc.) were used without further purification. De- ionized water (DI H2O) was purified from a Millipore system (18.2 MΩ).

Preparation and Characterization of Polymer Particles

The polymer capsules incorporating model release compounds were produced by electrospray. This formulation method employs high voltages to inject charge into a liquid, causing the liquid to break into a jet of fine aerosol droplets propelled towards

75 a metal plate collector. As the solvent evaporates in flight, dense, solid polymer particles are generated.259 A variety of parameters, such as applied voltage, solution injection rate, and plate collector height, can be tuned to control the size and morphology of the particles.260 This versatile method yields highly reproducible particles and entraps payloads at high encapsulation efficiencies. The dye-doped particles were obtained as follows: For most formulations, the polymer (100 mg) was dissolved in 0.75 mL of CHCl3 and diluted with a solution containing the active compound in DMF (40 mg/mL, 0.25 mL) at 10% w/v. The prepared solutions were electrosprayed at 20 kV (Gamma High Voltage, ES30) at a flow rate of 0.5 mL/hr (KD

Scientific) using a 25 gauge needle. However, for polyester 1 particles encapsulating

FDA molecules, 35 mg of the polymer was mixed with 3.5 mg of FDA, and dissolved in 0.5 mL of 3:1 DMF/THF (7% w/v). This solution was electrosprayed at 28 kV

(Gamma High Voltage, ES30) at a flow rate of 0.1 mL/hr (KD Scientific) using a 25 gauge needle. The duration of the spray was kept the same between samples in order to yield the same final electrosprayed mass of polymer. Samples were collected onto microscope glass slides on an aluminum plate collector at a distance of 30 cm. The particles (~ 2 mg) were removed from their glass slide substrate by sonication in DI

H2O, washed with DI H2O and finally dispersed in 5 mL of DI H2O.

The morphology of the polymer particles was examined by fluorescence microscopy (Nikon, Eclipse, NIS Elements software) and SEM (Agilent, 8500).

Particle diameter distributions were extracted from recorded fluorescence images and

SEM photographs using NIS Elements (Nikon) and ImageJ software (NIH). The degree of polymer degradation induced by irradiation at 980 nm was studied by gel

76 permeation chromatography (Waters), comparing a particle solution exposed to NIR light (980 nm, pulsed laser, 1 W) for one hour to an un-irradiated sample. The amount of dye incorporated into the various particles samples was determined by dissolving the particles with CHCl3 to release the dye from the particles. Fluorescein- and Nile blue-loaded PLGA particles (5 mL, 0.4 mg/mL) were solubilized in 0.5 mL of CHCl3 first and then diluted in DI H2O to a volume of 50 mL, whereas Nile red- and IR780- loaded PLGA particles were completely solubilized in 50 mL CHCl3. The fluorescence of the mixtures was measured and the dye concentration quantified by linear calibration with matrix-matched standards. Steady-state fluorescence measurements were performed using a Fluorolog spectrofluorimeter (Horiba Jobin-

Yvon) and quartz cuvette (volume: 1.5 mL, optical path length: 1.0 cm). Differential scanning calorimetry (DSC, Perkin Elmer DSC-7) was used to measure the glass transition temperatures of PLGA.

Release Experiments in Aqueous Environment

In a typical experiment, release from polymer particles was photo-initiated by irradiating aliquots (0.4 mg/mL) in a micro quartz cuvette at 50 nm increments from

780 nm to 1030 nm, for specified periods of time using either a Ti:Sapphire laser (Mai

Tai HP, Spectra Physics, 100 fs pulse width, 80 MHz repetition rate, 1 W, beam dia.: 2 mm) or a CW laser diode (Thorlabs, 980 nm only, 170 – 350 mW, beam dia.: 2 mm).

A wave plate/polarizer combination was used to ensure an equal output power at 980,

900 and 800 nm. Release of the dyes was followed by fluorescence spectroscopy and an emission spectrum was recorded immediately after every irradiation period. The

77 release experiments regarding the hydrophilic dyes (fluorescein, Nile blue) were performed in pure water, whereas for the hydrophobic dyes (Nile red, IR780), a surfactant (pluronic F127, 0.1% w/v) was added to the particle solutions to establish a sink condition to facilitate release. The morphology of the irradiated particles was investigated by optical microscopy and SEM. Solution temperatures were measured using a thin wire thermocouple (J-Kem Scientific) immersed in the particle solutions while irradiating with the NIR light sources and connected to a temperature controller

(J-Kem Scientific). Digital photographs were acquired using a Panasonic DMC-ZS5.

The fraction of released material was quantified using fluorescence spectroscopy.

Typically, after varying periods of irradiation, irradiated samples and non-irradiated controls were spun down (4500 rpm, 15 min) to separate the released material in the supernatant from the payload still encapsulated in the particles. The supernatant and particle pellet were dried (rotovap, vacuum) and subsequently dissolved in organic solvent (CHCl3 or DMF). The fluorescence spectra were acquired and the fluorescence intensity of the supernatant was compared to the total fluorescence intensity, i.e., supernatant and particle pellet, to obtain the percentage of released material. The dye concentration was quantified by linear calibration with standards.

Nanoparticle Uptake and Photocontrolled Release in Cells

MV-4-11 (ATCC® CRL-9591TM) human lymphoblast macrophage cells from peripheral blood were purchased from ATCC and cultured in Iscove’s Modified

Eagle’s Medium (IMEM) supplemented with 10% fetal bovine serum. Cells were cultured in suspension and passaged three times before the experiment. Cells (5 x 105

/mL) were incubated at 37 °C, 5% CO2, with FDA-loaded polyester 1 particles (100

µg/mL) suspended in IMEM for 3 h to accomplish cellular uptake. The cells were then isolated by centrifugation and washed three times with IMEM to remove particles that were not internalized and remained free in media. To induce intracellular release and subsequent hydrolysis of FDA molecules, aliquots (100 µL) of cells containing FDA- loaded particles were irradiated at 980 nm (pulsed laser, 0.5 W) for 15 and 30 min to maximize exposure of the sample to the laser. The sample was subsequently diluted eight-fold and quantified using a guava EasyCyte flow cytometer (Millipore). Non- irradiated cells incorporating FDA-loaded particles were similarly prepared and analyzed as a control to probe the effect of nonspecific FDA release, whereas cells incubated without polymeric particles were used to set the background threshold. A commercial cell viability assay (Guava ViaCount, Millipore) was performed on non- irradiated and irradiated cell suspensions (980 nm, pulsed laser, 0.5 W, 30 min), 48 hours post irradiation to determine the effect of the laser on cells.

Internal Temperature Measurement using Fluorescence Lifetime

All lifetime experiments were performed on fluorescein-doped particles or free dye embedded in polyacrylamide gels. The hydrogels containing the particles were obtained by mixing a 250-µL particles aliquot (0.4 mg/mL) with acrylamide (60 mg), bis-acrylamide (1.4 mg) and lithium acylphosphinate salt (58 mg/mL, 10 µL). The gelation was photo-initiated under UV irradiation (30 s, Luzchem). To obtain the free dye embedded in hydrogels, 250-µL fluorescein-doped PLGA particles aliquot (0.4 mg/mL) was first heated at 65 °C for 15 min to release the dye molecules from the

PLGA particles and the empty polymer particles were removed by centrifugation before applying the same gelling process.

Fluorescence lifetime spectroscopy was done using a time-correlated single photon counting (TCSPC) system (Horiba) equipped with a NanoLED excitation source (488 nm, 1 MHz impulse repetition rate, Horiba) and a R928P detector

(Hamamatsu Photonics, Japan). The detector was set to 520 nm for detection of fluorescein. The slit width varied between 2 and 20 nm to achieve an appropriate count rate. The instrument response function (IRF) was obtained by using a scattering solution of Ludox-40 (Sigma-Aldrich) in water (prompt) at 480 nm emission. The samples and the prompt were measured in a semi-micro quartz cuvette. The lifetime was recorded on a 450 ns scale. A total of 4094 channels were used with a time calibration of 0.110 channel/ns. All decay curves were fitted with one exponential.

The fluorescence lifetimes were extracted using DAS6 v6.6 decay analysis software

(Horiba). The goodness of fit was judged by χ2 values, Durbin–Watson parameters, and visual observation of the fit line and residuals, which should be distributed randomly about zero.

Fluorescence lifetime imaging was performed on a SliceScope two-photon microscope (Scientifica, UK) using a 60X water immersion objective (LUMPLFLN

60XW, NA=1.0, Olympus). A Chameleon Ultra ΙΙ IR laser (Coherent) (80 MHz repetition rate, 100 – 150 fs pulses) tuned at 980nm was used for the excitation of both fluorescein and confined water. ScanImage r3.8 was used to control the scanning mirrors.261 Fluorescence emission was detected with a hybrid PMT detector (HPM-

100-40, Becker and Hickl, Germany) between 490 – 540 nm by means of a

GFP emission filter (ET 515/50, Chroma). The acquisition of fluorescence lifetimes was synchronized by a TCSPC module (SPC-150, Becker and Hickl). The following parameters were kept constant for all acquired images: pixel size (30 nm; all 512 x 512 pixels), pixel dwell time (3.2 µs), laser excitation intensity (10 mW after the microscope objective), and FLIM acquisition time (60 s/image). Fluorescence lifetime images were analyzed with SPCimage (Becker and Hickl). To minimize lifetime calculation errors, we used a minimum threshold of 15 photons at the peak

(corresponding to ~1000 photons per pixel), and a binning factor between 2 and 10 pixels to assure sufficient photons in the regions of interest. The same calculated IRF was used for all experiments. The control images of free fluorescein in polyacrylamide gels were analyzed with a single exponential model, whereas the PLGA particles images were much better fitted with a double exponential model. The goodness of fit was evaluated with χ2 values and visual observation of the fit line and residuals.

1.6 ACKNOWLEDGMENTS

We are grateful to the King Abdulaziz City for Science and Technology

(KACST-UC San Diego Center of Excellence in Nanomedicine) and the NIH New

Innovator Award (1 DP2 OD006499-01) for financial support. This work was also supported in part by NIH R01AG032132 (RM, KD). The authors also thank J. Beak and S. Jin for their assistance with the electrospraying method, K. Brzezinska for differential scanning calorimetry (the MRL Shared Experimental Facilities are supported by the MRSEC Program of the NSF under DMR 1121053; a member of the

NSF-funded Materials Research Facilities Network), M. Chan and J. Karpiak for their help with the hydrogel preparation, Jason Olejniczak and Amy Moore for their help with the gel permeation chromatography measurements, and D. Boudreau for useful advice and recommendations. Chapter 1, with the exception of the preamble, is in full a reprint of the material as it appears in ACS Nano 2014. Mathieu L. Viger,

Wangzhong Sheng, Kim Doré, Ali H. Alhasan, Carl-Johan Carling, Jacques Lux,

Caroline de Gracia Lux, Madeleine Grossman, Roberto Malinow, and Adah Almutairi,

ACS Press 2014. The dissertation author and M.L. Viger were the primary co- investigators and co-authors of this work and contributed equally.

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34. Fomina, N.; McFearin, C.; Sermsakdi, M.; Edigin, O.; Almutairi, A. UV and Near-IR Triggered Release from Polymeric Nanoparticles. J Am Chem Soc 2010, 132, 9540- 9542.

35. Siegel, S. J.; Kahn, J. B.; Metzger, K.; Winey, K. I.; Werner, K.; Dan, N. Effect of Drug Type on the Degradation Rate of PLGA Matrices. Eur J Pharm Biopharm 2006, 64, 287-293.

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37. Lei, Y.; Child, J. R.; Tsavalas, J. G. Design and Analysis of the Homogeneous and Heterogeneous Distribution of Water Confined within Colloidal Polymer Particles. Coll Polymer Sci 2013, 291, 143-156.

38. Tsavalas, J. G.; Sundberg, D. C. Hydroplasticization of Polymers: Model Predictions and Application to Emulsion Polymers. Langmuir 2010, 26, 6960-6966.

39. Sandor, M.; Enscore, D.; Weston, P.; Mathiowitz, E. Effect of Protein Molecular Weight on Release from Micron-Sized PLGA Microspheres. J Controlled Release 2001, 76, 297-311.

40. Zhao, A. Y.; Hunter, S. K.; Rodgers, V. G. J. Theoretical Prediction of Induction Period from Transient Pore Evolvement in Polyester-Based Microparticles. J Pharm Sci 2010, 99, 4477-4487.

41. Ikada, Y.; Tsuji, H. Biodegradable Polyesters for Medical and Ecological Applications. Macromol Rapid Comm 2000, 21, 117-132.

42. Ihre, H. R.; De Jesus, O. L. P.; Szoka, F. C.; Frechet, J. M. J. Polyester Dendritic Systems for Drug Delivery Applications: Design, Synthesis, and Characterization. Bioconj Chem 2002, 13, 443-452.

43. Mohamed, F.; van der Walle, C. F. Engineering Biodegradable Polyester Particles with Specific Drug Targeting and Drug Release Properties. J Pharm Sci 2008, 97, 71- 87.

44. Breitenbach, A.; Li, Y. X.; Kissel, T. Branched Biodegradable Polyesters for Parenteral Drug Delivery Systems. Journal of Controlled Release 2000, 64, 167-178.

45. Lux, C. D.; Joshi-Barr, S.; Nguyen, T.; Mahmoud, E.; Schopf, E.; Fomina, N.; Almutairi, A. Biocompatible Polymeric Nanoparticles Degrade and Release Cargo in Response to Biologically Relevant Levels of Hydrogen Peroxide. J Am Chem Soc 2012, 134, 15758-15764.

46. Vert, M. Aliphatic Polyesters: Great Degradable Polymers that Cannot Do Everything. Biomacromolecules 2005, 6, 538-546.

47. Messager, L.; Portecop, N.; Hachet, E.; Lapeyre, V.; Pignot-Paintrand, I.; Catargi, B.; Auzely-Velty, R.; Ravaine, V. Photochemical Crosslinking of Hyaluronic Acid Confined in Nanoemulsions: Towards Nanogels with a Controlled Structure. J Mater Chem B 2013, 1, 3369-3379.

48. Bhattarai, N.; Gunn, J.; Zhang, M. Q. Chitosan-Based Hydrogels for Controlled, Localized Drug Delivery. Adv Drug Delivery Rev 2010, 62, 83-99.

49. Broaders, K. E.; Cohen, J. A.; Beaudette, T. T.; Bachelder, E. M.; Frechet, J. M. J. Acetalated Dextran is a Chemically and Biologically Tunable Material for ParticulateI. Proc Nat Acad Sci USA 2009, 106, 5497-5502.

50. Broaders, K. E.; Grandhe, S.; Frechet, J. M. J. A Biocompatible Oxidation-Triggered Carrier Polymer with Potential in Therapeutics. J Am Chem Soc 2011, 133, 756-758.

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57. Blaiszik, B. J.; Kramer, S. L. B.; Olugebefola, S. C.; Moore, J. S.; Sottos, N. R.; White, S. R. Self-Healing Polymers and Composites. Ann Rev Mater Res 2010, 40, 179-211.

58. Naczynski, D. J.; Tan, M. C.; Zevon, M.; Wall, B.; Kohl, J.; Kulesa, A.; Chen, S.; Roth, C. M.; Riman, R. E.; Moghe, P. V. Rare-Earth-Doped Biological Composites as In Vivo Shortwave Infrared Reporters. Nat Comm 2013, 4:2199, 1-10.

59. Hong, G. S.; Lee, J. C.; Robinson, J. T.; Raaz, U.; Xie, L. M.; Huang, N. F.; Cooke, J. P.; Dai, H. J. Multifunctional In Vivo Vascular Imaging using Near-Infrared II Fluorescence. Nat Med 2012, 18, 1841-+.

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62. Jayakumar, M. K. G.; Idris, N. M.; Zhang, Y. Remote Activation of Biomolecules in Deep Tissues using Near-Infrared-to-UV Upconversion Nanotransducers. Proc Nat Acad Sci USA 2012, 109, 8483-8488.

63. Fairbanks, B. D.; Schwartz, M. P.; Bowman, C. N.; Anseth, K. S. Photoinitiated Polymerization of PEG-Diacrylate with Lithium Phenyl-2,4,6- Trimethylbenzoylphosphinate: Polymerization Rate and Cytocompatibility. Biomaterials 2009, 30, 6702-6707.

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66. Almeria, B.; Deng, W.; Fahmy, T. M.; Gomez, A. Controlling the morphology of electrospray-generated PLGA microparticles for drug delivery. J Colloid Interface Sci 2010, 343, 125-133.

67. Pologruto, T. A.; Sabatini, B. L.; Svoboda, K. ScanImage: Flexible Software for Operating Laser Scanning Microscopes. Biomed Eng Online 2003, 2, 13.

1.8 SUPPORTING INFORMATION

SEM images, fluorescence microscopic images, particle size distributions of the electrosprayed particles, release profiles, cytotoxicity, and lifetime data.

Photographs of particle suspensions before and after irradiation. Fluorescence release experiments. This material is available free of charge via the Internet at http://pubs.acs.org.

Supplemental Figures

Figure 1.S1. Electrosprayed PLGA particles are uniform. (A) Dark field microscopic images and (B) size distribution histogram of fluorescein-loaded PLGA particles; (C) Dark field microscopic images and (D) size distribution histogram of Nile Blue-loaded PLGA particles; (E) Dark field microscopic images and (F) size distribution histogram of Nile Red-loaded PLGA particles; (G) Dark field microscopic images and (H) size distribution histogram of IR780-loaded PLGA particles.

Figure 1.S2. Electrosprayed PLGA particles efficiently encapsulate the dye molecules. Optical fluorescence microscopic images of (A) Fluorescein-loaded PLGA particles, (B) Nile Blue-loaded PLGA particles, and (C) Nile Red-loaded PLGA particles.

Figure 1.S3. (A) Digital photographs of a solution of fluorescein-loaded PLGA particles in water after 0, 5, 10, 15, 30, and 45 min laser irradiation (980 nm, pulsed laser, 1 W); Fluorescence microscopy images of fluorescein-loaded PLGA 1 particles (B) before and (C) after laser irradiation (980 nm, pulsed laser, 1 W, 45 min).

Figure 1.S4. SEM images of fluorescein-loaded PLGA particles (A) before and (B) after 980 nm irradiation (pulsed laser, 1 W, 90 min).

Figure 1.S5. GPC chromatograms of the PLGA before (black trace) and after irradiation (red trace) at 980 nm for 1 hr. (pulsed laser, 1 W).

Figure 1.S6. Photographs displaying laser-induced particles agglomeration over time (980 nm pulsed laser, 1 W). Initially, particles are well dispersed (t = 0). After 5 min of exposure to laser irradiation, minimal amounts of aggregates are observed. Agglomeration becomes significant after 10 min irradiation. Aggregates can be easily broken up and suspended in solution by sonication.

Figure 1.S7. NIR irradiation at a water-resonant wavelength induces release of both hydrophobic and hydrophilic compounds. (A) Cumulative Nile Blue release after irradiation of a suspension of Nile Blue-doped PLGA particles with pulsed laser light (1 W) at 980 nm (open circles). Control (no irradiation, solid circles). λex = 460 nm, λem = 475 – 750 nm; (B) Cumulative Nile Red release after irradiation of a suspension of Nile Red-doped PLGA particles with pulsed laser light (1 W) at 980 nm (open circles). Control (no irradiation, solid circles). λex = 560 nm, λem = 570 – 750 nm; (C) Cumulative IR780 release after irradiation of a suspension of IR780-doped PLGA particles with pulsed laser light (1 W) at 980 nm (open circles). Control (no irradiation, solid circles). λex = 740 nm, λem = 750 – 900 nm. [dye] = ~10% w/w.

Figure 1.S8. (A) Fluorescence microscopy images and (B) size distribution histogram of Nile Red-loaded polyester 1 particles; (C) Fluorescence microscopy images and (D) size distribution histogram of Nile Red-loaded polyester 2 particles; (E) Fluorescence microscopy images and (F) size distribution histogram of Nile Red-loaded acetalated dextran particles.

Figure 1.S9. NIR irradiation at a water-resonant wavelength can induce release from capsules made of different polymeric materials. Cumulative Nile Red release after irradiation of a suspension of Nile Red-doped (A) polyester 1 (MW = 51 kDa), (B) polyester 2 (MW = 17 kDa) and (C) acetalated dextran (MW = 9-11 kDa) particles with pulsed laser light (1 W) at 980 nm (open circles). Controls (no irradiation, solid circles). Molecular structures of the polymers are shown as insets on the graphs. λex = 560 nm, λem = 570 – 750 nm. [dye]polyester 1 = ~20% w/w, [dye]polyester 2 = ~10% w/w, and [dye]acetalated dextran = ~10% w/w.

Figure 1.S10. Changes in average lifetime of fluorescein encapsulated in PLGA particles and embedded in a polyacrylamide gel (N = 6) upon alternating cycles of irradiation at 980 nm (5 min) and non-irradiation/relaxation (5 min).

Figure 1.S11. Electrosprayed FDA-doped polyester 1 particles. (A) Dark field microscopic image and (B) size distribution histogram.

Figure 1.S12. Cytotoxicity study on the effect of NIR-irradiation: (A) Viability and (B) cell density with or without irradiation at 980 nm (pulsed laser, 0.5 W). Cell type: MV-4-11 (ATCC® CRL-9591TM) human lymphoblast macrophage cells.

Supplementary Tables

Table 1.S1. Fluorescence lifetime data for fluorescein molecules in a polyacrylamide gel heated at various temperatures.

2 T (°C) τav. (ns) Error (ns) ∆τav. (ps) Error (ps) χ

20 3.680 0.005 0 5 1 30 3.605 0.004 -75 4 1.2 40 3.52 0.01 -156 13 1 50 3.44 0.02 -237 23 0.9 60 3.359 0.004 -321 4 0.8 70 3.286 0.003 -394 3 0.8

2 τav. is the averaged lifetime value and χ is the “quality of fit” parameter. ∆τav. is the change in lifetime and is calculated using the following equation: ∆τav. = τav. (X°C) – τav. (20°C).

Table 1.S2. FLIM data of fluorescein within particles embedded in a polyacrylamide gel after different irradiated time at 980 nm.

Replica 1 Replica 2 Replica 3 Replica 4 Replica 5

Time (min) τav. (ps) ∆τav. (ps) τav. (ps) ∆τav. (ps) τav. (ps) ∆τav. (ps) τav. (ps) ∆τav. (ps) τav. (ps) ∆τav. (ps) Mean ∆τav. (ps) Error (ps)

0 1930 0 1888 0 1628 0 1903 0 2272 0 0 0 5 1893 -37 1655 -233 1557 -71 1756 -147 2209 -63 -110 30 10 1840 -90 1657 -231 1536 -92 1694 -209 2177 -95 -143 26 15 1688 -242 1636 -252 1527 -101 1721 -182 2052 -220 -199 22 20 1635 -295 1654 -234 1476 -152 1645 -258 2055 -217 -231 22 25 1594 -336 1563 -325 1425 -203 1683 -220 2019 -253 -267 31

τav. is the averaged lifetime value. ∆τav. is the change in lifetime and is calculated using the following equation: ∆τav. = τav. (X°C) – τav. (20°C).

Table 1.S3. FLIM data of fluorescein free in a polyacrylamide gel after different irradiated time at 980 nm.

Replica 1 Replica 2 Replica 3 Replica 4 Replica 5

Time (min) τav. (ps) ∆τav. (ps) τav. (ps) ∆τav. (ps) τav. (ps) ∆τav. (ps) τav. (ps) ∆τav. (ps) τav. (ps) ∆τav. (ps) Mean ∆τav. (ps) Error (ps)

0 2732 0 2724 0 2705 0 2667 0 2638 0 0 0 5 2785 53 2799 75 2602 -103 2608 -59 2640 2 -6 39 10 2785 53 2644 -80 2612 -93 2789 122 2537 -101 -20 47 15 2791 59 2624 -100 2613 -92 2906 239 2531 -107 0 71 20 2836 104 2688 -36 2602 -103 2963 296 2594 -44 43 79 25 2823 91 2674 -50 2583 -122 2982 315 2577 -61 35 86

τav. is the averaged lifetime value. ∆τav. is the change in lifetime and is calculated using the following equation: ∆τav. = τav. (X°C) – τav. (20°C).

Chapter II:

Selective Photothermolysis Theory and Gold-Nanoparticle Laser-Induced

Photothermal-Assisted Liposuction (Nanolipo) in a Porcine Model

2.0 CHAPTER TWO PREAMBLE

In the Introductory Chapter, I presented an overview of the field of photothermal therapy and argued that, in order to practically apply the photothermal effect in biological systems, the focus should not be the development of more photothermal agents, but rather the confinement of heat to specific targets of interest.

The work presented in Chapter I was valuable in exposing me to how selective heat confinement might be achieved, but also in helping me realize the limitations of using wavelengths of light outside of the biological optical window. In this chapter, I explore this concept further and present a detailed derivation of the fundamental principles underlying selective photothermolysis, citing seminal works in the field by

R.R. Anderson, John A. Parrish, and later, G.B. Altshuler. I also present my own work in developing a bulk-scale proof-of-concept selective photothermolysis model using a thermoresponsive polymer, which undergoes a change in opacity after being heated through its lower critical solution temperature (LCST), as an indicator for selective photothermal heating. Working with one of the most widely studied photothermal agents, I went on to explore the feasibility and demonstrate proof of concept for a

99 gold-nanoparticle laser-induced photothermal-assisted liposuction technology in a porcine model, as well as investigate the biosafety, biodistribution, and clearance of the gold nanoparticles from the bodies of the animals tested.

2.1 ENGINEERING ANALYSIS OF SELECTIVE PHOTOTHERMOLYSIS

AND MODELING

The concept of selective photothermolysis, originally described by Anderson and Parrish, was briefly introduced in the Introductory Chapter. In essence, it is a corollary of the photothermal effect in which the parameters of light exposure are carefully controlled to achieve specific and targeted heating in pigmented biological structures within tissue. In this section, an engineering analysis of selective photothermolysis is briefly presented.

The Heat Equation

As the photothermal effect deals with the conversion of light to heat, the fundamental equation governing it is the heat equation. This famous partial differential equation (PDE) has been extremely well-studied and its derivation can be readily found in many engineering textbooks, since its application is quite broad. In short, the heat equation takes the form:

훿푇 휌퐶 − κ∇2푇 = 퐻 푝 훿푡

Where, ρ is the mass density, Cp is the specific heat, κ is the thermal conductivity and

H is the heat flux. This equation is commonly rearranged into the form:

100

훿푇 퐻 − 퐷∇2푇 = 훿푡 휌퐶푝

Where D = κ/ 휌퐶푝 is a term called the thermal diffusivity.

This equation basically says that as long as there is a heat flux, H, into the system, there will be a rise in the temperature over time in an arbitrarily defined volume of tissue, which makes sense. Although the heat equation can be solved using numerical methods for general cases, there are simpler, more fundamental solutions in the form of Green’s functions. For example, the solution for a point source heat flux with an infinitesimally short duration in an infinite, three dimensional piece of homogenous tissue is:

퐸 1 |푥|2 ( ) 푇 푥, 푡 = ( 3/2) exp (− ) 휌퐶푝 8(휋퐷푡) 4퐷푡

Where E is the amount of optical energy deposited into the tissue at the point source.

Obviously, many assumptions have to be made in order to for this solution to hold, and although this solution does not quite describe real-world phenomenon, it is a good place to start in terms of modeling thermal conduction behavior in tissue.

By implementing different boundary conditions, such as a heat source with one dimensional plane geometry, applied to a three dimensional block of homogenous tissue, the solution then becomes:

퐸 1 푧2 푇(푥, 푡) = ( ) exp (− ) 휌퐶푝 2√휋퐷푡 4퐷푡

101

Where z is the penetration depth. Although this solution still relies on several major assumptions, it is much more representative of the geometry in photothermal applications.

Heat Deposition from a Light Source

In the traditional sense, heat is transferred to tissue through conduction. In the heating source with a plane geometry, one might imagine a hot plate placed over the surface of the tissue to transfer heat through the established contact surface. However, in photothermal applications, electromagnetic energy in the form of photons must first be absorbed, and then relax vibrationally to generate heat. For a continuous-wave light source, the thermal energy is deposited in the form of:

퐻 = 휇푎훷

Where 휇푎 is the absorption coefficient and 훷 is the light fluence rate, also known as irradiance. In this form of photothermal heating, when the laser is turned on, the light is assumed to arrive at the tissue surface instantaneously. Subsequently, for as long as the laser is kept on, the energy will be continuously deposited into tissue and converted into heat through tissue vibrational relaxation or thermal relaxation.

In contrast to laser light of a continuous-wave nature, a pulsed laser could be used instead. In the current context, the term “pulsed” is used to loosely describe light that is applied in a discontinuous manner, including continuous-wave light that is applied in a gated manner. The main difference between a continuous-wave and pulsed laser is the manner in which the energy is deposited. In the case of the continuous-wave laser, the parameter that is most relevant is the amount of energy

102 deposited per unit of time, since the longer the laser is kept on, the more energy is deposited. In the case of a pulsed laser, however, the parameter that is most relevant becomes the energy deposited per pulse, since the laser is not left on indefinitely. The thermal energy deposited can now be expressed as:

퐻(푥, 푡) = 휇푎(푥)훷(푥)푓(푡)

Where 푓(푡) is a dimensionless parameter that describes the temporal shape of the pulse. By integrating the heat flux with respect to the time, the total energy deposited can be calculated:

∞

퐸 = 휇푎(푥)훷(푥) ∫ 푓(푡) 푑푡 0

If the laser pulse has some duration tp, and it is assumed that energy deposition is instantaneously achieved and maintained for the duration of the pulse, then the energy deposited by each pulse is 퐸 = 휇푎(푥)훷(푥)푡푝.

Another name for the heat flux is the power density. One caveat is that the power density should be low enough as to not induce cavitation effects as a result of heating water above its boiling point, especially for applications involving discontinuous energy exposure.

Heat Diffusion and Penetration Depth of Heat

Following conversion from the absorbed light energy, heat will diffuse into the tissue. The two primary factors influencing heat dissipation in tissue are the heat capacity and thermal conductivity. Although these terms are not constant, they are

103 often treated as such since complex biological tissue is comprised mostly of water. By doing so, the thermal diffusivity in biological tissue, and thus, the thermal penetration depth can be estimated. The thermal penetration depth is expressed as:

푧 = √4퐷푡

Where z = 0 is defined as the surface of the tissue. By making the appropriate estimations, the thermal diffusivity in tissue is calculated to be approximated to be 20

µm/ms. This value is important in determining the thermal relaxation time in tissues.

Thermal Relaxation Time of Tissue and Selective Photothermolysis

The thermal relaxation time (TRT or τ) is defined as the time it takes for the temperature at depth z to reach 1/e of the temperature at z = 0, where the heat source is located. Ultimately, it is an indicator for how fast heat is lost in an object. As it was explained briefly in the Introductory Chapter, and as the reader might envisage, if subsequent pulses of light are used to photothermally heat the target structure before it has time to lose its heat (t < TRT), the target structure will be heated up continuously and selectively compared to the surrounding areas since energy is deposited and built- up in the target structure before diffusion can happen. This concept is in essence, selective photothermolysis. In the original work by Anderson and Parrish, one might expect a heating profile such as the one shown in Fig. 2.1.1. A prerequisite, however, was that the target structure had to have an absorption coefficient at least twice as high as that of the surrounding tissue. Since the major component in tissue is water, it would be difficult to achieve selective photothermolysis relying on water as the

104 absorbing pigment. Additionally, they must be spaced sparsely enough such that their heat diffusional areas do not overlap, as indicated qualitatively in Fig. 2.1.1.

Figure 2.1.1. Schematic showing various achievable temperature profiles during selective photothermolysis. Pigmented target structures with sufficient spacing can be heated selectively if pulses of light are applied sequentially such that the time between pulses is less than the thermal relaxation rate of the absorbing structure. From Anderson, RR and Parrish, JA, “Selective photothermolysis: precise microsurgery by selective absorption of pulsed radiation.” Science, 1983. Re-printed with permission from AAAS.

Extended theory of Selective Photothermolysis

105

Although this initial model was adequate in describing the selective photothermal phenomenon for the heating of specific pigmented structures, it did not do a good job of accounting for conduction to surrounding tissue, because after all, over time heat will diffuse into the surrounding tissues. Working with Anderson, G.B.

Altshuler proposed an alternate, extended theory of selective photothermolysis that adds a buffer zone into consideration, physically represented as interstitial spacing between adjacent target structures, in calculating the extent of the selective photothermal effect. In essence, this allowed for more accurate calculations of the selective photothermal effect. The governing equations for the TRT, as well as the thermal damage time, input power density, and input energy fluence were all calculated by Altshuler for planar, cylindrical, and spherical geometries with the dimensional parameter d (Tab. 2.1.1,). There is actually a mistake for the “TRT of the heater” set of terms, where all of the d1 terms in the numerator should be squared.

Table. 2.1.1. Calculation of Parameters for Achieving Selective Photothermolysis. From Altushuler, G.B. et al., “Extended Theory of Selective Photothermolysis.” Lasers in Surgery and Medicine. 2001. Re-printed with permission from Wiley.

106

Based on these equations, the necessary powers and exposure times can be estimated for the selective photothermolysis of target structures with specific geometries. As mentioned previously, these fundamental principles have been successfully applied for the treatment of port-wine stains, as well as for laser hair removal. pNIPAM Selective Photothermolysis Model

With these fundamental principles in mind, I developed a bulk-scale model in order to visualize the selective photothermolysis. Poly N-isopropylacrylamide

(pNIPAM) is a thermoresponsive polymer that undergoes a hydrophobicity change upon heating through its lower critical solution temperature (LCST), changing from hydrophilic to hydrophobic as it is heated from below its LCST to above, respectively.

To model hypothetical selective photothermal heating in tissue, I designed a density gradient pNIPAM model consisting of three layers of pNIPAM solution. At the bottom was 5 wt% pNIPAM mixed with 10 wt % sucrose in deionized water to allow this layer to remain situated at the bottom. The middle, target layer consisted of 5 wt% pNIPAM mixed with a neodymium-based photothermal agent that I developed in the lab, at a concentration of 1 OD. On the top was simply a 5 wt% pNIPAM solution. pNIPAM in its native state (without modifications or substitutions) has a LCST of about 32 °C. When the density gradient solution is placed in a 50 °C water-bath as control, the whole solution becomes opaque, indicating that the 5 wt% pNIPAM in all of the layers undergo a transition simultaneously (data not shown). When a continuous-wave laser at 1 W/cm2 is used to illuminate the density gradient solution from above, as might be the case in heating of tissues, a small opaque region begins to

107 form in the center of the middle layer containing the PT agent, which begins to expand and grow. As it grows, it does not seem to be inhibited by the boundaries of the different layers, except maybe slightly at the middle-to-bottom layer interface. As the exposure time is continued, the whole solution eventually turns opaque (Fig. 2.1.2).

Figure. 2.1.2. Continuous light exposure induces heating of the photothermal layer. As exposure is continued, heat expands into none-photothermally absorbing layers (above and below) until pNIPAM in the whole volume undergoes a LCST transition, turning cloudy.

In contrast to the continuous wave heating, when light was applied in a discontinuous manner, specifically using higher power densities applied as shorter

108 pulses, the change in opacity, indicative of photothermal heating, was selectively restricted to the middle layer, even as the number of pulses increased (Fig. 2.1.3).

Figure. 2.1.3. Discontinuous light exposure with sufficient cool-down times in between exposure events enables selective photothermal heating of only the middle layer of the density gradient pNIPAm solution.

This model shows that it is very much feasible to confine heating to a target region in a selective manner by applying the principles of selective photothermolysis.

Armed with this knowledge, we proceeded to test this concept in an animal tissue model. We proceeded carefully, since we know that getting light through biological tissue is wholly different from passing it through water, which is achieved quite readily in the pNIPAM model. The next section of this chapter, is in part, a reprint of the material as it appears in Plastic and Reconstructive Surgery - Global Open.

Wangzhong Sheng, Ali H. Alhasan, Gabriella DiBernardo, Khalid M. Almutairi, J.

Peter Rubin, Barry E. DiBernardo, Adah PhD Almutairi. LLW Journals 2015. The dissertation author was the primary investigator and author of this work.

109

2.2 ABSTRACT

Background: Conventional suction-assisted lipectomy (SAL) often results in contour irregularity. Selective photothermal heating of adipose tissue by polymer-coated gold nanorods (AuNRs) energized by an external near-infrared (NIR) exposure at 800 nm is introduced in this work to facilitate fat removal. Materials and methods: The effects of NanoLipo were examined in food grade porcine abdominal tissue (skin, fat, and fascia) by histology. The efficacy of NanoLipo was compared to that of conventional

SAL in vivo in Yucatan mini pigs by quantification of removed subcutaneous tissue and fatty acids and ultrasound measurement of adipose layer thickness. Results:

NanoLipo led to the appearance of disruptions in adipose tissue that were not apparent in control groups in ex vivo samples. NanoLipo allowed removal of more subcutaneous tissue (~33% vs. ~25% of removed material, p < 0.05) and approximately twice as much free fatty acids (~60% vs. ~30% of removed tissue, p <

0.05) than conventional SAL. Most importantly, NanoLipo led to a greater decrease in adipose layer thickness at one month post-surgery (p < 0.001). Conclusions:

NanoLipo facilitates removal of a greater quantity of fat and requires less suction time

(4 min vs. 10 min) than conventional SAL. As the safety of PEG-coated AuNRs is well-established, a clinical trial is currently being organized.

110

2.3 INTRODUCTION

According to the most recent statistics by the American Society for Aesthetic

Plastic Surgery, liposuction, including conventional suction-assisted lipectomy (SAL), ultrasound-assisted liposuction (UAL), and laser-assisted liposuction (LAL), is the most common aesthetic procedure performed by plastic surgeons in the United

States.(1) However, these procedures are often associated with secondary complications due to irregular fat removal such as contour deformities, irregular lumpy appearance, and excess skin, leading to patient dissatisfaction.(2-4) In this article, a surgical device technology termed NanoLipo is introduced. This technology employs a gold nanorod (AuNR) solution energized by external near-infrared (NIR) laser exposure (Fig. 2.2.1) to overcome these challenges by uniformly and selectively heating adipose tissue while sparing surrounding tissue.

NanoLipo heats adipose tissue by surface plasmon resonance (SPR), through which AuNRs absorb laser energy of specific wavelengths in the near infra-red (NIR) region.(5) The wavelength absorbed can be tuned by altering particle shape, size, geometry, and aspect ratio.(6) This absorption causes gold electrons to oscillate with the frequency of the electromagnetic field, generating heat with extremely high efficiency.(7) Photothermal conversion through SPR takes advantage of the difference in thermal relaxation rates between fat and surrounding tissues to allow very rapid, localized heating of adipose tissue.(8, 9) Because fat has a lower specific heat capacity

-1 -1 -1 -1 (2.3 kJ g K for fat vs. 4.18 kJ g K for water),(10, 11) as well as a lower thermal

-1 -1 -1 -1 conductivity (0.23 W m K for fat vs. 0.631 W m K for water) than water, it heats faster and dissipates heat slower. Heating by this mechanism selectively softens and

111 loosens adipose tissue, facilitating removal with minimal trauma. Only surgeon- defined regions, where the AuNR solution is infused, absorb laser energy, minimizing the potential for damage to surrounding tissues.

NanoLipo employs gold nanorods that absorb a wavelength within the tissue transparency window (800-900 nm), where competing endogenous chromophores including water and hemoglobin have lower absorption.(9) Current FDA-approved technologies for LAL rely predominantly on wavelengths around or beyond 1000 nm, where water absorbs and emits heat.(12) Consequently, these methods require the insertion of laser probes into the subcutaneous tissue to liquefy small volumes of fat.(13) The point source nature of the heating device makes uniform results difficult to achieve. Surrounding subcutaneous tissue, such as muscle and fibrous connective tissues, will also be heated significantly.(14, 15) In NanoLipo, an externally applied benign laser source could provide more uniform planar heating than an internally inserted fiber optic laser, as well as increase the safety margin by selective heating of adipose tissue. Finally, an important advantage of the NanoLipo system is the ability to precisely control temperature increases through the concentration of the exogenous solution (Fig. 2.2.2), in addition to laser energy, pulsing sequence, and wavelength.(7)

The amount of heating can thus be finely tuned to achieve the target temperature, while keeping the skin surface temperature lower than 42 °C to mechanically weaken adipose tissue without damage to skin.(13)

2.4 MATERIALS & METHODS

Gold nanorod (AuNR) solution

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AuNR solution was produced, packaged and released in accordance with the

FDA’s current Good Manufacturing Practice guidelines by NanoSpectra BioSciences,

Inc. (Houston, TX) following the literature.(16) AuNRs (10 nm x 40 nm) were functionalized with poly(ethylene-glycol) (PEG) (5 kDa) via displacement of hexadecylcetyltrimethylammonium bromide (CTAB), a detergent used in the synthesis of AuNRs. Proposed specifications for AuNRs in the commercially produced

NanoLipo system are summarized in Table 2.2.1.

Laser source

The Lumenis LightSheer® DuetTM laser system (Lumenis, Yokneam, Israel), a commercially available, FDA-approved device with an 800 nm pulsed diode, was used for all studies. The laser probe, whose application area is 3.5 cm x 2.2 cm, was set to generate three consecutive 30 ms pulses of 6 J/cm2 (46 J/ pulse) each pass (i.e. 138 J per pass before accounting for any attenuation by the tissue). In animal studies, the skin was cooled using a damp towel to keep skin surface temperatures below 42 °C.

Ex vivo studies

Ex vivo studies were performed on food-grade porcine abdominal tissue (pork belly). AuNR or saline solution was injected into the experimental and control areas, respectively. Another region was treated with the laser only. The experimental and laser-only regions were exposed to 20 passes of the Lumenis laser. Frozen sections

(10 µm) were stained with hematoxylin and eosin, (8) prior to imaging using a

Hamamatsu NanoZoomer 2.0HT slide scanning microscope.

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Animal studies

All animal studies were approved by the University of California, San Diego

(UCSD) Institutional Animal Care and Use Committee (IACUC). Yucatan mini pigs

(45-55 kg) were purchased from S&S Farms, Ramona, CA, housed at the UCSD pig facility at Elliot Farms and fed a standard diet (three meals per day). Zucker rats were purchased from Harlan Sprague Dawley, housed in the UCSD vivarium, and fed an unrestricted normal diet to maximize body fat content.

NanoLipo procedure

Abdominal hair was trimmed using clippers and removed using NairTM.

Following sterilization with surgical betadine, two regions were tattooed for conventional SAL or NanoLipo (Fig. 2.2.S1). During all procedures, surface skin temperature was monitored by a FLIR E50 infrared thermal camera.

NanoLipo solution was mixed with anesthetic tumescent solution (Ringer’s solution saline with 0.1% lidocaine, 1 ppm epinephrine) to a concentration of 2.5 ×

1011 AuNR/mL (14 μg AuNR/mL). 100 mL of the solution was injected into adipose tissue through a small stab incision in a systematic fan pattern to ensure uniform permeation and distribution in the target region (5 cm x 5 cm). The laser was applied to the marked area over the course of 5 ± 1 min to deliver 1000-2000 J of energy over multiple passes, alternating the orientation of the laser application probe to ensure complete coverage of the area. The skin was cooled using a wet towel every four passes to maintain a safe skin temperature, as verified by the thermal imaging camera

(FLIR E50, FLIR, Wilsonville, OR). Subcutaneous tissue was removed by suction-

114 assisted liposuction (Gomco OptiVac® G180, Allied Healthcare Products, St. Louis,

MO) and the incision was closed using an absorbable suture, while the operated areas were marked with a non-absorbable suture.

Processing of removed subcutaneous tissue

Lipoaspirates were centrifuged at 1000 rpm for 5 min to separate the liquid phase, including injection solution, from solid subcutaneous tissue. Both phases were weighed, and subcutaneous tissue was examined under dark field microscopy at 10X magnification. Cell diameters (along the longest axis, all cells in each field of three representative images, totaling ~50 cells) were measured using ImageJ software. To quantify the proportion of removed subcutaneous tissue consisting of free fatty acids and glycerol, 1 g of each subcutaneous tissue sample was digested with 3 mg collagenase/dispase following manufacturer protocol for 1 h at 37 °C (Roche) and centrifuged (2000 rpm, 5 min) to produce three distinct layers (from top to bottom: free fatty acids and glycerol, adipocytes, and fibrous matter). The percentage of removed tissue consisting of free fatty acids and glycerol was determined by measuring the volume of the upper layer containing secreted fatty acids and glycerol, converting it to mass (assuming 1 mL = 1 g), and dividing by the total mass of removed tissue.(17)

Ultrasound and skin appearance assessment

Ultrasound measurements were taken using a Biosound MyLab30Vet machine with a LA435 Linear Probe 18-10 MHz transducer through a thick layer of ultrasound gel before, immediately after, and at 10 d, one month, two months, and three months

115 post-procedure to monitor changes in tissue depth. The effect of user pressure was accounted for by applying maximum pressure and slowly relaxing, acquiring the measurement the moment just before the transducer detached from the surface.

Additionally, images were consistently acquired during exhalation to account for any changes in depth due to breathing. Images parallel and perpendicular to the spinal axis of the animal were acquired to obtain complete coverage of the operated area. The machine was set to measure the same depth for each region across all time points.

In each image, the distance from the top of the deep fascial membrane to the top of the superficial fascial membrane, which appears white on ultrasound,(18) was measured at four cross-sections spaced 1 cm apart in each ultrasound image (five images per treated region) using ImageJ. Images in which resolution was too low to identify fascial membranes (fewer than 10% of images) were not analyzed. Each measured distance is plotted to illustrate the change in average depth over time.

Interpretation of ultrasound was aided by compression testing during image collection; fat layers compressed more than fibrous layers.

Statistics

Continuous variables, except for adipose thickness measured by ultrasound, are reported as means and standard errors. Groups were compared by two-tailed student’s t-test in Microsoft Excel 2010.

2.5 RESULTS

Ex vivo fat liquefaction

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In the first proof-of-concept experiments, food-grade porcine fatty abdominal tissue was used to determine the minimum concentration of NanoLipo AuNR solution to facilitate removal of the adipose layer at a safe laser power and short duration with minimal heating. Porcine tissue was subcutaneously injected with NanoLipo AuNR solution (0.1 g/L) and irradiated with a laser (800 nm, 2.5 kJ total, 30 ms/pulse), injected with AuNR solution only, irradiated without AuNRs, or left untreated as a control. In the absence of AuNR solution, histology showed no signs of significant lipolysis (Fig. 3 A, C). Skin surface temperature was monitored and did not exceed 45

°C in NanoLipo treated samples. NanoLipo-treated regions appeared more translucent than regions treated with laser alone, suggesting liquefaction of fat (Fig. S2). The tissue was also mechanically softer when probed with tweezers. Histology revealed disruption, apparent as voids, in subcutaneous adipose tissue but not in the dermal layer or connective tissue (superficial to the adipose layer) of NanoLipo-treated samples (Fig. 3 B, D). Disruption of adipose tissue in these samples made them noticeably more fragile and difficult to section. While deep connective tissue appears to be disrupted by NanoLipo, insertion of the cannula might also cause such effects.

NanoLipo efficacy in Yucatan mini pigs

We next examined whether NanoLipo enhances fat removal relative to standard liposuction techniques using Yucatan mini pigs. NanoLipo and conventional

SAL were performed on two abdominal regions on each pig. NanoLipo allowed removal of considerably more subcutaneous tissue (Fig. 2.2.4A) and fat than conventional SAL in a comparable amount of time. The NanoLipo method required

117 far less time (4 min vs. 10 min) to remove a similar volume of lipoaspirate and caused less bruising than SAL. Collagenase digestion and centrifugation revealed that the fatty content of tissue removed following NanoLipo was nearly twice that following

SAL (Fig. 2.2.4B). Dark field microscopy of adipocytes in the lipoaspirate following

NanoLipo showed that the adipocytes from NanoLipo-treated areas were significantly smaller than those from SAL-treated areas (Fig. 2.2.4C, D).

Assessment of tissue depth using ultrasound

The thickness of the adipose layer before and immediately after the procedure, as well as at 1, 2, and 3 months post-operation was measured using ultrasound (Fig.

2.2.5). Analysis of ultrasound images reveals comparable depth changes between

NanoLipo and SAL immediately post-operation (Fig. 2.2.5C). However, the change in adipose tissue layer thickness at one month post-surgery is significant in NanoLipo- treated areas (p < 0.001) but not in those treated using SAL. Similarly, reductions in adipose tissue layer thickness at 3 months post-surgery were greater in NanoLipo- treated than SAL-treated areas (Fig. 2.2.5A, B).

2.6 DISCUSSION

Surface plasmon resonance (SPR) in metal nanostructures has been used extensively for photothermal therapy of cancer.(19),(20) In NanoLipo, we sought to apply SPR-based heating to improve on conventional SAL by expanding on the principle of selective photothermolysis.(9) Selecting an appropriate laser pulse length, among other parameters, allows NanoLipo to precisely increase the amount of energy

118 conferred to the target, i.e., adipose tissue. The major advantage of this technology is a consequence of the addition of exogenous energy absorbers (AuNRs) rather than relying on endogenous elements, such as water.

In this work, ultrasound was used to assess whether NanoLipo enhances reductions in adipose tissue thickness and yields more uniform results relative to conventional SAL. A comparison of depth changes between NanoLipo and SAL groups immediately post-operation correlated well with the total volume of subcutaneous tissue removed (Fig. 4A). Ultrasound results at one month post-op suggests that NanoLipo’s localized thermally-aided fat removal resulted in less swelling or that such swelling subsides faster (Fig. 5C). NanoLipo appears to lead to greater reductions in thickness at 3 months post-surgery. Tissue depth reduction appears to be more uniform in NanoLipo-treated regions than in those treated with

SAL across all time points, as evidenced by a tighter distribution of thicknesses (Fig.

5C).

The results presented herein strongly suggest that NanoLipo may aid in removal of adipose tissue while maintaining the integrity of overlying tissues.

Although the NanoLipo system includes one more step than SAL, the amount of time saved during lipoaspiration (4 min vs 10 min to obtain the same volume) more than makes up for the time spent to apply the laser. This accelerated fat removal may result from release of fatty acids and glycerol from adipocytes, evidenced by reductions in the diameter of cells from lipoaspirate following NanoLipo procedures.

The NanoLipo technique has a high safety margin. PEG-coated AuNRs have been studied extensively and are essentially innocuous.(21, 22) The injected dose is

119

also well below the LD50 of AuNRs (23). Furthermore, since AuNRs do not bind to tissue, a large portion of injected AuNRs are immediately removed by aspiration, and since the concentrations delivered (0.01-0.05 g/kg body weight) are well below the expected toxicity limit (3.2 g/kg) for gold,(23) long-term AuNR exposure is not expected. A long-term investigation of the distribution of AuNRs in pigs is underway.

We anticipate the NanoLipo procedure to be well-tolerated, as the mechanical changes induced by selective photothermolysis are temporary.

NanoLipo would not be significantly more expensive than conventional liposuction. Compatible laser systems, including the Lumenis LightSheer employed in this study, are widely available in cosmetic surgery clinics, and AuNRs at the concentration used here cost less than a dollar per liter of lipoaspirate.

2.7 CONCLUSIONS

In the present work, we introduce a new system, termed NanoLipo, that facilitates removal of a greater quantity of fat (double) and requires less time (4 min vs. 10 min) than conventional SAL. NanoLipo appeared to yield more uniform reductions in adipose layer thickness. A greater proportion of removed tissue consisted of fatty acids, which agrees with our hypothesis that NanoLipo causes fat liquefaction.

The presented data suggest that NanoLipo offers potential advantages over conventional SAL, warranting longer-term studies.

2.8 ACKNOWLEDGEMENTS

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We thank the Karen Christman lab (Department of Bioengineering, UC San Diego) for access to cryosectioning equipment, Jennifer Santini and the Light Microscopy

Facility (Department of Cellular and Molecular Imaging, UC San Diego) for access to microscopes for histological analysis, Dr. James Day for access to the inductively coupled plasma mass spectrometry facility at the Scripps Isotope Geochemistry

Laboratory (Scripps Institute of Oceanography, UC San Diego), and Arnold V. Garcia and the UC San Diego Animal Care team for the contributions to the pig studies.

Chapter 2, is in part, a reprint of the material as it appears in Plastic and

Reconstructive Surgery - Global Open. Wangzhong Sheng, Ali H. Alhasan, Gabriella

DiBernardo, Khalid M. Almutairi, J. Peter Rubin, Barry E. DiBernardo, Adah PhD

Almutairi. LLW Journals 2015. The dissertation author was the primary investigator and author of this work.

2.9 REFERENCES

1. Ahmad, J., Eaves, F. F., Rohrich, R. J., Kenkel, J. M. The American Society for Aesthetic Plastic Surgery (ASAPS) Survey: Current Trends in Liposuction. Aesthetic Surgery Journal 2011;31:214-224.

2. Illouz, Y. G. Complications of liposuction. Clinics in Plastic Surgery 2006;33:129-+.

3. Teimourian, B., Adham, M. N. A national survey of complications associated with suction lipectomy: What we did then and what we do now. Plastic and Reconstructive Surgery 2000;105:1881-1884.

4. Teimourian, B. Complications Associated with Suction Lipectomy. Clinics in Plastic Surgery 1989;16:385-394.

5. Maestro, L. M., Haro-Gonzalez, P., Coello, J. G., Jaque, D. Absorption efficiency of gold nanorods determined by quantum dot fluorescence thermometry. Applied Physics Letters 2012;100.

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6. Jain, P. K., Lee, K. S., El-Sayed, I. H., El-Sayed, M. A. Calculated absorption and scattering properties of gold nanoparticles of different size, shape, and composition: Applications in biological imaging and biomedicine. Journal of Physical Chemistry B 2006;110:7238-7248.

7. Govorov, A. O., Richardson, H. H. Generating heat with metal nanoparticles. Nano Today 2007;2:30-38.

8. Anderson RR, Farinelli W, Laubach H, Manstein D, Yaroslavsky AN, Gubeli J 3rd, Jordan K, Neil GR, Shinn M, Chandler W, Williams GP, Benson SV, Douglas DR, Dylla HF. Selective photothermolysis of lipid-rich tissues: A free electron laser study. Lasers in Surgery and Medicine 2006;38:913-919.

9. Anderson, R. R., Parrish, J. A. Selective Photothermolysis - Precise Microsurgery by Selective Absorption of Pulsed Radiation. Science 1983;220:524-527.

10. El-Brawany, M. A. Measurement of thermal and ultrasonic properties of some biological tissues. Journal of Medical Engineering & Technology 2009;33:8.

11. Bowman, H. F., Balasubramaniam, T. A. New Technique Utilizing Thermistor Probes for Measurement of Thermal-Properties of Biomaterials. Cryobiology 1976;13:572- 580.

12. Kou, L. H., Labrie, D., Chylek, P. Refractive-Indexes of Water and Ice in the 0.65- Mu-M to 2.5-Mu-M Spectral Range. Applied Optics 1993;32:3531-3540.

13. DiBernardo, B. E. Randomized, Blinded Split Abdomen Study Evaluating Skin Shrinkage and Skin Tightening in Laser-Assisted Liposuction Versus Liposuction Control. Aesthetic Surgery Journal 2010;30:593-602.

14. Goldberg, S. N., Gazelle, G. S., Halpern, E. F., Rittman, W. J., Mueller, P. R., Rosenthal, D. I. Radiofrequency tissue ablation: Importance of local temperature along the electrode tip exposure in determining lesion shape and size. Academic Radiology 1996;3:212-218.

15. Thomas, L. W. The chemical composition of adipose tissue of man and mice. Journal of Experimental Physiology 1962;47:10.

16. Perez-Juste, J., Pastoriza-Santos, I., Liz-Marzan, L. M., Mulvaney, P. Gold nanorods: Synthesis, characterization and applications. Coordination Chemistry Reviews 2005;249:1870-1901.

17. Erdim, M., Tezel, E., Numanoglu, A., Sav, A. The effects of the size of liposuction cannula on adipocyte survival and the optimum temperature for fat graft storage: an experimental study. Journal of Plastic, Reconstructive, and Aesthetic Surgery 2009;62:1210-1214.

18. Leahy, S., Toomey, C., McCreesh, K., O'Neill, C., Jakeman, P. Ultrasound Measurement of Subcutaneous Adipose Tissue Thickness Accurately Predicts Total

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and Segmental Body Fat of Young Adults. Ultrasound in Medicine and Biology 2012;38:28-34.

19. Arvizo, R., Bhattacharya, R., Mukherjee, P. Gold nanoparticles: opportunities and challenges in nanomedicine. Expert Opinion on Drug Delivery 2010;7:753-763.

20. Lukianova-Hleb, E. Y., Ren, X. Y., Zasadzinski, J. A., Wu, X. W., Lapotko, D. O. Plasmonic Nanobubbles Enhance Efficacy and Selectivity of Chemotherapy Against Drug-Resistant Cancer Cells. Advanced Materials 2012;24:3831-3837.

21. Rayavarapu RG, Petersen W, Hartsuiker L, Chin P, Janssen H, van Leeuwen FW, Otto C, Manohar S, van Leeuwen TG. In vitro toxicity studies of polymer-coated gold nanorods. Nanotechnology 2010;21.

22. Dykman, L. A., Khlebtsov, N. G. Gold Nanoparticles in Biology and Medicine: Recent Advances and Prospects. Acta Naturae 2011;3:34-55.

23. Hainfeld, J. F., Slatkin, D. N., Focella, T. M., Smilowitz, H. M. Gold nanoparticles: a new X-ray contrast agent. British Journal of Radiology 2006;79:248-253.

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2.10 FIGURE LEGENDS AND TABLES

A B C D E Figure 2.2.1. Schematic of NanoLipo procedure. A solution of NanoLipo AuNRs is A. infused into the adipose tissue and B. permeates it. C. An external 800 nm laser is applied to the epidermis; the epidermis temperature is carefully controlled by contact cooling (not shown). D. Adipose tissue loses its mechanical integrity as triglycerides stored in adipocytes are secreted as free fatty acids and glycerol. E. Adipose tissue and the NanoLipo solution are removed using standard liposuction procedures.

Table 2.2.1. Device specifications for NanoLipo solution

Test Procedure Specifications Absorption peak UV/vis spectrophotometry 800 ± 10 nm Particle Transmission electron microscopy 1.6 x 10^13 NR/ml concentration (TEM) Optical density at UV/vis spectrophotometry 50 ± 5 OD 800 nm Dimensions Transmission electron microscopy Length : 40 ± 5 nm (TEM) Width : 10 ± 2 nm CTAB ISO 2871-2: determination of ≤ 4 µM concentration cationic-active matter content

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Figure 2.2.2. Tunable absorption of AuNR solution used in NanoLipo, which absorbs efficiently in the 800 nm region. Water and other endogenous chromophores do not absorb at this wavelength.

Figure 2.2.3. Histological effects of NanoLipo in ex vivo porcine skin and subcutaneous tissue. A. H&E stained section of AuNR-injected (left) and untreated (right) samples and B. H&E stained section of NanoLipo-treated (left) and laser- treated samples. C-D. Higher magnification power image of C. untreated region and D. NanoLipo-treated region; arrows indicate intact connective tissue and * indicates disruptions. Scale bars = 2.5 mm (A-B), 250 µm (C-D).

125

Figure 2.2.4. NanoLipo allows removal of more tissue and more fat than standard suction-assisted lipectomy (SAL). A. Subcutaneous tissue was weighed upon removal. B. Free fatty acids and glycerol were separated from subcutaneous tissue by digestion and centrifugation. C. Dark field microscopy and D. quantification of average diameter of adipocytes in lipoaspirates. Scale bar = 100 µM. *, p < 0.05; ***, p < 0.001. n = 3 procedures in 3 pigs.

126

Figure 2.2.5. Ultrasound imaging analysis reveals that NanoLipo enhances reductions in tissue thickness. Representative ultrasound images prior to and 3 months post-surgery of A. NanoLipo treated region and B. SAL-treated region. Red lines indicate adipose layer targeted by liposuction. C. Quantification of tissue thickness for NanoLipo (left) and SAL (right). **, p < 0.01; ***, p < 0.001 relative to baseline. n = 3 procedures in 3 pigs; 20 measurements per procedure.

127

Figure 2.2.S1. Treatment regions were marked with non-absorbable sutures.

Figure 2.2.S2. Details of ex vivo experiments. Top and right, photograph of skin and subcutaneous fat following injection of AuNRs and laser exposure. Below, temperature measurement by thermal camera immediately following laser exposure of AuNR-injected area.

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2.11 BIODISTRIBUTION AND CLEARANCE OF GOLD

NANOPARTICLES FROM PIGS

Study Background and Design

A biodistribution and clearance study was conducted in pigs to evaluate the safety and efficacy of these nanomaterials in living organisms. Relative to photothermal therapy of cancer, the amount of material used for NanoLipo is nearly two orders of magnitude lower (0.6 OD vs 50 OD). However, it was still important to examine whether the gold nanorods used for NanoLipo results in any toxicity.

Animals from the NanoLipo efficacy study were divided into two groups based on time-point: one month (n=3; Pigs #464, 469, 473) and six months (n=3; Pigs 391,

392, 394), with one untreated pig serving as control (Pig #475). Urine and blood were periodically collected to construct the clearance profile. Animals were sacrificed at their respective time points and major organs were harvested and processed for ICP-

MS analysis. The UCSD Animal Care Program Staff aided in the collection of specimens for analysis, and a ThermoQuest Element 2 high-resolution ICP-MS machine at the Scipps Isotope Geochemistry Lab was used to perform the ICP-MS

Analysis.

Gold is Cleared from Animals after Six Months

Injected gold nanoparticles were found to mostly clear from the bodies of pigs after six months. The peak concentration in the blood was measured between 10 days and one month after treatment. However, this peak concentration was less than 1% of the injected dose. This clearance profile is expected, especially for PEGylated

129 nanomaterials that are able to achieve host immune evasion.1 The gold in the blood was largely cleared after one month of residence time.

Clearance Profile of Gold from Blood

Concentration Concentration (ppb) 1

0 Pre Post 10D 1Mo 2Mo 3Mo 4Mo 5MO 6Mo

Figure 2.3.1. Clearance profile of AuNRs from pig blood. The concentration of gold was maximum at approximately one month after treatment. This profile is not unexpected for PEGylated nanoparticles, which can avoid immune detection and remain in circulation for extended periods of time. This concentration was less than 1% of the injected dose.

Gold nanoparticles significantly accumulated in the liver at both one month and six months, which is expected since this class of materials is known to clear through the reticuloendothelial (RES) system.1 Surprisingly, gold nanoparticles also seemed to accumulate in the lungs of pigs. This observation may be explained by the aspect ratio of the nanorods, as nanomaterials with high aspect ratios are known to preferentially distribute to the lungs as a result of their entrapment in alveolar structures.2 Despite the presence of gold nanoparticles in the bodies of these animals

130 for up to six months, the animals demonstrated no signs of toxicity in the form of sickness or behavior changes.

15.00 Gold in Major Organs

13.00 after One Month 11.00

9.00 Pig #464 Pig #469 7.00 Pig #473 Pig #475 5.00

3.00

1.00 Percent (%) Dose Injected of Percent -1.00 Spleen Kidney Liver Heart Lung Brain

Figure 2.3.2. Biodistribution of gold in major organs of pigs after one month. After one month, gold was found in the liver and lungs of the animals.

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Gold in Major Organs 15.00 After Six Months 13.00 11.00 Pig #391 9.00 Pig #392 7.00 Pig #394 5.00 3.00 1.00

-1.00 Spleen Kidney Liver Heart Lung Brain Treated Percent (%) Dose Injected of Percent Region

Figure 2.3.3. Biodistribution of gold in major organs of pigs after six month. After six month, most of the injected gold was cleared, although some gold was still found in the liver and lungs of the animals.

References

1. J. V. Jokerst, T. Lobovkina, R. N. Zare and S. S. Gambhir, Nanomedicine, 2011, 6, 715-728.

2. H. M. Braakhuis, M. V. Park, I. Gosens, W. H. De Jong and F. R. Cassee, Particle and fibre toxicology, 2014, 11, 18.

Chapter III:

A Single-Blind Study Evaluating the Efficacy of NanoLipo in an

ex vivo Human Tissue Model

3.0 PREAMBLE: FUNCTION AND PHYSIOLOGY OF HUMAN ADIPOSE

TISSUE

Long thought to be an inert organ whose sole function was energy storage, it is becoming increasingly clear that adipose tissue is much more complex. Adipose tissue has been found to play an incredibly important role in regulating metabolic processes as a vital part of the endocrine system1, 2, with adipose tissue remodeling at the cellular and tissue levels being implicated in the obesity epidemic,3 as well as related pathologies of insulin resistance, metabolic syndrome, inflammation, and even cancer4. In addition, the discovery of depots of adult stem cells5-7 in adipose tissue makes it invaluable as a readily available yet inexpensive source of cells for stem cell therapies. Decellularized forms of the underlying adipose tissue scaffolds were also recently applied for adipose8, 9 and broader tissue engineering applications.10, 11 These factors make adipose tissue research increasingly relevant in medicine. Since adipose tissue was not considered to be a functional organ until recently, there was a lack of terminology in describing and delineating the nuances between the different types of adipose tissue, or “fat”, as it was colloquially referred to. As researchers grasped a

132

133 better understanding of adipose tissue anatomy, physiology, and function, it became increasingly important to consolidate the various terms used in the literature to create a hierarchical classification to describe adipose tissue architecture.12, 13

Adipose Tissue at the Cellular Level

The basic unit of adipose tissue at the cellular level is the adipocyte. There are two main types of adipocytes, white and brown. White adipocytes constitute the majority of adipocytes and are primarily responsible for storing excess energy in the form of triglycerides by packaging them into a single large lipid droplet which can occupy up to 85% of the cell by weight, which would not be well-tolerated in other types of cells.14 Since the droplets are so large, other cellular constituents, including the nucleus and various organelles, are pushed to the periphery of the cell. It was also recently discovered that white adipocytes play a significant biochemical role in the cell biology of adipose tissue expansion.15 The less common brown adipocyte, on the other hand, is comprised of multiple, smaller lipid droplets, along with an abundance of mitochondria, accounting for their color as well as their important role in thermo- regulation16. In addition to white and brown adipocytes, adipose tissue is also comprised of smaller populations of endothelial cells, fibroblasts, macrophages, and various other cell types that play important roles in regulating and maintaining adipose tissue health. Generally, adipose tissues containing proportionally more white or brown adipocytes are referred to as white adipose tissue (WAT) and brown adipose tissue (BAT), respectively.

Visceral vs. Subcutaneous Fat

134

At the macroscopic level, total adipose tissue, not including those found in bone marrow, the head, hands, and feet, is divided into two main compartments:

Visceral adipose tissue (VAT) and subcutaneous adipose tissue (SAT), which are found inside and outside of the abdominal cavity, respectively. VAT can be further partitioned into various sub-compartments corresponding to the VAT immediately surrounding the various vital organs, as well as the VAT in the intraperitoneal space.

VAT is relatively more active metabolically17 owing to the proportionally larger percentage of BAT compared to WAT. SAT, on the other hand, is comprised of mostly WAT, and is partitioned into two main sub-compartments: A superficial

(sSAT) and a deep (dSAT) layer distributed thought various regions of the body and separated by a fibrous membrane referred to as Scarpa’s fascia, which has been observed with common imaging methods such as computer-aided tomography. The

SAT compartment is the most relevant in the context of liposuction, since VAT is located deep within the thoracic cavity.

Architectural Organization of Subcutaneous Adipose Tissue

Subcutaneous adipose tissue is organized into complex hierarchal levels.18 As stated earlier, the smallest unit of adipose tissue is the adipocyte. Immediately surrounding the adipocytes are thin fibrous septa rich with proteins such as elastin, fibronectin, and collagen, which help the cells organize locally into micro-lobules measuring approximately several micrometers in diameter. An analogy used by Klein to describe the micro-lobule is a raspberry, in which the adipocytes are the individual lobes of the raspberry.19 The next level of hierarchical organization is the secondary

135 lobule, also sometimes called a pearl because of the surrounding transparent collagenous membrane which gives this structure a pearl-like luster. The analogy used to describe a pearl is a cluster of grapes, in which each individual grape corresponds to the primary lobules while the vine corresponds to the connective tissue holding the clusters together. Fat pearls vary in size depending on its location inside the body, but can be as large as 1 cm in diameter. Visually, pearls can be sliced open to reveal lobules that appear to be small granules of a yellowish paste-like substance. Many pearls are packed together into a rhomboidal prism structure known as a fat section.

Finally, fat sections are organized in a crystal lattice–like fashion to create a fat compartment, such as the deep subcutaneous adipose tissue compartment. Different fat compartments are typically named after the area of the body in which they are located, e.g. abdominal or buttocks. Digital photographs of freshly human adipose tissue can be found in Fig 3.1, Fig. 3.3, and Fig. 3.5.

Obesity and The Biology of Adipose Tissue Expansion

Adipose tissue mechanical properties, specifically plasticity and expandability, may be key determinants in obesity-associated metabolic dysregulation.20 The primary response of the body to over-nutrition is lipogenesis, a process by which free fatty acids are esterified into triglyceride molecules which are then stored inside adipocytes as lipid droplets. The reverse of this process is lipolysis, in which adipocytes break down and release stored triglycerides in the form of free fatty acids and glycerol. The balance between lipogenesis and lipolysis is generally controlled metabolically, although various environmental factors may induce lipolysis.

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In general, during lipogenesis, white adipocytes located throughout the body in all adipose tissue types simultaneously increase in size by expanding their membranes in response to increasing amounts of triglycerides being deposited. When the cells reach a certain capacity, typically on the order of 100 μm or more, adipose tissue expansion as a whole occurs. Under healthy conditions, this process is accompanied by an acute inflammatory response associated with a net loss of adipose tissue collagen, allowing for adipocyte lipolysis (fat breakdown) followed by re-deposition into newly formed adipocytes. Meanwhile, dysfunctional adipose tissue expansion is associated with an increased number of extracellular matrix (ECM) deposits, owing to an inability by the body to effectively degrade ECM. The consequence of this limited capacity for expansion exacerbates insulin resistance as well as leads to chronic inflammation and hypoxia. Rutkowski et al. reviewed the mechanism of adipose tissue expansion in detail from a cell biology and metabolic perspective that focuses on the various biochemical factors involved.15 Nevertheless, the excess ECM deposits lead to changes in macroscopic mechanical properties, especially in SAT, as an indirect indication of adipose tissue health.

This process of pathologic adipose tissue expansion was recently implicated in breast cancer tumorgenesis4. Obesity is known to be one of the main risk factors for cancer, although the mechanism was not clear. Seo et al showed that obesity induced pathologic adipose tissue deposition in mammary tissue resulted in increased ECM stiffness, which in turn promoted the differentiation of adipose stromal cells into myofibroblasts. Measurements of the Young’s modulus of decullarized adipose tissue from obese mice compared to controls showed that the stiffness was nearly double in

137 the obese group, which was verified by the longer collagen and fibronectin fiber networks observed in the ECM of obese mice. The newly differentiated myofibroblasts would then deposit even more ECM to further increase the density and stiffness of the ECM, resulting in a positive feedback loop that not only promoted tumor malignancy, but also enhanced the tumorigenic potential of pre-malignant human breast epithelial cells.

All of these factors, taken together, help account for the various, complex sources of potential complications during liposuction. For example, the tissue history significantly influences the fibrousness of the tissue, and thus the difficulty of treatment during liposuction. Additionally, adipose tissue created during fat expansion, often called “resistant fat”, is extremely tough to be rid of. A specific form of liposuction called body contouring may aid in the removal of resistant fat. Human adipose tissue is also uniquely different from that of any other species, including porcine tissues, the best animal model substitute, so it is beneficial to model NanoLipo is a human adipose tissue model. The remainder of this chapter is comprised of work that is currently being prepared for submission for the American Society for Plastic

Surgeon’s Aesthetic Surgery Journal. Wangzhong Sheng, William J. Seare, Ali H.

Alhasan, Esther Cory, Paul Chasan, Robert Sah, Khalid Almutairi, Barry DiBernardo, and Adah Almutairi, “A Single-Blind Study Evaluating the Efficacy of Gold

Nanoparticle Photothermal-Assisted Liposuction in an ex vivo Human Tissue Model.”

Oxford Academic 2017. The dissertation author was the primary investigator and author of this work.

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3.1 ABSTRACT

Background: Liposuction is one of the most performed cosmetic surgery procedures.

In a previously reported study, gold-nanoparticle (GNP) laser-assisted liposuction

(NanoLipo) was shown to improve procedure parameters and outcomes in a porcine model. Objective: An ex vivo human liposuction model was developed to assess the ease, efficacy, and outcomes of NanoLipo, and to further explore its mechanism of action in facilitating liposuction. Methods: NanoLipo was compared to a Control without GNPs in sets of fresh, non-perfused, anatomically-symmetric, matched tissue specimens from twelve patients. A subset of three experiments was performed under single-blind conditions. Intra-operative assessments included lipoaspirate volume, percentage of free oil, ease of removal, and temperature rise. Specimens were palpated, visualized for evenness, and graded with and without skin. Post-operative assessment included viability staining of the lipoaspirate and remaining tissues. Micro- computed tomography was used to assess the distribution of infused GNPs within the tissues. Results: NanoLipo consistently removed more adipose tissue with more liberated triglycerides compared to Control. NanoLipo specimens were smoother, thinner, and had fewer and smaller irregularities. Infused solutions preferentially distributed between fibrous membranes and fat pearls. After NanoLipo, selective structural-tissue disruptions, indicated by loss of metabolic activity, were observed.

Thus, NanoLipo likely creates a bimodal mechanism of action whereby fat lobules are dislodged from surrounding fibro-connective tissue, while lipolysis is simultaneously induced. Conclusions: NanoLipo showed many advantages compared to Control

139 under blinded and non-blinded conditions. This technology may be promising in facilitating fat removal for well-trained and less-skilled liposuction practitioners alike.

3.2 INTRODUCTION

According to data gathered by the American Society of Plastic Surgeons

(ASPS) from 2015, liposuction surgery is one of the most popular cosmetic surgery procedures, and is arguably the number one performed procedure because of under- reporting by non-society affiliated practitioners. Despite its popularity, the number of annual procedures has fallen 37% since the year 2000.21 Part of the reason for this decline may be the relatively lower satisfaction rate reported among patients receiving liposuction surgery compared to other top cosmetic procedures. Rohrich, et al. reported that satisfaction rates were 73% and 82% for patients who either gained or did not gain weight post-surgery, respectively,22 with asymmetry, contour irregularities, inadequate or excessive fat removal, scarring, and residual skin laxity being some of the top reasons for dissatisfaction.23 Broughton, et al. similarly reported that only 79.7% of a cohort of 209 surveyed patients would elect to have the procedure again.24 These numbers leave some room for improvement, given that satisfaction rates as high as 98% have been reported for procedures such as breast augmentation.25 Both experienced liposuction doctors and other less experienced practitioners could benefit greatly if the procedure could be performed more quickly, while removing more fat with more forgiving and even results, especially in the patients with more fibrous tissue such as patients undergoing repeat procedures.

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The objective of this study was to assess a technology called NanoLipo that uses gold nanoparticles (GNPs) added to standard tumescent fluid, injected into the targeted fat, followed by the application of externally applied near-infrared photothermal laser energy to create lipolysis and focused disruptions to fibro-connective tissues, while assisting in producing superior aesthetic results. Concepts were borrowed from photothermal therapy (PTT) for cancer that similarly involves the use of various agents capable of capturing optimally-penetrating laser energy and creating focused heating to lethally damage cancerous tissues.26 Plasmonic GNPs were chosen for this application because it is one of the most extensively studied photothermal agents in oncology. However, one major distinction from cancer therapies is that in NanoLipo,

GNPs are injected at concentrations orders of magnitude lower than those used in

PTT, since the goal is not complete tissue destruction.27 Another important distinction between NanoLipo and cancer therapies that rely on bulk heating is that the laser energy in NanoLipo is applied for strategically predetermined pulse durations and optimized power outputs, leveraging known heating relaxation times for the GNPs and affected target fat structures. NanoLipo is thereby able to specificly disrupt the fibroconnective tissues and lipocytes through selevtive photothermolysis.28, 29 The ability of this novel photothermal-assisted liposuction procedure to produce superior outcomes in many aspects was previously reported in a porcine model.27 While this model is the best alternative to human skin and fat, porcine tissues still have significant biological, physiological, and mechanical differences from tissues of human origin.30 In this work, the first ex vivo human tissue liposuction model was developed, to the best of our knowledge, to was successfully used to evaluate the

141 limitations, benefits, and efficacy of the NanoLipo technology in improving liposuction before proceeding with clinical trials.

3.3 MATERIALS AND METHODS

Adipose Tissue Harvest and Preparation for NanoLipo Treatment

The use of human tissues conformed to the principles of the Declaration of

Helsinki, as well as all U.S. federal guidelines and regulations.31 Subjects voluntarily undergoing combined liposuction and abdominoplasty procedures at the private practices of participating plastic surgeons provided informed consent for the use of their tissues. An institutional review board was not used for this study because experiments were conducted on discarded tissue and did not involve direct experimentation on human subjects.32, 33 Twelve bilaterally symmetric tissue flaps containing skin and a sufficient amount of adipose tissue were surgically removed from the lower abdomen, lateral flank, and sometimes extending to the area above the buttocks of subjects, and were provided, without identification, to the researchers

(Table S1). All tissues were excised and tested on the same day and in most cases within a few hours to ensure tissue viability and similarity to in situ tissues. Samples were typically between 3 cm to 6 cm thick, 10-20 cm wide, and up to 30-36 cm in length per bilaterally symmetric sample when received. Except for one sample, a

Fitzpatrick skin type VI, all samples were type III or less. A treatment zone of approximately 4 cm x 7 cm was identified and marked in symmetrical areas selected for lack of stretch-marks, hair, and scars, with matched even consistency, texture, and pigmentation of the skin. A larger non-treatment region of 2-3 cm, at minimum, was

142 included on all sides of the treatment zone to minimize loss of infused tumescent solutions out of the exposed edges, to ensure even vacuum and application of the laser, and to serve as handles for the specimen during liposuction while keeping the tumescent and liposuction cannula tips in the treatment area. All samples were immediately post-processed and analyzed after treatment.

NanoLipo Procedure

In our model, NanoLipo involved the infusion of 40 cc of GNP-containing phosphate buffered saline (PBS) solution injected into the interstitial spaces of adipose tissue in the targeted treatment zone using standardized tumescent and instrumentation techniques. The Control contained no GNPs in the tumescent solution. The GNPs were surface-functionalized with a passivating poly(ethylene-glycol) (PEG, 5 kDa) polymer coating to prevent clumping and to improve tolerability by biological tissues

(NanoSpectra BioSciences, Inc., Houston, TX).34 The stock 50 optical density (OD)

GNPs provided was diluted with PBS to down to 1.2 OD before infusion into tissue specimens. GNP size, aspect ratio, and other physical, and optical properties were confirmed to be as indicated by the manufacturer.

A tumescent infiltration protocol was developed and instituted in an attempt to maximize the even disbursement of the GNP within the treatment zone. A 2 mm x 130 mm multi-port infusion cannula was inserted parallel to and approximately 1-3 cm below the skin surface, approximately the maximum depth of light penetration.35 The cannula was inserted fully into and through the tissues to the distal edge of the treatment zone. The tumescent or GNP solution was introduced into the tissues only

143 during withdrawal of the cannula to the proximal edge of the treatment zone, injecting approximately 1 cc per pass with 10 passes per side. This process was repeated as multiple and separate introductions from all four sides for a total tumescent infusion of

40 cc. Every attempt was made to infuse the solutions evenly while manually locating and digitally guiding the cannula within the specimen.

For the laser energy application, the 4 cm x 7 cm tumesced treatment zone was exposed to the laser through the skin in a pre-determined application pattern in an attempt to maximize the even distribution of the laser energy within the treatment zone. A commercially available 800 nm pulsed diode laser (Lumenis LightSheer®

DUETTM, Lumenis Ltd. Yokneam, Israel) was used for all studies. The laser was set to generate passes comprised of three consecutive pulses of 30 ms duration, at an energy density of 6 J/cm2 (46 J/ pulse; 138 J/pass) through its 22 mm by 35 mm exposure area. A total of twelve consecutive effective passes were administered, for a total of

~1650 J of energy per treatment area. Mild vacuum, a built-in function of the machine, was applied during the laser treatment. The skin surface temperature was monitored periodically with a thermal camera (FLIR E50; FLIR Systems, Wilsonville, OR). A cool dampened towel (~10 °C) was placed over the treated region every two passes for ten seconds, or more often if needed, to cool the skin surface. A thermocouple was periodically placed approximately 2 cm parallel to and below the skin surface, through one of the exposed edges of the tissue, to measure the internal bulk temperature. The same pre-liposuction treatment procedures were administered to both the Nanolipo and the Control group, with the only difference being the presence or absence of GNPs in the infusion solution, respectively.

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To perform the liposuction, a standard fat harvesting/transfer multi-port cannula (Tulip Medical, San Diego, CA) was connected to a commercially available liposuction vacuum pump (Grams, Newport Beach CA) and set to 28 mmHg.

Liposuction was first performed in the Control, then in the NanoLipo specimen.

Liposuction was performed for a defined amount of time (30 sec – 1 min), and varied depending on the initial amount of aspirate from the first sample from the Control group so that the target volume of about 40 cc would be harvested and be about the same as the amount used during tumescent infiltration. The time to suction the 40cc for the Control group was then used for the NanoLipo group. The liposuction procedure was performed similarly to the tumescent solution infiltration procedure, with suctioning from all four sides into the treatment zone as evenly as possible and with similar effort. For the last three specimens, practitioner 1, who had conducted the previous experiments under non-blinded conditions, performed both the NanoLipo and

Control pretreatments. Practitioner 2, an experienced board certified plastic surgeon with his primary practice in liposuction surgery, performed the liposuction procedures and made clinical determinations under blinded conditions between the NanoLipo and

Control specimens, regarding ease of liposuction, amount and speed of lipoaspirate harvest, and evenness of fat removal. Digital photographs detailing the protocol for this ex vivo human tissue liposuction model can be found in Fig. 3.S1.

Post-Treatment Assessments

The criteria used to assess clinical treatment outcomes in the tissue remaining after liposuction was manual palpation and visual inspection. During manual

145 palpation, the practitioner applied gentle digital pressure on the skin over the treatment zone to assess for the amount and evenness of fat removal. A comparison of evenness and any lumps or unexpected divots was made between each NanoLipo treated specimen and its bilaterally-matched Control. A pinch test was subsequently performed to further assess for the smoothness and evenness within the treatment region. After these initial assessments, the skin was surgically removed from the subcutaneous fat with a scalpel blade, much like harvesting a full-thickness skin graft, leaving only the liposuctioned fatty tissues behind. Inspection, palpation, visual determinations, and rank ordering of clinical criteria were completed using before and after skin-removal assessments by Practitioner 2. Digital photographs of the specimens and underlying adipose tissue were taken to document the achieved results. Surgical forceps were used to stretch and manipulate the post-liposuction specimens to aid in the inspection and visualization of the treatment outcomes.

Lipoaspirate was collected directly into conical centrifugal tubes within the suction canister and gently centrifuged at 1,000 RPM for 5 min to standardize the settling and separation of three distinct volume fractions. The top fraction contains the liberated triglyceride oils, the middle layer contains intact adipose tissue and cells with trapped free oil, and the bottom layer contains the tumescent fluid, blood cells, plasma, and stem cells. The middle layer was isolated and treated with collagenase type I

(powder, from Clostridium histolyticum) to digest the connective tissue and liberate the adipocytes and any remaining oil. The released fat cells were then stained with a

Calcein/Ethidium Homodimer live/dead assay to assess cell viability. This digestion and staining was performed by following standard protocols provided by the supplier

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(all reagents and kits purchased from Thermo Fisher Scientific, Waltham, MA).

Optical microscopy was used to image the cells and tissues.

In a related experiment, the viability of intact adipose tissue remaining after liposuction was assessed with Nitroblue Tetrazolium Chloride (NBTC) staining, as previously reported.28, 36 Briefly, NBTC is a dye that stains for mitochondrial activity such that metabolically active tissue will be stained a deep-blue color while photothermal-induced reduction of metabolic activity would lead to loss of staining. A thin tissue slice approximately 1.0 cm thick bisecting the treatment region along the vertical laser beam axis was isolated, stained, and photographed. To assess the distribution of infused GNP solution, a micro-Computed Tomography scanner (µCT,

Skyscan 1076, Kontich, Belgium) was used to image adipose tissue specimens that were either injected or not injected with solution. A calibration curve involving air, adipose tissue, saline, and various concentrations of GNPs was constructed to help identify the structures imaged. Imaging was performed at 18 μm isotropic voxel size at an electrical potential of 50 kVp and current of 200 μA using a 0.5 mm aluminum filter with a beam hardening correction algorithm applied during image reconstructions.

3.4 RESULTS

Assessment of Treatment Outcomes under Non-Blinded Conditions

During the tumescent step, no differences were seen between the NanoLipo and the Control saline only groups. During the laser treatment step, there was a minimal rise in the skin temperature of approximately 10°C in the Control group after

147 the complete 12 passes, while in the NanoLipo group, the skin surface temperature typically rose quickly after only a few passes (Fig. 3.2A). The most significant heating was observed at the skin surface in non-study pilot experiments for the NanoLipo group when towel cooling was not performed, with temperatures exceeding 40 °C or more at local hot spots matching the pattern of the diodes in the laser aperture after each pass that would quickly equilibrate with the surrounding tissue. The heating was significantly mitigated with towel cooling, although the hot spots were still observed.

In the Control group, skin surface temperature never exceeded 40 °C regardless of whether towel cooling was performed or not. Additionally, thermocouple readings for internal bulk temperature for the NanoLipo group barely exceeded 30 °C at the end of twelve passes, while the internal bulk temperature for the Control group was unchanged from the starting temperature.

During the liposuction step, in general, the NanoLipo treated specimen appeared to be better liposuctioned and more evenly reduced than the Controls. In thicker samples, the practitioners most often found a clinical difference in the ease with which the suction cannula moved to remove the fat for the NanoLipo group, as well as an increased rate at which fat was removed compared to the Control.

Practitioner 1, a non-plastic surgeon researcher, on average rated the NanoLipo and

Control procedures 5/10 and 7/10 in terms of difficulty, respectively, on a scale of 1 to

10 with one being very easy and ten being very difficult. Practitioner 2, a trained plastic surgeon, on average rated the NanoLipo and Control procedures 4/10 and 5/10 in terms of difficulty, respectively, on the same grading scale. Both practitioners noted

148 that the improvement in the ease of suction was markedly greater for more fibrous specimens.

The immediate appearance of the NanoLipo treatment area was the presence of a clearly defined depressed region approximately matching the shape and size of the 4 cm x 7 cm laser treatment area (Fig. 31A, dashed rectangle). This feature was only occasionally present, and to a much lesser extent, in the Control specimens (Fig.

3.1B). In some thinner specimens, nearly all of the adipose tissue within the treatment zone was removed for the NanoLipo group, resulting in a full-thickness void spanning most of the sample (Fig. 3.S2). In the Control treatments, this full thickness tissue removal was not seen and the indentations were always to a lesser extent. When visually examined from the deep side (turned up-side-down), or when the skin was surgically removed from the specimens, the NanoLipo treated specimens were more uniform, smoother, and had proportionally more fibrous tissue strands exposed, with fewer lumps and smaller, and much finer residual fat lobules compared to Control, as can be clearly seen in Fig. 1C and 1D, respectively.

After the lipoaspirate was centrifuged, three distinct volume fractions were observed. The free fatty triglyceride oil volume fraction was situated at the top in the

NanoLipo treated lipoaspirate. In some cases, it accounted for nearly half of the total lipoaspirate tissue and oil volume (not including the aqueous phase), while this fraction was practically non-existent in the Control lipoaspirate (Fig. 3.2B). Liberated triglycerides were consistently observed for NanoLipo treated specimens, as clearly seen in the liposuction tubing in Vid. 1. Liberated triglycerides in the NanoLipo treated specimen were even observed for the single high-Fitzpatrick skin type

149 specimen that was tested (Fig. 3.S4). Adipocytes isolated from the middle fatty tissue volume fraction were viable, as indicated by the green staining imparted to the cells

(Fig. 3.2C).

Assessment of Treatment Outcomes under Single-Blind Conditions

Clinical assessment of the NanoLipo versus the Control treated specimens was undertaken. In a single-blind series with three consecutive cases, Practitioner 2 noted that one of the specimens was markedly thinner and had significantly fewer lumps after suctioning. After skin removal, the specimen that was thinner was also much smoother within the treated region. Fig. 3 displays a representative comparison between two blinded specimens from one set of tissues. As shown, the NanoLipo specimen is smooth, even, and has smaller fatty lobules (Fig. 3.3A) while the Control specimen has many large fat lobules and pearls clearly visible throughout the region

(Fig. 3.3B). A demarcation (dotted line) that matched well with the edge of the treatment zone was observed in the NanoLipo treated specimen, beyond which distinctive large fat pearls similar to those found throughout the Control specimen are clearly observed (arrows). Practitioner 2 noted that the first blinded specimens

(NanoLipo) was marginally easier to suction than the second (Control) and that the former (NanoLipo) had regions where the difficulty of suction was significantly reduced. Although Practitioner 2 indicated that he would not have been comfortable making a determination based on the ease of suction alone, he was able to identify the

NanoLipo treated specimen correctly based on the obvious visual and physical differences after suctioning.

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NanoLipo Mechanism Investigations

Upon infusion into adipose tissue, photothermal solution expanded the interstitial spaces between the connective tissue membranes separating pearls of fat.

Fig. 3.4A, 3.4C, 3.4E, and 3.4B, 3.4D, 3.4F respectively show the sagittal, transverse, and coronal views of µCT images of tissue specimens that were or were not infused with GNP photothermal solution. The relatively darker spaces represent air, the relatively grey spaces represent fat-rich regions, and the relatively white spaces represent water-rich regions. Scarpa’s fascia can be seen in the transverse and sagittal views as a relatively thicker white line parallel to the skin surface. In Fig. 4A, 4C, and

4E, connective tissue regions have been expanded by infused solution while the same connective tissue separating pearls of fat in the un-injected Control specimen remains collapsed. The expanded regions were, in some cases, up to 2 mm in size. Consistent with the observed distribution of photothermal solution, there was a loss of staining in some regions of the NanoLipo treated specimen (Fig. 3.5A, arrows), whereas the whole specimen was stained a deep-blue color for the Control specimen.

3.5 DISCUSSION

We previously reported the ability of NanoLipo to produce aesthetically superior results compared to saline in our porcine model.27 These promising results motivated us to develop an ex vivo human liposuction model to investigate whether

NanoLipo could be successfully used as an adjunct for fat removal in human tissue specimens before proceeding with human clinical trials. One of the distinctive findings in the current study was the presence of liberated triglycerides in the lipoaspirate of

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NanoLipo treated specimens before any post-processing (Fig. 3.2B, left). We speculate that a specific form of the photothermal effect, called selective photothermolysis (SPT), may be the mechanism responsible.29, 37 Briefly, SPT posits that short bursts of light applied over a wide area can selectively and precisely heat up pigmented target structures based on their geometry without significant heating of the surroundings. For example, a 3 ms pulse of light can be used to preferentially heat spherical targets with diameters 100 µm or smaller, while a 300 ms pulse can heat structures up to 1 mm.37 Incidentally, adipocytes, as well as some of the fibrous membrane support structures in adipose tissue, are between 100200 µm in their native state. Thus, they may be readily targeted by this selective photothermal phenomenon. We believe that this principle, which has been successfully applied to target the melanin in hair follicles during laser hair removal,38 can be applied to liposuction, as well.

The first observation in support of this mechanism is the lack of significant bulk heating in the adipose tissue of the NanoLipo treated group (Fig. 3.2A). The most significant heating for both the NanoLipo and Control specimens was observed at the skin surface. For the measurements used to construct the heating profile in Fig. 3.2A, the thermal camera was aimed at the local hot-spots in the pattern generated by the laser aperture, which were typically at least 10 °C hotter, or more, than the surrounding tissue. This inhomogeneous heating was manageable when a damp towel was used to immediately cool the surface after laser exposure to quickly thermal- equilibrate the hot-spots with the surrounding tissue to prevent the temperature from exceeding 40 °C for an extended period of time. Heating of the skin surface may be

152 mitigated with the use of a chilled applicator. These results are consistent with what is historically regarded as the greatest limitation in laser-based biomedical applications, i.e., the limited penetration depth of light into tissue.35 Even wavelengths of light in the middle of the biological optical window around 800 nm, which was specifically chosen for this application, can penetrate only a few centimeters into tissue at maximum. Nevertheless, the light that does successfully penetrate the skin seems to be sufficient in producing a photothermal effect, as evidenced by the presence of liberated triglyceride even in the specimen with the highest Fitzpatrick skin type, which would have the lowest amount of light energy irradiating the GNPs within the fatty tissues among all of the tested specimens (Fig. 3.S4).

Another clue for the mechanism of NanoLipo may be the way that infused

GNP solution distributes within adipose tissue. Adipose tissue is organized into several hierarchical levels.39 At the lowest level are the adipocytes, which are packed into lobules that are further nested within larger structures, called pearls. Many pearls packed together to form a section, such as the superficial or deep subcutaneous adipose tissue sections that are separated by a thick membrane called Scarpa’s fascia.

In Fig. 4, µCT images showed that the infused GNP fluid preferentially distributed between and expanded the interstitial spaces between fat pearls. This result is expected, since it would be unfavorable for the hydrophilic GNP fluid to be well- mixed with the highly hydrophobic fat regions of adipose tissue. Upon exposure to external pulses of light through the skin, these photothermal fluid-filled compartments are heated up preferentially, which is supported by the loss of metabolic activity shown by NBTC staining precisely at this structural level (Fig. 3.5A). Additionally,

153 the deep blue color expressed throughout the remainder of the NanoLipo treated specimen indicates that full-thickness fat loss does not occur and that selective damage occurs only where the photothermal agents are deposited, while staining in the entirety of the Control specimen (Fig. 3.5B) indicates that the laser is well tolerated.

These results taken together suggest a mechanism by which thermal stress imparted to the connective tissue network surrounding and intertwined within adipose tissue weakens those structures, while also acting to disrupt the lipocytes themselves, thus facilitating the removal of both the fat and parts of the fibrous tissue support structure during liposuction. The photothermal heating imparted to the tissue likely extends beyond the fibrous tissue membrane level through conduction, and may induce thermal lipolysis. We believe that it is this lipolysis that is responsible for the difference in liberated triglycerides we observed between the NanoLipo and the

Control.40 This specific mechanism we propose is consistent with all of the observations presented in this study, i.e., the presence of liberated triglyceride and the lack of large fat pearls accompanied by the thinner and smoother appearance for the

NanoLipo treated specimens.

The photothermal phenomenon, which has been studied extensively in the context of cancer therapy26, has only been sparsely explored in humans or human derived specimens. In one study, Tuchin et al. stained thin human adipose tissue sections with a photothermal dye, indocyanine green, and exposed the specimen to over three hours of light exposure. The continuous and long-term exposure eventually resulted in “destructive engineering” of adipose tissue, but certainly cannot be applied realistically or practically in the context of liposuction.41 Another study involves the

154 investigation of a 1060 nm pulsed Nd:YAG laser administered through the tip of fiber optic cable attached to a liposuction cannula, for laser lipolysis through photothermal heating of resident water in tissues.42 The one-dimensional point-source nature of the laser aperture does not produce uniform heating when it is used to cannulate adipose tissue. Another limiting factor of this method is that the resonance wavelength is tuned to optical absorption of water molecules in tissue, which means that significant care must be taken not to overheat the targeted or non-targeted tissues. In fact, significant carbonization of adipose tissue was often observed as a result of this treatment.42

Therefore, although this technology has been approved by the FDA and marketed as

SmartLipo (Cynosure, Westford, MA),43 it has not gained popularity among plastic surgeons due to its lack of efficacy and no demonstrable improved result over traditional liposuction techniques, with the added potential for thermal injury. In contrast, NanoLipo treatment overcomes this shortcoming by applying the light energy in a two-dimensional nature over the target area that is infused with GNPs, resulting in uniform and even heating of the adipose tissue, as was demonstrated in this study.

Additionally, one question that was unanswered at the time of the porcine study was whether the injected nanoparticles induced any potentially harmful side-effects. We have since conducted a six-month biodistribution and biocompatibility study to assess any toxicity concerns based on the finding that the majority of nanoparticles are eliminated from the bodies of pigs, and that the minuscule amounts of residual particles exhibit no observable signs of toxicity in our porcine model and with other studies.44

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Ultimately, the outcome of liposuction surgery depends, first and foremost, on the provider experience, skill, and training.23 Other factors, such as patient tissue history, including repeat procedures, scarring in the treatment area, or the location on the body45, as well as the amount of time and care taken, and the amount of tissue to be removed, all play important roles. However, even for highly trained plastic surgeons, significant time, effort, and attention to detail are needed to carefully suction the tough to treat fibrous or large areas to produce aesthetically pleasing results. Many plastic surgeons opt to perform two or more rounds of suctioning within the same session, with latter rounds involving the use of finer cannulas to extract the tough fat residing within especially fibrous pockets of connective tissue that lead to the formation of lumps and bumps post-treatment if left unaddressed.39 In the current study, the practitioners were allotted at most only a few minutes to perform the simulated liposuction without multiple rounds of suctioning using finer cannulas. We believe and have substantiated with our data that NanoLipo treatment was able to generate superior outcomes compared to the Control. If these results can be extrapolated to larger treatment sizes in clinical settings, we believe the photothermal pre-treatment process of NanoLipo applied to liposuction may be a useful tool for liposuction practitioners.

3.6 CONCLUSION

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Our goal in this study was to investigate NanoLipo primarily as a technique to hasten, facilitate, and improve the liposuction procedure and outcomes. Based on the observed results, we have shown that photothermal pretreatment of adipose tissue is effective in assisting the more rapid and efficient removal of fat while producing aesthetically superior outcomes compared to conventional suction assisted lipectomy in excised human tissue specimens. We believe that with this current data in hand, our next step is to evaluate NanoLipo in human clinical pilot studies while examining its short-term and long-term indications, advantages, results, and potential complications.

3.7 CONFLICT OF INTEREST

This work was supported in part by eLux Medical Inc., a start-up company with financial interest in this technology.

3.8 ACKNOWLEDGEMENTS

The authors would like to thank the private practices of Mark M. Mofid and

William and Jeffrey Umansky for providing some of the tissue specimens used for this work. The authors would also like to thank NIH R01EY024134, Air Force office of

Scientific Research (AFOSR) FA9550-15-1-0273, and eLux Medical Inc. for financial support. This chapter is in part currently being prepared for submission with the

American Society for Plastic Surgeon’s Aesthetic Surgery Journal. Wangzhong Sheng,

William J. Seare, Ali H. Alhasan, Esther Cory, Paul Chasan, Robert Sah, Khalid

Almutairi, Barry DiBernardo, and Adah Almutairi, “A Single-Blind Study Evaluating the Efficacy of Gold Nanoparticle Photothermal-Assisted Liposuction in an ex vivo

157

Human Tissue Model.” Oxford Academic 2017. The dissertation author was the primary investigator and author of this work.

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3.10 FIGURES

Figure 3.1. NanoLipo procedure results in the removal of visibly more fat, indicated by A. A clearly visible depressed region within the underlying adipose tissue matching the shape and size of the laser aperture (dotted rectangle). B. The same feature is not present in the Control (saline + laser + liposuction) treated specimen. C. Proportionally more fibrous tissue is observed in the NanoLipo treated specimen with very few fat lobules remaining, while. D. many fat lobules adhering tightly to the connective tissue membranes are observed for the Control specimens.

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Figure 3.2. A. Heating profiles of tissue specimens for NanoLipo and Control pretreatments, measured externally and internally for NanoLipo, and externally only for Control. The NanoLipo external temperatures were recorded for local hotspots matching the pattern of the laser diode. Both NanoLipo internal measurements and Control skin surface measurements showed mild heating, while NanoLipo external measurements showed the most heating. B. After centrifugation, the liberated triglyceride phase accounted for nearly 50% of total lipoaspirated tissue volume for the NanoLipo group (left tube), while almost no liberated triglyceride was observed in the Control (right tube). C. Calcein/ethidium homodimer live/dead assay; tissue fluorescing green from the calcein dye shows that the majority of the tissue is viable, while minimal red staining, indicative of dead cells, was observed.

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Figure 3.3. (Screen Shot) Liberated triglyceride in lipoaspirated tissues are a hallmark of the NanoLipo procedure. Triglycerides liberated as a result of NanoLipo treatment flows through a suction tube for a NanoLipo treated specimen compared to lipoaspirate from conventional suction-assisted lipectomy.

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Figure 3.4. Smoothness within the treatment zone was used as a key marker to distinguish the NanoLipo treated specimen from Control specimens under blinded conditions. Digital photographs for one set of blinded experiments showing adipose tissue after liposuction and skin removal are shown. A. NanoLipo treated specimen was even in appearance and was absent of fat lobules, whereas in B. large and lumpy fat lobules are clearly visible and can be felt throughout the Control specimen. Near the edge of the NanoLipo specimen outside of the treatment zone (below the dotted line in A.), several large fat pearls are clearly visible (arrow).

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Figure 3.5. Micro computer-aided tomography images of the distribution of infused GNP solution within adipose tissue compared to non-injected Control. A., B. show the sagittal views, C., D. show the transverse views, and E., F. show the coronal views of injected and non-injected adipose tissue specimens, respectively. The relatively darker spaces represent air, the relatively grey spaces represent fat-rich regions, and the relatively white spaces represent water-rich regions of adipose tissue, such as connective tissue membranes. Infused solution preferentially disperses between and expands into the interstitial spaces between fat pearls (A., C., and E.), while the interstitial spaces between connective tissue of non-injected adipose tissue specimens remain collapsed in their native state (B., D., and E.). Scarpa’s fascia can be seen in the transverse and sagittal views as a thicker white line parallel to the skin surface.

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Figure 3.6. Absence of nitroblue tetrazolium staining (arrows) shows selective thermal damage of connective tissue membranes at the fat pearl level in adipose tissue due to photothermal inactivation of intracellular reductases, and is consistent with the way GNP solution distributes within iinterstitial spaces of adipose tissue. Specimens were subjected to A. NanoLipo or B. Control treatment, vertically bisected with a surgical scalpel through the optical exposure axis to isolate thin slices approximately 0.5 cm thick, and stained. Complete staining of the Control specimen, as well as staining of the overlying skin and deeper adipose tissue regions in the NanoLipo treated specimen, demonstrates that these tissues were unaffected and that photothermal damage was selectively restricted to the fibrous connective tissue membranes at the pearl level.

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A B

C D

Figure 3.S1. An ex vivo human tissue model was developed to test the ability of NanoLipo pretreatment to aid liposuction. The NanoLipo procedure involves infusion of gold nanorod solution into adipose tissue, followed by exposure to laser pulses matching the resonant wavelength of the nanorods at 800 nm through the surface of the skin, to heat and melt fat to facilitate its liberation and removal during liposuction. A. The set of tools, including infusion and suction cannulas, used for this study. B. A specimen approximately 3-4 cm thick with a treatment zone marked for treatment. C. An external laser was applied to the specimen; inset – surface temperature was monitored with a thermal camera to insure that temperatures did not exceed 40 ᵒC. D. Simulated liposuction was performed on excised tissue specimens for specified amounts of time for the NanoLipo and control groups.

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Figure 3.S2. Thin tissue specimens with skin removed after treatment. A. A void spanning the full thickness of the adipose tissue was observed for the NanoLipo specimen after skin removal, while B. the treated region was only slightly depressed for the control specimen.

Figure 3.S3. Thicker tissue specimens exhibited large indentations in the skin after treatment despite equal effort being exerted evenly throughout the treatment region. Indentations from two different tissue specimens are shown.

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Figure 3.S4. A. A set of high Fitzpatrick skin-type specimen were treated. Aggressive towel cooling had to be performed to ensure skin surface temperatures remained low. B. Liberated triglyceride (arrow) was observed in lipoaspirate of the NanoLipo treated specimen (left tube), while none was observed in the control lipoaspirate (right tube).

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Table 3.1. Physical characteristics of specimens tested.