This item is only available as the following downloads:
1 VITAMIN E LOADED SILICONE HYDROGEL CONTACT LENSES FOR EXTENDED OPHTHALMIC DRUG DELIVERY By CHENG CHUN PENG A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT OF THE REQUIRE MENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY UNIVERSITY OF FLORIDA 2011
2 2011 Cheng Chun Peng
3 To my dearest Mom and Dad, and all my family
4 ACKNOWLEDGMENTS It would have been next to impossible to finish this dissertation without the sup port, patience and guidance of the following people through these years. It is to them that I owe my deepest gratitude. First and foremost, I would like to thank my doct oral advis or, Dr. Anuj Chauhan for not only his guidance on my research but also all h is encouragement on my life in Gainesville. He taught me the impo rtance of keeping motivated and efficient on my work, giving me enough freed om to explore, fail and improve. Most of all he has been a good friend of mine for which I will cherish I also w ish to extend many thanks to my other doctoral committee members, Dr. Tanmay Lele, Dr. Peng Jiang, and Dr. Gregory S. Schultz for their insightful viewpoints and willingness to participate in my doctoral review process. In addition, I would like to thank D r. Tanmay Lele and Dr. Spyros A. Svoronos for the opportunity to serve as a teaching assistant under their guidance. I would also express my gratitude to Dr. Caryn Plummer for her kindly hel p in providing me the opportunity to conduct ani mal studies for my research I am extremely grateful to Dr. Yiider Tseng for his gracious support as both a scholar and a friend through my years in Gainesville. I would also like to thank Dr. Nae Lih Wu in Nat ional Taiwan University for encouraging me to continue doctoral research in United States group. First of all, I would like to extend my special thanks to Dr. Jinah Kim, wh o not only helped me establish the foundation of th is dissertation, but also served as my role model by showing me how to cherish and enjoy your life. I thank Dr. Yash Kapoor, Dr. Brett Howell and Dr. Chhavi Gupta for their pre cious help in my research. The demonstration of leadership, discipline, dedication and commitmen t toward work and life
5 they showed me are one of the invaluable treasu res I have obtained in my research career in Gainesville Hyun Jung Jung has been a great friend and shared my best an d worst time in my research for which I am extremely grateful I al so would like to thank Michael Burke for his help in the lidocaine delivery experiments. Finally I thank the current group members, Loki, Han and Ming, for sharing in my last days in Gainesville. Many current and former staff members of the Department of C hemical Engineering were also very accommodating during my time in Gainesville. I would like to gratefully acknowledge the technical support for this research by Sean Poole, James Hinnant and Dennis Vince. I would also like to thank Shirley Kelly, Deborah Sandoval, Melissa Fox, and Carolyn Miller for their assistance during these years. I am truly grateful to have so many wonderf ul friends in my life to stand by me through all the good times and bad on the way to what I am now. Thanks my friends and my fami ly in Gainesville: Rob, Akhil, Poom, Jun, Can, Derek, Wei Chiang, Tzung Hua, Jack and Hungta and everyone whom I have had the privilege to share my life with all these years. It is impossible to accomplish my work with out all of you r kindly support Fina lly, word s cannot describe my everlasting gratefulness to my big loving family in Taiwan, especially to my parents, Yueh Chin Tai and Kuo Yuan Peng for their unconditionally trust on every single decision I made in my whole life, and I hope I have made you proud.
6 TABLE OF CONTENTS page ACKNOWLEDGMENTS ................................ ................................ ................................ .. 4 LIST OF TABLES ................................ ................................ ................................ ............ 9 LIST OF FIGURES ................................ ................................ ................................ ........ 10 LIST OF ABBREVIATIONS ................................ ................................ ........................... 14 1 INTRODUCTION ................................ ................................ ................................ .... 18 2 CHARACTERIZATION OF VITAMIN E LOADED SILICONE HYDROGEL ............ 30 2.1 Materials and Methods ................................ ................................ ...................... 30 2.1.1 Materials ................................ ................................ ................................ .. 30 2.1.2 Vitamin E Loading into Pure Lenses ................................ ........................ 31 2.1.3 Ion Permeability Measurements ................................ .............................. 31 2.1.4 Oxygen Permeability Measurements ................................ ....................... 32 2.1.5 Transmittance Measurement of Vitamin E Loaded Contact Lens ............ 33 2.1.6 Preparation of Silicone Hydrogel ................................ ............................. 33 2.1.7 Mechanical Properties Measurements ................................ .................... 34 2.2 Results and Discussion ................................ ................................ ..................... 34 2.2.1 Vitamin Loadings in the Lenses ................................ ............................... 34 2.2.2 Transparency of the Vitamin E laden Contact Lenses ............................ 35 2.2.3 Water Content of Pure and Vitamin E Loaded Lenses ............................ 35 2.2.4 Size Change Due to Vitamin E Loading ................................ .................. 36 2.2.5 Ion Permeability of Vitamin E Loaded Lenses ................................ ......... 37 2.2.6 Oxygen Permeability of Vitamin E Loaded Lenses ................................ .. 38 2.2.7 Transmittance of Vitamin E Loaded Lenses ................................ ............ 41 2.2.8 Mechanical Properties of Vitamin E Loaded Silicone Hydrogel ............... 42 3 HYDROPHILIC DRUG DELIVERY BY VITAMIN E LOADED SILICONE HYDROGEL ................................ ................................ ................................ ............ 55 3.1 Materials and Methods ................................ ................................ ...................... 55 3.1.1 Materials ................................ ................................ ................................ .. 55 3.1.2 Drug Loading into Pure Lenses ................................ ............................... 56 3.1.3 Vitamin E Loading into Pure Lenses ................................ ........................ 56 3.1.4 Drug Loading into Vitamin E Loaded Lenses ................................ .......... 57 3.1.5 Drug Release Experiments ................................ ................................ ...... 57 3.2 Results and Discussi on ................................ ................................ ..................... 57 3.2.1 Dynamics of Drug Transport from Contact Lenses without Vitamin E ..... 57 3.2.2 Dynamics of Drug Transport from Vitamin E Loaded Lenses .................. 59 184.108.40.206 Timolol Vitamin E loaded lenses ................................ .................... 59
7 220.127.116.11 DXP Vitamin E loaded lenses ................................ ........................ 62 18.104.22.168 Fluconazole Vitamin E loaded lenses ................................ ............ 63 3.2.3 Model for Hydrophilic Drugs ................................ ................................ .... 64 3.2.4 Dif fusivities of Drugs in Vitamin E Loaded Lenses ................................ .. 66 4 HYDROPHOBIC DRUG DELIVERY BY VITAMIN E LOADED SILICONE HYDROGEL ................................ ................................ ................................ ............ 82 4.1 Mater ials and Methods ................................ ................................ ...................... 83 4.1.1 Materials ................................ ................................ ................................ .. 83 4.1.2 Drug Loading into Pure Lenses ................................ ............................... 83 4.1.3 Vitamin E Loading into Pure Lenses ................................ ........................ 83 4.1.4 Drug Loading into Vitamin E Loaded Lenses ................................ .......... 84 4.1.5 Drug Release Experiments ................................ ................................ ...... 84 4.1.6 Viscoelastic Measurement ................................ ................................ ....... 85 4.2 Results and Discussion ................................ ................................ ..................... 85 4.2.1 Dynamics of Drug Transport from Contact Lenses without Vitamin E ..... 85 4.2.2 Dynamics of Drug Transport from Vitamin E Loaded Lenses .................. 86 4.2.3 Diffusivities of Drugs in Vitamin E Loaded Lenses ................................ .. 90 4.2.4 Scaling Model for Effect of Vitamin E Loading on Extended DX Delivery ................................ ................................ ................................ ......... 92 5 ANESTHETICS DELIVERY BY VITAMIN E LOADED SILICONE HYDROGEL ... 105 5.1 Materials and Methods ................................ ................................ .................... 107 5.1.1 Materials ................................ ................................ ................................ 107 5.1.2 Drug Loading into Pure Lenses ................................ ............................. 107 5.1.3 Vitamin E Loading into Pure Lenses ................................ ...................... 108 5.1.4 Drug Release Experiments ................................ ................................ .... 109 5.1.5 Silicone Hydrogel Preparation ................................ ............................... 109 5.1.6 Par tition Coefficient ................................ ................................ ............... 110 5.1.7 Determination of Critical Micelle Concentration (CMC) of Lidocaine ..... 111 5.2 Results and Discussion ................................ ................................ ................... 111 5.2.1 Dynamics of Drug Release from Contact Lenses ................................ .. 111 22.214.171.124 Drug uptake through drug PBS solution ................................ ....... 111 126.96.36.199 Drug uptake through drug ethanol solution ................................ .. 114 5.2.2 Lidocaine Release Study (Surfactant Behavior) ................................ .... 115 5.2. 3 Mechanisms of Extended Drug Release by Vitamin E Loaded Contact Lens ................................ ................................ ................................ ............ 117 6 ION TRANSPORT OF SILICONE HYDROGEL ................................ .................... 138 6.1 Materials and Methods ................................ ................................ .................... 143 6.1.1 Materials ................................ ................................ ................................ 143 6.1.2 Preparation of Silicone Hydrogel ................................ ........................... 143 6.1.3 Water Fraction Measurements ................................ .............................. 144 6.1.4 Ion Permeability Measurements ................................ ............................ 145
8 6.1 .4.1 Salt release in perfect sink (kinetic desorption). ........................... 14 5 188.8.131.52 Ion transport in diffusion cell (direct permeation) ......................... 145 6.2 Results and Discussion ................................ ................................ ................... 146 6.2.1 Comparison of Transport Measurements from the Kinetic and Permeation Approaches ................................ ................................ .............. 146 184.108.40.206 Kin etics of salt release in perfect sink ................................ .......... 146 220.127.116.11 Ion transport through permeation in a diffusion cell ..................... 150 6.2.2 Effect of Compo sition of Silicone Hydrogel on Ion Permeability ............ 154 7 CYCLOSPORINE DELIEVERY BY SILICONE HYDROGEL FOR CRONIC DRY EYE SYMDROME ................................ ................................ ................................ 185 7.1 Materials and Methods ................................ ................................ .................... 187 7.1.1 Materials ................................ ................................ ................................ 187 7.1.2 Drug Loading into Contact Lens ................................ ............................ 187 7.1.3 Drug Release Experiments from Lenses Loaded with CyA ................... 188 7.1.4 Vitamin E Loading into Contact Lens ................................ ..................... 188 7.2 Results ................................ ................................ ................................ ............ 189 7.2.1 Drug Uptake by Pure Contact Lens ................................ ....................... 189 7.2.2 Drug Release by Pure Contact Lens ................................ ..................... 191 7.2.3 Drug Uptake by Vitamin E Loaded Contact Lens ................................ .. 191 7.2.4 Drug Release by Vitamin E Loaded Contact Lens ................................ 192 7.3 Discussion ................................ ................................ ................................ ...... 193 7.3.1 Release Mechanism and Model Fitting ................................ .................. 194 7.3.2 Therapeutic Release Rates ................................ ................................ ... 198 8 DRUG DELIVERY BY CONTACT LENS IN GLAUCOMATOUS DOGS ............... 211 8.1 Materials and Methods ................................ ................................ .................... 213 8.1.1 Materials ................................ ................................ ................................ 213 8.1.2 Drug and Vitamin E loading into Contact Lens ................................ ...... 213 8.1.3 Animal Model ................................ ................................ ......................... 214 8.1.4 Data Analysis ................................ ................................ ........................ 215 8.2 Results ................................ ................................ ................................ ............ 216 8.2.1 Contact Lens without Vitamin E ................................ ............................. 216 8.2.2 Contact Lens with Vitamin E ................................ ................................ .. 216 8.2.3 Eye Drop ................................ ................................ ............................... 217 8.2.4 Drug Administration Methods Comparison ................................ ............ 217 8.3 Discussion ................................ ................................ ................................ ...... 218 9 CONCLUSIONS ................................ ................................ ................................ ... 225 LIST OF REFERENCES ................................ ................................ ............................. 227 BIOGRAPHICAL SKETCH ................................ ................................ .......................... 240
9 LIST OF TABLES Table page 2 1 List of silicone hydrogel extended wear commercial contact lens (dipoter 6.50) explored in this study (n = 6). ................................ ................................ .... 54 3 1 Model parameters obtained by fitting experimental da ta to the model ................ 81 4 1 Partition coefficient ( K ) of DX in lenses soaked in DX PBS solution. ................ 104 6 1 Composition of silicone hydr ogel. ................................ ................................ ..... 179 6 2 Parameters of different silicone hydrogels. ................................ ....................... 180 6 3 Parameters for GEL A3 with various sodium chloride concentrati ons for salt loading. ................................ ................................ ................................ ............. 181 6 4 Fitting results of ion transport by diffusion cell for silicone hydrogels. All samples are 0.13 mm thick and contain no preloaded salt. .............................. 182 6 5 Fitting results of ion transport by diffusion cell for Gel A3 with various sodium chloride concentrations in the donor compartment. ................................ .......... 183 6 6 Fittin g results of ion transport by diffusion cell for silicone hydrogels which pre soaked in sodium chloride solution with various concentrations. ............... 184 7 1 Results of CyA uptake by silicone contact lens. ................................ ............... 209 7 2 Results of CyA uptake by Vitamin E loaded ACUVUE OASYS TM lenses. ....... 210 8 1 Summary of various drug delivery methods considered in this study. .............. 224
10 LIST OF FIGURES Figure page 1 1 Schematic illustration of ophthalmic drug delivery through eye drops. ............... 28 1 2 Schematic illustration of the microemulsion laden contact lens inserted in the eye. ................................ ................................ ................................ ..................... 29 2 1 Correlation of Vitamin E loading and concent ration of soaking solution for different lenses. ................................ ................................ ................................ .. 44 2 2 Images of Commercial NIGHT&DAY TM contact lens and NIGHT&DAY TM lens with 30%Vitamin E l oading ................................ ................................ ................ 45 2 3 Plot of water content ( Q ) and EW of Vitamin E loaded lenses versus Vitamin E loading. ................................ ................................ ................................ ........... 46 2 4 Percent incr ease in diameter of dry lenses and wet lenses before and after loading Vitamin E. ................................ ................................ ............................... 47 2 5 Effect of Vitamin E loading on ion permeability of lenses. ................................ .. 48 2 6 Effect of Vitamin E loading on oxygen permeability (Dk). ................................ ... 49 2 7 Transmittance spectrum for commercial contact lenses. ................................ .... 50 2 8 Transmittance spectrum for NIGHT& DAY TM and ACUVUE OASYS TM with different Vitamin E loading. ................................ ................................ ................. 51 2 9 Transmittance spectrum of NIGHT&DAY TM with Vitamin E loading .................... 52 2 10 Dependence of storage module of Vitamin E loaded silicone hydrogel on frequency.. ................................ ................................ ................................ .......... 53 3 1 Effect of timolol loading method on profile of timolol release by commercial contact lenses.. ................................ ................................ ................................ ... 70 3 2 Profiles of timolol release by Vitamin E loaded contact lenses. .......................... 71 3 3 Profiles of repeated timolol releases by Vitamin E loaded contact lenses. ......... 73 3 4 Profiles of DXP release by V itamin E loaded contact lenses ............................. 74 3 5 Profiles of fluconazole release by V itamin E loaded contact lenses ................... 76 3 6 Drug release duration increase by Vitamin E loaded contact lenses. ................. 78
11 3 7 Plot of % t imolol release by Vitamin E loaded NIGHT&DAY TM versus square root of time. ................................ ................................ ................................ ......... 80 4 1 Effect of DX loading method on profile of DX release by contact lenses... ......... 96 4 2 Profiles of experimental and model fitted DX uptake and release by Vitamin E loaded contact lenses.. ................................ ................................ ....................... 97 4 3 Plot of % drug release by Vitamin E loaded le nses against square root of time. ................................ ................................ ................................ .................. 99 4 4 Fitted DX diffusivity and partition coefficient for contact lenses with diffe rent Vitamin E volume fraction ................................ ................................ ................ 101 4 5 Effec t of Vitamin E volume fraction on increase in drug uptake times. ............. 102 4 6 Dependence of the loss modulus G" on frequency for pure Vitamin E ............. 103 5 1 Molecular structures of model drugs. ................................ ................................ 123 5 2 Lidocaine release in PBS by O 2 OPTIX TM with various Vitamin E loading. ........ 124 5 3 Vitamin E release from O 2 OPTIX TM during lidocaine release in PBS. .............. 125 5 4 Bupivacaine release in PBS by O 2 OPTIX TM with various Vitamin E loading. .... 126 5 5 Short time and long time tetracaine release in PBS by O 2 OPTIX TM with various Vitamin E loadings. ................................ ................................ .............. 127 5 6 Lidocaine bup ivacaine and tetracaine release in PBS by O 2 OPTIX TM with various Vitamin E loading. ................................ ................................ ................ 128 5 7 Lidocaine release by O 2 OPTIX TM with 0.36g Vitamin E/g pure lens. ................ 130 5 8 The relationship between surface tension and lidocaine concentration in PBS. ................................ ................................ ................................ ......................... 131 5 9 The calculated partition coefficient (K) of Vitamin E loaded silicone hydrogel at various lidocaine hydrochloride concentrations. ................................ ........... 132 5 10 The relationship between the lidocaine partition coefficient in Vitamin E (K VE ) and the bulk drug concentration. ................................ ................................ ...... 133 5 11 Model fitting for anesthetic drug release increase ratio on Vitamin E loading fraction in the silicone hydrogel. ................................ ................................ ....... 134 5 12 Lidocai ne release from pure 0.2 mm thick silicone hydrogel or gel with Vitamin E loading. ................................ ................................ ............................. 135
12 5 13 Lidocaine release by silicone hydrogel with or without Vitamin E loadin g with various thickness.. ................................ ................................ ............................ 136 5 14 Lidocaine release by silicone hydrogel (with or without Vitamin E loading with various thicknesses. ................................ ................................ ......................... 137 6 1 NaCl release profile and model prediction (solid line) of different silicone hydrogel in perfect sink. ................................ ................................ .................... 161 6 2 NaCl release profile of 0.13 mm thick Gel A3 with different pre soaking NaCl concentration in perfect sink. ................................ ................................ ............ 162 6 3 NaCl release profile of Gel A3 with different thickness in perfect sink. ............. 163 6 4 Ion permeability test for silicone hydrogels by diffusion cell.. ........................... 164 6 5 Ion permeability test of Gel A3 by diffusion cell with various NaCl concentration in the donor compartment.. ................................ ........................ 165 6 6 Ion permeability test by diffusion cell of silicone hydrogel that were pre soaked in sodium chloride solution with different concentration. ................... 166 6 7 Ion permeability test of Gel A3 with different thickness by diffusion cell. .......... 168 6 8 NaCl partition coefficient (K), diffusivity(D), KD and Water content (Q) for silicone hydrogel with different T RIS/Macromer compositions. ......................... 170 6 9 NaCl par tition coefficient (K), diffusivity(D), KD and Water content (Q) for silicone hydrogel with different EGDMA compositions. ................................ ..... 172 6 10 NaCl partition coefficient (K), diffusivity(D), KD and Water content (Q) for silicone hydrogel with different DMA compositions. ................................ .......... 174 6 11 The relat ionship between salt partition coefficient (K) and water fraction (Q) of silicone hydrogel. ................................ ................................ .......................... 176 6 12 The relationship between salt diffusivity (D) and the reciprocal of water fraction (1/Q) of silicone hydrogels. ................................ ................................ .. 177 6 13 The relationship between salt permeability (KD) and reciprocal water fraction (1/Q) of silicone hydrogels. ................................ ................................ ............... 178 7 1 Cumulative drug uptake by ACUVUE OASYS TM lens ................................ ..... 201 7 2 Cumulative CyA release by 1 DAY ACUVUE ................................ ................. 202 7 3 Cumulati ve CyA release from silicone contact lens. ................................ ......... 203
13 7 4 Cumulative drug release from Vitamin E loaded ACUVUE OASYS TM lenses. ................................ ................................ ................................ ......................... 204 7 5 Daily average CyA release rate from Vitamin E loaded ACUVUE OASYS TM lenses. ................................ ................................ ................................ .............. 205 7 6 Plot of % CyA release by silicone contact lenses versus square root of time. .. 206 7 7 Plot of % CyA release by Vitamin E loaded ACUVUE OASYS TM versus square root of time. ................................ ................................ ........................... 207 7 8 Compariso n of CyA and dexamethasone delivery by Vitam in E loaded ACUVUE OASYS TM ................................ ................................ ....................... 208 8 1 Effect of insertion of drug loaded contact lenses on the intrao cular pressure .. 220 8 2 Effe ct of drug administration through eye drops on the intraocular pressure.. .. 222 8 3 Comparison of the effect of various drug delivery methods on the differences in the IOP between the treated an d the untreated eyes. ................................ .. 223
14 LIST OF ABBREVIATION S BCL Bandage contact lens CMC Critical micelle concentration CyA Cyclosporine A DI Deionized DMA N, N Dimet hylac rylamide DX Dexamethasone DXP Dexamethasone 21 disodium phosp hate EGDMA Ethylene glycol dimethacrylate GVHD Graft versus host disease HEMA 2 hydroxyethyl methacrylate IOP Intraocular pressure LASI K Laser in situ keratomileusis MAA Methacrylic acid NVP 1 vinyl 2 pyrrolidone OU Both eyes PBS Phosphate buffered saline PLTF Pre lens tear film POLTF Post lens tear film PRK P hotorefractive keratectomy TRIS 3 Methacryloxypropyl tris(trimethylsiloxy)silane UV Ultraviolet
15 Abstract of Dissertation Presented to the Graduate School of the University of Florida in Partial Fu lfillment of the Requirements for the Degree of Doctor of Philosophy VITAMIN E LOADED SILICONE HYDROGEL CONTACT LENSES FOR EXTENDED OPHTHALMIC DRUG DELIVERY By Cheng Chun Peng August 2011 Chair: Anuj Chauhan Major: Chemical Engineering Ophthalmic drug delivery via eye drops is inefficient as only 1 5% of the applied drug enters the cornea and the rest is absorbed into the bloodstream. This absorbed drug then enters other organs where it can cause side effects. Furthermore, drug administration through ey e drops results in a rapid variation in drug delivery rates to the cornea that limits the efficacy of therapeutic systems and limit compliance The purpose of our study is to develop a novel soft contact lens system for long term and controlled drug delive ry to eliminate these problems. The aims are to characterize the lens system, to establish the drug transport mechanism models, and to evaluate the potential of practical applications for ocular disease treatments. Our approach focuses on creating transp ort barriers to increase release duration from commercial contact lenses. In the absence of diffusion barriers, drug molecules diffuse out of the lens in ab out a few hours In contrast, if diffusion barriers are present, molecules have to diffuse around th ese, resulting in an increase in the path length and the release duration. In this study the hydrophobic Vitamin E is loa ded into silicone hydrogel contact lens by being dissolved in ethanol that swells the lens and subsequently forming aggregates inside t he lens after solvent evaporation.
16 Several properties including geometry, ion permeability, oxygen permeability a nd UV transmittance are characterized to determine the pros and cons of loading Vitamin E into the lenses. The results indicate the property c hange s caused by Vitamin E loading do not disqualify these silicone hydrogels as extended wear contact lens. In addition, Vitamin E loading has a beneficial effect of blocking UV radiation which will reduce the corneal damage due to UV light. Among all the lens property changes, ion permeability demonstrated the strongest dependency on the Vitamin E loading in the lens, and thus a further securitization of the ion transport of silicone hydrogels with various compositions is discussed. In vitro drug s rele ase r esul ts show that the increase in release duration is significantly dependent on the interaction between Vitamin E and the drug of interest. F or hydrophilic drugs (timolol, fluconazole, dexamethasone phosphate) the drug release duration increases quad ratically in Vitamin E loading. For example, for NIGHT&DAY TM lens loadings of 10 and 40% Vitamin E increase release durations to 0.5 an d 16 days, respectively. For hydrophobic drugs dexame thasone and cyclosporine A the effect of the Vitamin E inclusion is smaller but still signifi cant for release. On the other hands, for some amphiphilic anesthetic drugs, including lidocaine, bupivacaine and tetracaine, the interfacial interaction between drug and Vitamin E aggregation plays the determinative role in drug transport through the Vitamin E/silicone hydrogel matrix. Ocular drugs delivery by contact lens can be viewed as a one dimensional transport by a flat thin film, and subsequent mathematical models based on the proposed mechanisms are established.
17 In addit ion, a case study was conduct ed to evaluate the use of Vitamin E loaded silicone hydrogel contact lenses for c hronic dry eye treatment I n vitro release experiments in perfect sink condition demonstrate that through incorporation of Vitamin E, the loaded cyclosporine A in the lens can be release d from 2 weeks to the total wear time of the lens which is about a month Vitamin E incorporation also renders the ther ap eutic window The long release duration along with the higher bioavailability compared to commercial eye drops (RESTASIS ) suggested that these lenses could potentially be useful for treatment for chronic dry eye and also for reducing the symptoms of cont act lens mediated dry eyes. Finally, a n in vivo study was conducted to investigate the feasi bility and effectiveness of timolol delivery to glaucomatous dog s via drug impregnated contact lenses. By utilizing contact lens to del iver timolol to the eye, th e intraocular pressure in the treated eye decreased effectively to similar degree compared to that by eye drop treatment while it significantly reduced the risk of systemic drug exposure. In conclusion, s ilicone hydrogel contact lenses with Vitamin E ar e promising candidates for extended ophthalmic drug delivery. These Vitamin loadings can significantly attenuate the drug delivery rate, reduce wastage and provide safer treatment route.
18 CHAPTER 1 INTRODUCTION While easy accessibility is clearly an advantage for ophthalmic drug delivery, several other features of the ocular physiology, anatomy and biochemistry render the ocular environment impervious to foreign substances, thus posing unique challenges for delivery drugs to eyes [1 4] The most common ophthalmic dosage forms are solutions, ointments and suspensions, which together account for nearly 90% of currently available formul ations in the United States  Amongst all the dosage forms for ophthalmic drug delivery, eye drop solutions are the preferred choic e since they are relatively simple to prepare, filter, and sterilized. Ophthalmic drugs are commonly delivered to the front of the eye through instillation of drug laden eye drops into the tear film, as shown in Figure 1 1 The human tear film is about 10 L in volume and 3 10 m in thickness [6, 7] The tear film is typically considered to comprise of three layers: a thin mucin layer lying on the corneal surface, which renders the corneal surface hydrophilic; the aqueous layer which comprises the bulk of t he tear film; and a very thin oily layer that minimizes evaporation of tears into air  The mucin layer is produced mainly by secretions from conjunctival goblet cells  ; the aqueous layer is secreted by the lacrimal glands and supplemented by conjunc tival secretion, and the lipid layer is produced by the tarsal meibomian glands [7, 8] The tear film is lined by the cornea and the conjunctiva, which is further classified as the bulbar or the palpebral conjunctiva. The volume and composition of the tea rs are maintained through a dynamic balance between the various secretion and elimination pathways for the fluid and the ions. The tear fluid is produced by the lacrimal glands, and from secretion from the conjunctiva, and eliminated through evaporation an d
19 drainage of tears through the canal iculi into the lacrimal sac  The transport of ions in the tear film includes inflow through the lacrimal secretions and outflow through the tear drainage. Additionally, active ion transporters, pumps, and osmotic an d electro osmotic transport across the cornea and the conjunctiva contribute to the ion transport in the tear film  The transport pathways for ions and tear fluid also contribute to the transport of drug delivered to the tear film through instillatio n of an eye drop. A typical eye drop is about 30 L in volume, which is too large to be accommodated in the tear film, and thus a fraction of the instilled drop typically rolls off after instillation  The remaining amount rapidly mixes with the tear volume due to blinking. Subsequently, a fraction of the drug penetrates the corneal and the conjunctival epithelium, and the remaining amount drains into the lacrimal sac and the nasal cavity, where it can get absorbed into the blood stream through the mucus membrane. A large fraction of the drug that is a bsorbed in the conjunctiva also enters systemic circulation because of the high blood perfusion in this tissue. The amount of drug absorbed by the conjunctiva is significantly more than that absorbed by the cornea because of the larger area and permeabilit y of the conjunctiva in compari son to those for the cornea  The drug absorbed by the cornea can diffuse across the three layers of the cornea, i.e., the epithelium, stroma, and the endothelium to reach the anterior chamber. A fraction of the drug is t hen cleared through the drainage of the aqueous humor, and a fraction can bind to ocular tissues such as the lens, ciliary bodies, etc. Due to the limited residence time of the drug in the tear film, the drug concentration at the target tissue could exhibi t a short duration burst above the toxicity threshold, followed by a long duration below the
20 therapeutic threshold. A few factors that aid in increasing the duration of action of the drug at the target tissues include a strong drug binding to the target ti ssue and/or adsorption desorption to other ocular tissue. The clearance mechanisms in eyes render ophthalmic drug delivery via eye drops relatively inefficient. Most ophthalmic drugs have a short residence time of approximately 2 minutes  due to the ra pid tear turnover. With such short residence time, only about 2 5 % of the applied drug penetrates the cornea to reac h the intraocular tissue [12, 13] The remaining drug is either absorbed in the conjunctiva or it drains with tears into the lacrimal sac, leading to drug wastage and som etimes adverse side effects  Additionally, the poor drug bioavailability and short residence time in tears results in the need for several daily administrations, reducing the patient compliance. To overcome the drawbacks of eye drops, several ophthalmic drug delivery systems aim for sustained drug release have been proposed such as suspension of nanoparticles, na nocapsules, liposomes and noisome ocular inserts like collagen shields and Ocusert, and therapeutic contact l enses Among these, contact lenses have been widely studied due to the high degree of comfort and biocompatibility. As illustr ated in Figure 1 2 i f drug loaded contact lenses are placed on the eye, the d rug is expected to diffuse through the lens matrix, and enter the po st lens tear film (POLTF), which is the thin tear film trapped between the cornea and the lens. In the presence of a lens, drug molecules will have a residence time of about 30 min utes in the post lens tear film, compared to about 2 minute s in the case of topical application as drops [12, 14, 15] The longer residence time will result in a higher drug flux thr ough the cornea and
21 reduce the drug inflow into the nasolacrimal sac, thus reducing the drug absorption into the blood stream Drug loaded contact lenses also have the potential to provide continuous drug release, because of the slow diffusion of the drug molecul es through the lens matrix compared to that in aqueous solution However, the maximum drug loading is limited by the solubili ty of the drugs in the gel matrix. Also the only additional resistance to drug transport is diffusion t hrough the gel matrix, which is significantly determined by the materials and microstructures of the contact lens and for most ocular drug s the release time is less than a few hours  The development of hydrogel based contact lens started in 1936 when the first hydrogel lens was int roduced by Wichterle by using po lymethylmethac rylate (PMMA), a resin that has greater clarity than glass [17, 18]. Sinc e then, researches of new hydrogel polymers were continuously motivated by the obvious pursuit for a competitive position in the new growing vision correction soft contact lens market  The evolution of new soft contact lens materials have also driven by an increased understanding of the physiological needs of the cornea, since the ocular environment places high demands on the performance of contact lenses as biomaterials. The contact lens on the eye must maintain a stable, continuous tear film for clea r vision; it needs to be resistant to deposition of tear film components and has proper mechanic strength to be nonirritating and comfortable for the wearer during blinking In addition, permeability to oxygen and ions are two key performance characteristi cs for contact lenses. Not like other tissues in human cornea obtains its oxygen directly from the air to maintain its clarity, structure and function, and hence a contact lens must be sufficiently permeable to oxygen to maintain normal corneal metabolism Effective ion permeation in the contact lens allows
22 the movement of the lens on the eye. Therefore, to satisfy all the criteria, the lens must have excellent surface characteristics being neither hydrophobic nor lipophilic, and have the appropriate bulk polymer composition and morphology at the same time. Design for new extended wearable contact lens has been widely studied since the negative physiological effect of cornea oxygen deprivation began to be realized [20, 21] Traditional hydrophilic hydrogel contact lens has been proved fails to provide the minimum oxygen transmissibility of 125 barrer/mm to completely avoid low oxygen related effects and to be safely worn overnight [22, 23] To overcome the performance hurdle, silicone hydrogel cont act lens was developed due to its high oxygen permeability since the first commercial one was produced in 1980 by Dow Corning. However, polysiloxane has negligible water content and manifest typically rubbery behavior. To offer the softness, wettability an d on eye comfort of a conventional hydrogel, while at the same time provide the higher oxygen transmission required by the cornea, new silicone hydrogels contact lens for extended wear were developed based on the combination of silioxane, hydrophilic co mo nomers, and bi functional macromer to find the optimized performance [24, 25] wear extended hydrogel contact lens which can safely and continuously provide ophthalmic drug delivery. In previou s studies on ocular drug delivery by contact lenses, d rug can be loaded in the contact lenses by either soaking the gels in drug solution [26 31] or by dissolving drug in the monomer solution before polymerization [32 34] The major problem of loading drug s by uptake from drug solutions is that in most cases the loading capacity of the soaked contact lenses is inadequate. An additional problem that can
23 occur when absorbing drugs in hydrophilic contact lens is that the preservatives included in the drug are often preferentially absorbed in the lens to a level that is toxic  For directly loading drug in the monomer solution, although it can allow higher loadings of the drugs in the lens, it can result in an activity loss during polymerization, and a majo rity of the drug diffuses from the lenses into the packaging medium. Moreover, even though providing higher bioavailability, commercial lenses release ophthalmic drugs rapidly in a time of only a few hour s and therefore limited their application for extend ed drug delivery. Recently Karlgard et al. measured the in vitro uptake and release of a number of ophthalmic drugs by commercially available p HEMA ba sed and silicone contact lenses . Drugs including cromolyn sodium, ketotifen fumarate, ketorolac trom ethamine and dexamethasone sodium phosphate, were loaded into these lenses by soaking these in drug solutions for a limited period of time. The release studies showed that a majority of the drug taken up by the gels was released in a short period of time f or less than a few hours, which is not suitable for extended drug delivery. To increase drug release durations, several different drug impregnated contact lens systems have been developed, including nanoparticle laden lenses [36 40] [41 48] and contact lens with layer structures  While these approaches are effective at extending the drug release duration from contact lenses, all studies cited above focused on lenses based on hydrophilic hydrogel materials, which are not suitable for exten ded wear due to insufficient oxygen permeability In addition, w hile the approaches listed above can increase the loading
24 capacity and the release duration from contact lenses, the lenses are still not suitable for extended release lasting a week or longer. The aim of my study is to develop a contact lens system containing diffusion barriers in the hydrogel matrix that can provide sufficient drug transport resistance, and this system should be still satisfy all the criteria as extended wear contact lens material. This extended wear extended release contact lens could serve as the drug delivery vehicle for long term ocular disease treatment, such as glaucoma and chronic dry eye, and/or alleviate common untoward syndromes for extended contact lens wearers. G laucoma, which is the second largest cause of blindness in the world after cataract affects about 60.5 million people, leaving 8.2 million with bilateral blindness . Most current glaucoma therapies are based on drug del ivery via eye drops that are inefficient and have lower than 50% compliance , with further lower compliance for multiple drops and/or multiple drugs therapies . For chronic dry eye, In 2007, there were about 35 million contact lens wearers in North America and about half of those reported some symptoms of dryness and discomfort, more commonly experienced at the end of the day  Our designed contact lenses can also be used to release topical a nesthetics for pain control after photorefractive kera tectomy (PRK) surgery The case studies of these potential applications are further discussed in Chapter s 5, 7 and 8. The release of a molecule from a contact lens can be considered as controlled by diffusion within the lens material. For a one dime n sional diffusion controlled process, the duration of release can be approximately calculated by l 2 / D where l is the path length that a compound needs to traverse and D is the molecular diffusivity. For a typical contact lens, l is the thickness of the lens, whi ch varies in the radial direction but is on
25 average approximately 80 to 100 m for a typical lens. The period of time over which a drug is released from a contact lens can be increased by either increasing l or by decreasing D In most diffusion controlled systems, augmentation of diffusivity has been performed by changing the bul k material to one of a different diffusivity. However, because of the strict requirements of a contact lens where many material properties cannot be compromised, there are practical limits to the selection of the bulk material. Furthermore, an effective s trategy to modifying the diffusion process must be applicable to a wide range of bioactive agents with a similar bulk material. Our proposed concept is directed towards controlling the diffusion of a bioactive agent in a contact lens matrix by the creation of diffusion barriers within the lens, such that an included bioactive agent is forced to take a long tortuous path to diffuse from the lens, resulting in extended release. The concept of using transport barriers has been explored extensively for designi ng membranes that retard gas transport [54 57] To our knowledge, this concept has not been applied to retard drug transport from a biomedical device. The diffusion barrier can be any solid or liquid material that is able to be dispersed within the lens ma terial in a manner that keeps the lens transparent. The most important requirement for an effective barrier material is that the barrier should be relatively impermeable to the molecules whose diffusion needs to be attenuated. A number of ophthalmic drugs are charged at physiological pH and so hydrophobic molecules will likely form effective barriers. It is also important to ensure that the barrier material is biocompatible so that diffusion of the compound forming the barrier into the tear film does not ca use toxicity.
26 The diffusion ba rrier material of interest in our study is Vitamin E which is a highly hydrophobic liquid It is a powerful antioxidant and has been shown in some animal studies that the topical application of Vitamin E inhibits a number of eye diseases including keratocyte apoptosis after surgery, ethanol induced apoptosis in the corneal epithelium, etc. [58, 59]. Also there have been a number of in vivo studies suggesting Vitamin E retard s cataract development [60 64] Because of the pot ential benefits of delivering Vitamin E to the eye, there have been several attempts to develop ophthalmic solutions containing Vitamin E [65, 66] O nce Vitamin E is trapped inside the gel matrix, it should be stable in the hydrated contact lens due to it s high hydrophobicity In Chapter 2 we explored the approaches to introducing Vitamin E into the silicone hydrogels, and characterized the properties of these Vitamin E loaded silicone hydrogels as extended contact lens, including ion permeability, oxygen permeability, geometry changes, elastic modules, water transport and light transmittance, etc. Among these properties, we especia lly focused the ion transport through the silicone hydrogels since first it can be correlated to some charged drugs transport in the gel matrix, which is very common at physiological pH, and the results are discussed in Chapter 6. The transport mechanisms of extended ocular drugs by Vitamin E loaded silicone hydrogel are further discussed throughout Chapter s 3 to 5. Chapter 3 discussed the release of three different hydrophilic ophthalmic d rugs by Vitamin E loaded silicone hydrogel contact lens, including timolol (beta blocker used for treating glaucoma), dexamethasone 21 disodium phospha te (anti inflammatory cortico steroid), and fluconazole (anti fungal). These drugs were chosen because they are hydrophilic at the
27 physiological p H, which should have negligible affinity to the desired Vitamin E barriers. In addition, we also explored the drug delivery these Vitamin E loaded si licone hydrogels for hydrophobic drug such as dexamethasone (anti inflammatory) and cyclosporine A (chronic dry eye treatment), and amphiphilic anesthetics drugs, including lidocaine, bupivacaine and tetracaine. The drug transport results and mechanisms by silicone hydrogel for these hydrophobic and amphiphilic drugs are discussed in Chapter s 4, 5 and 7. Finally, in Chapter 8 we demonstrate the in vivo evaluation of the safety and efficacy of glaucoma therapy through release of timolol from silicone hydrog el contact lenses. Timolol is a beta blocker that is widely used to treat glaucoma by reducin g IOP through decreasing the production of aqueous hu mor  We focus on this drug because of the large number of glaucoma patients in the world  and also because of the potential of serious side effects from systemic exposure to timolol  W e choose the colony of beagle dogs who are affected by or carriers of a hereditary form of primary open angle glaucoma, the most common form of glaucoma in human bei ngs . Beagle dogs have been used in several prior studies on glaucoma therapy [70 76]. Another benefit of using the Beagle dogs is that t he cornea shape and size of these dogs are similar to that of human beings, and therefore the commercially availab le contact lens for human can be used in this study without further modification.
28 Figure 1 1 Schematic illustration of ophthalmic drug delivery through eye drops. Te ar film Cornea Conjunctiva
29 Figure 1 2 Schematic illustration of the microemulsion laden contact lens inserted in the eye. Post lens tear film Cornea Drug loaded contact lens Pre lens tear film
30 CHAPTER 2 CHARACTERI ZATION OF VITAMIN E LOADED SILICONE HYDR OGEL In this c hapter we explored the possibility to introducing Vitamin E into the silicone hydrogel polymer matrix as contact lenses materials. H ydrophobic Vitamin E is loa ded into silicone hydrogel contact lens by being dissolved in ethanol that swells the lens and subsequentl y forming aggregates inside the lens after solvent evaporation. T his in situ approach to create diffusion barriers is particularly suited for biomedical applications in which the polymer processing steps could potentially damage some materials that are u sed as transport barriers. Most experiments i n this c hapter were done by loading Vitamin E into commercial silicone hydrogel contact lens to evaluate the properties chance resulted from the additional Vitamin E inside the gel matrix system. Important prope rties for the bulk material, including geometry, equilibrium water content, ion permeability, oxygen permeability, and light tran smittance. In addition, to understand the effect of Vitamin E on the elastic modulus of silicone hydrogel, a lab synthesized si licone hydrogel was prepared as a substation of commercial silicone hydrogel contact lenses. 2 .1 Materials and Methods 2. 1 .1 Materials Five commercial silicone contact lenses (diopter 6.50) that are used in this study are described in Table 2 1 2 hy droxyethyl methacrylate (97%), sodium hydroxide pellets (97+%), ethanol ( 99.5%), and p hosphate b uffered s aline (PBS) were purchased from Sigma Aldrich Chemicals (St. Louis, MO) and ethylene glycol dimethacrylate (EGDMA) from Sigma Aldrich Chem icals (Milwaukee, WI).
31 Sodium chloride (99.9+ %) were purchased from Fisher Chemical (Fairlawn, NJ). Vitamin E (D alpha tocopherol, Covitol F1370) was gifted by Cognis Corporation For preparation of silicone hydrogel, ethylene glycol dimethacrylate (EGD MA, 98%), N, N Dimet hylac rylamide (DMA, 99%) and 1 vinyl 2 pyrrolidone (NVP, 99+ %) were purcha sed from Sigma Aldrich Chemicals (Milwaukee, WI ). The macromer acryloxy(polyethyleneoxy) propylether terminated poly(dimethylsiloxane) (DBE U12, 95+ %) were purch ased from Gelest Inc. (Morrisville, PA). 3 Methacryloxypropyl tris(trimethylsiloxy)silane (TRIS) was supplied by Silar laboratories (Scotia, NY), and 2, 4, 6 trimethylbenzoyl diphenyl phosphineoxide (Darocur TPO) were kindly provided by Ciba Specialty Che micals (Tarrytown, NY). All chemicals were used as received without further purification if not specifically mentioned. 2. 1 2 Vitamin E L oading into P ure L enses Vitamin E was loaded into le nses by soaking the lens in 3 m L of a Vitamin E ethanol solution fo r 24 hours. Vitamin E ethanol solutions of various concentration s were prepared by simply mixing Vitamin E and ethanol with vortexing for a few seconds followed by moderate magnetic stirring for several minutes. After the loading step, the lens was taken o ut and excess Vitamin E ethanol solution on the lens surface was blotted and the lens was then dried in air overnight. 2. 1 3 Ion P ermeability M easurements Ion permeability of lenses was measured by using a homemade horizontal diffusion cell that consist s of a donor and a receiving compartment, which were both fabricated from Plexiglas. The ion permeability of the lens was determined by measuring the rate of transport of ions across the lens. To mount the lens in the diffusion cell, the circular edge of t he dried lens was glued to the outer edge of a 1 cm
32 hole cut into a plastic spacer. The spacer along with the lens was then soaked in deionized ( DI ) water for longer than three hours to fully hydrate the lens. The excess water on the spacer was wiped off a nd the spacer was subsequently placed in between the two compartments of the diffusion cell and clamped. Latex O rings were also inserted in between the spacer and each of the compartments to ensure sealing. The latex O rings were boiled in DI water for 4 0 min utes for three times before placing in the diffusion cell to leach out impurities. After assembling the diffusion chamber, t he receiving chamber was filled with 30 m L of DI water and the donor chamber was filled with 18 m L of 0.1 M NaCl solution The ion conductivity of the fluid in the receiving chamber was measured as a function of time with a conductivity meter with temperature sensor (Con 110 series, OAKTON ) and linear regression was applied to the data after reaching pseudo steady state (after 7 0 minutes) to obtain the best fit slopes The rate of conductivity change can be converted to the rate of ion transport, which can then be 2. 1 4 Oxygen P ermeability M easurements To measure the oxygen permeability, lenses were mounted in a horizontal diffusion cell by following the same procedure as d escribed in Section 2.2.3 To create oxygen gradients in the cell, the donor c ompartment was filled with 18 mL of DI water that was equilibrated with air, and the receiving chamber was f illed with 32 mL of DI water that was degassed by bubbling nitrogen for 10 minutes. Both compartments were kept well stirred with minimal boundary layer thicknesses adjacent to the lens by stirring at 900 rpm. The dissolved oxygen concentration in the receiving reservoir was measured every 12 seconds by an oxygen sensor (DO BTA, Vernier ) for a total
33 duration of 2 hours. The measured data was fitted to a mathematical model described later to determine the oxygen per meability of the lens. 2. 1 5 Transmittance Measurement of Vitamin E Loaded Contact L ens The transmittance of Vitamin E laden lenses was measured using UV VIS spectrophotometer (Thermospectronic Genesys 10 UV). The lenses were hydrated by soaking in PBS ove rnight, then cut into stripes and mounted on the outer surface of a quartz cuvette. The cuvette was placed in the spectrophotometer and the transmittance values were measured at wavelengths ranging from 200 nm to 500 nm. 2.1 .6 Preparation of Silicone Hydro gel To prepare the silicone hydrogel, hydrophilic monomers with high ion permeability are copolymerized along with the hydrophobic silicone monomer with high oxygen permeability, and a macromer is needed in the monomer mixture to ensure solubilization of a ll monomers. In this study, TRIS was used as the hydrophobic monomer, DMA was the hydrophilic monomers, and DBE U12 was selected as the macromer. Highly hydrophilic NVP monomer was also added to increase water content of the hydrogel and EGDMA was introduc ed in the monomer mixture for controlled crosslinking. To prepare the polymerizing mixture, 2.4 mL of a mixture that comprises 0.8 mL TRIS and 0.8 mL macromer and 0.8 mL of the hydrophilic DMA/MAA mixture was combined with 0.12 mL of NVP and 0.1 mL of EGDM A. After well mixed with vortexing for few second, the mixture was purged with bubbling nitrogen for 15 minutes to reduce the dissolved oxygen. To each monomer mixture, 12 mg of photoinitiator Darocur TPO was added with stirring for 5 minutes and the fina l mixture was immediately injected into a mold which is composed of two 5 mm thick glass plates. The plates were separated by a plastic spacer with various thicknesses. The mold was then
34 placed on ultraviolet transilluminator UVB 10 (UltraLum Inc.) and the gel mixture was cured by irradiating with UVB light (305 nm) for 50 minutes. The synthesized hydrogel was either cut into circular pieces (about 1.65 cm diameter) with a cork borer for subsequent experiments. Prior to conducting further tests, the prepare d hydrogel was soaked in ethanol for 3 hours then dried at ambient temperature overnight to remove the unreacted monomer within. 2.1 .7 Mechanical P roperties M easurements The mechanical properties of gels are analyzed in tensile mode by using a dynamic m echanical analyzer (DMA Q800, TA instruments). A 0.4 mm thick rectangular hydrated gel was mounted on the tension clamp while submerged in water at room temperature. A periodic tensile force was applied in the longitudinal direction with varied frequency a nd the response (storage modulus and loss modulus) of gel was determined. A static preload force of 0.01 N was applied and a 115% of force track was used to keep the sample taut on the tension clamps. Strain sweep tests were conducted at room temperature a t 1 Hz to determine linear viscoelastic range, and proper strain which was therefore confirmed within linear range by the strain sweep test will be chosen for subsequent strain controlled frequency sweep experiments. 2. 2 Results and Discussion 2. 2 1 Vitami n Loadings in the L enses Vitamin E loadings into each lens for different initial concentration of Vitamin E loading solutions are shown in Figure 2 1 Vitamin E loading has a linear dependency on the concentration of Vitamin E loading solutions. In additio n, ACUVUE OASYS and NIGHT&DAY have the highest and the lowest affinity for Vitamin E, respectively.
35 2 2 .2 Transparency of the Vitamin E laden Contact Lenses An image of a Vitamin E loaded contact lens is show n in Figure 2 2 As evident from the image, the Vitamin E loaded lenses are transparent irrespective of the Vitamin E loading, but attain a slightly yellowish color at high Vitamin E loadings. 2. 2 3 Water C ontent of P ure and Vitamin E L oaded L enses Water contents ( Q ) of lenses are listed on each l ens package and were also measured as ( 2 1 ) where W eq W l and W ve are mass of hydrated lens at equilibrium, mass of dry pure lens, and mass of Vitamin E loaded in the lens, r espectively. Both the listed and measured Q s are shown in Table 2 1 Additionally, the values of the equilibrium water content ( EW ) which is defined as mass of water absorbed by unit mass of pure lens i.e., ( 2 2 ) are also listed in Table 2 1 Results show that ACUVUE ADVAN C E has the high est EW (86.0 2.3) and NIGHT&DAY has a relatively low EW (31.1 5.5). The effect of Vitamin E loading on Q and EW are clearly seen in Fig ure 2 3 In Figure 2 3 A, water content of Vitamin E loaded lenses tends to decrease relatively linearly as Vitamin E loading increases. However, W eq of Vitamin E loaded lenses increases as Vitamin E loading which may be causing the decrease in the Q value s. To observe the effect of Vitamin E loading on water amount absorbed in lens polymers, EW was plotted verses Vitamin E loading in Figure 2 3 B. The EW for Vitamin E loaded lenses is also less than that for the pure lenses for each type of lens but the tre nds are different The EW s of
36 ACUVUE OASYS and PureVision lenses linearly decrease and the values of EW are 46% and 44% respectively for about 20% Vitamin E loading. The EW s of NIGHT&DAY and O 2 OPTIX lenses decrease by about 10% for Vitamin E loading s of about 10% but there is negligible decrease in EW s with further increase in Vitamin E loadings. The latter behavior for the NIGHT&DAY and O 2 OPTIX lenses suggests that at low loadings, the Vitamin E is solubilized in the lens and so it reduces the wa ter content of the gel because of its hydrophobicity but beyond a critical weight fraction the extra Vitamin E simply phase separates, and thus it has no further effect on the EW The critical Vitamin E loading which can be solubilized by the NIGHT&DAY an d O 2 OPTIX in the later sections based on drug transport data (6.2% for NIGHT&DAY and 9.7% for O 2 OPTIX EW for ACUVUE OASYS and PureVision len ses suggests that these lenses can either solubilize large amounts of Vitamin E or the Vitamin E that phase separates coats the polymer and thus continues to reduce the EW 2.2 4 Size Change D ue to Vitamin E L oad ing The sizes of the contact lenses are ex pected to increase due to Vitamin E uptake. The diameters of the lenses both with and without Vitamin E were measured both in dry and hydrated states, and the s ize changes of lenses after loading the Vitamin E are shown in Figure 2 4 The % dry and hydrate d diameter increase are the increase in the dry and hydrated diameter divided by the dry and hydrated diameter of the lens without Vitamin E, respectively. The solid lines in the figure are the best fit straight line s Figure 2 4 A shows that the dry diamet er change of lenses is about 30 % of the Vitamin E loading For example, about 30 % Vitamin E loaded lens shows increase of about 10%
37 in diameter in dry state, which suggests that the expansion of lens by Vitamin E loading is isotropic. In Figure 2 4 B, wet diameter change is much less than dry diameter change, which is expected because Vitamin E does not absorb water F or example, lenses with about 30% Vitamin E loaded lens expand about only 6.5 % in diameter. F rom application perspective, changes in wet di ameter should be small to preserve the power of the contact lens, and all the lenses show less than 8 % increase in wet diameter for about 40% of Vitamin E, which can likely be tolerated by eyes. There may be further changes to the corrective power due to refractive index changes in the lens. In any case, if there is a significant change in the power of the lens, the listed power for a lens can be modified from the original value. 2. 2 5 Ion P ermeability of Vitamin E L oaded L enses Ion permeability of contact lenses is a critical variable for lens motion on the eye according to Domscheke et al  The thickness of the lens varies in the radial direction and the exact profiles are not available in literature. To obtain the permeability, each lens was treated as a section of a sphere with radius equal to the known base curve of the lens and 80 m in thickness. The calculated values of ion permeability are plotted in Figure 2 5 A as a function of the Vitamin E loading for O 2 OPTIX NIGHT&DAY and ACUVUE OASYS l enses. The results show that the ion permeability of pure O 2 OPTIX is highest among three lenses and is about 3.4 fold and 2.5 fold that of the pure NIGHT&DAY and ACUVUE OASYS respectively. Also it is clearly seen that the ion permeability decreases as Vitamin E loading increases for all the lenses.
38 The decrease ion permeability for Vitamin E loaded lens can be seen more clearly in Figure 2 5 B in which the ratio of ion permeability of lens with and without Vitamin E is plotted as a function of the Vita min E loading Interestingly, the graphs are almost the same for O 2 OPTIX and NIGHT&DAY and the decrease in ion permeability by Vitamin E is much larger for ACUVUE OASYS for the same Vitamin E loadings compared to the other two lenses. D ion should be la rger than 6.0 10 6 mm 2 /min for sufficient on eye movement of lens according to Nicolson et al  Figure 2 5 indicate s that all Vitamin E loaded lenses in our study have adequate ion permeability to maintain on eye motion. 2. 2 6 Oxygen P ermeability of Vitamin E L oaded L enses The oxygen permeability of extended wear contact lenses must be sufficiently high to avoid deprivation of oxygen to cornea, which could cause adverse responses [20, 21] The lens permeability ( Dk ) is the product of the diffusivity D and the oxygen partition coefficient k and it is typically expressed in units of 10 11 (cm 2 /sec)( mL O 2 /( mL mmHg)) or 10 11 mL O 2 cm/(seccm 2 mmHg), which is also referred as a barrer or a Fatt. The oxygen permeability is an intrinsic property of a materi al to transport oxygen through its bulk and is independent of thickness. The oxygen transmissibility, Dk/t refers to the oxygen transport capacity of a specific contact lens with thickness t and it generally expressed in units of 10 9 cm mL O 2 /(sec mL mmH g) or 10 9 mL O 2 /(seccm 2 mmHg). To avoid hypoxia, an extended wearable contact lens must provide at least a minimum oxygen transmissibility ( Dk/t ) of 87, which cannot be achieved by traditional hydrophilic contact lens  Recently, the suggested minimum value of Dk/t to avoid hypoxia has been proposed to increase to 125  The reported values of Dk values of various
39 commercial contact lenses are 140 for NIGHT&DAY 110 for O 2 OPTIX 103 for ACUVUE OASYS for PureVision erage thickness of 80 m, these commercial silicone hydrogel contact lenses can provide sufficient oxygen transmissibility to be used for extend wear. The influence of Vitamin E loading on oxygen transport through the contact lenses was determined by mounting the lenses is a di ffusion cell with gradient in the dissolved oxygen concentration across the lens, and then measuring the oxygen concentration in the receiver chamber. Below a model is presented to fit the measured oxygen concentration data to determine the oxygen diffusiv ity through the lens. Overall mass balance of dissolved oxygen in the closed diffusion cell is given by ( 2 3 ) where V r and V d are the DI water volumes of the receiving and don or compartments, respectively, and C r and C d indicate the dissolved oxygen concentrations with initial concentration of C r0 and C d0 in the receiving and donor chambers, respectively. Since the lens volume is substantially less than the fluid volume, the sy stem reaches a pseudo steady state very rapidly and thus the oxygen flux through the lens can be expressed as ( 2 4 ) where A and h are the surface area and the average thi ckness of hydrated lens respectively; D is the oxygen diffusion coefficient of the lens material and k is the oxygen partition coefficient between lens and DI water. The above equation implicitly assumes negligible mass transfer resistance in the boundary layers in the receiver and donor compartments. This assumption was verified by showing that the measured
40 oxygen concentration profiles were not sensitive to stirring at stirring speeds of 900 rpm. Equation s 2 3 and 2 4 can be combined to give: ( 2 5 ) The solution to the above equation with the initial condition Cr ( t = 0) = Cr 0 is ( 2 6 ) The parameter DkA/h can be obtained by fitting the experiment data to the above equation using the fun ction fminsearch in MATLAB The exact value of D through various lenses could not be directly obtained because the detailed shapes of the lenses were not available in literature, but could be calculated by using the approximate surface area of these len ses descr ibed in Chapter 2.3.5. The validity of this approach was established by measuring oxygen diffusivity through pHEMA hydro gels prepared of two different thicknesses fro m the procedures reported in earlier study  The measured value of 14.61.3 f or the synthesized hydrogel with water content 41.1% was in good agreement with reported value of 12.9 for conventional hydrogel materials of which oxygen permeability is primarily determined by its water content  The effect of Vitamin E loading on Dk of silicone contact lenses is shown in Figure 2 6 The calculated Dk values were 148, 118, and 111 for NIGHT&DAY TM ACUVUE OASYS TM and O 2 OPTIX TM respectively, which were also in good agreement with the reported Dk values from the manufacturers and other research groups, providing the accuracy of the measurement methods . The results show Vitamin E loading in NIGHT&DAY TM slightly reduces the oxygen permeability when the Vitamin E amount goes up to about 75%. On the other hand, no significant change wa s observed for
41 ACUVUE OASYS TM and O 2 OPTIX TM up to about 35% of Vitamin E loading in the lens. While it is not feasible to quantitatively evaluate the effect of Vitamin E on Dk values due to the relatively large standard deviations in the measured values, it is clear that the Dk value of these Vitamin E loaded lenses with average thickness 80 m are still sufficiently high to meet the minimum requirements to avoid hypoxia. The results that Vitamin E loading in the lens has much higher influence on ion transport than on oxygen transport suggests that most Vitamin E aggregates exist in the hydrop hilic polymer region in the gel matrix. This is plausible because Vitamin E likely has a much lower solubility in the hydrophilic regions than in the hydrophobic silicone rich region in the gel matrix. Since ion transport occurs primarily through the hydro philic channels, the presence of Vitamin E aggregates significant reduces the ion permeability. On the other hand, oxygen transport occurs mainly through the silicone rich channels, which may not contain Vitamin E aggregates resulting in a minimal attenuat ion in oxygen permeability. 2 2 7 Transmittance of Vitamin E L oaded L enses In addition to correcting vision, a contact lens could potentially also prevent or minimize exposure of the corneal tissue to damaging effects of UV light. Currently ACUVUE is the only brand that claims the benefit of protection against UV radiation  Figure 2 7 shows the measured transmittance spectrum for three commercial contact lenses used in our study. NIGHT&DAY TM and O 2 OPTIX TM have no significant protection against UVB (28 0 315 nm) and UVA (315 400 nm), while ACUVUE OASYS TM completely blocks UVB and UVA radiation. These results match the reported UV transmittance characteristics of silicone hydrogel contact lenses reported by the manufacturers and other independent researc h group [80, 81]
42 The effect of Vitamin E loading on the transmittance spectrum of NIGHT&DAY TM and ACUVUE OASYS TM is shown in Figure 2 8 The results clearly show that Vitamin E loaded NIGHT&DAY TM lenses completely block UVB radiation and also partially b lock UVA radiation proportionally to the Vitamin E loading. Since ACUVUE OASYS TM block UV radiation, Vitamin E loading only marginally increases the UV protection for these lenses. The UV radiation is known to induce photo oxidation of Vitamin E transform ing Vitamin E into various photoproducts [82, 83] To explore the effect of photo oxidation on protection again UV radiation, the transmittance spectra of lenses was measured as a function of time while exposing the lenses to natural light. While the UVB b locking effect of Vitamin E was retained, the ability to absorb UVA radiation decreased for a few days and then reached equilibrium, as shown in Figure 2 9 2 2 8 Mechanical Properties of Vitamin E Loaded Silicone H ydrogel Kim et al. had found that the fre have significant dependency on the compositions of silicone hydrogels  It was viscoelastic properties also dep end on the frequency. The modulus of a lens has important consequences on vision correction and safety. The modulus of silicone hydrogels is larger than that of hydrogels and the most suitable lens is the one that balances the advantages of silicone hydrog els while not significantly influencing modulus. The modulus of commercial silicone hydrogel lenses lies in the range of 1 1. 5 MPa  The dynamic mechanical analyze results for hydrogel with different Vitamin E loading are shown in Figure 2 10 which i ndicates that the storage modulus of the silicone hydrogel decreased a s the amount of Vitamin E in the hydrogel increased. In
43 addition, with similar amount of Vitamin E loading, the storage modules of the gel by directly adding Vitamin E into the monomer m ixture before photo curing is smaller than that by soaking the polymerized silicon hydrogel in Vitamin E/ethanol solution. In this c hapter we have shown the approach of in situ creation of transport barriers of Vitamin E, and the material properties as e xtended wear contact lens were characterized. The results suggested that with proper amount of Vitamin E loading the Vitamin E loaded commercial silicone hydrogel lens can still be qualified for continuous wear. For example, Vitamin E loading in the NIGHT &DAY TM lens leads to slight increase in lens sizes ( 6.5% increase for 30% loading), a slight reduction in oxygen diffusion ( about 40% reduction for 75% loading), and a more significant reduction in the ion permeability (50% reduction for 10% loading) Howe ver, these changes are not sufficient to preclude use of the Vitamin E laden commercial lenses for extended wear Additionally, Vitamin E l oading has a beneficial effect of blocking UV radiation which will reduce the corneal damage due to UV light.
44 Fi gure 2 1 Correlation of Vitamin E loading and concentration of soaking solution for different lenses. The lines are the best fit straight line to data. The slope and R 2 of the line are 5.26, 0.9692 (ACUVUE 2 respectively.
45 Figure 2 2 Images of Commercial NIGHT&DAY TM contact lens (left in panel A ) and NIGHT&DAY TM le ns with 30%Vitamin E l oading (right image in panel A and panel B ) Photos courtesy of Cheng Chun Peng B A
46 Figure 2 3 Plot of A) water content ( Q ) B) EW of Vitamin E loaded lenses versus Vitamin E loadin g. B A
47 Figure 2 4 Percent increase in diameter of A) dry lenses B) wet lenses before and after loading Vitamin E. Lines are best fit straight lines passing zero to the data. A B
48 Figure 2 5 Effect of Vitamin E loading on ion permeability of lenses. The error bars denote 95% confidence intervals. The solid dash line in A) indicates the minimum requirement for sufficient on eye movement  A B
49 Figure 2 6 Effect of Vitamin E loading on oxygen permeability (Dk) Data are presented as mean S.D. with n 3. The reported values from manufacturers are shown in hollow marker. The solid dash line indicates the minimum requirement to avoid deprivation of oxygen to cornea 
50 Figure 2 7 Transmit tance spectrum for commercial contact lenses. All measurements were conducted within 24 hours after sample preparation, and data are presented as mean S.D. with n = 3.
51 Figure 2 8 Transmittance spectrum for a) NIGHT&DAY TM and b) ACUVUE OASYS TM with different V ita min E loading. All measurements were conducted within 24 hours after sample preparation, and data are presented as mean S.D. with n = 3. A B
52 Figure 2 9 Transmittance spectrum of NIGHT&DAY TM with Vitamin E loading A) 0.15 g Vitamin E/g pure lens and B) 0.28 g Vitamin E/g pure lens, and data are presented as mean S.D. with n = 3. B A
53 Figure 2 10 Dependence of storage module of Vitamin E loaded silicone hydrogel on frequency. Vitamin E was loaded by either soak the silicone hydrogel into Vitamin E ethanol solution or by directly adding Vitamin E into monomer mixture before polymeriz ation.
54 Table 2 1 List of silicone hydrogel extended wear commercial contact lens (dipoter 6.50) explored in this study (n = 6) Commercial name a (manufacturer) Material a Dry weight measured [mg] Water content, Q measured (listed a ) [%] EW measured [%] Diameter [mm] W et measured (listed a ) D ry measured ACUVUE ADVANCE (Johnson&Johnson Vision Care, Inc., Jacksonville FL) ACUVUE OASYS (Johnson&Johnson Vision Care, Inc., Jacksonville FL) NIGHT&DAY (Ciba Vision Corp., Duluth, GA) O 2 OPTIX (Ciba Vision Corp., Duluth, GA) PureVision (Bausch&Lomb, Inc., Rochester, NY) G alyfilcon A Senofilcon A Lotrafilcon A Lotrafilcon B Balafilcon A 19.7 0.3 21.7 0.1 22.2 0.3 25.9 0.2 21.0 0.2 46.2 0.7 (47) 36.9 0.9 (38) 23.6 0.3 (24) 31.5 1.3 (33) 35.0 0.7 (36) 86.1 2.3 58.4 1.5 27.3 0.6 46.0 2.7 53.9 1.7 14.40 0.31 (14.0) 14.12 0.26 (14.0) 13.92 0.07 (13.8) 14.43 0.23 (14.2) 14.18 0.15 (14.0) 11.46 0.34 12.18 0.29 12.85 0.15 12.78 0.12 12.49 0.17 a Referred from product packages
55 CHAPTER 3 HYDROPHILIC DRUG DEL IVERY BY VITAMIN E L OADED SILICONE HYDROGEL In Chapter 2 we successfully es tablished the approach to prepare silicone hydrogel contact lens that containing Vitamin E and evaluated its potential as continu ous wear contact lens. In this c hapter we focused on the investigation of the efficacy of Vitamin E aggregates in the gel matri x as diffusion barriers to extend the drug release duration through the gel matrix. Three different ophthalmic dr ugs were explored in this c hapter, including timolol (beta blocker used for treating glaucoma), dexamethasone 21 disodium phosphate ( DXP, anti inflammatory corticosteroid), and fluconazole ( anti fungal ). These drugs were chosen because they are hydrophilic at the physiological pH, which should have negligible affinity to the desired Vitamin E barriers. 3. 1 Materials and Methods 3. 1 .1 Material s Five commercial silicone contact lenses (diopter 6.50) that are used in this study are described earlier in Table 2 1 D examethasone 21 disodium phosphate (DXP, 99%), timolol maleate ( 98%), hydroxyethyl methacrylate (97%), sodium hydroxide pellets (97+%), ethanol ( 99.5%), and p hosphate b uffered s aline (PBS) were purchased from Sigma Aldrich Chemicals (St. Louis, MO) and ethylene glycol dimethacrylate (EGDMA) from Sigma Aldrich Chemicals (Milwaukee, WI). Sodium chlorid e (99.9+ %) were purchased from Fisher Chemical (Fairlawn, NJ). Darocur TPO was kindly provided by Ciba Specialty Chemicals (Tarrytown, NY) and Vitamin E (D alpha
56 tocopherol, Covitol F1370) was gifted by Cognis Corporation All chemicals were used as rec eived without further purification if not specifically mentioned. 3. 1 .2 Drug L oading into P ure L ense s The commercial silicone contact len ses were rinsed with DI water and then dried in air before further use. The drug timolol maleate was converted to timo lol base for further use by increasing the pH of timolol maleate solution, and then separating out the precipitated timolol base. All other drugs were used as supplied. The drug (timolol, DXP, fluconazole) was loaded into the lenses by soaking the lens eit her in 2 m L of a drug PBS solution for 1 or 7 days or in the same volume of a drug ethanol solution for 3 hours. During soaking the lens in either solution the dynamic concentration in the solution was not monitored since the absorbance of these drugs in this concentration range was beyond the measurement limit of the UV VIS spectr ometer. At the end of the loading stage the lens was taken out and excess drug solution was blotted from the surface. The lens was then dried in air overnight, and used for later release experiments. 3. 1 3 Vitamin E L oading i nto P ure L enses Vitamin E was loaded into le nses by soaking the lens in 3 m L of a Vitamin E ethanol solution for 24 hours. Vitamin E ethanol solutions of various concentration s were prepared by simply mixing V itamin E and ethanol with vortexing for a few seconds followed by moderate magnetic stirring for several minutes. After the loading step, the lens was taken out and excess Vitamin E ethanol solution on the lens surface was blotted and the lens was then dr ied in air overnight.
57 3. 1 4 Drug L oading into Vitamin E L oaded L enses The drug was loaded in Vitamin E loaded lenses either by directly adding drug in the Vitamin E ethanol solution before soaking the pure lens in the solution or by soaking the Vitamin E loaded lens in a drug / PBS solution. For the case of adding drug in a Vitamin E / ethanol solution the drug was dissolved in 3 m L of a Vitamin E / ethanol solution and then the pure lens wa s soaked in this drug/Vitamin E / ethanol solution for 24 hours. For th e case of soaking in drug / PBS solution, the Vitamin E loaded lens was soaked in 2 m L of a drug / PBS solution until equilibrium. 3. 1 5 Drug Release E xperiments The drug release experiments were carried out by so aking a drug loaded lens in 2 m L of PBS. Durin g the release experiments, the dynamic drug concentration in the PBS was analyzed by measuring the absorbance of solution with a UV VIS spectrophotometer (Thermospectronic Genesys 10 UV) The absorbance of solution was measured at wavelength of 241 nm for DXP 294 nm for timolol, and 210 nm for fluconazole. Control experiments were conducted to ensure that diffusion of Vitamin E from the lenses was negligible and so it did not interfere with the drug detection. 3. 2 Results and Discussion 3. 2 1 Dynamics of D rug Transport from Contact L enses without Vitamin E Figure 3 1 shows the dynamics of timolol release by each of five contact lenses soaked in 0.8 mg/ m L of timolol PBS solution or timolol ethanol solution. The soaking duration was either 24 hours or 7 days in PBS and 3 hours in ethanol, but the release profiles for 24 hours in PBS were not drawn in Figure 3 1
58 since they were identical to those for 7 days soaking in PBS. To observe the effect of different loading methods on timolol release dynamics, mass of d rug released divided by total drug released is plotted as a function of time. All the lenses release 90% of timolol in less than 1.5 hours. In addition, timolol release profiles for different loading methods overlap for each lens except for PureVision len s which shows a slightly faster release from the lens soaked in timolol ethanol solution than that soaked in PBS medium. ACUVUE OASYS lens releases 90% of timolol relatively slow ly for 1.2 hours compared to the other lenses. ACUVUE ADVANCE lens exhibi ts rapid timolol release lasting less than 0.5 hour and the other three lenses show comparable release duration s It is observed that the release durations of timolol are not correlated to the water content of the lenses. The total amount of drug released is the highest by PureVision (about 57 g), lowest by NIGHT&DAY (about 22 g), and those of the other lenses are similar ranging 26 30 g based on PBS medium soaking method. The amounts of timolol uptake and release are also uncorrelated to the water con tent, likely due to differences in the hydrophilic components of the lenses, which lead to differences in drug binding to the hydrophilic component rich phases in the lenses. It is interesting that all the lenses soaked in ethanol solution for 3 hours rele ase substantially high total amount of timolol ; about 2.5 3 times more than those soaked in PBS solution. For example, ACUVUE OASYS lens soaked in PBS solution for 7 days releases 28 g of timolol, but that soaked in ethanol solution for 3 hours release about 95.7 g. The increased uptake of timolol from ethanol soaking is likely due to the fact that timolol does
59 not ionize in ethanol and so it preferentially binds to the polymer. In P BS, the drug is almost entirely ionized, which leads to a very large solubility in water, and consequently to small binding to the gel. The drug release from control lenses, i.e., without Vitamin E, were also conducted with the other two drugs (DXP and f luconazole) but these are not presented here because the major conclusions are the same as those mentioned above in the context of timolol. The % release profiles were independent of the method of loading and the total release durations were all about 1 1 0 hours. These control data are presented in later sections while comparing the results with the release from the Vitamin E loaded lenses. 3. 2 2 Dynamics of D rug T ransport from V itamin E L oaded L enses 3. 2 2 1 Timolol Vitamin E loaded lenses Figure 3 2 show s timolol release dynamics by V itamin E loaded lenses for different loadings of Vitamin E. T imolol and V itamin E were loaded into lenses simultaneously by soakin g the lens in 0.8 mg/m L of timolol Vitamin E ethanol solution for 24 hours. For pure lenses (no Vitamin E loading), timolol was loaded by soaking in timolol ethanol solution of 0.8 mg/m L for 3 hours. It is clearly seen in the figure that the rate of timolol release by all the lenses except PureVision TM decreases as Vitamin E loading increases while the total drug release amount does not change significantly. Specifically, NIGHT&DAY shows 9.8 fold release time for 16% Vitamin E loading corresponding to r elease time of about 5.5 hours, 76 fold for 27% corresponding to 43 hours release and 341 fold f or 74% corresponding to 192 hours release The t otal amount of timolol released by NIGHT&DAY is lowest at about 50 g. The drug transport data for
60 PureVision lenses suggests that the Vitamin E simply dissolves in the matrix leading to negligible barrier effect. However the drug transport data for ACUVUE OASYS lenses shows a significant barrier effect, which in combination with the EW data suggests that the barrier effect in these lenses likely arises due to Vitamin E that coats polymer fibers rather tha n forming larger aggregates, which appears to be the mechanism for NIGHT&DAY and O 2 OPTIX To explore the effect of the loading method, t imolol was also loaded into Vitamin E containing lenses by soaking the lenses in timolol PBS solution for 7 day s. Timolol release profiles of the ACUVUE OASYS and O 2 lenses for sequential loading of Vitamin E and timolol are also shown in Figure 3 2 It can be clearly seen that this method also increases timolol release duration compared to the control lenses without Vitamin E. Additionally, there is an increase in the total amount of drug released for the higher Vitamin E loading (74% for and 97% for O 2 ). Therefore, loading timolol and Vitamin E at the same time through ethanol medium is much more efficient way for preparation of timolol Vitamin E loaded lenses. For O 2 with same amount of Vitamin E loading, the release profiles from the lenses where timolol and Vitamin E were loaded sequentially are almost the same as for the case where timolol and Vitamin E were loaded simultaneously. However, for ACUVUE OASYS from different loading methods are similar to each other, the release profiles are slightly different. The differ ence is lik ely to be resulted from the non homogeneous distribution of timolol inside the lens. Timolol loaded by
61 drug PBS solution goes into the gel matrix by diffusion for longer time, leading to a well distribution in the lens. On the other hand, timolo l uptake in drug ethanol solution might result in high drug concentration in the center region of lens after ethanol evaporation. The morphology of the Vitamin E laden lens could potentially change over time, which could impact the drug transport. To inve stigate this issue, drug release experiments were soaked in 2 mL PBS solution after the release experiment were over. The lenses were subsequently stored for 6 months a nd then furt her soaked in 250 mL DI water with moderate stirring for 48 hours to remove the residual timolol prior to be used in second release experiment. The cleaned Vitamin E lenses were dried and weighed to ensure that the Vitamin E loading was kept the same as th e initial loading. The dry weight of the lens was within 1% difference of that measured immediately after the initial Vitamin E loading, which proves that Vitamin E does not diffuse out into PBS during the storage. The lens es were then soaked in 0.8 mg/mL timolol PBS solution for 7 days to load the drug. After the drug loading, the drug releas e profiles were measured in 2 mL PBS ( Figure 3 3 ). The release profiles in this case were almost identical to the first release profiles; this proves that the morphol ogy of the Vitamin E laden lenses is stable even when soaked in PBS for 6 months, and thus the drug release behavior of these lenses will not degrade during packaging and shelf storage. The morphology of the Vitamin E laden lenses does not
62 change during PB S soaking likely because of the negligible solubility of Vitamin E in PBS. 3.2 .2. 2 DXP Vitamin E loaded lenses DXP release profiles for various Vitamin E loaded commercial lenses are shown in Figure 3 4 The dry Vitamin E loaded lenses were soaked in 0.7 m g/mL DXP PBS solution for sufficient time to reach equilibrium. In all experiments of DXP Vitamin E loaded lenses explored here, the uptake periods were longer than the release equilibrium time, suggesting that equilibrium was achieved during loading. Figu re 3 4 indicates that, similar to the release rates for timolol, the DXP release rates from all lenses decrease as the Vitamin E loading increases, while the total drug release amount is relatively independent of the Vitamin E loading. With similar Vitamin E loading, ACUVUE 2 releases about 27 g of DXP in 7 days for 10% Vitamin E loading and in 3 weeks for 23% Vitamin E loading, while 40 8 hours even with 36% Vitamin E loaded inside. In addition, even though the drug release duration is much longer, the duration release time increase ratio by Vitamin E loaded lens for DXP is similar to that for timolol with similar Vitamin E loading amount. This suggests that the attenuation in drug release rates is similar for all hydrophilic drugs even though the diffusivities of the drugs in the pure lenses may be vastly dif ferent, which will be further discussed later.
63 3.2 .2 3 Fluconazole Vitamin E loaded lenses To further validate the hypothesis that the attenuation in drug release rates is similar for all hydrophilic drugs, we explored transport of an antifungal drug fluco nazole in Vitamin E laden lenses. Figure 3 5 shows the fluconazole release O 2 of Vitamin E loading on transport rates of t imolol and DXP from this lens. To load drugs into lenses, the Vitamin E loaded lenses were soaked in 0.7 mg/mL fluconazole PBS solution for sufficient time to reach equilibrium. The results clearly show a significant reduction in release rates due to Vitam in E g of fluconazole in 10 hours for 17% Vitamin E loading, in 24 hours for 26%, 88 hours for 39 % and 227 hour for 66% Vitamin E loading, which is a 6.2, 14, 55 and 142 fold release duration increase, respectively. The total amount of fluconazole released by different lens is similar, with the exception of O 2 which has a slightly higher drug release of about 80 g. With similar Vitamin E loading, ACUVUE luconazole release period than 2 The effect of Vitamin E loading on hydrophilic drug transport is summarized in Figure 3 6 The increase in the release times from Vitamin E loaded lenses relative to release times from the control len ses without Vitamin E is relatively similar for the three hydrophilic drugs particularly for ACUVUE O 2
64 compa res to timolol and DXP. The data also clearly shows that for each drug, the release time is quadratic to the Vitamin E loading. These issues are discussed below in the model development section. 3.2 .3 Model for Hydrophilic D rugs The hydrophilic drugs hav e a negligi ble partitioning in Vitamin E. The increase in release times for charged drugs is likely due to the presence of Vitamin E aggregates inside the gel that act as diffusion barriers. These barriers lead to an increase in the length of the path that molecules take to diffuse from inside the gel to the fluid reservoir The path length of the tortuous path should scale as where h is the half thickness of lens, and depends on the microstructure, including particle size and aspect ratio, of the Vit amin E aggregates distribution in the gel; is the volume ratio of Vitamin E in the dry gel, and is the fraction that is present as the Vitamin E particles. The fraction is assumed to be either existing as bound to the polymer gel or as particles but in regions of the gel that do not contribute to drug transport. For a diffusion controlled release, the time for release can be scales as l 2 /D The gel thickness increases due to Vitamin E uptake, and by assuming isotropic expansion and small Vitamin E load ing, it can be written as where h 0 is a half thickness of pure lens The time of release thus scales as ( 3 1 )
65 The term (1+ / 3) 2 does not make a significant contribution to increase in release time as for as large as 1, this term is less than 2. By neglecting this term we get ( 3 2 ) where time is the duration in which 90% of release is completed and 0 is the corresponding duration for the le ns without Vitamin E. It is noted that E quation 3 2 is only valid for The parameters and can be obtained by fitting the data shown in Figure 3 6 to the above model. The error between the experimental data and model prediction was defined as where and are the predicted release time ratio by model and the experimental release time ratio respectively. The parameters and for timolol, fluconazole and DXP were obtained using the function fminsearch in MATLAB minimizing the error and are listed in Table 3 1 For a given lens, the same value of was imposed in all fits since this parameter should be the same for all the drugs as it only depends on the interaction of Vitamin E with the lens matrix. Also the values of should be similar for all drugs since this is a geometric parameter that only depends on the microstructure of the Vitamin E laden lenses. The good fits between the model and the data with identical and similar for each drug further substantiate the mechanisms and the model pr esented above.
66 3. 2 4 Diffusivities of D rugs in Vitamin E L oaded L enses Contact lenses have a complex geometry including curvature with variable thicknesses from center to edge depending on power. However, a diameter of a lens (about 14 mm) is much larger than its thickness (about 80 to 100 m) and so we can simplify the geometry of lens as thin flat film with variable thickness. Under this assumption, the mass transfer problem for transport in the contact lens can be described by the following equations: ( 3 3 ) where C is the drug concentration in the gel, D is the effective diffusivity and y and t denote the transverse coordinate and time, respectively. The boundary conditions for the drug release exper iment are ( 3 4 ) where h is the half thickness of the gel, which depends on the curved lateral coordinate x C w is the drug concentration in the release medium The first boundary co ndition assumes symmetry at the center of the gel and the second boundary condition assumes equilibrium between the drug concentration in the gel and that in the PBS phase. A mass balance on the PBS in the beaker yields ( 3 5 )
67 where V w is the PBS volume, P ( x ) is the perimeter of the lens at the coordinate x and S is a half of maximum arc length Finally the initial conditions for the drug release experiments are ( 3 6 ) T he fluid volume is much larger than lens volume and the solubility of timolol fluconazole and DXP is very high in PBS of this pH 7.4, which satisfies perfect sink condition. Under perfect sink conditions, the set of equations listed above can be solved analytically to give the following solution for the concentration profile in the lens: ( 3 7 ) In short time limit, the concentration profile can also be expressed as ( 3 8 ) This result is only valid for times shorter than the By using E quation 3 5 and E quation 3 8 we obtain the following equation: ( 3 9 ) where A sur face is the total surface area of the lens. E quation 3 9 can be integrated to give, ( 3 10 )
68 The fractional release can thus be expressed as ( 3 11 ) where is the mean thickness of the gel defined as The above equation is only valid for times shorter than where h min is the minimum gel thickness, which typically equals the center thickness for negative power contact lenses. Figure 3 7 plots % drugs release by Vitamin E loaded NIGHT&DAY TM lenses as a function of square root of time for timolol The lines in the figure are the best fit straight line to short time release data The fits are all good with R 2 values large r than 0.98 showing that the drug transport in these lenses is diffusion controlled. The short time data in the drug release profiles from Vitamin E laden lenses is linear for all drugs and all lenses (data only shown for timolol release from NIGHT&DAY TM ) proving that the transport is diffusion limited for all cases. The results reported in this c hapter conclusively show that Vitamin E loading in commercial silicone contact lens can substantially increase the release duration of hydrophilic drugs The mech anism of increase in duration is due to the barrier effect of Vitamin E. While it is reasonable to assume that the effect is caused by the presence of particles of Vitamin E, it is also possible that Vitamin E does not form macroscopic aggregates and is si mply adsorbed on the polymer gel. The surface adsorption could impede surface diffusion of the drug along the
69 polymer leading to a reduction in effective diffusion rates. Also, the release profiles from the Vitamin E laden contact lenses are not zero order and that may have significant clinician implications.
70 Figure 3 1 Effect of timolol loading method on profile of timolol release by commercial contact lenses. Drug release ( M ) divided by tota l amount released ( M f ) are plotted as a function of time. Timolol was loaded by soaking the lens in 0.8 mg/ mL of indicated medium for indicated duration of time. Total amount of drug released for each lens is marked in parenthesis on the legends.
71 Fi gure 3 2 Profiles of timolol release by Vitamin E loaded contact lenses. Timolol and Vitamin E were loaded together by soaking A) ACUVUE 2 lens in e ither timolol/Vitamin E ethanol solution (shown as solid markers), or in timolol PBS solution (shown as hollow markers). Vitamin E loadings are indicated. Some of data are presented as mean S.D. with n = 3. A B
72 Figure 3 2 Continued D C
73 Figure 3 3 Profiles of repeated timolol releases by Vitamin E loaded contact lenses. For the second releases timolol was loaded by used soaking Vitamin E loaded lens in timolol PBS solution (0.8 mg/ mL ) for 7 days. Vitamin E loadings are indicated. Some of data are presented as mean S.D. with n = 3.
74 Figure 3 4 Profiles of DXP release by Vitamin E loaded contact lenses A) ACUVUE 2 in Vitamin E ethanol solution and the lens was dried. And then DXP was loaded by soaking Vitamin E loaded lens in DXP PBS solution (0.7 mg/ mL ). Vitamin E lo adings are indicated. A B
75 Figure 3.4. Continued C D
76 Figure 3 5 Profiles of fluconazole release by Vitamin E loaded contact lenses A) ACUVUE 2 was loaded first by soaking pure contact lens in Vitamin E ethanol solution and the lens was dried. And then fluconazole was loaded by soaking Vitamin E loaded lens in fluconazole PBS solution (0.7 mg/ mL ). Vitamin E loadings are indicated. A B
77 Figure 3 5. Continued C
78 Figure 3 6 Drug release duration increase by Vitamin E loaded contact lenses. A) ACUVUE 2 nd order polynomial cu rves to data of each lens. Drug release time is the duration in release of 90 % of total drug released. A B
79 Figure 3 6. Continued. C D
80 Figure 3 7 Plot of % timolol release by Vitamin E loaded NIGHT&DAY TM versus square root of time. The lines are the best fit straight for short time data. All R 2 S.D. with n = 3.
81 Table 3 1 Model p arameter s obtained by fitting experimental data to the model Contact lenses Timolol Fluconazole DXP ACUVUE 0.0117 24.2 22.0 28.8 0.0621 47.6 31.5 42.8 O 2 0.0973 35.2 35.9 42.1 0.1019 1.06 10.95
82 CHAPTER 4 HYDROPHOBIC DRUG DELIVERY BY VIT AMIN E LOADED SILICO NE HYDROGEL In Chapter 3, we have showed that the release duration of hydrophilic drugs from commercial silicone hydrogel contact lenses can be significantly increased by incorpor ating Vitamin E into the lenses The Vitamin E loaded lenses exhibit slower release of hydrophilic drugs because Vitamin E is a hydrophobic solute and so hydrophilic molecules need to diffuse around the Vitamin E barriers leading to an effective increase in release times. On the other hand, for hydrophobic drugs the transport mechanism should be significantly different since now the drug can partition into the Vitamin E aggregates. We propose that hydrophobic molecules could partition and diffuse through Vitamin E and the relative high visco sity of Vitamin E compared to the gel matrix will lead to reduced drug diffusivity. To test this hypothesis, we explore the transport of dexamethasone (DX) a hydrophobic corticosteroid through Vitamin E laden silicone hydrogel contact lenses. DX is a glu cocorticoid steroid that relieves eye inflammation and swelling, heat, redness, and pain caused by chemicals, infection, and/or severe allergies. Prolonged systemic administration of steroid can cause serious side effects such as diabetes, hemorrhagic ulce rs, skin atrophy, myopathies, osteoporosis and psychosis  In view of the potential for side effects, con trolled release of DX from contact lenses could be clinically useful. Furthermore, there are several other ophthalmic drugs that are hydrophobic and ha ve size similar to DX and thus it can be considered as a test drug to explore transport of small, hydrophobic molecules through Vitamin E laden silicone hydrogel contact lenses. This study will lead to an understanding of the effect of Vitamin
83 E loa ding on extended drug delivery for hydrophobic drugs, which in turn will allow f or rational design for extended release of other drugs from the lenses. 4. 1 Materials and Methods 4. 1 .1 Materials F ive commercial silicone contact lenses (diopter 6.50) are used in this study, including ACUVUE ADVANCE TM and ACUVUE OASYS Johnson&Johnson Vision Care, Inc (Jacksonville FL) NIGHT&DAY O 2 OPTIX Ciba Vision Corp. ( Duluth, GA) and PureVision Bausch&Lomb, Inc. ( Rochester, NY) D examethason e (DX, 98%), ethanol ( 99.5%), and p hosphate b uffered s aline (PBS) were purchased from Sigma Aldrich Chemicals (St. Louis, MO). Vitamin E (D alpha tocopherol, Covitol F1370) was kindly provided by Cognis Corporation All chemicals were used as supplied without further purification. 4.1 .2 Drug L oading in to P ure L ense s The commercial silicone contact lenses were rinsed with DI water and then air dried before further use. To evaluate the effect of different loading approaches, DX was loaded into t he lenses by either soaking the lens in either 2 m L of a drug PBS solution for 1 or 7 days or in the same volume of a drug ethanol solution for 3 hours. While soaking the lens in either solution the dynamic concentration in the solution was not monitored At the end of the loading stage the lens was taken out and excess drug solution was blotted from the surface of the lens. The lens was then air dried and subsequently used for release experiments. 4. 1 3 Vitamin E L oading into P ure L enses Vitamin E was loa ded into le nses by soaking the lens in 3 m L of a Vitamin E ethanol solution for 24 hours. Vitamin E ethanol solutions of various concentration s
84 were prepared as reported in Chapter 3 After the loading step, the lens was taken out and excess Vitamin E etha nol solution on the lens surface was blotted out, and the lens was then air dried overnight. The Vitamin E loading amount was determined by measuring the weight of dry lens before and after loading Vitamin E into the len s 4.1 4 Drug L oading into Vitamin E L oaded L enses The drug was loaded in Vitamin E loaded lenses either by directly adding drug in the Vitamin E ethanol solution before soaking the pure lens in the solution or by soaking the Vitamin E loaded lens in a drug PBS solution. For the case of add ing drug in a Vitamin E ethanol solution the drug was dissolved in 3 m L of a Vitamin E ethanol solution and then the pure lens was soaked in th e drug/Vitamin E ethanol solution for 24 hours. For the case of soaking in drug PBS solution, the Vitamin E load ed lens was soaked in 2 m L of a drug PBS solution until equilibrium. While loading DX into lenses, changes in drug concentratio n of soaking solution were monitored. The total amount of drug loaded into the gel was determined by finding the total amount of drug loss from the aqueous solution by measuring the absorbance of final solution after soaking at 241 nm for DX with a UV VIS spectrophotometer (Thermospectronic Genesys 10 UV) 4. 1 5 Drug Release Experiments The drug release experiments were carried out by so aking a drug loaded lens in 2 m L of PBS. During the release experiments, the dynamic drug concentration in PBS was analyzed in the same manner as described above for the drug loading experiments. Control experiments were conducted to ensure that diff usion of Vitamin E from the lenses was negligible and thus did not interfere with the drug detection.
85 4. 1 .6 Viscoelastic Measurement The viscoelastic response of pure Vitamin E was measured in a as a function of frequency with 0.1% str ain in a cone and pla te rheometer (AR G2, TA Instruments, New Castle, DE) with 1000 m gap at 25 o C. 4. 2 Results and Discussion 4. 2 1 Dynamics of Drug Transport from Contact L enses without Vitamin E The DX release profiles from five diff erent contact lenses for three different loading methods are sh own in Figure 4 1 Since DX i s a hydrop hobic drug and has limited solubility in PB S, DX PBS solution of 0.08 mg/m L which is close to the maximum solubility of DX in PBS at room temperature, was used for DX loading into lenses. The concentration of DX ethano l was the same, i.e., 0.08 mg/m L as that of DX PBS solution for comparison, though the solubility of DX in ethanol is about 1 mg/m L For ACUVUE ADVANCE TM ACUVUE OASYS TM and O 2 OPTIX TM the DX release profiles of three different loading methods are identical However, the D X release behavio rs by NIGHT&DA Y TM and PureVision TM lenses exhibit a slight dependency on loading methods. For these lenses, there is not much difference in the total release amount of DX from the lenses soaked in DX PBS solution for two different soaking times but slower DX release is observed f rom lenses that were soaked for 7 days than that for 24 hours. This suggests that equilibrium time for DX loading for these two lenses could be longer than 24 hours. Among five lenses NIGHT&DAY lens shows the longest release time (16 hours for 90% of total release ) followed by ACUVUE OASYS TM (10.5 hours), O 2 OPTIX TM (9.5 hours), and PureVision TM (8.5 hours), and then ACUVUE ADVANCE TM has the shortest release time (4.5 hours) by loading the drug with DX PBS solution for 7 days Th ere is a good correlation between the water content of the lenses
86 reported by the manufacturers and the duration of release as shown in Figure 4 1 F with increasing water content resulting in shorter release durations. F or total release amount of DX, PureV ision TM and ACUVUE OASYS TM lens es release relatively smaller amount s ( c.a. 28 g and 35 g, respectively) compared to the other three lenses ( c.a. 38 to 41 g). There is no correlation between amount of drugs released and the water content, which is likel y because the hydrophobic drugs are expected to partition in the silicon e rich phases, and so the partition coefficients in the gels will be mainly influenced by the silicone composition of the gels. All the lenses soaked in DX ethanol solution release sub stantially low amount of DX (2 to 8 g). The solubility of DX in ethanol is very high and the partition coefficient of DX between lens and ethanol is very low in the drug loa ding step, which results in low loading of DX. 4. 2 2 Dynamics of D rug T ransport from V itamin E L oaded L enses The d yna m ics of DX uptake and release by Vitamin E loaded lenses for f our different Vitamin E loadings are shown in Figure 4 2 The inset s in the figure show the magnified view s of the plots for drug release during the initial hours. In these experiments, Vitamin E was loaded in the lens first then air dried, and subsequently DX was loaded by soaking the lens in the DX PBS solutions. The method of loading by direct addition of DX in Vitamin E ethanol was not used since DX loading through PBS medium was much more eff icient as shown earlier In the figure all the lenses exhibit increase in loading or release time as Vitamin E loading increases. With similar Vitamin E loadings in the lenses, DX loading time is longest for ACUVUE OASYS TM followed by NIGHT&DAY TM O 2 OPT IX TM and shortest for PureVision TM For DX loading, the effect of Vitamin E loading is similar for NIGHT&DAY and O 2 OPTIX with about 2
87 fold loading time increase for about 10% Vitamin E loading, and about 10 fold for about 30% loading. However, t he effec t of about 10% Vitamin E loading for PureVision lens on loading duration is negligible and even of about 40% loading shows only 6 fold increase. These behaviors are similar for DX release time increase even though the changes in release duration are slig htly less than in loading duration F or example, NIGHT&DAY TM lenses with 27% Vitamin E loading shows 6.5 fold increase in release duration compared to 9 fold increase in loading duration with the same Vitamin E loading. The difference between the measured DX delivery time of uptake and release is likely caused by the accumulation error of drug loss during the measurement process. The release experiment is conducted in a lower drug concentration range than the uptake experiment, and therefore contains larger relative error. The comparison for the DX uptake and release delivery to an estimated model in perfect condition will be discussed later. It is noted that the effect of Vitamin E on uptake or release duration increase for hydrophilic drugs as shown in Cha pter 3 is much larger than that for DX with comparable Vitamin E loading. For example, by comparing the hydrophilic drug release and DX uptake experiment results, NIGHT&DAY TM with 27% Vitamin E loading has 76 times increase in timolol delivery time while i t has only 8.8 times increase in DX even though actual delivery time is longer for DX (1 42 hours) than for timolol (43 hours). O 2 OPTIX with 34% Vitamin E loading also shows larger increase with 34.3 fold for timolol while 15.5 fold for DX. Furthermore, w hile there is no significant difference in drug delivery time for DX and dexamethasone 21 disodium phosphate (DXP) by pure lens (For example, 10.5 hours and 14 hours by ACUVUE
88 Vitamin E loaded lenses deliver DXP for longer dura tion compared to DX. With about fold increase in release time for DXP which is about 12 days, and only 8.8 fold delivery time (4.5 days) for DX. These results also support the theory for Vitamin E aggregates insid e lens serving as diffusion barriers. Since timolol and DXP are hydrophilic ionic drugs, it cannot diffuse through the highly hydrophobic Vitamin E particles while the hydrophobic DX can partition and diffuse through Vitamin E. The reduction in release rat es for hydrophilic drugs is thus likely due to presence of Vitamin E particles that act as diffusion barriers which create an extended tortuous diffusion path For DX, while it can diffuse through the Vitamin E barrier, the diffusivity may be reduced becau se of increased viscosity and/or altered adsorption to the polymer, and this reduction in diffusivity of DX through the Vitamin E barrier could lead to the reduction in drug uptake and release rates. This will be discussed further in the section on model d evelopment. To understand the mechanism of transport of hydrophobic drugs through the Vitamin E laden lenses, it is instructive to determine the partition coefficient of the drugs both in the pure gel without Vitamin E and in the lenses with various Vitam in E lo adings. These data can then be used to obtain the partition coefficient of the drug in the Vitamin E aggregates later. For loading experiment the partition coefficient of drug in the Vitamin E loaded lens ( K ) was defined as ( 4 1 ) where V w and V l are the volumes of the drug PBS solution and the dry lens (either with or without Vitamin E load ing ) respectively, and C l ,f C w, i and C w,f are the equilibrium
89 concentratio ns of the drug in the lens phase and the initial and equilibrium concentrations in the aqueous phase, respecti vely, in the loading experiment Partition coefficient of drug in the pure lens ( K pl ) can be also written as ( 4 2 ) where V pl and C pl,f are the volume of the dry pure lens and the equilibrium concentration of the drug in the pure lens phase, respectively. The mass balance of drug in the vial yields ( 4 3 ) wh ere Mi is total mass of drug in the vial and C ve,f is the equilibrium concentration of the drug in the Vitamin E aggregates. V ve is the volume of Vitamin E aggregates in the lens and is calculated by where is the volume ratio of Vitamin E in the dry le ns and is the Vitamin E loading the could either existing in the form that bounds to the polymer gel or as particles but in regions of the gel that do not contribute to drug transport, which we have obtained previously in Chapter 3 Partition coefficien t of drug in Vitamin E phase ( K ve ) can be obtained as ( 4 4 ) T he values of K and K ve are listed in Table 4 1 K and K ve are comparable for DX, which is due to the hydrophobic nature of th e drug and Vitamin E. These partition coefficient values will be utilized in the model presented below.
90 4. 2 .3 Diffusivities of D rugs in Vitamin E L oaded L enses The thi ckness of each commercial contact lens varies in the radial direction and depends on the base curve, but the average thickness is about 80 100 m, which is much smaller compared to the diameter of lens (about 14 mm). Thus, t he drug delivery by contact lens can be considered as a one dimensional diffusion transport. To confirm whether the DX up take and release by Vitamin E loaded lenses are controlled by one dimensional diffusion as expected, the drug release profiles can be plotted as percentage of drug release versus square root of time. For diffusion controlled transport, the percentage of dr ug release will be linear to the square root of time, and the results are shown in Figure 4 3 The lines in the figure are the best fit straight line to short time release data The fits are all good with R 2 values larger than 0.9 9 showing that the drug tr ansport in these lenses is diffusion controlled. Below we develop a model based on the one dimensional diffusion equation to fit the experiment results and obtain the diffusion coefficient of DX in the lenses. Due to the large aspect ratio, we assume that the geometry of contact lens can be modeled as a flat thin film with homogenous thickness 80 m, which is the typical average thickness of commercial contact lens. The thickness variation in the radial direction can easily be integrated into the model but is not presented here for simplicity. If the drug diffusivity ( D ) and partition coefficient ( K ) are independent of the drug concentration, the drug transport to the transverse y direction can be described as ( 4 5 ) where C g is the drug concentration in the lens gel matrix. The boundary conditions for the drug release experiment are
91 ( 4 6 ) where h is the half thickness of the gel, which is about 40 m for pure contact lens without Vitamin E loading. The half thickness is adjusted with Vitamin E loading amount by isotropic expansion assumption. The first boundary co ndition assumes symmetry at the center of the gel and the second describes equilibrium of DX concentration between the gel and the aqueous phase. A mass balance on the aqueous reservoir in the beaker yields ( 4 7 ) where V w is the water volume in the beaker and A g is the cross sectional area of the lens. In addition, the initial conditions for the DX delivery are ( 4 8 ) For uptake, C g,i is zero and C w, i is the initial concentration of DX solution for loading (0.08 mg/mL ). For release, C g,i is the final equilibrium DX concentration in the lens after drug uptake process, and C w,i is zero. The equations were solved by finite difference method with MATLAB and the fitted D and K are determined by fitting the The fitting results were s hown in Figure 4 2 as solid lines. The good fits between the experiment and model results suggest the validity of our proposed model. It is noted that the fits are better for the uptake profiles compared to the release profiles,
92 particularly in the long ti me period, where the observed DX release amount are less than predicted value. This is very likely caused by the accumulated drug loss during the experiments, which also explains the observation that the experimental partition coefficients for all lenses e xplored in this study are larger for release than those for uptake. Therefore, the diffusivity values fitted to the uptake data are expected to be more reliable than those from release. Figure 4 4 shows the fitted D and K for ACUVUE OASYS TM NIGHT&DAY TM a nd O 2 OPTIX TM with different Vitamin E loading. For all lenses, while the diffusivity decreases significantly as the amount of Vitamin E in the lens increases, the drug partition coefficient almost remains the same regardless of the Vitamin E loading. The r esults suggest that while Vitamin E has a similar partition coefficient to the lens gel matrix, the DX diffusivity for Vitamin E is much smaller than that f or the lens matrix, likely due to the high viscosity of Vitamin E. 4. 2 4 Scaling Model for Effect of Vitamin E L oad ing on Extended DX D elivery The scaling mod el proposed in Chapter 3 for hydrophilic drugs delivery by Vitamin E loaded silicone hydrogel contact lens is likely not valid for hydrophobic drugs that can partition into the Vitamin E phase. For these hydrophobic drugs the transport occurs partially by diffusion around the Vitamin E aggregates and partially by dissolution and diffusion through these aggregates. Accordingly, the increase in release time is much larger for hydrophilic drugs such as timolol compared to hydrophobic drugs such as DX. The hydrophobic drugs can partition into the Vitamin E aggregates, diffuse through these, and then diffuse into the gel matrix. Thus the transport of hydrophobic drugs through the Vitamin E laden gels can b e considered as diffusion through regions of the gel matrix and regions of Vitamin E arranged in series. Since the diffusivity of DX is much smaller for Vitamin E than that for the gel matrix, the drug transport time will be
93 determined mainly by the diffus ion through the Vitamin E region when the Vitamin E loading amount increases. For one dimensional drug diffusion in a pure lens without Vitamin E loading with average thickness h the drug transport duration 0 can be estimated as h 2 / D G where D G is the drug diffusivity in the gel matrix. For Vitamin E loaded contact lens, the time it takes for the drug diffuse through the Vitamin E aggregates region can be scaled as ( h ( )) 2 / D V where D V is drug diffusivity in the Vitamin E aggregates. Thus, the ratio of the transport time increase by Vitamin E loaded lenses ( / 0 ) is given by the following expression: ( 4 9 ) The valu es of were obtained by fitting the drug transport data for the hydrophilic drugs, which is 0.0117, 0.0621, and 0.0973 for ACUVUE OASYS TM NIGHT&DAY TM and O 2 OPTIX TM respectively The only unknown parameter D G /D V can then be obtained by fitting the expe rimental data to the above equatio n. The fitting results for DX uptake duration increase by V itamin E loaded commercial lenses are shown in Figure 4 5 and the fitted D G /D V values are 330 for ACUVUE OASYS TM 395 for NIGHT&DAY TM and 405 for O 2 OPTIX TM respe ctively. The fitted results are satisfied with the assumption in our model that D G >> D V The reduced diffusivity of DX through the Vitamin E barrier is likely due to the high viscosity of Vitamin E. The diffusivity is inversely related to the viscosity a nd thus the ratio D G /D V may b e related to the ratio of the viscosity of Vitamin E and water. To test this speculation, the dynamic viscosity of Vitamin E was measured by cone and plate
94 rheometer. The slope of the log log plot of loss modulus (G") versus t he angular frequency is one, as shown in Figure 4 6 suggesting that Vitamin E can be characterized which is about 2100 fold to water at 25 o 20% of the viscosity ratio, which is encouraging. The differences between the diffusivity and the viscosity ratios could perhaps be attributed to channeling of drug through specific paths, viz. silicone rich hydrophobic channel, and thus a fraction of the Vitamin E loaded in the gel may not function as a barrier. If one assumes that only about 50% of the precipitated Vitamin E acts as barriers, the ratio of D G /D V obtained by fitting the data will increase to 4 times the values reported above bringing it in reasonable agreement with the viscosity ratio. In this c hapter we show that the drug del ivery duration for DX from contact lenses can be significantly increased to more than a week by incorporation of Vitamin E into the contact lenses. The mechanism for the extended release is likely related to the reduced diffusivity of DX through the Vitami n E barriers due to its high viscosity. A mathematical model based on diffusion controlled transport fits the uptake and release profiles from the Vitamin E loaded lenses well showing that the transport is diffusion controlled, and a scaling model fits th e dependence of effective diffusivity on the Vitamin E loading. While in vi vo studies are necessary to explore the efficacy of Vitamin E loaded lenses for ophthalmic drug delivery, the results of this study along with those from our prior studies in Chapt er s 2 and 3 strongly suggest that Vitamin E loaded contact lenses
95 could be very useful vehicles for extended drug delivery of both hydr ophobic and hydrophilic drugs.
96 Figure 4 1 Effe ct of DX loading method on profile of DX release by A) ACUVUE ADVANCE TM B) ACUVUE OASYS TM C) NIGHT&DAY TM D) O 2 OPTIX TM E) 1 versus water content of contact lenses. Drug release ( M ) divided b y total amount released ( M f ) are plotted as a function of time. DX was loaded by soaking the lens in 0.08 mg/ mL of indicated medium for indicated duration of time. Total amount of drug released for each lens is marked in parenthesis on the legends. A B C E F D
97 Figure 4 2 Profiles of experimental and model fitted DX uptake and release by Vitamin E loaded contact lenses A) ACUVUE O 2 ented by solid and hollow markers for uptake and release, respectively, and model fitted results are presented in solid line. Vitamin E was loaded first by soaking pure contact lens in Vitamin E ethanol solution and the lens was dried. And then DX was load ed by soaking the Vitamin E loaded lens in DX PBS solution (0.08 mg/ mL ). The data are presented as mean S.D. with n = 3. A B
98 Figure 4 2. Continued. D C
99 Figure 4 3 Plot of % drug release by V itamin E loaded lenses versus square root of time. The lines are the best fit s traight for short time data of DX release by A ) B C ) O 2 and D ) 2 9 Some of data are presented as m ean S.D. with n = 3. B A
100 Figure 4 3. Continued. C D
101 Figure 4 4 Fitted DX diffusivity and partition coefficient for contact lenses with different Vitamin E volume fraction ( ).
102 Figure 4 5 Effect of Vitamin E volume fraction ( ) on increase in drug uptake times. The solid lines are best fits to the data based on Equation 4 9.
103 Figure 4 6 Depen dence of the loss modulus G" on frequency for pure Vitamin E (as supplied). The slope of the log one, suggesting that Vitamin E can be characterized as a Newtonian fluid. The value of Viscosity ( ) estimated f rom the linear fit of G" to frequency was 1.918 Pa.S.
104 Table 4 1 Partition coefficient ( K ) of DX in lenses soaked in DX PBS solution. Contact lenses Vitamin E loading [g V itamin E /g pure lens] K f or loading K ve for loading K for release K ve for release ACUVUE 0 77.4 105.7 0.11 89.7 211.6 116.9 234.4 0.24 76.7 80.3 96.2 61.5 0.42 94.9 154.3 149.6 294.7 0.7 80.2 89.0 120.4 152.4 0 119.2 137.7 0.1 120.3 137.8 0.17 111.7 112.2 126.5 98.6 0.28 106.7 82.1 137.7 250.6 0.35 110.4 109.5 162.8 299.6 O 2 0 131.3 146.3 0.12 131 143.9 0.21 120.7 140.9 134.5 153.9 0.34 119.4 154.0 141.8 205.4 0.46 115.9 113.8 149.1 214.5 0 290.7 326.5 0.13 159.6 241.7 0.39 181.3 203.8
105 CHAPTER 5 ANESTHETICS DELIVERY BY VITAMIN E LOADED SILICONE HYDROGEL Excimer laser vision correction has been widely accepted in our daily life since the first proceduce was approved b y FDA in 1995, and now more than one million procedures are performed annually in the United States  Well developed excimer refractive surgery techniques for low to moderate refraction errors include myopia, hyperopia and astigmatism, but currently la ser in situ keratomileusis (LASIK) is the most preferred surgery for the treatment, followed by photorefractive keratectomy (PRK) which comprises a much smaller fraction [87 89] LASIK is the procedure of choice for most patients in the civilian community mainly because of the significantly less postoperative discomfort, faster visual recovery, and maintenance of an intact [90, 91] However, the higher risk of several serious potential complications associated with LASIK, including corneal flap loss, tear or striae, and keratectasia limits its general applications [91 95] For people with thin corneas, anterior basement membrane dystrophy and significant dry eye [96, 97] PRK remains the preferred procedure to LASIK. PRK is also preferred b y active people who are subject to trauma, such as those in the military or involved in contact sports, because the potential problems with flap stability after LASIK could lead to flap dislocation with trauma  In the United States military healthcare system, PRK is the preferred refractive surgical procedure, while LASIK has not been approved  Thus, the current focus on PRK research is to improve the postoperative pain control as well as reduce the visual recovery time after surgery. Patients who have PRK generally receive a bandage contact lens (BCL) postoperatively. Several studies have shown that the BCL protects the deepithelialized
106 cornea, leads to a faster reepithelialization, and reduces pain [98 102] Lenses are generally worn for 4 to 5 days after surgery, though typically the corneal reepithelializes in 2 to 4 days with BCL [98, 99] Contact lens with higher oxygen permeability are preferred because lens w ith low DK/t may lengthen postoperative healing time because of the decreased oxy gen exposure [98, 103] After PRK is preformed, BCL is placed on the treated eye, followed by post medication including antibiotics, anti inflammatory, lubricant eye drops, and oral and topical anesthetics. For example, the patient might need to apply one drop every 2 hours as needed for as long as the first 72 hours of topical nonpreserved 0.5% tetracaine hydrochloride to control the pain after PRK  Reports indicate that pain starts approximately 3 hours after PRK and reaches its maximum at about 7 ho urs, and usually is over about 24 hours following surgery [104, 105] The frequent dosage requirements interfere with the patients daily activities and can lead to the potential risk of drug overdose. The aim of the study in this chapter is to in vitro in vestigate the potential of using Vitamin E loaded contact lens for postoperative treatment after PRK to obtain better pain control for patients. Three common topical anesthetic drugs are explored here, including lidocaine, bupivacaine and tetracaine, and t he molecular structures of these drugs are shown in Figure 5 1 The pKa values are 7.4, 8.1 and 8.4 for lidocaine, bupivacaine and tetracaine, respectively, and thus these three drugs present in ionized forms at physiological pH. To clarify the drug trans port mechanism inside these composite hydrogel systems, we also prepared lab synthesized silicone hydrogel to obtain further understanding for future model prediction.
107 5. 1 Materials and Methods 5. 1 .1 Materials Commercial silicone hydrogel contact lenses O 2 OPTIX (Lotrafilcon B, diopter 6.50 ) from Ciba Vision Corp. (Duluth, GA) were used in this study Lidocaine hydrochloride, bupivacaine hydrochloride, tetracaine hydrochloride, e thanol (>99.5%) sphate buffered saline (PBS) were purchased from Sigma Aldrich Chemicals (St. Louis, MO), and Vitamin E (D alpha tocopherol, Covitol F1370) were kindly gifted by Cognis corporation (Kankakee, IL). For preparation of silicone hydrogel, e thylene glycol dimethacryla te (EGDMA, 98%), N, N Dimethylac ry lamide (DMA, 99%) and 1 vinyl 2 pyrrolidone (NVP, 99+ %) were purcha sed from Sigma Aldrich Chemicals (Milwaukee, WI ). The macromer acryloxy(polyethyleneoxy) propylether terminated poly(dimethylsiloxane) (DBE U12, 95+ %) were purchased from Gelest Inc. (Morr isville, PA). 3 Methacryloxypropyltris(trimethylsiloxy)silane (TRIS) was supplied by Silar laboratories (Scotia, NY) and 2, 4, 6 trimethylbenzoyl diphenyl phosphineoxide (Darocur TPO) were kindly provid ed by Ciba Specialty Chemical s (Tarrytown, NY). All chemicals in this study were reagent grade and used as supplied without further purification. 5.1 .2 Drug L oading into P ure L ense s O 2 OPTIX lens was rinsed with DI water and then air dried before further use. Drugs were loaded into the lenses by soaking th e lens in 3 m L of a drug PBS solution for at least 7 days until reaching equilibrium The initial drug concentrations were 5, 2.5 and 1 mg/mL for lidocaine, bupivacaine and tetracaine, respectively. While soaking the lens in either solution the dynamic co ncentration in the solution was not monitored At the end of the loading stage the lens was taken out and excess drug solution was
108 blotted from the surface of the lens. The lens was then air dried and subsequently used for release experiments. Since all th e drugs explored in this study have high solubility in ethanol, these drugs were also loaded into the contact lens by soaking the lens in a 3 mL of drug/ethanol solution for 24 hours. T he initial soaking drug concen trations were 10, 10 and 1 mg/mL for lido caine, bupivacaine and tetracaine, respectively. At the end of the loading stage the lens was taken out and excess drug solution was blotted from the surface of the lens and subsequently used for release experiments. 5.1 3 Vitamin E L oading into P ure L ens es Vitamin E was l oaded into contact lens by soaking a lens in 3 mL of a Vitamin E ethanol solution for 24 hours. After the load ing step, the lens was withdrawn and blotted to remove excess solution on the surface The lens was then dried in air overnight. The Vitamin E loading amount was det ermined by measuring the increase in lens weight. The mass of Vitamin E loaded into a lens is directly proportional to the concentration of Vitamin E loading in the ethanol Vitamin E solution  The concentrations of the Vitamin E in the loading solution in this study were 0.05, 0.10 and 0.15g Vitamin E/g ethanol, which leaded to about 0.18, 0.37 and 0.55 g Vitamin E/g pure lens, respectively. Subsequently, these Vitamin E loaded contact lens were soaked into drug /PBS solutions through the same drug loading procedure for pure lens, as de scribed in Section 5.2.2 The drugs were also loaded into lens by directly adding drug into Vitamin E/ethanol solution. E ach drug was dissolved in a 3 mL of 0.05, 0.10 or 0.15g Vit amin E/g ethanol solution for 24 hours, and the drug concentration was designed to be 10, 10 and 1 mg/ mL for lidocaine, bupivacaine and tetracaine, respectively.
109 5.1 .4 Drug Release Experiments The drug release experiments we re carried out by soaking drug impregnated lens in 2 mL of fresh PBS. Since all these three anesthetic drugs have high solubility in water and the volume of aqueous medium is much larger than that of the hydrated contact lens, the drug release can be thus viewed as drug transport at per fect sink condition. Therefore, the amount of residue drug in the lens at final equilibrium is negligible, and the initial drug loading is equal to the total amount of drug release. The dynamic drug concentration in aqueous solution was determined by measu ring the absorbance in the wavelength range with a UV VIS spectrophotometer (Thermospectronic Genesys 10 UV) UV VIS absorbance was converted to the concentration of drug by following the absorbance spectra deconvolution method reported previously to detec t both drug and potential Vitamin E release  The absorb ance was measured in from 231 to 291 nm for lido caine and bupivacaine, and 195 to 255 nm for tetracaine. The absorbance was measured in scanning range rather than at a single wavelength to ensur e that the experimental methods did not lead to drug degradation which will manifest as changes in the absorption spectrum. In this study, the drug release duration by the lens is defined as the time it takes to complete 90% of final drug release amount a t equilibrium. 5. 1 5 Silicone H ydrogel P reparation To prepare the silicone hydrogel, hydrophilic monomers with high ion permeability are copolymerized along with the hydrophobic silicone monomer with high oxygen permeability, and a macromer is needed in t he monomer mixture to ensure solubilization of all monomers. In this study, TRIS was used as the hydrophobic monomer, DMA was the hydrophilic monomers, and DBE U12 was selected as the
110 macromer. Highly hydrophilic NVP monomer was also added to increase wate r content of the hydrogel and EGDMA was introduced in the monomer mixture for controlled crosslinking. To prepare the polymerizing mixture, 2.4 mL of a mixture that comprises 0.8 mL TRIS and 0.8 mL macromer and 0.8 mL of the hydrophilic DMA/MAA m ixture was combined with 0.12 mL of NVP and 0.1 mL of EGDMA. After well mixed with vortexing for few second, the mixture was purged with bubbling nitrogen for 15 minutes to reduce the dissolved oxygen. To each monomer mixture, 12 mg of photoinitiator Darocur TPO wa s added with stirring for 5 minutes and the final mixture was immediately injected into a mold which is composed of two 5 mm thick glass plates. The plates were separated by a plastic spacer with various thicknesses. The mold was then placed on ultraviolet transilluminator UVB 10 (UltraLum Inc.) and the gel mixture was cured by irradiating with UVB light (305 nm) for 50 minutes. The synthesized hydrogel was either cut into circular pieces (about 1.65 cm diameter) with a cork borer for subsequent experiments Prior to conducting further tests, the prepared hydrogel was soaked in ethanol for 3 hours then dried at ambient temperature overnight to remove the unreacted monomer within. 5. 1 6 Partition C oefficient The synthesized silicone hydrogel was used for investigation of the dependency of lidocaine partition coefficient in silicone hydrogel on drug concentration. To load Vitamin E into the hydrogel, each circular piece of hydrogel was soaked in a 5 mL of 0.35 g/mL Vitamin E/ethanol for 24 hours, which resu lted in a loading of 0.28 0.01 g Vitamin E/ g pure gel. The hydrogel, with or without Vi tamin E, was then soaked in 5 mL of lidocaine/PBS solution with v arious concentration until reached equilibrium. The
111 equilibrium drug loading in the hydrogel was subs equently determined by conducting drug release in fres h PBS The volume of the release medium was adjusted with the drug loading to assure the drug concentration in the aqueous medium maintained in measurable range. 5. 1 7 Determination of Critical Micelle Concentration (CMC) of L idocaine Surface tension isotherm of lidocaine was measured at room temperature by creating a pendant drop of lidocaine/PBS solution against ambient atmosphere. The drop shape was digitally imaged and then fitted to the Young Laplac e equation by using the Drop Shape Analysis System DSA100 (KRSS) to calculate the surface tension. The concentration of lidocaine sol ution was varied from 0.01 mg/mL to 360 mg/mL 5. 2 Results and Discussion 5. 2 1 Dynamics of Drug Release from Contact L ens es 5.2 .1.1 Drug uptake through drug PBS solution The results of lidocaine release by O 2 OPTIX TM with various Vitamin E loadings were shown in Figure 5 2 The drug release duration increases as the Vitamin E loading increases. For instance, pure O 2 OPTIX TM released 90% of its initial drug loading in 1.8 hours, while lenses with 27% and 36% of Vitamin E can extend the release duration to 6.2 and 10.8 hours, respectively. In addition, Vitamin E inside the lens can enhance the total lidocaine loading when soak ed in drug/PBS solution. Since Vitamin E is highly hydrophobic, when the Vitamin E loaded lens is equilibrium with drug/PBS solution it generally does not affect hydrophilic drug loading, such as timolol, dexamethasone phosphate, and fluconazole  Th e pH of the lidocaine/PBS solutions ranges from 6.0 to 7.4 based on the drug concentration in this study, which is lower than the pKa of lidocaine. Thus, the majority of lidocaine should present in the
112 hydrophilic ionized form, which is highly unlikely to partition into the Vitamin E aggregates The increase of drug loading could be resulted from the surface adsorption of lidocaine on the interface between Vitamin E and the gel matrix, and we will further examine this assumption later. Another support of t he interaction between lidocaine and Vitamin E were found, as we observed tiny amount of Vitamin E (less than 1% of the Vitamin loading inside the lens) release into the aqueous reservoir during the drug release experiment, as shown in Figure 5 3 Since no Vitamin E release were detected for other hydrophilic drugs b y the same system it is reasonable to assume that the existence of lidocaine enhanced the solubility of Vitamin E in PBS. In practical, the Vitamin release should not cause significant differen ce on the performance of these Vitamin E loaded lens because first the loss amount is negligible and second in practice use the drug concentration is much lower than our experiment condition, which should lead to even lower Vitamin E loss. Figure 5 4 exhib ited the bupivacaine release by O 2 OPTIX TM Similar to lidocaine, the release duration of bupivacaine extended with the amount of Vitamin E loading in the lens. Pure O 2 OPTIX TM released 90% of its initial drug loading in 3.2 hours, while lenses with 27% and 36% of Vitamin E loading can release in 10.2 and 20.7 hours, respectively. The drug uptake by the lens was also enhanced with Vitamin E, while similar Vitamin E loss were detected during bupivacaine release as well (data not shown). The results of tetraca ine release by O 2 OPTIX TM with various Vitamin E loadings were shown in Figure 5 5 Again, higher Vitamin E loading in the lens resulted in longer drug release duration. Pure O 2 OPTIX TM released 90% of its initial drug loading in 2.4 hours, while lenses wit h 27% and 36% of Vitamin E loading can release the loaded
113 tetracaine in 13.9 and 22.6 hours, respectively. However, unlike lidocaine and bupivacaine, the total drug uptake does not significantly increased as the Vitamin E loading amount increased. This is possibly due to the higher light sensitivity of tetracaine compared to lidocaine and bupivacaine, which leads to drug degradation during release experiment. The degradation can be observed by the gradual drug loss in long time, as shown in Figure 5 5 B It was also observed by the absorbance spectrum change in the long time release (data not shown). Vitamin E release is not discussed during tetracaine release since its absorbance spectrum is not overlapped with that of tetracaine in our study. Even though V itamin E inside the lens can effectively extend the release duration of these anesthetic drugs, the effect is not as significant as those on other model drugs. For example, with ca. 0.36 g Vitamin E/g pure lens loading, the drug release time increased to f rom 1.8 hour by pure lens to 6.2 hours, which is a 3.5 fold increase. With similar amount of Vitamin E loading, the lens can release the hydrophilic drug timolol for 28 hours, a 40 fold increase compare d to pure lens ; it can also release the hydrophobic dr ug dexamethasone for 150 hours, which is a 15 fold increase. For hydrophilic drugs, such as timolol, the Vitamin E loading inside the lens act as diffusion barriers for drug transport due to the negligible affinity between drug and Vitamin E aggregates; fo r hydrophobic drugs, such as dexamethasone, the drug can freely partition into the highly viscous Vitamin E aggregates. The fact that we observed both hydrophobic and hydrophilic behavior of lidocaine in the Vitamin E loaded silicone hydrogel composite sys tem implies that lidocaine could act as a surfactant like molecule in this system, and thus the transport mechanism of these anesthetics cannot be simply considered as
114 either hydrophilic or hydrophobic drug alone, but have to be controlled by other mechani sm based on the unique interaction between Vitamin E and these anesthetics drugs, which will be discussed later. 5.2 .1.2 Drug uptake through drug ethanol solution The drug release results by lens which the drug uptake was through drug/ethanol solution w ere shown in Figure 5 6 With same amount of Vitamin E in the lens, the drug release duration significantly increased by drug/ethanol uptake compared to that by drug/PBS uptake. For instance, for O 2 OPTIX TM with 27% of Vitamin E loading, the drug release d uration of lidocaine and bupivacaine are 70 and 32 hours, respectively, which is much higher than 6.2 and 10.2 hours by drug/PBS uptake. In addition, the total drug uptake amount by the lens is independent of Vitamin E loading, which is different than what we observed in the drug/PBS uptake. On the other hand, similar characterizations such as degradation of tetracaine ( Figure 5 6 C ) and Vitamin E release from lidocaine and bupivacaine (data not shown). The possible mechanisms that result in these differenc es will be discussed later. In this study we explore two different approaches to load drug into the contact lens. Drug uptake through drug/PBS solutions requires longer time to reach equilibrium, and it need two steps to load Vitamin E and drug separately. However, since the lens was kept in the drug/PBS solution, this method eliminates the drug loss in the package solution. Since the anesthetic drug release time can be increased to about 20 hours with 35% Vitamin E through drug/PBS uptake, these lenses can be used to provide sustained topical anesthetics release based on daily replacement, which still has its practical potential as the postoperative pain usually felt by the patient within the first day after PRK surgery. On the other hand, drug uptake throu gh drug/ethanol requires much
115 less time and can directly load Vitamin E and drug in one step. However, these lens need to be kept in a package condition for practical use, with could lead to drug loss and reduce the effect of extension of drug release dura tion. To explore the packaging effect on the lidocaine release by contact lenses, O 2 OPTIX TM lens was first soaked in a 3 mL of 10 mg/mL lidocaine / ethanol Vitamin E solution for 24 hours, where the Vitamin E concentration is 0.1 g Vitamin E/ mL ethanol, an d the drug loaded lens was subsequently removed into a 3 mL of 10 mg/mL drug PBS solution for 1 and 7 days prior to in vitro release. As shown in Figure 5 7 while through drug/ethanol uptake the lens is able to release 90% of the included lidocaine in 70 hours, the release duration decreases to 40 hour s after soaking the lens in drug/PBS solution for 1 day, and further reduced to about 7 hours after 7 days, which is similar to the duration by lens through drug/PBS uptake. 5. 2 2 Lidocaine R elease S tudy ( S urfactant B ehavior) The observed behavior of lidocaine transport in Vitamin E loaded contact lens implied that the lidocaine molecular could act as a surfactant like molecule between the hydrophobic Vitamin E aggregation and hydrophilic regions inside the hydrogel. To further verify this hypothesis, the surface tension isotherm of lidocaine hydrochloride in PBS was measured, a nd the results were plotted against drug concentration in logarithmic scale in Figure 5 8 The result clearly demonstrated that the l idocaine affects the surface tension of the drop when the concentration is above 0.2 mg/mL and the CMC of lidocaine in PBS is above 360 mg/ mL which is much higher than th e concentrations explored in our loading release studies. The drug partition coeffi cient in a pure silicone hydrogel (K gel ) based on the mass balance in the uptake experiment can be defined as:
116 ( 5 1 ) Where C gel,f and C w,f is the final drug concentration in the hydrogel a nd in the aqueous medium when the uptake experiment reached equilibrium, and the volume of hydrogel and aqueous medium were notated as V gel and V w respectively. The only unknown parameter C gel,f can be determined by the drug extract experiment at perfect sink. Similarly, the overall apparent drug partition coefficient in a Vitamin E loaded silicone hydrogel (K ve gel ) can be defined as: ( 5 2 ) In addition, the drug uptake by the Vitamin E loaded silico ne hydrogel is the summation of the amount of drug uptake by pure hydrogel and by Vitamin E aggregates in the gel, and thus the drug partition coefficient in Vitamin E aggregates can be determined as: ( 5 3 ) The results of K gel K ve and K ve gel at different initial soaking lidocaine concentration are shown in Figure 5 9 Since lidocaine is highly likely acting as a surfactant like molecule in our Vitamin E loaded silicone hydrogel system, it is reasonable to assume that the majority of lidocaine uptake by Vitamin E aggregates inside the hydrogel is through surface binding. According to general Langmuir adsorption model at equilibrium, which relates the adsorbed surface concentration of the drug on the Vitamin E aggregates ( ) to the free drug concentration in the aqueous phase( C ) by the following equation
117 ( 5 4 ) where is the surface conce ntration at the maximum packing on the surface, and k ad and k d is the rate constants for adsorption and desorption of the drug on the Vitamin E surface, respectively. In addition, the previously obtained K ve can be related as: ( 5 5 ) where V and S are the volume and surface area of Vitamin E aggregates, respectively. By Equation 5 4 and 5 5 we can derive the following relation: ( 5 6 ) Therefore, if the Langmuir adsorption model successful describes the interaction between lidocaine and Vitamin E, the inverse of K VE should be linear to the bulk drug concentration, and the parameter V/S can be obtained as the value of the slope. The good fitting of experimental results to this linear model, as shown in Figure 5 10 highly supported the validity of our assumption, and the value of V/S is determined as 0.0447 mL /mg. 5. 2 3 Mechanisms o f E xtended D rug R elease by Vitamin E L oaded C ontact L ens The anesthetic drugs are in charged form at or below physiological pH in our study, and thus these drugs should ha ve a negligi ble partitioning into the Vitamin E aggregates inside the lens matrix. Si nce it is observed that these anesthetic drugs have strong interfacial interaction at the surface of Vitamin E aggregates, it is reasonable to assume that the Vitamin E aggregates inside the gel matrix act as diffusion barriers for the drug transport When encountering the Vitamin E aggregates, drugs in the gel matrix need to
118 take a detoured route around the surface of Vitamin E with specific surface diffusivity to diffuse out of the lens. The increased path length of the tortuous path should scale as w here h is the half thickness of lens, and is the parameter that depends on the microstructure, including particle size and aspect ratio, of the Vitamin E aggregates distribution in the gel; is the volume ratio of Vitamin E in the dry gel, and is the fraction that is present as the Vitamin E p articles. The fraction is assumed to be either existing as bound to the polymer gel or as particles but in regions of the gel that do not contribute to drug transport. For a diffusion controlled release, t he release duration from a pure lens can be scale d as: ( 5 7 ) whe re h o is the half thickness of pure lens, and D gel is the diffusion coefficient of the drug inside the pure gel matrix. Where Vitam in E was loaded into the lens, the gel thickness h i ncreases due to Vitamin E uptake, and by assumin g isotropic expansion it can be written as Thus, we can estimate the drug release duration from Vitamin E loaded contact lens as ( 5 8 ) whe re D s is the surface diffusivity of drug on the interface between Vitamin E aggregates and the gel matrix. Due to the surfactant like behavior of these anesthetics, we expected that D s should be higher than D gel By combin ing Equation 5 7 and 5 8, we can obtain:
119 ( 5 9 ) From ou r previous studies in Chapter 3 for O 2 OPTIX TM is about 35 and is 0.0937. Equation 5 9 were fitted to the experimental data by using the function fminsearch in MATLAB to obtain the parameter D gel / D s for each drug. The model fitting results were shown in Figure 5 11 and the fitted D gel / D s are 0.0291, 0.0319 and 0.0460 for lidocaine, bupivacai ne and tetracaine, respectively. Earlier studies have shown that the release behavior of most drugs by the contact lens, with or without Vitamin E, is independent on the drug loading approaches [106, 108] However, the drug release duration and the drug loading capacity of these three anesthetic drugs by Vitamin E loaded lens through drug/ethanol uptake are significantly different than those through drug/PBS uptake. When lidocaine is loaded into the lens through drug/ethanol solution, since lidocaine shou ld be in uncharged form in ethanol, it is reasonable to assume that the uncharged lidocaine is mostly contained in the hydrophobic regions of the silicone hydrogel matrix after ethanol evaporation, and the drug loading capacity is simply determined by the equilibrium ethanol uptake of the lens. When Vitamin E is also included in the drug/ethanol solution, while the total loading capacity is not affected, the uncharged lidocaine loading now should distribute in both the hydrophobic silicone region of the gel matrix and the Vitamin E aggregates based on the respective partition coefficient. Since the uncharged lidocaine will changed to charged form once it encounter the PBS during drug release which can be explained by Equation 5 8, the extra resistance of li docaine transport observed form the lenses through drug/ethanol uptake should be arose from the drug diffusion inside the hydrophobic regions inside the V itamin E/hydrogel matrix.
120 Therefore, the effect on lidocaine release extension by loading Vitamin E and lidocaine simultaneously could results from two possible mechanisms. First, the majority of drug could partition into the hydrophobic region of silicone hydrogel matrix, and the Vitamin E aggregates on the boundary between hydrophobic and hydrophilic region serves as diffusion bar riers to hinder the drug release from hydrophobic region to hydrophilic region. In this case, the additional drug release time can be scaled as where and D silicone are the respective aspect ratio parameter and diffusion c oefficient for lidocaine diffusion around the Vitamin E aggregates from silicone rich hydrophobic region in the gel matri x. Another alternative hypothesis is that the additional drug release duration is controlled by the drug partitioned in the Vitamin E a ggregates with relatively smaller diffusivity compared to that of hydrogel matrix. In other words, Vitamin E works as a drug reservoir that provides sustained release in this case, and the one dimensional diffusion inside Vitamin E aggregate can be scaled as where r is th e effective radius of the Vitamin E aggregates. Thus, when drug loaded gel is soaked in fresh PBS, the overall lidocai ne transport is the combination of the drug distribution between gel matrix and Vitamin E, the thermodynamic equilibriu m between charged and uncharged form of lidocaine at the interface between Vitamin E and the hydrated hydrogel, and the drug diffusion inside the Vitamin E aggregates. To further examine the above assumptions on the mechanism of lidocaine transport by Vita min E loaded silicone hydrogel, the lab synthesized silicone hydrogel with different thickness were used to conduct the drug uptake/release experiments. As shown in Figure 5 12 the synthesized silicone hydrogel (with and without Vitamin E)
121 demonstrated si milar lidocaine transport behaviors as those by commercial O 2 OPTIX TM While the 0.2 mm thick pure gel released 90% of the loaded lidocaine in about 5 hours, the hydrogel with 0.25 g Vitamin E/g pure gel loading can extend the release duration to 15 and 48 hours through drug/PBS and drug/ethanol uptake, respectively. The loading capacity increased with Vitamin E loading through drug/PBS uptake, but kept the same through drug /ethanol uptake, which is the same as we observed from O 2 OPTIX TM Figure 5 13 prese nted the lidocaine release results by silicone hydrogel with various thicknesses, of which the drugs were loaded through drug/PBS uptake. The drug release time increased when the Vitamin E loading or the gel thickness increased. If Equation 5 8 holds to ex plain the lidocaine release transport here, than the kinetic release time should be proportion to the square of gel thickness; i.e., if we define a scaled time as time/(gel thickness/0.1 mm) 2 then the release results from gels with different thickness sho uld be overlapped when plotted against the scaled time if the drug transport is controlled by one dimensional diffusion, and the results were shown in Figure 5 13 B For both the gels with or without Vitamin E loading, the % releases overlapped with differe nt thicknesses, which support the validity of our proposed transport mechanisms. For the lidocaine release by gel through drug/ethanol uptake, if the additional drug release resistance comes from the detour around Vitamin E aggregates from hydrophobic regi ons to hydrophilic regions of hydrogel matrix, the drug release time should be still proportional to the square of thickness when the Vitamin E loading inside the gel kept the same. However, if the drug release was mainly controlled by the drug diffusion i n the Vitamin E aggregates, it is only affected by the size of Vitamin E
122 aggregates in the gel matrix instead of overall gel thickness. As shown in Figure 5 14 the significant difference between the scaled releases of Vitamin E loaded gel with different t hicknesses suggested that the latter should be the more appropriate assumption to describe the lidocaine transport by Vitamin E loaded silicone hydrogel through drug/ethanol uptake. In this study we investigate the potential to provide extended anesthetics delivery by Vitamin E loaded silicone hydrogel contact lenses for postoperative pain control, especially for patients who accepted PRK for vision correction. The thermodynamic properties of these anesthetic drugs, including the amphiphilic behaviors and t he dependency on the pH of environment, significantly affects the mechanisms of drug transport by Vitamin E loaded silicone hydrogel contact lenses at different drug loading conditions. The Vitamin E loaded silicone contact lens can provide continuous anes thetics release for about 1 day through drug/PBS uptake. The release duration by lenses through drug/Vitamin E ethanol uptake can be further increased, while the packaging effect needs to be overcome for future practical use. Future in vivo studies are als o needed to further evaluate the feasibility of sustained anesthetic release by contact lenses. Furthermore, the potential complexity of drug interactions between the anesthetics and other ophthalmic drugs during the invol ved in postoperative treatment sho uld be taken into consideration in the future work.
123 Bupivacaine Lidocaine Tetracaine Figure 5 1 Molecular structures of model drugs
124 Figure 5 2 Lidocaine release in PBS by O 2 OPTIX TM with various Vitamin E loading.
1 25 Figure 5 3 Vitamin E release from O 2 OPTIX TM during lidocaine release in PBS.
126 Figure 5 4 Bupiv acaine release in PBS by O 2 OPTIX TM with various Vitamin E loading
127 A B Figure 5 5 A) Short time and B ) long time tetr acaine release in PBS by O 2 OPTIX TM with var ious V itamin E loadings.
128 Figure 5 6 A ) L idocaine B) bupivacaine and C ) tetr acaine release in PBS by O 2 OPTIX TM with various Vitamin E loading. Drugs were loaded by soaking in drug/ethanol Vitamin E solution for 24 hours. A B
129 Figure 5 6. Continued. C
130 Figure 5 7 Lidocaine release by O 2 OPTIX TM with 0.36g Vitamin E/g pure lens. Lidocaine was loaded into the lens by soaking in 10 mg/ mL drug/ethanol for 24 hours (Ethanol) and subsequently soaked in 10 mg/ mL drug/PBS for 1 day (Ethanol +PBS (1 day)) or 7 days (Ethanol +PBS (7 days) prior to release in 2 mL PBS. Drug was also loaded by soaking the lens in 10 mg/ mL drug/PBS soluti on for 7 days.
131 Figure 5 8 The relationship between surface tension and lidocaine concentration in PBS.
132 Figure 5 9 The calculated partition c oefficient (K) of Vitamin E loaded silicone hydrogel at various lidocaine hydrochloride concentrations.
133 Figure 5 10 The relationship between the lidocaine partition coefficient in Vitamin E (K VE ) and the bulk drug concentration.
134 Figure 5 11 Model fitting for anesthetic drug release increase ratio on Vitamin E loading fraction in the silicone hydrogel.
135 Figure 5 12 Lidocaine release from pure 0.2 mm thick silicone hydrogel or gel with Vitamin E loading (0.25 g Vitamin E/g pure gel). Lidocaine were loaded into hydrogels through drug/PBS solution (10 mg/ mL ) or drug/(eth anol+Vitamin E) solution (10 mg/ mL ).
136 Figure 5 13 Lidocaine release by silicone hydrogel with or without Vitamin E loading (0.25 g Vitamin E/g pure gel) with various thickness. Lidocaine was loaded i nto the sample through soaking the gel into 10 mg/ mL drug/PBS solution.
137 Figure 5 14 Lidocaine release by silicone hydrogel ( with or without Vitamin E loading (0.25 g Vitamin E/g pure gel)) with vario us thicknesses Lidocaine was loaded into the sample through soaking the gel into 10 mg/ mL drug/(ethanol + Vitamin E) solution.
138 CHAPTER 6 ION TRANSPORT OF SIL ICONE HYDROGEL Contact lenses for correcting vision are available in several different wear modaliti es including extended continuous wear for a period of 1 4 weeks, depending on the type of the lens. Extended wear lenses are required to allow rapid oxygen transfer because cornea is an avascular organ, and so it gets its oxygen supply directly from air [ 22, 23] Silicone based materials were explored as potential candidates for achieving the high oxygen transport due to the very high oxygen solubility in these materials. However, silicone based contact lenses were not useful as contact lens materials bec ause the lenses made of such materials adhered to the cornea. It may be speculated that the hydrophobic nature of the silicone lenses leads to the adherence. It was however determined that surface treatment of a lens to render the surface hydrophilic is not enough to prevent adherence to cornea. If a particular lens material did not move on the eye, application to the surface of contact lens did not change the outcome significantly  It was later discovered that the extended wear contact lenses are required to also allow sufficient ion transport to maintain on eye movement and not adhere to the cornea. The importance of ion permeability of contact lens material for maintaining lens motion was first described by Domschke et al., in a 1997 ACS present ation  The need of ion transport through the lens is attributed to the requirement of a fluid hydrodynamic boundary layer between the lens and the cornea  In the absence of the fluid layer, the lens can adhere to the cornea. Nicolson et al. re ported in a US patent to claim t he ionoflux diffusion coefficient (D ion ) i.e. the ion permeability of the silicone hydrogel material, should be at least larger than 1.5 10 6 mm 2 /min for sufficient on eye movement of lens  While ion permeability is critical to
139 lens motion, an increase in the permeability beyond a critical value does not lead to a further increase in on eye movement of the lens  To compensate the drawback of the negligible water content from pure silicone, hydrophilic monomers are copolymerized along with the hydrophobic silioxane monomer. In general, the polysiloxanes and the hydrophilic polymers are immiscible and thus a proper macromer is needed in the monomer mixture to ensure solubilization of all monomers. The final sili cone hydrogel matrix can be best described as isotropic structures, either as a dispersed system with only one continuous phase, or as an interpenetrating polymer network that both the hydrophilic and hydrophobic phases are continuous throughout the materi al. A number of the commercial extended wear contact lenses are reported to possess a bicontinuous microstructure that facilitates rapid exchange of oxygen through the silicone rich phase and ions through the hydrophilic phase  Thus, fine tuning of ea ch composition in the hydrogel mixture with proper microstructure is critical for ensure the balance of all the key properties of the extended wear contact lenses, including oxygen permeability, ion permeability, water content, elastic modulus, surface wet tability, etc. While the importance of ion permeability has been known for extended wear contact lenses, surprisingly there are a very limited number of studies focusing on ion transport through these silicone hydrogels. Most prior studies on ion transpo rt in contact lenses have merely reported the ion permeability, i.e., the product of the diffusivity and the partition coefficient, by direct permeation approach in a diffusion cell, which is a better mimic to the real physiological environment [19, 25, 41 84, 106, 110] Ion permeability was established by measurement of the flux of ions from the donor
140 reservoir, across the lens and into the receiver. The ion permeability could then be eved. While ion permeability is important because it directly determines the net ion flux across the lenses at pseudo steady state, independent measurements of the diffusivity and partition coefficient are important as well because these two relate to diff erent aspects of the lenses. Since partition coefficient is a thermodynamic property, it likely depends only on the total fraction of the two phases (silicone and hydrophilic) in the hydrogel matrix, while the diffusivity depends strongly on the connectiv ity of these phases (dispersed or bicontinuous morphology). Thus, a further scrutinized study on the salt transport through these silicone hydrogel materials is needed. The diffusion of solutes in conventional hydrogels has been widely studied due to the interest of its wide application, including separation process such as chromatography  water purification [112 114] and hemodialysis as an artificial kidney operation [115, 116] Salt transport in hydrogels is typically explored through direct memb rane permeation and kinetic sorption/desorption methods. In the direct membrane permeation study, the permeability of a solute is calculated based on the measured permeation flux [112, 117 123] In the sorption/desorption experiments, the overall sorption or desorption kinetics are analyzed to characterize the diffusion coefficients of the solutes in the hydrogel [112, 113, 119 121, 123, 124] Both models diffusion model to extract the transport paramet ers, and thus is suitable for homogeneous hydrogel systems where the transport of solute is merely controlled by self diffusion  In both simplified model only the salt release in the initial short time period (kinetic sorption/desorption) or
141 the pseu do steady state ion flux (direct permeation) were used to obtain the transport parameters. Since silicone hydrogel could have more complicated microstructure then thes e conventional hydrogels, the validity of these models on silicone hydrogel systems need ed to be verified. There are a number of models describing the diffusion of solutes in hydrogels, and an overview of these models and experimental data was earlier made by Amsden  Since the solute transport occurs primarily within the water filled re gions in the space bounded by the polymer chains, any factor which affects the size of these spaces will have an effect on the transport of the solute through the hydrogel matrix, including the relative size of solute and of the openings between polymer ch ains, polymer chain mobility, and the existence of charged groups on the polymer which may have strong interaction with the solute molecule  One of the most popularly used model to explain sodium chloride transport in hydrated homogenous hydrogels wa s first proposed by Yasuda et al.  The model is based on the free volume theory by Cohen and Turnbull which explains the process of solute diffusion in a pure liquid composed of hard molecular spheres  Their derivation is based on the concept t hat molecular transport occurs by the movement of molecules into voids, with a size greater than v f In such a system the solute molecules move with the gas kinetic velocity u but m ost of t he time is confined to a cage delineated by their immediate neighbors. Occasionally, there is a fluctuation in density which opens up a hole within a cage large enough to permit a considerable displacement of the molecule contained by it. Such a displaceme nt leads to diffusive motion only if another molecule jumps into the hole before the first can return to its
142 original position. Therefore, by calculating the statistical redistribution of the free volume, Cohen and Turnbull derived the relation between the diffusion constant D in a liquid of hard spheres and the free volume as D = A exp( v / v f ), where v is the minimum required volume of the void; A is some constants that related to the average thermal velocity and the solute diameter, and is a numerical factor used to correct for overlap of free volume available to more than one molecu le [126, 127] Yasuda et al. incorporated the free volume theory to hydrated hydrogel systems and derived that the salt diffusion coefficient D in the hydrogel should be exponentially proportional to the reciprocal water fraction of the hydrated hydrogel. While this free volume theory model has successfully described the sodium chloride transport in a variety of homogeneous hydrogels [112, 113, 117 120] to our knowledge, currently the salt transport in the silicone hydrogels as extended wear contact lens m aterials has not been scrutinized with similar rationale. In this study, we use both permeation measurement and kinetic sorption/desorption approach to explore the sodium chloride transport in the silicone hydrogel contact lens materials. To confirm the m echanism of salt transport in the hydrogel matrix, in addition to direct measure ion permeability with ion flux at pseudo steady state, we propose to model the early transients in the permeation measurement data to determine both diffusivity and the partit ion coefficient of the lens as well. Additionally, we propose to utilize the kinetics sorption/desorption approach to extract the diffusivity and the partition coefficient by loading the lens with salt by soaking and then measuring the release rates of th e salt under perfect sink conditions. The salt release data is then utilized in a one dimensional diffusion controlled transport model [84, 106, 107, 128] to determine the partition coefficient and diffusivity of the lens.
143 Results are compared from both approaches (diffusion cell and release in perfect sink). Also, fundamental issues related to mechanisms of transport are addressed and the dependence of the partition coefficient and the diffusivity on composition and microstructure are explored. 6. 1 Mat erials and Methods 6. 1 .1 Materials Ethylene glycol dimethacryla te (EGDMA, 98%), N, N Dimethylac rylamide (DMA, 99%) and 1 vinyl 2 pyrrolidone (NVP, 99+ %) were purcha sed from Sigma Aldrich Chemicals (Milwaukee, WI ). Timolol maleate ( 98%), e thanol (99.5+ %) phosphate buffered saline (PBS) were obtained from Sigma Aldrich Chemicals (St. Louis, MO). Sodium chloride ( NaCl, 99.9+ %) were purchased from Fisher Chemical (Fairlawn, NJ). The macromer acryloxy(polyethyleneoxy) propylether terminated po ly(dimethylsiloxane) (DBE U12, 95+ %) were purchased from Gelest Inc. (Morrisville, PA). Methacrylic acid ( MAA 99.5%) were obtained from Polysciences, Inc. (Warrington, PA) and 3 Methacryloxypropyltris(trimethylsiloxy)silane (TRIS) was supplied by Silar l aboratories (Scotia, NY) 2, 4, 6 trimethylbenzoyl diphenyl phosphineoxide (Darocur TPO) were kindly gifted by Ciba Specialty Chemical s (Tarrytown, NY). All chemicals were used without further purification. 6. 1 .2 Preparation of Silicone Hydrogel TRIS was used as the hydrophobic monomer; DMA and MAA were the hydrophilic monomers; and DBE U12 was selected as the macromer. Highly hydrophilic NVP monomer was also added to increase water content of the hydrogel and EGDMA was introduced in the monomer mixture f or controlling the crosslinking. These six components were mixed in several different ratios listed in Table 6 1 to prepare the
144 polymerization mixture. Four different series of mixtures were designed in this study. In Series A the ratio of DMA to MAA is v aried while keeping the fraction of hydrophilic monomers (DMA+MAA) fixed; in series B the ratio of TRIS to Macromer is varied while keeping the fraction of silicone component (TRIS + Macromer) fixed; in series C the amount of crosslinker is varied while ke eping amount of all other components fixed, and finally in Series D the composition is randomly chosen. A s an ex ample, to prepare Gel A1, 2.4 mL of a mixture that comprised 0.8 mL TRIS and 0. 8 mL macromer and 0.8 mL of the hydrophilic DMA/MAA m ixture was c ombined with 0.12 mL of NVP and 0.1 mL of EGDMA. After vortexing for few second, the mixture was purged with bubbling nitrogen for 15 minutes to reduce the dissolved oxygen. To each monomer mix ture, 12 mg of photoinitiator Darocur TPO was added with stir ring for 5 minutes and the final mixture was immediately injected into a mold which is composed of two 5 mm thick glass plates. The plates were separated by a plastic spacer of a desired thickness, which is 0.13 mm, 0.26 mm or 0.40 mm in this study. The mo ld was then placed on Ultraviolet transilluminator UVB 10 (UltraLum Inc.) and the gel mixture was cured by irradiating with UVB light (305 nm) for 50 minutes. The synthesized hydrogel was either cut into circular pieces (about 1.65 cm diameter) with a cork bore r or other desired size and shape by scissors for subsequent experiments. Prior to conducting further tests, the prepared hydrogel was soaked in ethanol for 3 hours then dried at ambient temperature overnight to remove the unreacted monomer. 6.1 .3 Wa ter Fraction M easurements To determine the weight fraction of water (Q) in the hydrated gel, a dry gel of mass W d is soaked in DI water overnight or longer to ensure equilibrium. The hydrated lens is then weighted and the equilibrium water fraction in the lens is calculated as
145 ( 6 1 ) where W eq is the mass of hydrated gel at equilibrium. 6.1 .4 Ion Permeability Measurements 6.1 .4.1 Salt release in perfect sink (kinetic d esorption). The circular hydrogel with 1.65 cm diameter was soaked in 0.5 M, 0.75 M or 1.0 M NaCl solution until equilibrium was achieved. The salt loaded gel was then transferred into a DI water sink with a constant stirring at 300 rpm. The NaCl concent ration of the aqueous medium is determined by measuring the conductivity by Con 110 series sensor OAKTON followed by calculation from pre established calibration curve. In this study, the NaCl concentration is within the range that is linear to the measur ed conductivity, with a slope of 8.5810 6 M/ s. Fo r salt loading process, the volume ratio of the solution to the gel was maintained about 80. For example, a piece of 0.13 m m thick gel was soaked in 3.5 mL sodium chloride solution; for salt release experiment, t he volume ratio of the DI water and th e gel was kept about 600. For example, a piece of 0.13 mm thick gel was soaked in 27.5 mL DI water, and the amount of aqueous medium was proportional to the thickness of the gel to maintain the liquid/gel volume ratio. Because the volume of aqueous medium is much larger than that of the gel, the total amou nt of NaCl loaded i n the gel can be viewed as equal to the total salt release in DI water reservoir, and the partition coefficient of NaCl in the gel can be thus determined. 6.1 .4 .2 Ion transport in diffus ion c ell (direct p ermeation) The ethanol extracted sample gel was soaked in DI water or in NaCl solutions with various concentrations overnight at room temperature. The fully hydrated gel was subsequently mounted in a horizonta l diffusion cell, and then 18 mL of NaCl solution
146 and 30 mL of DI water were placed into the donor and receiver compartments with a constant stirring at 300 rpm, respectively. After the ion transport starts, the NaCl concentration in the receiving compartment is determined by measu ring the conductivity of the solution for 3 hours. The conductivity increases linearly in time after pseudo steady state is attained, and the slope of the increase in conductivity with time is a measure of the ion permeability. The value of the ion permeab ility can also be obtained by solving the diffusion equations for ion transport in the hydrogel, which will be discussed later. 6. 2 Results and Discussion Below we first compare the transport measurements from the two different methods and then discuss t he dependency of the transport parameters on composition. The two method s are compared only for Gel A1 4 and then the composition dependency of transport parameters measured through the kinetic approach is explored for all the gels prepared. 6. 2 .1 Compari son of Transport Measurements from the Kinetic and Permeation A pproach es 6. 2 .1.1 Kinetics of salt release in perfect sink In this approach the lens is soaked in a salt solution till equilibrium. The salt loaded lens is then soaked in DI water and the rel ease dynamics are measured. The salt partition coefficient K in silicone hydrogel can be determined by ( 6 2 ) where V w and V g are the volume s o f the aqueous phase and th e fully hydrated gel, respectively, and C g,f C w, i and C w,f are t he equilibrium concentration s of NaCl in the gel,
147 and the initial and equilibrium concentration s in the aqueous phase in the loading step, respectively. The dynamic ion release mechanism of silicone hydrogels can be viewed as one dimensional diffusion problem since the diameter of the hydrogel sample is much larger than its thickness. Therefore, the ion transport can be described by the diffusion equation, i.e., ( 6 3 ) The boundary conditions for the ion release experiment are ( 6 4 ) where h is the half thickness of the gel. The first boundary condition assumes symmetry at the center of the gel and the second describes equilibrium of salt concentration between the gel and the aqueous phase. A mass balance on the aqueous reservoir in the beaker yields ( 6 5 ) where V w is the water volume in the beaker, A gel is the cross sectional area of the gel, and C is the sodium chloride concentration in the gel. In addition, the initial conditions for the drug release experiments are ( 6 6 ) Since the aqueous reservoir volume in the beaker is much larger than the gel volume, it is reasonable to assume that the concentration C w is negligible in this perfect
148 sink condition to simplify our calculation. The above s et of equations can thus be solved analytically to yield ( 6 7 ) The error between the experiment data and model predicted value was defined as where N is the number of data points; Y ex and Y are the ratio o f measured NaCl release amount to that at equilibrium obtained by experiment and model calculation, respectively. In addition, a parameter was introduced in this model to compensate the offset error of time zero in salt release experiments, i.e. t in the fitting model is equal to t app where t app is the apparent experiment time. This imaginary time offset is a parameter that present s th e effects of diffusion boundary layer on the surface, interaction between polymer chains and charged ions, and the default errors from experimental setup on the initial reading of salt release. The above model was used to fit the experiment data to determi ne the D and of the silicone gel by using the and the results are summarized in Table 6 2 along with the values of the water content an d the salt partition coefficient. The comparison of the experiment and model predicted release pro file for different gels are plotted in Figure 6 1 It is clear that the model prediction in Figure 6 1 is well matched with the experiment release profiles, with a small time offset of time zero less tha n 40 seconds. The model we proposed above based on tw o assumptions: the partition coefficient and the diffusivity of hydrogel are independent of ionic concentration, and the dominating
149 mechanism of ion transport here is Fickian diffusion. These two assumptions are needed to be confirmed prior to further disc ussion. The effect of ionic strength on the ion transport of silicone hydrogel was studied by conducting salt release in perfect sink with gels that have different initial NaCl amounts. The gels were soaked in NaCl solution with different co ncentration un til equilibrium, and the following release profiles and fitted parameters of these gels are shown in Figure 6 2 and Table 6 3 Within the concentration range we explored here, the final release amount of a given silicone hydrogel is proportional to the ini tial soaking concentration, i.e. the ion partition coefficient keeps constant. The normalized release profiles for different loading are overlapping each other implied that the diffusivity is independent on ion concentration, as shown in Figure 6 2 B These results also su ggest that the electrostatic interactions of the double layer on the gel surface do not cause significant impact on ion transport within explored ion strength range. Another assumption in our propo sed mechanism is that the ion transport of silicone hydrogel here is diffusion controlled. From Equation 6 7 we know that if this process is diffusion controlled with homogeneous diffusivity, the release time will proportional to h 2 The NaCl release profiles from gels with different thickness were conducted in perfect sink and the dynamic release profiles w ere obt ained as the method for 0.13 mm thick samples. The volume of aqueous reservoir in both uptake and release experiments varied with the gel thickness so that the gel fluid volume ratios were maintained at the same as tho se of 0.13 mm thick samples, and the measured release profiles from gel with various thickness were plotted versus scaled time, which is defined as time/(gel thickness/0.1 mm) 2 and the results were shown in Figure 6 3 It is
150 obvious that within the margins of experimental error, the release profiles of the tested silicone hydrogels with different thickness are overlapped, proving that the dominating mechanism of ion transport of silicone hydrogel in perfect sink is the ion dif fusion through the gel phase. 6.2 .1. 2 Ion t ransport through permeation in a diffusion cell Ion permeability of silicone hydrogel was also measured by diffusion cell, and the results were shown in Figure 6 4 To obtain the ion permeability of the hydrogel the slope of the increase in conductivity with ti me is calculated after pseudo steady state is attained. The ionoflux diffusion coefficient, D ion is then determined by solving the diffusion flux equation for ion transport in the lens as : ( 6 8 ) where D ion is ionoflux diffusion coefficient (mm 2 /min) n is the rate of ion transport (mol/min), A is the area of ion transport (mm 2 ), dc is concentration difference between donor and receiving compartm ent (mol/mm 3 ), and dx is the thickness of lens (mm). The values of D ion for the gels were listed in Table 6 4 In additional to use the diffusion flux equation at pseudo steady state, we attempt to model the ion transpo rt of the hydrogel in the diffusion cell by the same diffusion equation as Equation 6 3, with the following bound ary conditions: ( 6 9 ) where h is the thickness of the gel, C is NaCl concentration in the donor compartment, and K is the salt partition coefficient of the gel. The first boundary condition is based on the assumption that the salt concentration in the rec eiver
151 compartment is much smaller than that in the donor compartment, which can be considered as perfect sink condition. The second boundary condition assumes equilibrium between the salt concentration in the gel and that in the donor solution. A mass bala nce on the receiving compartment yields ( 6 10 ) where V w is the DI water volume of the receiving compartment, A is the cross sectional area of the gel and C w is the NaCl concentration in the receiving medium. Finally the initial conditions for the drug release experiments are ( 6 11 ) where C i is the concentration of NaCl solution in which the gel was soaked overnight prior to conducting the diffusion cell experiment, and thus KC i is the initial salt concentration in the gel. The above set of equations can be solved analytically to yield ( 6 12 ) When the pseudo steady stat e of ion transport is attained, Equation 6 12 will reduced to Equation 6 8, where KD=D ion The comparison of the experiment and model predicted diffusion profile for different gels are plotted in Figure 6 4 and the error be tween the experiment data and model predicted value was defined as the same for previous fitting of salt release in perfect sink. was also introdu ced to compensate the offset error of time zero, and K of each gel was calculated from previous perfect sink
152 test results. D and of each gel were subsequently obtained by using the function Table 6 4 summarized the diffusivity fitting results for ion transport of silicone hydrogel in diffusion cell. For each sample, the ionoflux d iffusivity obtained by Equation 6 9 is consistent with KD the product of salt partition coefficient and the fitted diffusivity from ou r diffusion controlled model. By comparing the fitting results of the identical sample gel from salt release in perfect sink and from transport in diffusion cell, we can conclude that the fitted diffusivity matches in both approaches, which confirms the va lidity of both proposed model above. However, we noticed that a significant offset to time zero in the diffusion cell experiment for gels that have no pre loaded NaCl within, which will be discussed later. Similar to the model for salt release in perfect s ink, the effect of ion strength also need to be taken into concerned for our proposed diffusion limited transport mechanism in diffusion cell. A clean Gel A3 sample that contains no pre loaded salt in the sample was mounted to the diffusion cell and the io n tra nsport with various NaCl concentrations in the donor compartment was explored, as shown in Figure 6 5 Figure 6 5 b is the transport profile of normalized concentration versus time, and the normalized concentration is defined as the measured NaCl conce ntration in t he receiving compartment divided by the NaCl concentration in the donor compartment. The results from various donor concentrations overlapped within the margins of experiment error, suggesting that the ion transpo rt of silicone hydrogel in dif fusion cell is linear in the explored concentration range. The fitted parameters were listed in Table 6 5 and the fitted D ion D and KD from experiments of various donor concentrations are similar to
153 each other respectively, which satisfy the assumption in our model that the ion diffusivity of silicone hydrogel is independent of salt concentration. A significant large time offset for around 10 minutes was also observed here. The fact that is not apparently dependent on the donor concentration suggests that the large offset of time zero of the ion transport in diffusion cell might not be caused by the surface diffusion resistance boundary layer at the interface of silicone hydrogel and the NaCl solution in donor compartment, since the thickness of this boundary layer is significantly dependent on the ionic strength in the bulk solution. Silicone gels with various amounts of pre loaded NaCl were prepared by soaking in solution with different NaCl concentration overn ight prior to release experiment in diffusion cell, and the release results were shown in Figure 6 6 with the fitted constants listed in Table 6 6 The results show that the fitted D ion D and K has no significant dependency on th e amount of pre loaded NaCl in the gel, but only depends on the composition of silicone gel. However, the fitted for each gel significantly decreases when the gel contains some NaCl initially prior to the diffusion cell experiment. The fact that the larg e offset from time zero for the ion transport is only observed when the sample gel has no pre loaded NaCl implies that an adsorption/desorption barriers of ions on the gel matrix could be involved. For pure hydrated gel without pre loaded salt, the initial ion transport could be limited by this adsorption/desorption process, and subsequently controlled by the ion diffusion through the gel matrix after the ion adsorption/desorption p rocess on the gel matrix reached equilibrium.
154 The hydrogels with different thickness were also used in the ion transport experiment in diffusion cell to confirm the assumption in our proposed model. From Equation 6 12 we can observe that for long period of time, if the ion transport of silicone gel is diffusion controlled with ho mogeneous diffusivity, the release time will proportional to gel thickness. Thus we can define the scaled time for ion transport in diffusion cell here as apparent release time divided by the ratio of gel thickness to 0.1 mm, and the measured conductivity change in the receiving compartment for Gel A3 with various thickness were plotted versus scaled time in Figure 6 7 All the ion transport profiles of Gel A3 with different thickness overlap within the marg ins of experiment error, as we observed in the sal t release experiments in perfect sink. The results again confirm that the rate limiting step of ion transport of silicone hydrogel in diffusion cell is the ion diffusion through the gel matrix, which is the basic assumption in our proposed model. 6.2 .2 Ef fect of Composition of Silicone Hydrogel on Ion Permeability We used the developed methods above to further investigate the effect of silicone hydrogel composition on ion permeability. All the experiments results described below were conducted by using ki netic sorption/desorption method at perfect sink. The thickness of sample hydrogel is 0.13 mm and the soaking NaCl concentration is 0.75 M if not fur ther specified. Figure 6 8 shows the ion release experiment results for hydrogels with different TRIS and Macromer ratios. Because the amount of hydrophilic monomer (DMA and NVP) is fixed, it is expected that the water content and NaCl partition coefficient have only a slight dependency on TRIS/macromer composition. For example, when we increase the TRIS/(TRIS +macromer) amount from 25% to 75% it will only leads to less than 30% decrease on water content and NaCl partition coefficient. The decrease is
155 caused by the higher hydrophobility of silicone than the macromer after polymerization. However, the fitted D exhibited a more complicated dependency on TRIS/(TRIS+macromer) composition, as shown in Figure 6 8 b. For the gel with TRIS/(TRIS+macromer) composition less than 50%, D decrease in a slower rate than that observed for gel with TRIS/(TRIS+macromer) ratio hi gher than 50%. The difference might be caused by the structure difference of these gels. For the gels with higher macromer content, it is easier to form bicontinuous polymer structure, which could provide more diffusion channels for ions to diffuse through the gel matrix. On the other hand, without enough macromer in the mixture the synthesized hydrogel could contains a lot of phase separated hydrophilic or hydrophobic regions, and therefore retards the ion transport. The effect of the crosslinker, EGDMA, o n ion transport properties of the synthesized silicone hydrogel was also studied. Figure 6 9 indicates that both water content and NaCl partition coefficient decreases as the amount of EGDMA increased. This is expected since when the crosslinking inside th e gel matrix increase, the mobility of the synthesized polymer chain decreases, and thus leading to lower water content and salt solubility. However, with about 2% of EGDMA in the monomer mixture will help the synthesized silicone hydrogel to obtain the op timal value of D as shown in Figure 6 9 b. With proper amount of EGDMA in the mixture as crosslinker, it is possible for the monomers to be polymerized in structures that enhance the ion permeability; when the amount of EGDMA over the critical value, the e xtra EGDMA could form higher crosslinked region inside the gel matrix that retards the ion transport.
156 Finally, the effect of hydrophilic DMA composition in the hydrogel on ion permeability were studied by using gel with various DMA, TRIS and macromer comp osition (Gel D1 D6), while the rest ingredient amount in the gel mixture (NVP, EGDMA and initiator) were kept the same, and the results are shown in Figure 6 10 If we neglect the effect of non hydrophilic monomers in the composition and focused on the con tent of DMA of the synthesized silicone gel, we can clearly observed the linear dependency of water content to the DMA composition in the gel, as shown in Figure 6 10 d. The NaCl partition coefficient shows the similar linear dependency to DMA component; on the other hand, even the fitted D increased as the DMA amount increase, the slight non linearity suggests other factors were involved as well. As a result, the KD of these silicone gels shows a much stronger dependency on the DMA amount. The summary of t he composition effect on ion transport properties, including salt partition coefficient, salt diffusivity and salt permeability, can be discussed based on the water fraction of the synthesized silicone hydrogels. It is noted that even though the water frac tion is better to be measured in salt solution rather than DI water for these comparisons, the variation within the salt concentration range explored here is not significant, as observed in both our studies and previous studies on conventional hydrogels sy stems [112, 113] Ideally, if the salt only partition in the water phase inside the gel matrix, the salt partition coefficient should be equal to the water fraction inside the hydrogel, as the case in pure water. However, as shown in Figure 6 11 even thou gh the salt partition coefficient is strongly correlated to the water content of these silicone hydrogels, the correlation is under the line where partition coefficient and water fraction are equal. Similar results were observed in different hydrogel syste m as well
157 [112, 113, 119, 120, 124] This deviation is considered to be resulted from the interaction between the polymer chain and the solute. Even though this interaction is generally much smaller compare to that between the water molecule and sodium chl oride, this effect becomes more dominant as the water content of the hydrated hydrogel decreases. When the water fraction increases, the partition of sodium chloride in the water phase becomes more significant, and thus the salt partition coefficient gradu ally approaches to the ideal assumption as pure water. In the free volume theory model of hydrated polymer membranes proposed by Yasuda et al., the volume fraction of diluent is expressed by the hydration H. In the consideration of permeation of salt, it s eems to be quite obvious that salt alone will not permeate through most pure polymer matrix. Consequently, it is reasonable to assume that salt can permeate only through the hydrogel system in which the diluent of the polymer is also a good solvent of the salt considered. The free volume of a hydrated hydrogel system can be expressed by ( 6 13 ) Since salt alone will not permeate through the polymer matrix, the free volume in the poly mer chains is negligible, and thus the free volume available for salt permeation in the hydrogel system is approximately equal to Hv f,H2O Thus, combined with the Cohen Turnbull relation, the diffusion constant D of salt through the hydrated membrane can b e written as a function of H as, ( 6 14 ) w here D 0 is the diffusion constant of sodium chloride in pure water at the temperature of the experiments, and K is a proportion ality constant related to the characteristic volume
158 v a nd v f,H2O Then, the plot of log D versus 1/H is expected to be linear starting from the diffusion constant of NaCl in pure water (1/H=1.0), which is 1.510 5 cm 2 /s at 25 o C  Figure 6 12 plotted the ion diffusivi ty ( D ) on logarithmic scale versus the reciprocal of weight water fraction (1/ Q ) of these silicone hydrogels. It is noted that here assume that the difference between the polymer density of these silicone hydrogel matrix is negligible, and thus the weight volume fraction can be viewed as the volume fraction. Since the water content of commercial silicone hydrogel contact lenses is between 20 to 60% [106, 130] here we only applied linear regression analysis on the hydrogel of which the wate r fraction is greater than 15%, and the results showed that with proper water content, the diffusion of sodium chloride through these silicone hydrogels can be predicted by the free volume theory. However, when the water fraction decreases, the free volume theory failed to relate the ion diffusivity to the water fraction. This is reasonable since the water fraction is low, it is highly possible that the hydrophilic phase of the hydrogel cannot form a continuous phase throughout the matrix, but dispersed ins ide the continuous silicone phase. The ion permeability is the product of salt partition coefficient and salt diffusion coefficient. Even though the salt partition coefficient is dependent on the water fraction, as we discussed earlier, this dependence is much smaller compared to the exponential relationship between the ion diffusivity and water fraction. Thus, in most case the relationship between ion permeability and water fraction can be also described by the free volume theory model. As shown in Figure 6 13 the free volume theory model linearly fitted the data from these silicone hydrogels with sufficient water fraction. When water fraction becomes lower, in addition to the tendency to form phase separated
159 regions inside the hydrogel matrix, the intera ction between polymer matrix and salt also becomes more significant. Both conditions violate the assumptions of the free volume theory, which requires the hydrogel to be homogeneous and the interaction between polymer and solute should be negligible. Some previous data of other silicone hydrogel contact lens materials were also compared to our study. It is noted that for the commercial continuous wear contact lenses NIGHT&DAY TM (Lotrafilcon A) and O 2 OPTIX TM (Lotrafilcon B), the relationship between its ion permeability and water content can be predicted by our regression results. Since both of these contact lens were made by TRIS and DMA as the main components  it is reasonable to consider our silicone hydrogels should have similar microstructures as t hese commercial contact lenses. Therefore, these silicon e hydrogels can be used to as a substitute to further investigate other import transport properties of extended wear contact lens, such as oxygen permeability and drug diffusivity. In brief summary, c ritical properties of silicone hydrogels can be manipulated by changing composition. Improvements in one property frequently occur at expense of other properties, and so a careful optimization is needed to design contact lenses for extended drug release. I n this study a simple ion permeability measurement for silicone hydrogel is established. While water content and ion partition coefficient of the hydrogel generally depend on the total composition only, the ion diffusivity will also significantly depend on the structures of different phases of the hydrogel. Thus, water content and salt partition coefficient can be predicted by the bulk composition, but the diffusivity and thus ion permeability are much more complicated, since it depends on the structure. Wh en the silicone hydrogel has proper composition to form bi continuous
160 microstructures in the matrix, the ion transport through the gel can be described by the free volume theory model. These results provided a solid foundation to further explore other solu te transport properties of silicone hydrogel contact lens, especially some ionized hydrophilic ophthalmic drugs. In the future studies, other properties such as elastic modulus and oxygen permeability should also be included to evaluate the design of new s ilicone hydrogel as extended wear extended release contact lens.
161 Figure 6 1 NaCl release profile and model prediction (solid line) of different silicone hydrogel in perfect sink. All samples are 0.1 3mm thick and pre soaked in 0.75 M NaCl (aq) prior to conducting release experiment. Data are presented as
162 Figure 6 2 NaCl release profile of 0.13 mm thick Gel A3 with different pre soaking NaCl concentration in perfect sink. The solid lines are the model prediction
163 Figure 6 3 NaCl release profile of Gel A3 with different thickness in perfect sink. Th e pre soaked concentration of NaCl (aq) is 0.75M. The solid lines are the model
164 Figure 6 4 Ion permeability test for silic one hydrogels by diffusion cell. All samples are 0.13 mm thick with no preloaded salt. The NaCl concentration in the donor compartment is 0.5 M. The solid lines are the model prediction and all
165 Figure 6 5 Ion permeability test of Gel A3 by diffusio n cell with various NaCl concentration in the donor compartment. All samples are 0.13 mm thick with no preloaded salt. The solid lines are the mode l prediction and all experiment
166 A B Figure 6 6 Ion permeability test by diffusion cell of silicone hydrogel that were pre soaked in sodium chl oride solution with different concentration. The samples and the sodium chloride concentratio n in the donor compartment are A ) Gel A 1 0.5M, B ) Gel A2 0.5M, C ) Gel A3 0.5M and D ) Gel A3 0.25 M, respectively. All the samples are 0.13 mm thick. The solid lines are the model prediction and all experiment data are presented
167 C D Figure 6 6. Continued.
168 A B Figure 6 7 Ion permeability test of Gel A3 with different thickness by diffusion cell. The sodium chloride concent ratio n in the donor compartment is A ) 0.25 M, with no preload salt in the gel ; B ) 0.5 M, with no preloaded salt in the gel and C ) 0.25 M, the gel was pre soaked in 0.25 M NaCl (aq) respectively. All experiment
169 C Figure 6 7. Continued.
170 A B Figure 6 8 A ) NaCl partition coefficient (K) B) diffusivity (D) C) KD and D ) Water content ( Q ) for silicone hydrogel with different TRIS/Macromer composi tions
171 C D Figure 6 8. Continued.
172 A B Figure 6 9 A ) NaCl partition coefficient (K) B) diffusivity(D) C) KD and D ) Water content ( Q ) for silicone hydrogel with different EGDMA compositions
173 C D Figure 6 9. Continued.
174 A B Figure 6 10 A ) NaCl partition coefficient (K) B ) diffusivity (D) C ) KD and D ) Water content (Q ) for silicone hydrogel with diffe rent DMA compositions
175 C D Figure 6 10. Continued.
176 Figure 6 11 The relationship between salt partition coefficient (K) and water fraction (Q) of silicone hydrogel. The dash line indicates K =Q, and the solid line is draw for visual guidance.
177 Figure 6 12 The relationship between salt diffusivity (D) and the reciprocal of water fraction (1/Q) of silicone hydrogels. The solid line is the linear best fit for data with Q greater than 0.15.
178 Figure 6 13 The relationship between salt permeability (KD) and reciprocal water fraction (1/Q) of silicone hydrogels. The solid line is the bes t linear fit for data with Q greater than 0.15. Data were compared with selected literature data for commercial silicone hydrogel contact lenses  and other silicone hydrogels as contact lens material 
179 Table 6 1 Composition of silicone hydrogel Each monomer mixture was mixed with additional 0.12 mL of NVP before polymerization. Sample TRIS ( mL ) Macromer ( mL ) DMA ( mL ) MAA ( mL ) EGDMA ( L ) A1 0.80 0.80 0.80 100 A2 0.80 0.80 0.68 0.12 100 A3 0.80 0.80 0.56 0.24 100 A4 0.80 0.80 0.40 0.40 100 B1 1.20 0.40 0.80 100 B2 1.00 0.60 0.80 100 B3 0.60 1.00 0.80 100 B4 0.40 1.20 0.80 100 C1 0.80 0.80 0.80 5 C2 0.80 0.80 0.80 10 C3 0.80 0.80 0.80 20 C4 0.80 0.80 0.80 50 C5 0.80 0.80 0.80 75 C6 0.80 0.80 0.80 300 C7 0.80 0.80 0.80 500 D1 1.37 0.34 0.69 100 D2 1.37 0.69 0.34 100 D3 0.6 0 0.6 0 1.2 0 100 D4 0.48 0.48 1.44 100 D5 0.96 0.48 0.96 100 D6 0.4 0 0.4 0 1.6 0 100
180 Table 6 2 Parameters of different silicone hydrogels. Sample Equilibrium Water Content, Q (%) Partition coefficient, K Diffusivity, D (10 6 mm 2 /min) (min) A1 22.5 0.1024 1957.7 0.160 A2 19.7 0.0834 1012.1 0.00 7 A3 16.6 0.0571 533.7 0.28 3 A4 11.1 0.0338 36.1 0.018
181 Table 6 3 Parameters for GEL A3 with various sodium chloride concentrations for sal t loading. Soaking NaCl concentration (M) Partition coefficient, K Diffusivity, D ( 10 6 mm 2 /min) (min) 1.0 0.0585 563.2 0.750 0.75 0.0531 534.7 0.283 0.5 0.0570 583.0 0.698
182 Table 6 4 Fitting results of ion transport by diffusion cell for silicone hydrogels. All samples are 0.13 mm thick and contain no preloaded salt. Sample Ionoflux Diffusivity, D ion ( 10 6 mm 2 /min) Diffusivity, D ( 10 6 mm 2 /min) KD ( 10 6 mm 2 /min) (min) A1 193.5 1673.0 171 .3 9.246 A2 112.3 1163.2 97.1 7.251 A3 39.6 596.6 34.3 13.786 A4 1.5 43.8 1.5 9.750
183 Table 6 5 Fitting results of ion transport by diffusion cell for Gel A3 with various sodium chloride concentrations in the donor compartment. All samples are 0.13 mm thick and contain no preloaded salt. Donor NaCl concentration (M) Ionoflux Diffusivity, D ion ( 10 6 mm 2 /min) Diffusivity, D ( 10 6 mm 2 /min) KD ( 10 6 mm 2 /min) (min) 0.1 38.3 588.9 33.9 8.827 0.25 39.7 631.8 36.3 10.757 0.5 39.6 582.9 33.5 13.787 0.75 42.2 596.6 34.3 10.580
184 Table 6 6 Fitting results of ion transport by diffusion cell for silicone hydr ogels which pre soaked in sodium chloride solution with various concentrations. All samples are 0.13 mm. Sample Donor NaCl conc. (M) Soaking NaCl conc. (M) Ionoflux Diffusivity, D ion ( 10 6 mm 2 /min) Diffusivity, D ( 10 6 mm 2 /min) KD ( 10 6 mm 2 /min) (mi n) A1 0.5 0 193.5 1673.0 171.3 9.246 0.5 0.5 199.0 1997.0 204.5 3.848 A2 0.5 0 112.3 1163.2 97.0 7.251 0.5 0.5 111.3 1290.5 107.6 3.703 A3 0.5 0 39.6 582.9 33.5 13.786 0.5 0.125 38.5 550.4 31.6 2.701 0.5 0.5 42.9 633.8 36.4 3.897 0.25 0 39.7 6 31.8 36.3 10.756 0.25 0.125 38.7 600.3 34.5 0.952 0.25 0.5 38.0 719.4 41.4 3.184
185 CHAPTER 7 CYCLOSPORINE DELIEVE RY BY SILICONE HYDRO GEL FOR CRONIC DRY E YE SYMDROME Cyclosporine A (CyA, also known as cyclosporin or cyclosporine) is an immunosuppressant dru g used for treatment of many ocular diseases including keratoconjunctivites sicca (dry eye syndrome)  uveitis in children and adolescent [132, 133] vernal keratoconjunctivitis  and peripheral ulcerative keratitis  It also can be used to p revent allograft rejection  CyA likely mitigates dry eyes symptoms by reduci ng inflammation in eyes and tear producing glands, potentially 25C)  CyA is delivered through eye drops of an oil in water emulsion containing 0.05% cyclospori ne  While emulsion based eye drops is FDA approved, this delivery approach suffers from low bioavailability with more than 95% of drug reaching systemic circulation through transnasal or conjunctival absorption. Also, CyA delivery through eye drops i s particularly complicated for moderate dry eye patients who are also contact lens wearers. In 2007, there were about 35 million contact lens wearers in North America and about half of those reported some symptoms of dryness and discomfort, more commonly e xperienced at the end of the day  CyA emulsion can also be used to alleviate these dryness and discomfort symptoms, but the patients need to remove the lens prior to applying the CyA eye drop and then reinsert the lens 15 minutes afterwards  To address the disadvantages of CyA delivery through eye drops, researchers have been focused on the development of sustained release devices. For example, biodegradable silicone based ocular implants have been developed to continuously
186 release CyA for sever al years [139, 140] Preclinical evaluation and animal studies suggest significant potential of these implants for life long treatment of ocular graft versus host disease (GVHD) such as severe keratoconjunctivites sicca  The device can also be used f or long term treatment of uveitis  and can prevent high risk penetration keratoplasty (PKP) rejection  These devices have higher bioavailability for deep ocular tissues compared to topical delivery [144, 145] Insertion and replacement of these CyA implants require surgery and thus these appear to be more suitable for severe ocular diseases that need long term treatment. For the treatment of moderate dry eye syndrome, which is one of the most common ocular ailment that affect more than 30 millio n people in United States [146, 147] other CyA delivery devices could be useful. For example, Gupta et.al proposed a novel combination therapy based on extended CyA delivery lasting about 3 months from punctual plug made of pHEMA  In this Chapter w e aim to develop contact lenses for providing sustained delivery of CyA To our knowledge, this study is the first one on CyA drug delivery through contact lenses. The specific goal of this study is to develop silicone hydrogel contact lenses that ca n del ivery CyA for the entire duration of wear time, which is about 1 month for several lenses. The approach explored here focuses on first measuring the CyA release profiles from commercial Silicone hydrogel contact lenses and then incorporating Vitamin E into the lenses to exte nd the release time to about 1 month. Such lenses will particularly benefit dry eye sufferers who also wear contact lenses as they will not need to remove lenses to deliver CyA eye drops. CyA has also shown promising results in treating contact lens mediated dry eyes and thus the CyA releasing
187 contact lenses could also potentially mitigate or minimize the discomfort and the dryness symptoms that arise due to contact lens use  7.1 Materials and Methods 7.1 .1 Materials Five commercial contact lenses (diopter 6.50) were used in this study, including 1 DAY ACUVUE and ACUVUE 2 (Rochester, NY). Cyclosporine A (CyA) was phosphate buffered saline (PBS) was purchased from Sigma Aldrich Chemicals (St. Louis, MO), and Vitamin E (D alpha tocopherol, Covitol F1370) were kindly gifted by Cognis corporation (Kankakee, IL). All chemicals were reagent grade and used as supplie d without further purification. 7.1 .2 Drug Loading into Contact L ens The commercial contact len ses were rinsed with DI water, dried in air and then weighed. The stock CyA PBS solution was prepared by dissolving 2.5 mg CyA in 100 mL PBS wi th moderate stirring at about 5 o C for 24 hours. The stock solution was later diluted with PBS to the desired drug concentration. The lenses were loa ded with CyA by soaking in 10 mL of a 15 g/mL CyA PBS solution for 7 days. At the end of the loading stage, the lenses were withdrawn from the solution, and excess drug solution was blotted from the surface of the lens. The dynamic changes of drug concentration of the soaking solution were not monitored during the drug loading process. The total amount of CyA loaded into a lens was calculated at the end of the loading phase by determining the total amount of drug loss from the aqueous solution. The drug concentrat ion in
188 aqueous solution was determined by measuring the absorbance in the wavelength range from 208 to 220 nm with a UV VIS spectrophotometer (Thermospectronic Genesys 10 UV) and then using a pre established calibration curve. The abso rbance was measured in the 208 to 220 nm range rather than at a single wavelength to ensure that the experimental methods did not lead to drug degradation which will manifest as changes in the absorption spectrum. As shown in the latter sections, the 7 day duration was insuf ficient for equilibration, and thus the duration of the uptake phase was increased to achieve equilibrium. The equilibrium studies were conducted only for ACUVUE OASYS TM lenses (dry weight = 22.3 0.3 mg ). A lens was soaked into 10 mL of 17 g/mL CyA/PBS solutions for a sufficiently long period of time to reach equilibrium. The mass of drug loaded into the lens was determined by measuring the reduction in the drug concentration of the soaking solution. 7.1 .3 Drug Release Experiments from Len ses L oaded with CyA The drug release experiments were carried out by soaking a CyA loaded l ens in 1.75 mL of PBS. During the release experiments, the dynamic drug concentration in the PBS was analyzed with the method described above. The release medi um was replaced by fresh 1.75 mL PBS after each measurement to maintain perfect sink conditions. 7.1 .4 Vitamin E Loading into Contact L ens Vitamin E was loaded into lenses by soaking a lens in 3 mL of a Vitamin E ethanol solution for 24 hours. The concentration of V itamin E in etha nol varies from 0.05 to 0.4 g/mL in the loading solution. After the loading step, the lens was withdrawn and blotted to remove excess solution on the surface. The lens was then dried in air overnight. The
189 Vitamin E loading amount was d etermined by measuring the increase in lens weight. CyA was loaded into the Vitamin E containing contact len ses by soaking the lens in 10 mL of a 17 g/mL C yA PBS solution The loading duration was increased with increasing Vitamin E loading to provide sufficient time for equilibration. Subsequently, drug release profiles were measured from the lenses that containing CyA and Vitamin E by following the procedures described earlier. 7.2 Results 7.2 .1 Drug Uptake by Pure Contact L ens The data from the loading experiments was utilized to calculate the fractional amount of drug absorbed ( f ) by the lenses during the 7 day soaking by using the following relati onship: ( 7 1 ) where C w, i and C w,f are the initial and final concentrations in the aq ueous phase, respectively The calculated f by the lenses during the 7 day loadin g is listed in Table 7 1 For 1 DAY ACUVUE about 94% of CyA in the initial drug solution remained in PBS medium after soaking for 7 days, while Silicone hydrogel lenses abso rbed a majo rity of CyA (51.6% to 75.6%) from the PBS solution into the gel matrix. The partition coefficient ( K ) of CyA contact in lenses is an important parameter as it determines the loading capacity. The value of K can also be calculated from the data collected in the drug loading experiments through the following relationship: ( 7 2 ) where V w and V pl are the volumes of the drug PBS solutio n and the dry pure lens
190 respectively, and C pl,,f is the final concentrations of the drug in the lens phase It is noted that the above relation defines the partition coefficient only if the concentration in the gel and the fluid is in equilibrium at the end of the loading experiment. To confirm whether the 7 day duration of the drug loading experiments is sufficiently long to reach equilibrium, drug uptake experiments were conducted with ACUVUE OASYS TM for longer times. In these the dynamic drug concen tration in the loading solution was also measured. Figure 7 1 indicates that CyA concentrations between PBS medium and ACUVUE OASYS TM reaches equilibrium after about 15 days, and the final equilibrium partition coefficient is 677.5 48.9. The shape of a bsorbance spectrum for CyA at various times remains unchanged (data not shown), suggesting that the drug is stable during the duration of the experiment. Since the equilibration time for CyA loading is larger than 7 days, the data from the 7 day drug loadi ng experiments cannot be strictly used to determine the partition coefficients. However the data can still be used to obtain approximate values of the partition coefficients. By comparing the K value for ACUVUE OASYS TM it is evident that the partition coefficients reported in Table 7 1 which are based on 7 day soaking are slightly lower than the true values. The estim ates for K for other Silicone hydrogel lenses are still useful as these allow comparison of the drug uptake potential for various lenses. The calc ulated K values for different commercial contact lenses after 7 days loading process are listed in Table 7 1 Th e drug partition coefficients are large for each s ilicone hydrogel lens and are at least one order of magnitude larger than that for th e 1 DAY ACUVUE lens.
191 7.2 .2 Drug Release by Pure Contact L ens Figure 7 2 shows the results of CyA release by 1 DAY ACUVUE under perfect sink condition. The release duration from the 1 DAY ACUVUE is around 24 hours, which is sufficient for this lens beca use it is intended for 1 day use only. Also, the 1 day release duration implies that 7 days is sufficiently long for establishing equilibrium in the loading phase, and thus the partition coefficient value for this lens reported in Table 7 1 is accurate. T he cumulative mass of drug released under perfect sink conditions is plotted as a function of time for various s ilicone hydrogel contact lenses in Figure 7 3 A, and the % Release profiles are plotted in Figure 7 3 B. The data clearly shows that each of the four commercial s ilicone hydrogel contact lenses release CyA for extended period lasting more than 7 days, which is significantly longer than the release duration by p HEMA hydrogel lenses (1 DAY ACU VUE ). The data shows that after 7 days, ACUVUE OASYS TM lens releases about 82% of the loaded CyA, and the other three types of lenses release about 50% of the loaded drug. In first 3 days, the ACUVUE OASYS TM lens releases about 15 g of drug each day and the other three types of lenses release about 10 g Cy A each day. 7.2 .3 Drug Uptake by Vitamin E Loaded Contact L ens The results of drug uptake by ACUVUE OASYS TM lenses with various Vitamin E loading are listed in Table 7 2 It is clearly seen that Vitamin E incorporation into the lens significantly increas es the mass of CyA absorbed. For example, after 120 days, 95% of the drug in the initial soaking solution was absorbed into the contact lens with about 40% Vitamin E loading compared to 60% for the lens without Vitamin E. The
192 shape of the CyA spectra agai n is unchanged with ti me suggesting that the drug is stable during the experimental duration of 4 months (data not shown). In addition, the concentration of Vitamin E in the release medium was undetectable. Also, there was negligible weight loss from the Vitamin E loaded contact lens during the release experiment. Both of these observations suggest that the Vitamin E is not released into the solution due to its very small solubility in aqueous solutions. By using the previously determined equilibrium CyA partition coefficient of pure ACUVUE OASYS TM and using a mass balance  we estimated the CyA partition coefficient of the Vitamin E phase in the lens ( K ve ) ( Table 7 2 ). The average value of K ve ( 1.27 0.1 8 10 4 ) is relatively independent of the amou nt of Vitamin E loaded into the lenses validating the accuracy of the measured value, and also suggesting that the duration of the loading phase for the Vitamin E containing lenses was sufficient to achieve equilibrium. The average K ve for CyA is about 20 fold higher than the pa rtition coefficient of pure ACUVUE OASYS TM which implies that the hydrophobic CyA has much higher affinity to the Vitamin E phases than to the silicone gel matrix. 7.2 .4 Drug Release by Vitamin E Loaded Contact L ens Figure 7 4 sh ows the results of CyA release from Vitamin E loaded ACUVUE OASYS TM lenses under perfect sink condition. As shown in Figure 7 4 A, the initial drug release rate decreases as the Vitamin E loading amount in the contact lenses increases, even though the tota l drug uptake amount increases as we discussed earlier. For example, after the first day of release, ACUVUE OASYS TM with 0% and 20% of Vitamin E loading released at an average rate of 10.2 and 6.1 g/day, r espectively. It is also noted that the drug relea se du ration by these contact lenses are significantly
193 extended with Vitamin E loading as shown in Figure 7 4 B. For example, for pure ACUVUE OASYS TM lens, 60% of the inclusive CyA in the lens was released in 7 days, while it took 16 and 46 days for ACUVUE OASYS TM with 10% and 20% Vitamin E loading to release the same percentage of loaded drugs, respectively. Moreover, Vitamin E loading reduces the variations in the drug release rates with time. As shown in Figure 7 5 the daily CyA release rate from AUCUV E OASYS TM lenses decrease rapidly with time from 10.2 g/day in day 1 to 1.3 g/day in day 15, while AUCUVE OASYS TM with 20% of Vitamin E releases CyA 6.1 g in day 1 and 1.8 g in day 15. The solid lines in the figure are fits to the diffusion model w hich will be described later. 7.3 Discussion CyA is a highly hydrophobic drug with very limited solubility in PBS, so its partition c oefficient is high in polymeric contact lenses. The partition coefficients are higher in s ilicone hydrogel contact lenses compared to the hydrophilic p HEMA lens (1 DAY ACUVUE ) because CyA has significantly higher affinity for the hydrophobic silicone rich phases compared to that for the hydrophilic p HEMA phase. The release time of CyA from p HEMA based contact lenses is m uch shorter than that from the s ilicone hydrogel lenses due to the smaller partition coefficient. The ratio of the release times is roughly equal to the ratio of the partition coefficients. The release duration from the 1 DAY ACUVUE is about 1 day, whic h is consistent with the studies of Kapoor et al., who showed that 100 m thick pHEMA hydrogel releases CyA for about 1 day  The release duration of CyA from extended w ear s ilicone hydrogel co ntact lenses is about 15 days. The CyA release duration from contact lenses is considerably longer than the release duration for other ophthalmic drugs such as timolol, dexamethasone,
194 dexamethasone phosphate, and fluconazole as we observed in Chapter 3 and 4 The reason for the long release time for CyA is its higher molecular weight and very high partit ion coefficient in the Sili cone hydrogel contact lens The larger size and the higher binding of the drug to the polymer in the lens reduce the effective diffusivity leading to long release times. While 15 day release is longer than that for other drugs, it is not adequate because extended wear lenses are prescribed for about 1 month use. Vitamin E incorporation in silicone contact lens leads to an increase in release duration of a number of ophthalmic drugs, without a significant impact on any critical lens property. Vitamin E inc orporation also increases the release duration of CyA from ACUVUE OASYS TM For example, t he release duration increases to more than a month on 10% Vitamin E incorporation into the lens. Incorporation of Vitamin E also increases the effective partition co efficient of CyA into t he lens which will have the additional benefit of reducing the drug loss from the lens into the packaging solution. The impact of Vitamin E incorporation on other s ilicone hydrogel lenses is expected to be similar to that on ACUVUE OASYS TM 7.3 .1 Release Mechanism and Model F itting To understand the mechanism of CyA release from the contact lenses, it is instructive to compare the release profiles with a one dimensional diffusion controlled model which yields the following equation for % Release at short times [106, 108] ( 7 3 ) where M t is the accumulated mass of drug released at time t M the accumulated mass of drug release as time approaches infinity and for perfect sink condition M = M 0 (initial drug loading). The above equation predicts that the plot of % Release with
195 square root of time should be linear at short times. Figure 7 6 plots % drug release by s ilicone hydrogel contact lens as f unction of square root of time for CyA. T he best fit straight lines are also shown in the figure. The straight lines fit the data well with R 2 values larger than 0.9 6, showing that the drug transport in these lenses is diffusion controlled. The slope is clearly larger for ACUVUE OASYS TM showing the drug diffusivity is highest for these lenses amongst those explored here. It is again noted that Equation 7 3 is not strictly valid because the system did not reach equilibrium during loading however the short time data should still satisfy Equation 7 3 as the concentration in the region of the lens close to the surface was near equilibriu m concentration. The CyA % release from Vitamin E loaded ACUVUE OASYS TM versus square root of time and the best fit straig ht l ines for short time release are also plotted in Figure 7 7 All the R 2 v alues of the fitting results for these ACUVUE OASYS TM lenses with various Vitamin E loaded lens are about 0.99, suggesting that the CyA release from the Vitamin E loaded Silicone hydrogel lenses are controlled by diffusion as well. The increased release times are due to the partitioning of the drug and slow diffusion through the Vitamin E aggregates in the lens. This effect is similar to the proposed mechanism of extended release o f hydrophobic dexamethasone by the same systems in Chapter 4 The comparison of CyA and dexamethasone delivery duration increase by Vitamin E loaded ACUVUE OASYS TM are plotted in Figure 7 8 In this figure, the ratio of the release times after inclusion of Vitamin E ( 0 ) and the release time from lenses without Vitamin E ( ) are plotted as a function of the Vitamin E loading in the lenses. While the trends are similar, Vitamin E loading has a slightly larger impact on
196 dexamethasone transport compared to CyA. This could perhap s indicate that the diffusivity of CyA through the Vitamin E regions is attenuated to a lesser extend compared to dexamethasone. It is also noted that the profile of % drug release vs. square root of time is slightly curved at short times and then becomes linear. The flux of drug transport from the gel to the surrounding fluid is determined by two resistances in series; resistance in the gel and then that in the fluid. The resistance in fluid is typically very small and can frequent ly be neglected. However in instances of low gel diffusivity, particularly at short times when the boundary layer thickness in the gel is very small, the fluid resistance can dominate. For such cases, the drug flux is independent of time because the flui d boundary layer thickness is unchanged in time [148, 150] A constant drug flux results in a linearly increasing cumulative release with time, which implies that the % Release vari es as the squ Figure 7 6 and Figure 7 7 Since the dynamic CyA release mechanism of contact lens can be viewed as one dimensional diffusion, it can be described by the diffusion equation i.e., ( 7 4 ) For simplicity, we assume the diffusivity of CyA in the lens is independent of drug concentration, which is usually val id in low drug concentration range. The diffusivity is also assumed to be independent of position, which may be not precisely correct because the distribution of Vitamin E in the lens may not be uniform. Since the exact distribution of Vitamin E in the le ns is not measured, we adopt the simplifying assumption of uniform and fixed diffusivity.
197 The boundary conditions for the CyA release experiment are ( 7 5 ) where h is the half thickness of the lens, which can be assumed a s 40 m. The thickness of the lens is position dependent due to the lens curvature, but this variation is also neglected for simplification of the model, and the lens is treated as a flat film with thickness equal to the average thickness of the curved lens. T he first boundary condition assumes symmetry at the center of the gel and the second describes equilibrium of drug concentration between the gel and the aqueous phase. Since the aqueous reservoir in the vial was replaced with fresh PBS during the release, it is reasonable to assume that the release occurs under perfect sink conditions, i.e., C w can be assumed to be negligible In addition, the initial conditions for the drug release experiments are ( 7 6 ) The above set of equation s can thus be solved analytically to yield ( 7 7 ) Therefore, at the surface of the lens, the flux of drug is: ( 7 8 ) We can thus relate the accumulated drug release rate measured by experiments to these equations by a mass balance in the vial:
198 ( 7 9 ) where A lens is the cross sectional area of the contact lens The mass of drug release in period of time from t 1 to t 2 can be determined by computing M t (t 2 ) M t (t 1 ). The data for daily amount released as a functio n of time is plotted in Figure 7 5 and the fitted data is plotted as the solid lines. The best fit values of diffusivity are 3.34 and 0.92 10 6 mm 2 /hour for ACUVUE OASYS TM with 0% and 20% of V itamin E loading, respectively. It is noted that this model n eglects the mass transfer resistance in the fluid phase, which is a reasonable assumption except at very short times. 7.3 .2 Therapeutic Release Rates Currently, CyA is delivered through 2 drops per day of oil in water emulsion (Restasis Allergan) that de liver about 28 g (assuming a drop volume of 28 L ) of drug to the eye for the treatment of moderate dry eye  Gupta et al. recently determined the bioavailability of CyA delivered through Restasis to be 2.8%, which indicating that about 0.78 g/day of CyA is delivered to cornea and conjunctiva through this treatment route  CyA delivery through contact lenses will likely have a much higher bioavailability due to the increase in the residence time of the drug in the tears. The bioavailab ility needs to be determined through animal experiments but initial estimate based on mathematical models supporte d by clinical data is about 50%  Based on the 50% bioavailability for contact lenses and 2.8% for Restasis a release rate of 1.6 g/d ay of CyA by contact lens should be able to provide equivalent therapeutic effects. According to the clinical studies of CyA emulsion eye drops (later Restasis ), the concentration of CyA in the emulsion can be increased to 0.4% (Phase II) or 0.1% (Phase III) without significant adverse effect after 12 weeks treatment [152,
199 153] While the higher concentrations are non toxic, no additional benefits were observed with the higher concentrations. Thus, it can be anticipated that the therapeutic window for da ily dose of CyA delivered via eye drops is between 28 to 224 g. Accordingly the therapeutic window for CyA delivery through contact lenses can be estimated to be between 1.6 g/day and 12.8 g/day. Based on these estimates, a suitable contact lens needs to deliver about 12.8 g/day on the first day and 1.6 g/day on the last day of the wear duration. The amount of drug released from the contact lenses in a day is plotted as a function of time (days) in Figure 7 5 for lenses with and without Vitamin E. T he data shows that the lens without Vitamin E can maintain delivery rates within the therapeutic window for 14 days, while ACUVUE OASYS TM with 20% of Vitamin E loading can maintain CyA release within the safe region at the rate above the equivalent Restas is release rate for 20 days. The duration of therapeutic release cannot be increased any further for the lens without Vitamin E without causing toxicity in the early phase of the release. However, the duration of therapeutic release from ACUVUE OASYS TM with 20% of Vitamin E can be further increased by increasing the drug loading in the lens. The drug loading in the lens can be increased by soaking the lens in solution with higher drug concentration since the lens was soaked in CyA solution at 60% of th e solubility limit in this study. To brief summarize the study i n this chapter : CyA is loaded into commercial silicone hydrogel contact lenses by soaking the lenses in CyA PBS solution. Subsequent in vitro release experiments in perfect sink condition dem onstrate that the loaded CyA in the lens can be release fo r about 2 weeks. The release duration can be further increased to the total wear time of the lens which is about a month, through
200 incorporation of Vitamin E into the lens. In addition to increasin g the release duration, the release rates are within the estimated therapeutic window for the entire 1 month period The long release duration along with the higher bioavailability compared to eye drops suggests that Vitamin E loaded s ilicone hydrogel lenses could be very useful for extended and controlled release of CyA. These lenses could potentially be useful for treatment for chronic dry eye and also for reducing the symptoms of contact lens mediated dry eyes. It is noted though that in vivo release and toxicity studies are needed to fully determine the benefits of CyA release from extended wear contact lenses.
201 Figure 7 1 Cumulative drug uptake by ACUVUE OASYS TM lens soaked in 17 g/ mL of 10 mL CyA/PBS solutions.
202 Figure 7 2 Cumulative CyA release by 1 DAY ACUVUE D rug was loa ded in the lens by soaking in 15 g/ mL of 10 mL CyA/PBS solutions for 7 days. The release prof iles were measured in 1.75 mL fresh of PBS that was replaced after every measurement. Data are plotted as mean SD (n=3).
203 F igure 7 3 Cumulative CyA release from silicone contact lens. The drug was loaded in the lenses by soaking in 10 mL of 15 g/ mL CyA/PBS solution for 7 days. The release profiles were measured in removed to 1 .75 mL fresh of PBS that was replaced after every measurement Data are plotted as mean SD (n=3).
204 Figure 7 4 Cumulative drug release from Vitamin E loaded ACUVUE OASYS TM lenses. The drug was loaded in the lenses by soaking in 10 mL of 17 g/ mL CyA/PBS solution for 30 to 60 days. The release pr ofiles were measured in 1.75 mL fresh of PBS that was replaced after every measurement Data were plotted as mean SD (n=3).
205 Figure 7 5 Daily average CyA release rate from Vitamin E loaded ACUVUE OASYS TM lenses. The drug was loaded in the lenses by soaking in 10 mL of 17 g/ mL CyA/PBS solution for 30 to 60 days. The experimental data is plotted as solid markers (mean SD (n=3)). The solid lines are the model fits based on diffusion model ( Equation 7 9 ). The dash lines represent the estimated therapeutic window on the basis of Phase II and Phase III studies for Restasis along with estimated b ioavailability from drops and lenses [152, 153] Phase II Phase III
206 Figure 7 6 Plot of % CyA release by silicone contact lenses versus square root of time. The lines are the best fit straight lines. The fitted slope and R 2 are 6.7967 and 0.9908 for ACUVUE OASYS TM 3.7416 and 0.9823 for O 2 OPTIX TM 3.6036 and 0.9940 for Pure Vision TM and 3.3205 and 0.9660 for NIGHT&DAY TM respectively. Data are presented as mean S.D. with n = 3.
207 Figure 7 7 P lot of % CyA release by Vitamin E loaded ACUVUE OASYS TM versus square root of time. The lines are the best fit straight line. The fitted slope and R 2 are 4.3901 and 0.9894, 3.0685 and 0.9891, 1.9404 and 0.9882 for lens with 0%, 10% and 20% of Vitamin E lo ading, respectively. Data are presented as mean S.D. with n = 3.
208 Figure 7 8 Comparison of CyA and dexamethasone  delivery by Vitamin E loaded ACUVUE OASYS TM
209 Table 7 1 Results of CyA uptake by silicone contact lens. Each lens was soaked in 10 mL of 15 g/ mL CyA PBS solution for 7 days. Data are shown as mean SD (n=3) V pl ( mL ) CyA Uptake ( g) Fraction of CyA Uptake by Contact Lens ( f ) Partition Coefficient, K 1 Day ACUVUE 0.0224 0.0004 9.0 2.7 0.060 0.018 31.6 10.3 NIGHT&DAY TM 0.0224 0.0 005 98.7 4.9 0.658 0.033 910 118 O 2 OPTIX TM 0.0249 0.0003 100.4 2.3 0.700 0.015 858 59 ACUVUE OASYS TM 0.0227 0.0002 77.4 2.4 0.516 0.016 486 30 Pure Vision TM 0.0224 0.0003 110.4 2.5 0.736 0.017 1330 104
21 0 Table 7 2 Results of CyA uptake by Vitamin E loaded ACUVUE OASYS TM lenses. Each lens was soaked in 10 mL of 17 g/ mL CyA PBS solution for various drug uptake durations. Data are shown as mean SD (n=3) Vitamin E loading (g Vitamin E/ g pure lens) Drug uptake duration (days) Drug uptake ( g) K K ve 0 30 103.4 3.1 6.7810 2 0.106 45 134.2 1.3 1.5810 3 1.1510 4 0.200 60 151.1 5.4 2.6210 3 1.3310 4 0.426 120 160.4 4.7 3.7010 3 1.1210 4 0.653 120 166.4 0.7 6.1010 3 1.4810 4
211 CHAPTER 8 DRUG DELIVERY BY CON TACT LENS IN GLAUCOM ATOUS DOGS While there is now extensive literature on in vitro studies for release of ophthalmic drugs through contact lenses, there are very few animal or human studies. A number of human studies were reported in 1970s focusing on management of glaucoma with hydrophilic lenses soaked in pilocarpine. The lenses used in these studies were mostly afocal, 0.2 mm thick, 13.5 14 mm of dia meter and radius of curvature between 7.8 and 8.6 mm Sauflon lenses [27 29] Hillman et al. compared clinical response of Sauflon lenses soaked in 1% pilocarpine solution with that of intensive pilocarpine 4% therapy, which comprises instilling 1 or 2 drops per min utes for 5 min utes every 5 min utes for half an hour and then every 15 min utes for 90 min utes  Hillman et al. reported that the clinical response to the contact lens soaked in 1% solution was bet ter than that for pilocarpine therapy  In another study with the same type of lenses and the same treatment methodologies, Hillman observed a 54.8% drop in intraocular pressure (IOP) with the contact lenses and a 49.7% reduction with the 4% pilocarpin e regimen  The encouraging results of these studies clearly prove the potential of glaucoma therapy through contact lenses. Also the amount of drug loaded in the contact lenses was substantially less than that delivered through eye drops supporting the predictions of Li and Chauhan regarding higher bioavailability of contact lenses compared to drops. These studies were however conducted with lenses that release the drug in a short burst much like the profiles from eye drops. In view of the current renewed interest in ophthalmic drug delivery by extended wear contact lenses, in vivo animal and human studies with extended wear contact lenses are much needed. Such studies are
212 necessary to demonstrate that extended and continuous release of drugs can a chieve the same or better therapeutic efficacy as eye drops. The goals of this study are to demonstrate the efficacy of glaucoma therapy through release of timolol from silicone hydrogel contact lenses. Timolol is a beta blocker that is widely used to tre at glaucoma by reducin g IOP through decreasing the production of aqueous hu mor  We focus on this drug because of the large number of glaucoma patients in the world  and also because of the potential of serious side effects from systemic exposure to timolol  For this in vivo study, we choose the colony of beagle dogs who are affected by or carriers of a hereditary form of primary open angle glaucoma, the most common form of glaucoma in human beings  Beagle dogs have been used in several pr ior studies on glaucoma therapy [70 76] Another benefit of using the Beagle dogs is that t he cornea shape and size of these dogs are similar to that of human beings, and therefore the commercially available contact lens for human can be used in this st udy without further modification. In this c hapter we compare the pharmacodynamics of IOP reduction for contact lenses with that from eye drops. Two different drug loadings are considered for contact lenses to explore the effect of drug loadings and also t o compare bioavailability of contact lenses with that for eye drops. Since silicone hydrogel contact lenses release timolol relatively rapidly, studies are also conducted with contact lenses loaded with Vitamin E, which have longer release times compared to control lenses without Vitamin E. In fact the in vitro results have shown that the timolol release time can be extended from 1 hour by pure NIGHT&DAY TM contact lens to 70 hours by NIGHT&DAY TM with ca. 25% of Vitamin E loading as shown in Chapter 3.
213 8 .1 Materials and Methods 8.1 .1 Materials NIGHT&DAY TM (Lotrafilcon A, Ciba Vision Corp., Duluth, GA) contact lenses (diopter 6.50) are used in this study. T imolol maleate ( 98%) ethanol ( 99.5%), and p hosphate b uffered s aline (PBS) were purchas ed from Sigma Aldrich Chemicals (St. Louis, MO ). Vitamin E (D alpha tocopherol, Covitol F1370) was gifted by Cognis Corporation All other chemicals were of rea gent grade and used without further purification. 8.1 .2 Drug and Vitamin E loading into Conta ct Lens NIGHT&DAY TM lenses were rinsed with DI water before further use. To load V itamin E into the contact lens, each rinsed lens was soaked in a 3 mL of 0.1 g/mL V itamin E ethanol solution for 24 hours. After the loading step, the lens was taken out and excess V itamin E ethanol solution on the lens surface was gently blotted out, and then t he lens was removed into a 30 mL of DI water to extract the residual ethanol in the lens. After 3 hours the lens was removed into fresh DI water to repeat the extracti on process. After extraction, the lens was taken out and the excess water on the surface was blotted out, followed by quickly dipping the lens in ethanol for few seconds to remove the V itamin E on the lens surface. The lens was then transferred into PBS so lution before further use. The V itamin E loading in the lens is c a. 25% w/w acco rding to previous studies, which will release ca. 75% of concluded timolol into tear film within 24 hours at perfect sink condition  To load timolol into the lens, NIGHT& DAY TM lens with or without V itamin E loading was soaked in a 3.5 mL of timolol maleate PBS solution for at least 5 days. The concentrations of timolol mal eate PBS solution are 2.67 mg/mL and 8 mg/mL which are
214 chosen according to our previous in vitro stud ies to provide the capacity of the lens with ca. 25% Vitamin E loading to release 20 g and 60 g of timolol into the eye within 24 hours, respectively [106, 128] The drug release capac ity of each delivery method was summarized in Table 8 1 During the w hole preparation process all lenses are kept in hydrated state to maintain their original shape. In the remainder of this Chapter the notation CON_H and CON_L denote control lenses without Vitamin E that are loaded with the higher (200 g) and the lower ( 67 g) amounts of drug, respectively. Similarly, the notation VIT_H and VIT_L refer to Vitamin E loaded lenses (25% w/w) loaded with the higher (200 g) and the lower (67 g) amounts of drug, respectively. 8.1 .3 Animal Model Before investigating the effec t of these timolol loaded contact lens on glaucomatous dogs, each enrolled study dog (12 adult Beagle dogs with inherited open angle glaucoma) had their IOP estimated via applanation tonometry (Tono Pen XL (Mentor O and O, Norwell, MA)) in both eyes (OU), determined 3 times daily at the same times of day for 5 days to establish a baseline for each individual animal for each of these parameters. A topical anesthetic (proparacaine hydrochloride 0.5%) was applied to each eye prior to the measurement of IOP OU After one week period of drug washout, 10 dogs were used in the investigation of using eye drop to delivery timolol to the eye. E ach study animal received one drop of timolol maleate 0.5% ophthalmic solution to one eye (randomly selected) daily for 5 days IOP OU was measured at time zero and then three times daily through the duration of drug administration. Control experiments were also done to establish that
215 Tono Pen XL can be used to accurately measure IOP even after insertion of a contact lens. After one week period for drug washout, contact lens es with or without Vitamin E loading which designed to release timolol 20 g or 60 g within 24 hours was placed in one eye (randomly selected) in each of the study dogs and IOP OU were measured at time zero an d then three times daily through the duration of drug administration (3 days). 12 dogs were randomly separated into 2 groups with 6 dogs eac h, and in the first week CON_L and CON_H were conducted in each g roups, then followed by VIA_L and VIA_H During the experiment, freshly loaded contact lenses replaced the previous 24 hours. The designed doses in this study have been tested and believed safe in the dog via applicati on. All experiments involving animals in this study were approved by the Institutional Animal Care and Use Committee at University of Florida and were performed in compliance with the ARVO Statement for the Use of Animal in Ophthalmic and Vision Research. 8.1 .4 Data Analysis Each of the measured parameters w as compared between each administration method and the untreated controls to determine if there was a difference in the eff ects among these therapeutic methods. The drug delivery methods comparisons wer e performed using SPSS programs utili zing one way ANOVA tests for multi comparison and Games Howell tests for Post Hoc test since the sample size are not equal. Within each test week, th e average measurements for IOP for each day were compared with subsequ ent measurements to detect significant changes (P < 0.05) using the Games
216 Howell tests and ANOVA for repeated measurements. Each measured parameter for drug treate d eyes was compared both to baseline and to the values for untreated eye. Measured parameters for treated eyes were compared between methods as well. 8.2 Results 8.2 .1 Contact Lens without Vitamin E The mean SEM changes in IOP for each of the tested timolol delivery methods are summarized in Figure 8 1 T imolol delivered by CON_L significantly d ecreas ed IOP from the baseline in the treated eye by 3.18 0.71 mmHg (P=0.002) and insignificantly increased IOP (P=0.167) in the untreated control eye by 2.27 0.79 mmHg from the baseline as indicate d in Figure 8 1 A The difference in IOP decline betwe en the treated and untreated eyes was significant (5.45 0.98 mmHg, P<0.001). As shown in Figure 8 1 B timolol delivered by CON_H significantly decreased IOP in the treated eye by 5.02 0 .83 mmHg (P<0.001) from the baseline, but the IOP difference in the untreated control eye was not significant from the baseline (P=1.000). The difference in IOP decline between the treated and untreated eyes was significant after one day ( 5.20 1.26 mmHg, P= 0.004 ). There was no significant difference between CON_L and CO N_H in treated eye IOP (1.84 1.01 mmHg, P=0.761 ) 8.2 .2 Co ntact Lens with Vitamin E T imolol delivered by VIT_L decreas ed IOP in the treated eye by 4.80 0.63 mmHg (P<0.001) from the baseline, but s howed no significant change in the untreated control eye (P=0.996), as shown in Figure 8 1 C The difference in IOP decline between the treated and untreated eyes was significant (4.18 0.80 mmHg, P<0.001). As shown in Figure 8 1 D timolol delivered by VIT_H significantly decreased IOP in the treated eye by 3. 27 0.71 mmHg from the baseline (P=0.002). The IOP decrease
217 in the untreated control eye from the baseline was not significant (P=0.998). The difference in IOP decline between the treated and untreated eyes was significant (3.90 0.91 mmHg, P=0.003). The re was no significant difference between VIT_L and VIT_H in treated eye (1.53 0.83 mmHg, P=0.777). 8.2 .3 Eye Drop As shown i n Figure 8 2 timolol delivered by eye drops decreased IOP in the treated eye by 4.64 0.41 mmHg from the baseline and in the unt reated control eye by 3.17 0.41 mmHg. The IOP decrease from baseline was significant for both the treated eye (P<0.001) and the untreated eye (P<0.001). The difference in IOP decline between the treated and untreated eyes by eye drops was significant (1. 47 0.43 mmHg, P=0.011). 8.2 .4 Drug Administration Methods Comparison Timolol delivery by different methods was compared here to evaluate the effects. Values for IOP in treated eye were not significantly different among all 5 different tested timolol trea tment methods (P=0.075). The decline of IOP in treated eye from baseline was significant for all these five methods (P<0.001). Figure 8 3 com pares the difference of IOP between treated and untreated eye from different drug administration methods. The resu lts indicated that there was no significant difference among all 4 contact lens delivery methods (P=0.068), while eye drops leaded to significantly smaller IOP difference between treated and untreated eye than CON_L ( 4.24 0.73 mmHg, P<0.001), CON_H ( 3. 47 0.90 mmHg, P=0.003) and VIT_L ( 2.46 0.71 mmHg, P=0.010), but not significantly lower than VIT_H ( 1.84 0.76 mmHg, P=0.134)
218 8.3 Discussion All the five different methods tested in this study effectively lower the IOP in treated eyes from the base line, and the effects are similar among these methods. The IOP reduction with contact lenses was not improved on increasing the drug loading likely due to saturation of the target sites with the drug. In fact, future studies are warranted with further low ering of drug loadings in the lens to determine if IOP reductions comparable to those with eye drops can be achieved with even smaller loadings in the lenses. By comparing the IOP reduction in the lenses with the smaller drug loading with those from the e ye drops, we can deduce that contact lenses deliver a larger fraction of the loaded drug into the cornea compared to eye drops. The IOP reduction in the untreated eye is significantly larger for the eye drops compared to the lenses. The reduction in IOP i n the untreated eye is commonly attributed to the drug transport into the other eye through systemic circulation. For the because of the smaller loss of the drug to the systemic circulation. This data further supports the hypothesis that contact lenses deliver a larger fraction of the loaded drug to the cornea. In this study we did not observe any significant difference in IOP reduction between contact lenses with or wi thout Vitamin E. Inclusion of Vitamin E in the contact lens increases the release duration but the total amount of drug loaded into the lenses with and without Vitamin E was equal. In this study lenses were replaced each day and the lenses loaded with Vit amin E release only about 75% of the drug in the wear time of a day. Thus the total release of drug was actually smaller for the Vitamin E loaded lenses. Furthermore, due to the larger release duration, the rate of drug released from
219 the Vitamin E loaded lenses was substantially smaller that the lenses without Vitamin E. It is thus encouraging that the lenses with Vitamin E can achieve the same therapeutic effect as eye drops in spite of the significantly smaller release rates albeit over a longer durati on. This shows that continuous release of ophthalmic drugs from contact lenses could be efficacious. To further evaluate the effect of extended release by Vitamin E loaded contact lens, in the future we can increase the lens wear interval from daily repl acement to 3 5 day replacement, which will provide further understanding for the effect of extended release from Vitamin E loaded contact lens in animal models. This pilot in vivo study demonstrates the potential advantages of delivering ophthalmic drug s through contact lenses. Contact lenses achieved same efficacy as eye drops but with one third of the drug loading, and also resulted in smaller IOP reduction in the untreated eye signifying reduced drug loss to the systemic circulation. Incorporation o f V itamin E into the lenses does not lead to toxicity but it also does not improve efficacy in this case likely because t lenses were replaced every day. In future, in vivo studies with continuous lens wear should be conducted to further investigate the f easibility of extended drug release by cont act lens both with and without V itamin E.
220 A B Figure 8 1 Effect of insertion of drug loaded contact lenses on t he intraocular pressure (IOP), A) CON_L, B) CON_H, C) VIT_L and D ) VIT_H. Data is presented as (mean SEM). Lenses are inserted in the treated eye at initial time and replaced every 24 hours.
221 C D Figure 8 1. Continued.
222 Figure 8 2 Effect of drug administration through eye drops on the intraocular pressure. Data is prese nted as (mean SEM). Drug loaded drops are instilled in the treated eye at the initial time and then every 24 hours.
223 Figure 8 3 Comparison of the effect of various drug delivery methods on the differences in the IOP between the treated and the untreated eyes. Data is presented as (mean SEM).
224 Table 8 1 Sum mary of various drug delivery methods considered in this study Methods Description Drug Release Capacity ( g) Estimate release within 24 hours ( g) Estimate uptake by eye ( g) CON_L Pure NIGHT&DAY TM soaked in 2.67 mg/ mL timolol ma leate solution 67 67 27 VIT_L NIGHT&DAY TM with 25% Vitamin E, soaked in 2.67 mg/ mL timolol ma leate solution 67 50 20 CON_H Pur e NIGHT&DAY TM soaked in 8 mg/ mL timolol ma leate solution 200 200 80 VIT_H NIGHT&DAY TM with 25% Vitamin E, soaked in 8 mg/ mL timolol ma leate solution 200 150 60 Eye drop 0.5% ophthalmic solution, one drop 150 150 7.5
225 CHAPTER 9 CONCLUSIONS Our study has conclusi vely shown the feasibility of extended ophthalmic drug delivery by silicone hydrogel contact lens containing Vitamin E as diffusion barriers. In Chapter 2 s everal properties including geometry, ion permeability, oxygen permeability a nd UV transmittance are characterized to determine the pros and cons of loading Vitamin E into the lenses. The results indicate the property change s caused by Vitamin E loading do not disqualify these silicone hydrogels as extended wear contact lens. In addition, Vitamin E loadi ng has a beneficial effect of blocking UV radiation which will reduce the corneal damage due to UV light. For extended wear, the most effected critical property of silicone hydrogel contact lens with Vitamin E is ion permeability, and thus a further securi tization of the ion transport of silicone hydrogels with various compositions is discussed in Chapter 6 Chapter s 3 to 5 demonstrated the In vitro drugs release by the Vitamin E loaded silicone hydrogel contact lenses, and the results indicated that the increase in release duration is significantly dependent on the interaction between Vitamin E and the drug of interest. F or hydrophilic drugs (timolol, fluconazole, dexamethasone phosphate), the drug release duration increases quadratically in Vitamin E lo ading; f or hydrophobic drugs dexamethasone and cyclosporine A the effect of the Vitamin E inclusion is smaller but still signifi cant for release. On the other hands, for some amphiphilic anesthetic drugs, including lidocaine, bupivacaine and tetracaine, t he interfacial interaction between drug and Vitamin E aggregation plays the determinative role for drug transport through the Vitamin E/silicone hydrogel matrix. Ocular drugs delivery by contact lens can be viewed as a one dimensional transport by a flat thin film, and
226 subsequent mathematical models based on the proposed mechanisms were established. In addition, a case study in Chapter 7 explored the potential of silicone hydrogel contact lens for the treatment of chronic dry eye syndrome, and the in vitro results suggests that CyA delivery by silicone hydrogel contact lens can provide much safer and more efficient drug delivery route compared to traditional commercial eye drops. Finally, in Chapter 8 we demonstrated that timolol can successfully delivered to glaucomatous dog s via drug impregnated contact lenses without irritation of eye or any other unwanted safety concern By utilizing contact lens to del iver timolol to the eye, the intraocular pressure in the treated eye decreased effectively to similar degree compared to that by eye drop treatment while it significantly reduced the risk of systemic drug exposure. In conclusion, s ilicone hydrogel contact lenses with Vitamin E are promising candidates for extended ophthalmic drug delivery. The Vitamin l oading inside the silicone hydrogel matrix can significantly attenuate the drug delivery rate, reduce wastage an d provide safer treatment route. W hile the results presented have focused on drug contact lenses, the novel approach of in situ creation of tra nsport barriers in silicone hydrogels could be used in other areas where extended release of solutes is desired, such as pucta plugs, ophtha coils, retinal implants, transdermal patches, wound healing patches, cornea replacement materials, etc.
227 LIST O F REFERENCES  S. Duvvuri, S. Majumdar, A.K. Mitra, Drug delivery to the retin a: challenges and opportunities, Exp ert Opin. Bio. Ther. 3(1) (2003) 45 56.  S.H. Kim, R.J. Lutz, N.S. Wang, M.R. Robinson, Transport barriers in transscleral dru g delivery for retinal diseases, Ophthalmic Res 39(5) (2007) 244 254.  S. Dey, B.S. Anand, J. Patel, A.K. Mitra, Transporters/receptors in the anterior chamber: pathways to explore ocular drug delivery strategies, Exp ert Opin Bio Ther 3(1) (2003) 23 44.  S. B. Koevary, Pharmacokinetics of topical ocular drug delivery: potential uses for the treatment of diseases of t he posterior segment and beyond, Curr Drug Metab 4(3) (2003) 213 222.  J.C. Lang, O cular drug delivery conventional ocular formulations Adv Drug Deliv Rev 16(1) (1995) 39 43.  F.J. Holly, Formatio n and rupture of the tear film, Exp Eye Res 15(5) (1973) 515 525.  R.L. Farris, Tear analysis in contact lens wearers. Transactions of the American Ophthalmological Society 83 (1985) 501 5 45.  M. Rolando, M. Zierhut, The ocular surface and tear film and their dysfunction in dry eye disease, Surv. Ophthalmol 45 (2001) S203 S210.  C.W. Chao, J.P. Vergnes, I.L. Freeman, S.I. Brown, Biosynthesis and partial characteriza tion of tear film glycoproteins: Incorporation of radioactive precursors by human l acrimal gland explants in vitro, Exp. Eye Res 30(4) (1980) 411 425.  J. Ruiz Ederra, M.H. Levin, A.S. Verkman, In situ fluorescence measurement of tear film [Na+],[K+],[Cl ], and pH in m ice shows marked hyperton icity in aquaporin 5 deficiency, Investig Ophthalmol Vis Sci 50(5) (2009) 2132.  K. Hosoya, V.H.L. Lee, K.J. Kim, Roles of the conju nctiva in ocular drug delivery a review of conjunctival transport mechanisms and their regu la tion, Eur J Pharm Biopharm 60(2) (2005) 227 240.  C. Le Bourlais, L. Acar, H. Zia, P.A. Sado, T. Needham, R. Leverge, Ophthalmic drug delivery systems Recent advances Prog Retin. Eye Res 17(1) (1998) 33 58.  S.K. Sahoo, F. Diinawaz, S. K rishnakumar, Nanotec hnology in ocular drug delivery, Drug Discov Today 13(3 4) (2008) 144 151.
228  J.L. Creech, A. Chauhan, C.J. Radke, Dispersive mixing in the posterior tear film under a soft contact lens, Ind Eng Chem Res 40(14) (2001) 3015 3026.  N.A. McNamara, K.A. Polse, R.J. Brand, A.D. Graham, J.S. Chan, C.D. McKenney, Tear mixing under a soft contact lens: Effects of lens diameter, Am J Ophthalmol 127(6) (1999) 659 665.  C.C.S. Karlgard, N.S. Wong, L.W. Jones, C. Moresoli, In vitro uptake and release studies of ocular pharmaceutical agents by silicon containing and p HEMA hydrog el contact lens materials, Int. J Pharm 257(1 2) (2003) 141 151.  Y. Lai, A. Wilson, S. Zantos, Contact Lens. Kirk Othmer Encyclopedia of Chemical Tech nology, Vol. 7, 4th ed., New York: Wiley, 1993, pp. 191 218.  O. Wichterle, D. Lim, Hydr ophilic gels for biological use, Nature 185 (1960) 117 118  P.C. Nicolson, J. Vogt, Soft cont act lens polymers: an evolution, Biomaterials 22(24) (2001) 3273 3 283.  E.C. Poggio, R.J. Glyn, O.D. Schein, J. Seddon, M.J. Shannon, V.A. Scardino, K.R. Kenyon, The incidence of ulcerative keratitis among users of daily wear and ex tended wear soft contact lenses, New Eng J. Med. 321(12) (1989) 779 783.  O.D. Sc hein, R.J. Glynn, E.C. Poggio, J.M. Seddon, K.R. Kenyon, The relative risk of ulcerative keratitis among users of daily wear and extended wear soft contact lenses. A case control study. Microbial Keratitis Study Group, New Eng J. Med. 321(12) (1989) 773 7 78  B.A. Holden, G.W. Mertz, C ritical oxygen levels to avoid corneal edema for daily and extended wear contact lenses Investig Ophthalmol Vis Sci 25(10) (1984) 1161 1167.  D.M. Harvitt, J.A. Bonanno, Re evaluation of the oxygen diffusion mode l for predicting minimum contact lens Dk/t values needed to avoid corneal anoxia, Optom Vis Sci 76(10) (1999) 712 719  W.G. Deichert, K.C. Su, M.F.V.A.N. Buren, Polysiloxa ne composition and contact lens, US Patents 4,153,641, 1979.  P. Nicolson R.C. Baron, P. Chabrecek, J. Court, A. Domscheke, H.J. Criesser Extended wear ophthalmic lens, US Patent No. 5760100, 1998.  R.M. Ramer, A.R. Gasset, Ocular penetration of pilocarpine: the effect on concentration on the oc ular penetration of pilocar pine, Ann Ophthalmol 6(11) (1974) 1160 1162
229  J.S. Hillman, M anagement of acute glaucoma with pilocarpine soaked hydrophilic lens Br. J Ophthalmol 58(7) (1974) 674 679.  J.S. Hillman, J.B. Marsters, A. Broad, P ilocarpine delivery by hydrophili c lens in management of acute glaucoma, Trans Ophthalmol Soc United Kingd 95 (APR) (1975) 79 84.  M. Ruben, R. Watkins, P ilocarpine dispensatio n for soft hydrophilic contact l ens Br. J Ophthalmol 59(8) (1975) 455 458.  B.W. Arthur, G.J. Hay S.M. Wasan, W.E. Willis, Ultrastructural effects of topical timolol on the rabbit cornea. Outcome alone and in conjunction wi th a gas permeable contact lens, Arch Ophthalmol 101(10) (1983) 1607 1610  C.L. Schultz, I.M. Nunez, D.L. Silor, M.L. Neil Contact lens containing a leachable absorbed mater ial, US Patents 5,723,131, 1998.  J.H. Ward, N.A. Peppas, Preparation of controlled release systems by free radical UV polymeriza tions in the presence of a drug, J Control Release 71(2) (2001) 183 1 92.  P. Colombo, R. Bettini, N.A. Peppas, Observation of swelling process and diffusion front position during swelling in hydroxypropyl methyl cellulose (HPMC) mat rices containing a soluble drug, J Control Release 61(1 2) (1999) 83 91.  M.T. am E nde, N.A. Peppas, Transport of ionizable drugs and proteins in crosslinked poly (acrylic acid) and poly (acrylic acid co 2 hydroxyethyl methacrylate) hydrogels. II Diffusion and release studies, J Control Release 48(1) (1997) 47 56.  C.L. Schultz, J .M. Mint, Drug delivery system f or anti glaucomatous medication, US Patent 6,410,045, 2002.  D. Gulsen, A. Chauhan, Dispersion of microemulsion drops in HEMA hydrogel: a potential o phthalmic drug delivery vehicle, Int. J Pharm 292(1 2) (2005) 95 117  D. Gulsen, C.C. Li, A. Chauhan, Dispersion of DMPC liposomes in contact lens es for ophthalmic drug delivery, Curr Eye Res 30(12) (2005) 1071 1080.  Y. Kapoor, A. Chauhan, Drug and surfactant transport in Cyclosporine A and Brij 98 laden p HEMA hydrogels, J Colloid Interface Sci 322(2) (2008) 624 633.  Y. Kapoor, A. Chauhan, Ophthalmic delivery of Cyclosporine A from Brij 97 microemulsion and su rfactant laden p HEMA hydrogels, Int J Pharm 361(1 2) (2008) 222 229.
230  Y. Kapoor, J.C. Th omas, G. Tan, V.T. John, A. Chauhan, Surfactant laden soft contact lenses for extend ed delivery of ophthalmic drugs, Biomaterials 30(5) (2009) 867 878.  M. Ali, S. Horikawa, S. Venkatesh, J. Saha, J.W. Hong, M.E. Byrne, Zero order therapeutic release f rom imprinted hydrogel contact lenses within in vitro physiological ocular tear flow, J Control Release 124(3) (2007) 154 162.  M.E. Byrne, K. Park, N.A. Peppas, Molecular imprinting within hydrogels Adv Drug Deliv Rev 54(1) (2002) 149 161.  S. Venkatesh, S.P. Sizemore, M.E. Byrne, Biomimetic hydrogels for enhanced loading and extended release of ocular therapeutics, Biomaterials 28(4) (2007) 717 724.  C. Alvarez Lorenzo, H. Hiratani, J.L. Gomez Amoza, R. Martinez Pacheco, C. Souto, A. Con cheiro, Soft contact lenses capable of sustained delivery of timolol J Pharm Sci 91(10) (2002) 2182 2192.  H. Hiratani, C. Alvarez Lorenzo, Timolol uptake and release by imprinted soft contact lenses made of N,N diethyl acrylamide and methacrylic ac id, J Control Release 83(2) (2002) 223 230.  H. Hiratani, C. Alvarez Lorenzo, The nature of backbone monomers determines the performance of imprinted soft contact lenses as timolol drug delivery systems Biomaterials 25(6) (2004) 1105 1113.  H. H iratani, A. Fujiwara, Y. Tamiya, Y. Mizutani, C. Alvarez Lorenzo, Ocular release of timolol from molecularly imprinted soft contact lenses Biomaterials 26(11) (2005) 1293 1298.  H. Hiratani, Y. Mizutani, C. Alvarez Lorenzo, Controlling drug release fr om imprinted hydrogels by modifying the characteristics of the imprinted cavities Macromol Biosci 5(8) (2005) 728 733.  J.B. Ciolino, T.R. Hoare, N.G. Iwata, I. Behlau, C.H. Dohlman, R. Langer, D.S. Kohane, A Drug Eluting Contact Lens Investig Oph thalmol Vis Sci 50(7) (2009) 3346 3352.  H.A. Quigley, A.T. Broman, The number of people with glaucoma worldwide in 2010 and 2020 Br J Ophthalmol 90(3) (2006) 262 267.  D.S. Friedman, B. Nordstrom, E. Mozaffari, H.A. Quigley, Glaucoma manage ment among individuals enrolled in a sing le comprehensive insurance plan, Ophthalmology 112(9) (2005) 1500 1504.
231  M.M. Hermann, A.M. Bron, C.P. Creuzot Garcher, M. Diestelhorst, Measurement of Adherence to Brimonidine Therapy for Glaucoma Usin g Electro nic Monitoring, J. Glaucoma Publish Ahead of Print : ( 10.1097/IJG.1090b1013e3181f1093eb1094a )  D. Fonn, Targeting contact lens induced dryness and discomfort: What properties will make lenses more comfortable Optom Vis Sci 84(4) (2007) 279 285. [ 54] W.R. Falla, M. Mulski, E.L. Cussler, Estimating diffusion through flake filled membranes, J Membr Sci 119(1) (1996) 129 138.  N.K. Lape, E.E. Nuxoll, E.L. Cussler, Polyd isperse flakes in barrier films, J Membr Sci 236(1 2) (2004) 29 37.  S.E. Nam, K.H. Lee, Preparation and characterization of palladium alloy composite membranes with a diffusion barrier for hydrogen separation Ind Eng Chem Res 44(1) (2005) 100 105.  A.A. Dameron, S.D. Davidson, B.B. Burton, P.F. Carcia, R.S. McLea n, S.M. George, Gas diffusion barriers on polymers using multilayers fabricated by Al2O3 and rap id SiO2 atomic layer deposition, J Phys Chem C 112(12) (2008) 4573 4580.  pentoxify lline and aprotinin on light induced retinal injury, Ophthalmologica 221(3) (2000) 159 166.  K. Bilgihan, U. Adiguzel, C. Sezer, G. Akyol, B. Hasanreisoglu, Effects of topical vitamin E on keratocyte apoptosis after traditio nal photorefractive keratect omy, Ophthalmologica 215(3) (2001) 192 196.  M. Kojima, Y.B. Shui, H. Murano, K. Sasaki, Inhibition of steroid induced cataract in rat eyes by administration of vitamin E ophthalmic solution, Ophthalmic Res 28(2) (1996) 64 71.  M. Nagata, M. Kojim a, K. Sasaki, Effect of vitamin E eye drops on napht halene induced cataract in rats, J Ocul Pharm Ther 15(4) (1999) 345 350.  Y. Ohta, Possibility of Clinical Application of Vitamin E to Cataract Prevention J Clin Biochem Nutr 35(1) (2004) 35 4 5.  Y. Ohta, T. Yamasaki, T. Niwa, Y. Majima, Preventive effect of vitamin E containing liposome instillation on cataract progression in 12 month old rats fed a 25% galactose diet J Ocul Pharm Ther 16(4) (2000) 323 335.
232  Y. Ohta, T. Yamasaki, T. Niwa, Y. Majima, I. Ishiguro, Preventive effect of topical vitamin E containing liposome instillation on the progression of galactose cataract. Comparison between 5 week and 12 week ol d rats fed a 25% galactose diet, Exp Eye Res 68(6) (1999) 747 755.  R.F. Hofmann, D.J. Bottoni, Gl utathione antioxidant eye drops, US Patents 5,817,630, 1998.  A.G. Braswell, K.J. Absher, A. Duar te, Liquid eye drop composition, US Patents 6,194,457 B1, 2001.  T.J. Zimmerman, W.P. Boger, B eta adrenergic blocki ng agents and the treatment of glaucoma Surv Ophthalmol 23(6) (1979) 347 362.  F.T. Fraunfelder, S.M. Meyer, S ystemic side effects from ophthalmic timolol and their prevention J Ocul Pharm 3(2) (1987) 177 184.  R.N. Weinreb, P.T. Kh aw, Prima ry open angle glaucoma, Lancet 363(9422) (2004) 1711 1720.  K.N. Gelatt, E.O. MacKay, Effect of different dose schedules of latanoprost on intraocular pressure and pupil size in the glaucomatous Beagle, Vet Ophthalmol 4(4) (2001) 283 288.  K.N. G elatt, E.O. MacKay, Effect of different dose schedules of travoprost on intraocular pressure and pupil size in the glaucomatous Beagle, Vet Ophthalmol 7(1) (2004) 53 57.  K.N. Gelatt, E.O. Mackay, T. Dashiell, A. Biken, Effect of different dose sched ules of 0.15% unoprostone isopropyl on intraocular pressure and pupil size in the glaucomatous beagle, J Ocul Pharm Ther 20(5) (2004) 411 420.  R.M. Gwin, K.N. Gelatt, G.G. Gum, R.L. Peiffer, L.W. Williams, E ffect of topical pilocarpine on intraocu lar pressure and pupil size in normo tensive and glaucomatous beagle, Investig Ophthalmol V is Sci 16(12) (1977) 1143 1148.  C.E. Plummer, E.O. MacKay, K.N. Gelatt, Comparison of the effects of topical administration of a fixed combination of dorzola mide timolol to monotherapy with timolol or dorzolamide on IOP, pupil size, and heart rate in glaucomatous dogs, Vet Ophthalmol 9(4) (2006) 245 249.  N. Takiyama, S. Shoji, I. Habata, S. Ohba, The effects of a timolol maleate gel forming solution on normotensive beagle dogs J Vet Med Sci 68(6) (2006) 631 633.
233  S. Volopich, M. Mosing, U. Auer, B. Nell, Comparison of the effect of hypertonic hydroxyethyl starch and mannitol on the intraocular pressure in healthy normotensive dogs and the effect of hypertonic hydroxyethyl starch on the intraocular pressur e in dogs with primary glaucoma, Vet Ophthalmol 9(4) (2006) 239 244.  A. Domschke, D. Lohmann, L. Winterton, On eye mobility of soft oxygen permeable contact lenses. Proceedings of the ACS Spring Nationa l Meeting, San Francisco: PMSE, 1997.  M.D. Young, W.J. Benjamin, Calibrated oxygen permeability of 35 conventional hydrogel materials and correlation with water content Eye Contact Lens 29(2) (2003) 126 133  N. Efron, P.B. Morgan, I.D. Cameron, N.A. Brennan, M. Goodwin, Oxygen permeability and water content of silicone hydrogel contact lens materials. Optometry & Vision Science 84(4) (2007) E328 E337  ACUVUE OASYS, Johnson & Johnson Vision Care, Inc., http: // www.acuvue.com/products acuvue oasys.htm  L. Moore, J.T. Ferreira, Ultraviolet (UV) transmittance characteristics of daily disposable and s ilicone hydrogel contact lenses, Contact Lens Anterior Eye 29(3) (2006) 115 122.  K. Kramer Stickland, E.S. Krol, D.C. Liebler, UV B induced photooxidation of vitamin E in mouse skin, Chem R es T oxicol 12(2) (1999) 187 191.  E.S. Krol, K.A. Kramer Stickland, D.C. Liebler, P hotoprotective action s of topically a pplied vitamin E, Drug M etab R ev 32(3 4) (2000) 413 420.  J. Kim, A. Conway, A. Chauhan, Extended delivery of ophthalmic drugs by s ilicone hydrogel contact lenses, Biomaterials 29(14) (2008) 2259 2269.  J.C. Melby, Systemic cortiscosteroid therap y: pharmacology a nd endocrinology considerations, Ann I nt M ed 81(4) (1974) 505 512  S.E. Wilson, Use of lasers for vision correction of nearsightedness and farsightedness New Eng J Med 351(5) (2004) 470 475.  H.P. Sandoval, L.E.F. de Castr o, D.T. Vroman, K.D. Solomon, Refractive surgery survey 2004, J Cataract Refract Surg 31(1) (2005) 221 233.  R.J. Duffey, D. Leaming, US trends in refractiv e surgery: 2003 ISRS/AAO survey, J Refract Surg 21(1) (2005) 87 91.  A. Reynolds, J.E Moore, S.A. Naroo, C.B. Moore, S. Shah, Excimer la ser surface ablation a review, Clin Exp Ophthalmol 38(2) (2010) 168 182.
234  H.V. Gimbel, E.E. Anderson Penno, J.A. van Westenbrugge, M. Ferensowicz, M.T. Furlong, Incidence and management of intrao perative and early postoperative complications in 1000 consecutive laser in situ keratomileusis cases Ophthalmology 105(10) (1998) 1839 1848.  H.V. Gimbel, S.G. Levy, Indications, resu lts, and complications of LASIK, Curr Opin Ophthalmol 9(4) (1998 ) 3 8  E.A. Paysse, Photorefractive keratectomy for anisometropic amblyopia in children Trans Am Ophthalmol Soc 102 (2004) 341 372  E.A. Davis, D.R. Hardten, R.L. Lindstrom, LASIK complications Int O phthalmol C lin 40(3) (2000) 67 75 [94 ] G.S. Schwartz, D.H. Park, S. Schloff, S.S. Lane, Traumatic flap displacement and subsequent diffuse lamellar keratitis aft er laser in situ keratomileusis, J Cataract R efract Surg 27(5) (2001) 781 783.  R.J. Smith, R.K. Maloney, Diffuse lamellar k eratitis: A new syndrome in lamellar refractive surgery Ophthalmology 105(9) (1998) 1721 1726.  S.A. Nissman, R.E. Tractenberg, A. Babbar Goel, J.F. Pasternak, Oral gabapentin for the treatment of postoperative pain after photorefractive keratectomy Am J Ophthalmol 145(4) (2008) 623 629  R. Autrata, J. Rehurek, Laser assisted subepithelial keratectomy for myopia: two year follow up J Cataract Refract Surg 29(4) (2003) 661 668  A.T. Engle, J.M. Laurent, S.C. Schallhorn, S.D. Toman, J.S Newacheck, D.J. Tanzer, J.L. Tidwell, Masked comparison of silicone hydrogel lotrafilcon A and etafilcon A extended wear bandage contact lenses after photorefractive keratectomy J Cataract Refract Surg 31(4) (2005) 681 686.  J.D. Edwards, K.S. Bo wer, D.A. Sediq, J.M. Burka, R.D. Stutzman, C.R. VanRoekel, C.P. Kuzmowych, J.B. Eaddy, Effects of lotrafilcon A and omafilcon A bandage contact lenses on visual outcomes after photorefractive keratectomy J Cataract & Refract Surg 34(8) (2008) 1288 129 4.  H.S. Brilakis, T.A. Deutsch, Topical tetracaine with bandage soft contact lens pain control after photorefractive keratectomy J Refract. Surg 16(4) (2000) 444 447.  P.M. Cherry, The treatment of pain following excimer laser photorefractive keratectomy: additive effect of local anesthetic drops, topical diclo fenac, and bandage soft contact, Ophthalmic surg lasers 27(5 Suppl) (1996) S477 S480
235  P. Demers, P. Thompson, R.G. Bernier, J. Lemire, P. Laflamme, Effect of occlusive pressure pa tching on the rate of epithelial wound healing af ter photorefractive keratectomy, J Cataract Refract Surg 22(1) (1996) 59 62.  B.A. Holden, P.R. Sankaridurg, D.F. Sweeney, S. Stretton, T.J. Naduvilath, G. N. Rao, Microbial keratitis in prospective studies of extended wear with dis posable hydrogel contact lenses, Cornea 24(2) (2005) 156 161  S. Tomas Barberan, P. Fagerholm, Influence of topical treatment on epithelial wound healing and pain in the early postoperative period follow ing photorefra ctive keratectomy, Acta Ophthalmol Scandinavica 77(2) (1999) 135 138.  S. Verma, M.C. Corbet, J. Marshall, A prospective, randomized, double masked trial to evaluate the role of topical anesthetics in controlling pain af ter photorefractive keratectom y, Ophthalmology 102(12) (1995) 1918 1924.  C.C. Peng, J. Kim, A. Chauhan, Extended delivery of hydrophilic drugs from silicone hydrogel contact lenses containi ng Vitamin E diffusion barriers, Biomaterials 31(14) (2010) 4032 4047.  J. Kim, A. Cha uhan, Dexamethasone transport and ocular delivery from poly ( hydroxyethyl methacrylate) gels, Int J Pharm 353(1 2) (2008) 205 222.  J. Kim, C.C. Peng, A. Chauhan, Extended release of dexamethasone from silicone hydrogel cont act lenses containing vi tmain E, J Control Release 148 (2010) 110 116.  B. Tighe, in: S. D (Ed.), Silicone hydrogels: the rebirth of continuous wear contact lenses, Betterworth Heinemann, Oxford, 2000, pp. 1 21.  S.L. Willis, A novel phosphorylcholine coated con tact l ens for extended wear use, Biomaterials 22(24) (2001) 3261 3272.  T. Tanaka, Encyclopedia of polymer science and engineering. Wiley: New York 6 (1986) 514.  H. Ju, A.C. Sagle, B.D. Freeman, J.I. Mardel, A.J. Hill, Characterization of sodium chlor ide and water transport in crosslinked poly (ethylene oxide) hydrogels J Membr Sci 358(1 2) (2010) 131 141.  A.C. Sagle, H. Ju, B.D. Freeman, M.M. Sharma, PEG based hydrogel membrane coatings Polymer 50(3) (2009) 756 766.  R.J. Petersen, Com posite reverse osmos is and nanofiltration membranes, J Membr Sci 83(1) (1993) 81 150.
236  W.R. Clark, R.J. Hamburger, M.J. Lysaght, Effect of membrane composition and structure on solute removal and b iocompatibility in hemodialysis, Kidney int 56(6) (1999) 2005 2015.  S.M. Murphy, C.J. Hamilton, M.L. Davies, B.J. Tighe, Polymer membranes in clinical sensor applications:: II. The design and fabrication of permselective hydrogels for electrochemical devices Biomaterials 13(14) (1992) 979 990. [117 ] P. Lopour, V. Janatova, Silicone rubber hydrogel composites as polymeric biomaterials:: VI. Transport proper ties in the water swollen state, Biomaterials 16(8) (1995) 633 640.  P. Lopour, P. Vondracek, V. Janatova, J. Sulc, J. Vacik, Silicone rubber hydrogel composites as polymeric biomaterials:: II. Hydrophilicity and permeability to water soluble low molecular weight compounds, Biomaterials 11(6) (1990) 397 402.  H. Yasuda, L.D. Ikenberry, C.E. Lamaze, Permeability of solutes through hydrated polymer membranes. Part II. Permeability o f water soluble organic solutes, Die Makromolekulare Chemie 125(1) (1969) 108 118.  H. Yasuda, C.E. Lamaze, L.D. Ikenberry, Permeability of solutes through hydrated polymer membranes. Part I. Diffusion of sodi um chloride, Die Makromolekulare Chemie 118(1) (1968) 19 35.  H.K. Lonsdale, U. Merten, R.L. Riley, Transport properties of cellulose acetate os motic membranes, J Appl Polym Sci 9(4) (1965) 1341 1362.  M.L. Cheng, Y.M. Sun, Observation of the solute transport in the permeation through hydrogel mem branes by using FTIR microscopy, J Membr Sci 253(1 2) (2005) 191 198.  Y.M. Sun, J.N. Chang, Solute transport in poly (2 hydroxyethyl m ethacrylate) hydrogel membranes, J Polym Res 2(2) (199 5) 71 82.  K. Nagai, S. Tanaka, Y. Hirata, T. Nakagawa, M.E. Arnold, B.D. Freeman, D. Leroux, D.E. Betts, J.M. DeSimone, F.A. DiGiano, Solubility and diffusivity of sodium chloride in phase separated block copolymers of poly (2 dimethylaminoethyl meth acrylate), poly (1, 1' dihydroperfluorooctyl methacrylate) and poly (1, 1, 2, 2 tetrahydroperfluorooctyl acry late), Polymer 42(25) (2001) 09941 09948.  R.W. Baker, Membrane technology and applications, Wiley, 2004.  B. Amsden, So lute diffusion wi thin hydrogels: mechanisms and models, Macromolecules 31(23) (1998) 8382 8395.
237  M.H. Cohen, D. Turnbull, Molecular transport in liquids and glasses J Chem Phys 31 (1959) 1164.  C.C. Li, A. Chauhan, Modeling ophthalmic drug delivery by soaked contact lenses Ind Eng Chem Res 45(10) (2006) 3718 3734.  C.J.D. Fell, H.P. Hutchison, Diffusion coefficients for sodium and potassium chlorides in water at elevated temperatures, J Chem Eng Data 16(4) (1971) 427 429.  L. Jones, K. Dumbl eton, M.M. Hom, A. Bruce, Manual of Contact Lens Prescribing and Fitting, Butterworth Heinemann, Boston, 2006, pp. 393 441.  M. Ca longe, The treatment of dry eye, Surv Ophthalmol 45 (2001) S227 S239.  R.C. Walton, R.B. Nussenblatt, S.M. Whitcup Cyclosporine therapy for severe sight threatening uvei tis in children and adolescents, Ophthalmology 105(11) (1998) 2028 2034.  P. Kulkarni, Review: Uvei tis and immunosuppressive drugs, J Ocul Pharm Ther 17(2) (2001) 181 187.  N. Pucci, E. N ovembre, A. Cianferoni, E. Lombardi, R. Bernardini, R. Caputo, L. Campa, A. Vierucci, Efficacy and safety of cyclosporine eyedrops in vernal keratoconjunctivitis, Ann Allergy Asthma Immunol 89(3) (2002) 298 303.  D.M. Wilson, G.R. John, J.P. Callen Peripheral ulcerative keratitis an extracutaneous neutrophilic disorder: Report of a patient with rheumatoid arthritis, pustular vasculitis, pyoderma gangrenosum, and Sweet's syndrome with an excellent r esponse to cyclosporine therapy, J Am Acad Der matol 40(2) (1999) 331 334.  J.C. Hill, T he use of cyclosp orine in high risk keratoplasty, Am J Ophthalmol 107(5) (1989) 506 510.  G. Ismailos, C. Reppas, J.B. Dressman, P. Macheras, U nusual solubility behavior of cyc losporine a in aqueous me dia, J Pharm Pharmacol 43(4) (1991) 287 289.  Drug Approval Package, U.S. Food and Drug Administration, http://www.accessdata.fda.gov/drugsatfda_docs/nda/200 3/21 023_Restasis.cfm.  J.L. Davis, B.C. Gilger, M.R. Robinson, Novel approaches to ocular drug delivery Curr Opin Mol. Ther 6(2) (2004) 195 205.  M.R. Robinson, K.G. Csaky, R.B. Nussenblatt, J.A. Smith, P. Yuan, C. Sung, M.P. Fronheiser, H. Kim, O cular therapeutic agent delivery devices and methods for making and using such devices US Patent 20070190111 A1 2007.
238  H. Kim, K.G. Csaky, B.C. Gilger, J.P. Dunn, S.S. Lee, M. Tremblay, F. de Monasterio, G. Tansey, P. Yuan, P.M. Bungay, R.J. Lutz, M.R. Robinson, Preclinical evaluation of a novel episcleral cyclosporine implant for ocular graft versus host disease Investig Ophthalmol V is Sci 46(2) (2005) 655 662.  B.C. Gilger, J.H. Salmon, D.A. Wilkie, L.P.J. Cruysberg, J. Kim, M. Hay at, H. Kim, S. Kim, P. Yuan, S.S. Lee, S.M. Harrington, P.R. Murray, H.F. Edelhauser, K.G. Csaky, M.R. Robinson, A novel bioerodible deep scleral lamellar cyclosporine implant for uveitis Investig Ophthalmol Vis Sci 47(6) (2006) 2596 2605.  S.S. Lee, H. Kim, N.S. Wang, P.M. Bungay, B.C. Gilger, P. Yuan, J. Kim, K.G. Csaky, M.R. Robinson, A pharmacokinetic and safety evaluation of an episcleral cyclosporine implant for potential use in high risk keratoplasty rejection Investig Ophthalmol Vis Sc i 48(5) (2007) 2023 2029.  M.R. Robinson, S.S. Lee, B.I. Rubin, A.S. Wayne, S.Z. Pavletic, M.R. Bishop, R. Childs, A.J. Barrett, K.G. Csaky, Topical corticosteroid therapy for cicatricial conjunctivitis associated with chronic graft versus host disea se, Bone Marrow Transpl 33(10) (2004) 1031 1035.  D. Benezra, G. Maftzir, O cular penetration of cyclosporine a in the rat eye Arch Ophthalmol 108(4) (1990) 584 587.  M.J. Doughty, D. Fonn, D. Richter, T. Simpson, B. Caffery, K. Gordon, A pat ient questionnaire approach to estimating the prevalence of dry eye symptoms in patients presenting to opt ometric practices across Canada, Optom Vis Sci 74(8) (1997) 624 631.  C.G. Begley, R.L. Chalmers, G.L. Mitchell, K.K. Nichols, B. Caffery, T. Simpson, R. DuToit, J. Portello, L. Davis, Characterization of ocular surface symptoms from optome tric practices in North America, Cornea 20(6) (2001) 610 618.  C. Gupta, A. Chauhan, Ophthalmic delivery of c yclosporin A by punctal plugs, J. Control. R elease 150 (2010) 70 76  M.M. Hom, Use of cyclosporine 0.05% ophthalmic emulsion for contact lens intolerant patients Eye Contact Lens 32(2) (2006) 109 111.  C. Gupta, A. Chauhan, Drug transport in HEMA conjunctival inserts containing precipita ted drug particles J Colloid Interf Sci 347(1) (2010) 31 42.  Restasis Official Website Allergan, Inc. www.restasis.com  K. Sall, O.D. Stevenson, T.K. Mundorf, B.L. Reis, A.P.S.G. Cs, Two multicenter, ran domized studies of the efficacy and safety of cyclosporine ophthalmic emulsion in mod erate to severe dry eye disease, Ophthalmology 107(4) (2000) 631 639.
239  D. Stevenson, J. Tauber, B.L. Reis, A.P.S.G. Cyclosporin, Efficacy and safety of cyclosporin A ophthalmic emulsion in the treatment of moderate to severe dry eye disease A dose ranging, randomized trial, Ophthalmology 107(5) (2000) 967 974.
240 BIOGRAPHICAL SKETCH Cheng Chun Peng was born in Hsinchu, Taiwan in 1981. After graduating from Taipei Muni cipal Chieh Kuo High School (Taipei, Taiwan) in 1999, he began his undergraduate studies in National Taiwan University (NTU) in Taipei, Taiwan in September of 1999. He received his Bachelor of Science degree in chemical engineering in June of 2003. Shortly thereafter, h e continued his graduate studies in NTU in the Fall of 2003 and received his Master of Science degree in chemical engineering in June of 2005. After serving as emergency medical technician in the Fire Department of Taichung County in Taichung Taiwan for 17 months, he joined the Department of Chemical Engineering at the University of Florida in the Fall of 2007. In worked to complete his doctoral research on exten ded ophthalmic drug delivery by silicone hydrogel contact lens.