Nanostructured Contact Lenses for Drug Delivery and UV Blocking

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Title:
Nanostructured Contact Lenses for Drug Delivery and UV Blocking
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1 online resource (154 p.)
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english
Creator:
Jung, Hyun-Jung
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University of Florida
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Gainesville, Fla.
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Thesis/Dissertation Information

Degree:
Doctorate ( Ph.D.)
Degree Grantor:
University of Florida
Degree Disciplines:
Chemical Engineering
Committee Chair:
Chauhan, Anuj
Committee Members:
Svoronos, Spyros
Jiang, Peng
Plummer, Caryn

Subjects

Subjects / Keywords:
contactlens -- egdma -- emulsion -- hema -- hydrogel -- inserts -- nanoparticles -- pgt -- timolol -- uvblocking
Chemical Engineering -- Dissertations, Academic -- UF
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Chemical Engineering thesis, Ph.D.
bibliography   ( marcgt )
theses   ( marcgt )
government publication (state, provincial, terriorial, dependent)   ( marcgt )
born-digital   ( sobekcm )
Electronic Thesis or Dissertation

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Abstract:
Our approach focused on dispersing highly crosslinked triggered nanoparticles in polymeric contact lenses to increase the duration of drug release from 1-2 hours to about 1-2 months. The nanoparticles were prepared by emulsion polymerization of a monomer with multivinyl functionalities of EGDMA (ethylene glycol dimethacrylate) and PGT (propoxylated glyceryl triacylate). In this study, glaucoma drug used, which is the base form of timolol was added to the monomer and was trapped in the nanoparticles.  The nanoparticles were about 3.5 nm in size and encapsulated 48-66% of the drug depending on the composition. The drug loaded particles were then dispersed in various polymers suitable for contact lenses including hydroxy methyl methacrylate (HEMA) and silicone hydrogel, which are common contact lens materials. In addition, Chapter 3 focused on developing a cylindrical (1-mm diameter 7.5 mm long) insert that can be inserted in the fornix for extended release of glaucoma drug timolol. Furthermore, this study has developed a novel approach of incorporating UV blocking feature into contact lenses by dispersing nanoparticles in contact lenses.  The nanoparticles encapsulate a UV blocking compound which is trapped in the particles due to the very high crosslinking.  The nanoparticles are loaded in silicone hydrogels by adding the particles to the polymerization mixture, followed by thermal or extended UV curing.  Also particles are loaded into polymerized silicone hydrogels and commercial contact lenses by soaking in a solution of particles in ethanol.  The conclusion of this study provide evidences that HEMA and silicone hydrogel loaded with nanoparticles could be very useful for extended delivery of ophthalmic drugs and that addition of highly crosslinked nanoparticles to polymers could be an effective method for manipulating release profiles with minimal impact on bulk properties such as modulus, transparency and water content. In addition, UV blocking compound loaded highly crosslinked particles can be loaded into contact lenses either before or after polymerization to prepare lenses with excellent UV blocking characteristic.
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In the series University of Florida Digital Collections.
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Includes vita.
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Includes bibliographical references.
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Description based on online resource; title from PDF title page.
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This bibliographic record is available under the Creative Commons CC0 public domain dedication. The University of Florida Libraries, as creator of this bibliographic record, has waived all rights to it worldwide under copyright law, including all related and neighboring rights, to the extent allowed by law.
Statement of Responsibility:
by Hyun-Jung Jung.
Thesis:
Thesis (Ph.D.)--University of Florida, 2012.
Local:
Adviser: Chauhan, Anuj.
Electronic Access:
RESTRICTED TO UF STUDENTS, STAFF, FACULTY, AND ON-CAMPUS USE UNTIL 2013-08-31

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lcc - LD1780 2012
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UFE0044533:00001


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1 NANOSTRUCTURED CONTACT LENSES FOR DRUG DELIVERY AND UV BLOCKING By HYUN JUNG JUNG A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY UNIVERSITY OF FLORIDA 2012

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2 2012 Hyun Jung Jung

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3 To my love husband, Seung Hwan, my parents, and all my family

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4 ACKNOWLEDGMENTS First of all, I thank my advisor and supervisory committee chair, Dr. Anuj Chauhan, for his continuous guidance and encouragement on my research. It would have been impossible to finish this dissertation without his support, patience and guidance throughout these years. He has given me lots of heartwarming advice both professionally and p ersonally. He is the best advisor that I have ever had and, at the same time, he is a great friend of mine. I also wish to extend many thanks to my other doctoral committee members, Dr. Spyros Svoronos, Dr. Peng Jiang, and Dr. Caryn Plummer for their insig htful viewpoints and willingness to participate in my doctoral review process. In addition, I would like to thank Dr. Yiider Tseng for the opportunity to serve as a teaching assistant under his guidance and his support. I also have had excellent group memb ers; Dr. Jinah Kim was especially helpful and an excellent mentor during my first semester in the lab, as she taught me many important lab procedures and techniques, as well as the consequence of efficiency. I thank Dr. Yash Kapoor, Dr. Brett Howell and Dr Chhavi Gupta for their invaluable assistance with my research. Also, I am very much thankful to Dr. Chen Chun Peng for being extremely caring for and considerate of the lab members. Lokendra Bengani has been an excellent friend and a great support in the last stages of my research. I would like to thank Michelle Abou Jaoude for her help in the timolol delivery experiments and making nanoparticles by using several different crosslinker s. I also thank the current group members, Sam Gause, Rob Damitz, Vincent Hsu for spending my last year with me in Gainesville. Lastly, I thank my parents, my sister and brother who always supported me with their love and prayer throughout my life, and my lovely husband, Seung Hwan Hong,

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5 who made me survive through the PhD program. I am grateful to God (who is always with me) for His sincere love and protection.

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6 TABLE OF CONTENTS page ACKNOWLEDGMENTS ................................ ................................ ................................ .. 4 LIST OF TABLES ................................ ................................ ................................ ............ 9 LIST OF FIGURE S ................................ ................................ ................................ ........ 11 LIST OF ABBREVIATIONS ................................ ................................ ........................... 14 ABSTRACT ................................ ................................ ................................ ................... 15 CHAPTER 1 INTRODUCTION ................................ ................................ ................................ .... 17 2 TIMO LOL LADEN NANOPARTICLES LOADED IN HEMA HYDROGEL ............... 26 2.1 Materials and Methods ................................ ................................ ...................... 27 2.1.1 Materials ................................ ................................ ................................ .. 27 2.1.2 Preparation of Highly Crosslinked PGT and EGDMA Nanoparticles ....... 27 2.1.3 Preparation of Nanoparticle Laden pHEMA Gels ................................ .... 28 2.1.4 Preparation of Highly Crosslin ked PGT with Surfactant Nanoparticles ... 28 2.1.5 Particle Size Distribution ................................ ................................ .......... 29 2.1.6 Equilibrium Water Content (EWC) ................................ ........................... 29 2.1.7 Transmittance ................................ ................................ .......................... 29 2.1.8 Me chanical Properties ................................ ................................ ............. 30 2.1.9 Drug Release from the Nanoparticles ................................ ...................... 30 2.1.10 Drug Release from the Nanoparticles Laden pHEMA Gels ................... 30 2.1.10 Drug Release Rates from the Nanoparticles Laden pHEMA after Packaging ................................ ................................ ................................ ..... 31 2.2 Results and Discussion ................................ ................................ ..................... 31 2.2.1 Particle Size Distribution ................................ ................................ .......... 31 2.2.2 Proposed Mechanism for Particles Formation ................................ ......... 32 2.2.3 Drug Release from Particles ................................ ................................ .... 33 2.2.4 Optical Clartity of PGT Nanoparticle Loaded pHEMA Gels ..................... 34 2.2.5 Mechnical Properties of PGT Nanoparticle Loaded pHEMA Gels ........... 34 2.2.6 Equilibrium Water Content of PGT Nanoparticle Loaded pHEMA Gels ... 35 2.2.7 Timolol Release of PGT Nanoparticle Loaded pHEMA Gels ................... 35 2.2.8 Effect of Timolol/PGT Ratio on Drug Release Profiles of PGT Nanoparticle Loaded pHEMA Gels ................................ ............................... 37 2.2.9 Effect of Temperatures on Drug Release Profiles of PGT Nanoparticle Loaded pHEMA Gels ................................ ................................ .................... 38 2.2.10 Release Mechanisms of PGT Nanoparticle Loaded pHEMA Gels ........ 38 2.2.11 Effect of Packaging of PGT Nanoparticle Loaded pHEMA Gels ............ 42 2.2.12 Effect of Surfactant Incorporation on Nanoparticle Laden Gels ............. 44

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7 2.2.13 Comparison of Lenses Loaded with EGDMA and PGT Particles .......... 45 2.3 Conclusion ................................ ................................ ................................ ........ 46 3 EXTENDED RELEASE OF NANOPARTICLE LOADED SILICONE HYDROGEL CONT ACT LENSES ................................ ................................ ............................... 71 3.1 Materials and Methods ................................ ................................ ...................... 71 3. 1.1 Materials ................................ ................................ ................................ .. 71 3.1.2 Preparation of Nanoparticles and Silicone Hydrogels .............................. 72 3.1.2.1 Preparation of drug containing PGT Nanoparticles ........................ 72 3.1.2.2 Preparation of Nanoparticle laden silicone hydrogels by adding particles to the polymerization mixture ................................ .................... 72 3.1.2.3 Loading Timolol PGT Nanoparticles into commercial contact lenses ................................ ................................ ................................ ..... 73 3.1.3 Particle Size Distribution ................................ ................................ .......... 74 3.1.4 Transmittance ................................ ................................ .......................... 74 3.1.5 Equilibrium Water Content (EWC) ................................ ........................... 74 3.1.6 Mechanical Properties ................................ ................................ ............. 74 3.1.7 Ion Permeability ................................ ................................ ....................... 75 3.1. 8 Oxygen Permeability ................................ ................................ ............... 75 3.1.9 Drug Release from the Nanoparticles Laden Silicone Hydrogels ............ 76 3.1.10 Drug Release Rates from the Nanoparticles Laden Silicone Hydrogels after Packaging ................................ ................................ ............ 76 3.1.11 In vitro and In vivo studies from Nanoparticles Loaded Commercial Lesnes ................................ ................................ ................................ ........... 76 3.1.11.1 In vitro Timolol PGT Nanoparticles into commercial contact lenses ................................ ................................ ................................ ..... 76 3.1.11.2 In vivo pharmacodynamics studies in Beagle dogs ...................... 77 3.2 Results and Discussion ................................ ................................ ..................... 78 3.2.1 Equilibrium Water Content (EWC) ................................ ........................... 78 3.2.2 Transmittance ................................ ................................ .......................... 78 3.2.3 Mechanical Properties ................................ ................................ ............. 79 3.2.4 Drug Release from the Nanoparticles Laden Silicone Hydrogels ............ 80 3.2.5 Effect of Temperatures on Drug Release Profiles ................................ ... 80 3.2.6 Effect of Packaging ................................ ................................ .................. 82 3.2.7 Maximizing Drug Loading in Nanoparticles ................................ ............. 84 3.2.8 Drug Release from Nanoparticles Loaded Commercial Lens .................. 85 3.2.9 Safety and Pharmacodynamic Efficacy of the Nanoparticles Loaded Commercial Lenses ................................ ................................ ...................... 86 3.3 Conclusion ................................ ................................ ................................ ........ 87 4 EXTENDED RELEASE OF TIMOLOL FROM NANOPARTICLE LOADED FORNIX INSERT FOR GLAUCOMA THERAPY ................................ .................... 99 4.1 Materials and Methods ................................ ................................ ...................... 99 4. 1.1 Materials ................................ ................................ ................................ .. 99 4.1.2 Preparation of Highly Crosslinked PGT Nanoparticles .......................... 100

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8 4.1.3 Preparation of Nanoparticles Laden Inserts ................................ .......... 100 4.1.4 Drug Release Experiments ................................ ................................ .... 101 4.1.5 Packaging Tests ................................ ................................ .................... 101 4.2 Results and Discussion ................................ ................................ ................... 102 4.2.1 Drug Release from Control p HEMA Inserts ................................ .......... 102 4.2.2 Effect of Timolol PGT particles Loading in Inserts on Drug Release ..... 102 4.2.3 Comparison of Release Rates from Inserts to Therapeutic Doses ........ 103 4.2.4 Effect of Timolol Loading in the Particles on Drug Release ................... 104 4.2.5 Mechanism of Release ................................ ................................ .......... 104 4.2.6 Effect of Packaging ................................ ................................ ................ 105 4.3 Conclusion ................................ ................................ ................................ ...... 105 5 UV BLOCKIN G NANOPARTICLE LOADED IN SILICONE HYDROGLE CONTACT LENSES ................................ ................................ ............................. 113 5.1 Materials and Methods ................................ ................................ .................... 114 5.1.1 Materials ................................ ................................ ................................ 114 5.1.2 Preparation of Highly Crosslinked PGT Nanoparticles with DP ............. 114 5.1.3 Preparation of Silicone gels ................................ ................................ ... 11 5 5.1.4 Preparation of Nanoparticle Loaded in Silicone Gels ............................ 116 5.1.5 Particles Added to Polymerization Mixture ................................ ............ 116 5.1.6 Particles Loaded by Soaking the Contact Lens in a Solution of Particles in Ethanol ................................ ................................ ..................... 117 5.1.7 Transmittance Measurements ................................ ............................... 117 5.2 Results and Discussion ................................ ................................ ................... 118 5.2.1 Incorporation of Particles into Lenses by Adding Particles to the Polymerization Mixture ................................ ................................ ................ 118 5.2.1.1 Effect of Particle Loading on UV blocking ................................ .... 118 5.2.2 Incorporation of Particles into Lenses by Soaking the Lenses in Solution of Particles in Ethanol ................................ ................................ .... 120 5.2.2.1 Mass of Particles Loaded ................................ ............................. 120 5.2.2.2 Effect o f UV Blocking Particles on Transmittance Spectra of the Lenses ................................ ................................ ................................ .. 121 5.2.2.3 Effect of Autoclaving ................................ ................................ .... 123 5.2. 3 UV Absorption in Commercial Contact Lenses Loaded with DP Containing Nanoparticles ................................ ................................ ............ 124 5.3 Conclusion ................................ ................................ ................................ ...... 124 6 CONCLUSION ................................ ................................ ................................ ...... 142 LIST OF REFERENCES ................................ ................................ ............................. 146 BIOGRAPHICAL SKETCH ................................ ................................ .......................... 154

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9 LIST OF TABLES Table page 2 1 Zero frequency storage modulus of pHEMA hydrogels loaded with nanoparticles ................................ ................................ ................................ ...... 68 2 2 Equilibrium water content (EWC) of HEMA hydrogels loaded with nanoparticles ................................ ................................ ................................ ...... 68 2 3 Dependency of total mass of drug released and the rate constant on temperature 100m gels ................................ ................................ ..................... 69 2 4 Dependency of total mass of drug released and the rate constant on temperature 200 m gels ................................ ................................ .................... 69 2 5 Effect of the composition of the particles on encapsulation efficiency ................ 70 4 1 Dependency of total mass of drug released (M ) and the rate constant at different particle loading. ................................ ................................ .................. 112 5 1 Experimental design of particles in polymerization mixture .............................. 135 5 2 Experimental design of soaking the lenses in solution of particles in ethanol ... 135 5 3 The percentage of UVC (below 280 nm) blocking range in nanoparticle laden silicone gels ................................ ................................ ................................ ...... 136 5 4 The percentage of UVB (280 315 nm) blocking range in nanoparticle laden silicone gels ................................ ................................ ................................ ...... 136 5 5 The percentage of UVA (315 380 nm) blocking range in nanoparticle laden silicone gels ................................ ................................ ................................ ...... 137 5 6 The percentage of UVA (315 400 nm) blocking range in nanoparticle laden silicone gels ................................ ................................ ................................ ...... 137 5 7 The weight changes of silicone lenses after soaked in 10.2% DP in nanoparticles with ethanol solutions ................................ ................................ 138 5 8 The weight changes of commercial contact lenses after autoclaving: 5.9 % of UV blocking dye loading in particles and then 2% particles in EtOH ................ 138 5 9 The percentage of UVC (below 280 nm) blocking range in nanoparticle laden silicone gels ................................ ................................ ................................ ...... 139 5 10 The percentage of UVB (280 315 nm) blocking range in nanoparticle laden silicone gels ................................ ................................ ................................ ...... 139

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10 5 11 The percent age of UVA (315 380 nm) blocking range in nanoparticle laden silicone gels ................................ ................................ ................................ ...... 140 5 12 The percentage of UVA (315 400 nm) blocking range in nanoparticle laden silicone gels ................................ ................................ ................................ ...... 140 5 13 The percentage of UV blocking range in 19.6% DP nanoparticle in Nigh&Day and Pu reVision ................................ ................................ ................................ 141

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11 LIST OF FIGURES Figure page 1 1 Schematic illustration of the nanoparticle laden contact lens inserted in the eye ................................ ................................ ................................ ...................... 25 2 1 Molecular structures of Timolol, PGT, and EGDMA ................................ ........... 49 2 2 Particle size distribution of PGT nanoparticles dispersed in ethanol and aqueous solutions with 10% pluronic F127 ................................ ........................ 50 2 3 Schematic illustration of the mechanism of formation of highly crosslinked nanoparticles due to the presence of diluents ................................ .................... 51 2 4 Cumulative drug release from the nanoparticles in a diffusion cell. Data is shown as mean std (n = 3) ................................ ................................ .............. 52 2 5 Storage modulus of gels loaded with PGT nanoparticles prepared with or without surfactant. Data is shown as mean std (n = 3) ................................ .... 53 2 6 Cumulative drug release from 100 and 200 m thick PGT nanoparticle laden pHEMA gels. Data is shown as mean std (n = 3) ................................ ... 54 2 7 Cumulative drug release from 100 and 200 m thick pHEMA gels without nanoparticles ................................ ................................ ................................ ...... 55 2 8 Cumulative drug release from 100 and 200 m thick gels at room temperature Data is shown as mean std (n = 3) ................................ .............. 56 2 9 Effect of temperature on drug release from the nanoparticle loaded gels. The solid lines are fits to the first order reaction model ................................ .............. 58 2 10 Arrhenius fit between the rate constant and temperature. The gel thickness for each case is 200 m. Data is shown as mean std (n = 3) .......................... 61 2 11 Potential mechanism for the formation of the ester link between timolol and PGT polymer ................................ ................................ ................................ ...... 62 2 12 Dependence of the total mass released from the gel per unit weight of the gel on 1/T. The y intercept represents the mass of drug loaded into the gels ... 63 2 13 Amount of timolol release from 100 and 200 m thickness after packaging. Data is shown as mean std (n = 3) ................................ ................................ .. 64 2 14 Cumulative release profiles from gels loaded with particles prepared in presence of F127 surfactant. Data is shown as mean std (n = 3) .................... 66

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12 2 15 Cumulative release profiles from pHEMA gels with 21.6% EGDMA particles. Data is shown as mean std (n = 3) ................................ ................................ .. 67 3 1 Transmittace spectra and photo graphic image ................................ .................. 89 3 2 Mechanical properties of gels loaded with PGT nanoparticles prepared and PGT nanoparticles including Vitamin E ................................ .............................. 90 3 3 Cumulative drug release from 100 and 200 m with 16.7% PGT nanoparticle and 20% Vitamin E laden silicone hydrogel ................................ ........................ 91 3 4 Cumulative drug release from 100 and 200 m thick PGT nanoparticle laden silicone hydrogel. Data is shown as mean std (n = 3) ................................ ...... 92 3 5 Effect of temperature on drug release from the nanoparticle loaded gels. The solid lines are fits to the first order reaction model ................................ .............. 93 3 6 Amount of timolol release from 100 and 200 m thickness after packaging ..... 94 3 7 Cumulative drug release from 5% PGT nanoparticle with timolol 1000 mg laden silicone hydrogel. Data is shown as mean std (n = 3) ............................ 95 3 8 Drug release at room temperature from Acuve Oasys lenses loaded with particles by soaking the lenses in a 3% solution of p articles in ethanol .............. 96 3 9 Drug release at room temperature after packing particle loaded Acuve Oasys lenses in refrigerator for 2 weeks. Data is shown as mean std (n = 3) ............ 97 3 10 Pharmacodynamic effect of extended wear of the particle loaded Acuve Oasys lenses. The in vitro drug release data at about 37 0 C .............................. 98 4 1 Image of the nanopaticle laden fornix insert ................................ ..................... 107 4 2 Cumulative drug release profiles from control pHEMA (without nanoparticles) inserts. Data is shown as mean std (n = 3) ................................ .................... 108 4 3 Cumulative drug release profiles from nanoparticle loaded inse rts .................. 109 4 4 Cumulative drug release profiles from nanoparticle loaded inserts. The two curves correspond to different drug loading s in the particle ............................. 110 4 5 Cumulative drug release profiles from nanoparticle loaded inserts with 25% particle loading after 3 mont h packaging at 4 C ................................ .............. 111 5 1 Transmittance (%) spectra for 100 m thick gels loaded with particles by direct addition to the polymerization mixture ................................ .................... 127 5 2 Transmittance (%) spectra for 100 m thick gels loaded with particles by soaking in a solution of ethanol with particles ................................ ................... 130

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13 5 3 Effect of autoclaving on transmittance of gels loaded with particles contained 15.8% DP by soaking in a solution of particles in ethanol ................................ 133 5 4 Transmittance spectra from commercial lenses of the particles contained 19.6% DP ................................ ................................ ................................ ......... 134

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14 LIST OF ABBREVIATION S AIBN Azobisisobutylonitrile BP Benzoyl peroxide Darocur TPO 2,4,6 Trimethylbenzoyl diphenyl phophineoxide DI Deionized DMA N, N D imet hylac rylamide DP 1,3 Diphenyl 1,3 propanedione EGDMA Ethylene glycol dimethacrylate EWC Equilibrium Water Content HEMA Hydroxyethyl methacrylate IOL Intraocular lens IOP Intraocular pressure Macromer Mega ( methacryloxypropyl) polydimethylsiloxane MAA Methacrylic acid MMA MethylMethacrylate NVP 1 vinyl 2 pyrrolidone OD Treated right eye OS Untreated left eye PBS Phosphate buffered saline PGT P ropoxylated glyceryl triacylate TRIS 3 Methacryloxypropyl tris(trimethylsiloxy)silane UV Ultraviolet

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15 Abstract of Dissertation Presented to the Graduate School of the University of Florida in Partial Fulfillment of the Requirements for the Degree of Doctor of Philosophy NANOSTRUCTURED CONTACT LENSES FOR DRUG DELIVERY AND UV BLOCKING By Hyun Jung Jung August 2012 Chair: Anuj Chauhan Major: Chemical Engineering Our approach focused on dispersing highly crosslinked triggered nanoparticles in polymeric contact lenses to increase the duration of drug release from 1 2 hours to about 1 2 months. The nano particles were prepared by emulsion polymerization of a monomer with multivinyl functionalities of EGDMA (ethylene glycol dimethacrylate) and PGT (propoxylated glyceryl triacylate). In this study, glaucoma drug used which is the base form of timolol was added to the monomer and was trapped in the nano particles. The nanoparticles were about 3.5 nm in size and encapsulated 48 66% of the drug depending on the composition The drug loaded particles were then dispersed in various polymers suitable for contact lenses including hydroxy methyl methacrylate (HEMA) and silicone hydrogel which are common contact lens materials. In addition, Chapter 3 focused on developing a cylindrical (1 mm diameter 7.5 mm long) insert that can be inserted in the fornix for extended release of glaucoma drug timolol. Furthermore this study has developed a novel approach of incorporating UV blocking feature into contact lenses by dispersing nanoparticles in contact lenses. The nanoparticles encapsulate a UV blocking compound which is trapped in the particles due to the very high crosslinking. The nanoparticles are loaded in silicone hydrogels by

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16 adding the particles to the polymerization mixture, followed by thermal or extended UV curing. A lso particles are loaded into polymerized silicone hydrogels and commercial contact lenses by soaking in a solution of particles in ethanol. The conclusion of this study provide evidences that HEMA and silicone hydrogel loaded with nanoparticles could be very useful for extended delivery of ophthalmic drugs and that addition of highly crosslinked nanoparticles to polymers could be an effective method for manipulating release profiles with minimal impact on bulk properties such as modulus, transparency and water content. In addition, UV blocking compound loaded highly crosslinked particles can be loaded into contact lenses either before or after polymerization to prepare lenses with excellent UV blocking characteristic.

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17 CHAPTER 1 INTRODUCTION Ophthalmic diseases are becoming increasingly prevalent due to aging of the population in the United States and several other countries. Typically, most ocular diseases for the anterior eye are treated by eye drops which have a low bioavailability for the corneal tissue because of the tear turnover non productive absorption in conjunctiva and nasal cavity and impermeability of corneal epithelium [1,2]. Such po or bioavailability results in only a small amount of the drug dosage, less than 5% of the initial dose being absorbed into the eye, with even a further smaller amount reaching target tissues In order to increase bioavailability, a large number of researc h groups have explored several approaches including developing novel eye drop formulations such as in situ forming gels [ 3 ] based on pH [ 4 7 ], temperature [ 8 10 ], and ionic [1 1 ,1 2 ] trigg ing colloidal particles and collagen shields [1 3 ,1 4] However, most o f these modifications prolonged the drug resident time to only a few hours and minimal ly increase d bioavailability. Several groups have also focused on using contact lenses for delivering ophthalmic drugs. Contact lenses can increase the residence time o f the drug to more than 30 min [ 1 5 ,1 6] which is significantly longer than the residence time of drugs delivered via eye drops [17]. The increased residence time leads to an increase in bioavailability to potentially as high as 50%, which will lead to reduc ed drug wastage and side effects [18]. There have been a number of attempts in the past to use contact lenses for ophthalmic drug delivery based on soaking hydrophilic lenses in a drug solution followed by insertion into the eye [19 24]. The major problem of loading drug by this

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18 method is that the loaded drug diffuses out in a very short time of a few hours, which is inadequate for extended drug delivery applications. To address this issue Gulsen and Chauhan [25 27] and Kapoor and Chauhan [28,29] have pro posed incorporating microemulsions, liposomes, or micelles into the lenses, and a number of researchers 34]. While the approaches listed above are effective at increasing the release duration from contact lenses, the lenses are still not suitable for extended release lasting a week or longer. Also these soft particles are destabilized during extended storage and lens processing. In a recent study, researchers sandwiched a lay er of drug (ciprofloxacin) containing PLGA polymer in between p HEMA sheets as a mimic of a contact lens [35]. The system delivered the drug for longer than a month but suffers from several deficiencies including high thickness (1 mm), low oxygen and ion permeability due to thickness and the PLGA layer, loss of transparency in the peripheral region which contains PLGA, etc. Additionally, these lenses cannot be autoclaved or packaged in PBS because of PLGA degradation. To summarize, most of the approaches listed above are unable to prevent drug loss from the lenses during packaging and the drug release duration is shorter than the desired duration of 1 2 weeks. This research is focusing on developing a lens for glaucoma therapy because of the growing number of glaucoma patients worldwide. G laucoma is an ocular disease characterized by an increase in the intraocular pressure (IOP) and eventual damage of the optic nerve and loss of vision The IOP is set by a balance between production of the aqueous humor fluid in the eye and its outflow through the trabecular and the uveoscleral pathways. Glaucoma sufferers are prescribed medications that reduce the

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19 IOP by either reducing the productio n of the aqueous humor or by increasing its outflow. Timolol a non selective beta blocker is one such medication that decreases the IOP by reducing the production through blocking of the beta receptors in the ciliary body. A contact lens that improves vis ion and also delivers glaucoma drugs may lead to high compliance for glaucoma patients that also require vision correction. Timolol which is the pilot drug in this study is commonly used in glaucoma treatment because it is a non selective beta blocker t hat decreases the intraocular pressure of aqueous humor fluid inside the eye. Timolol is known to cause toxic side effects such as arrhythmia and thus it is an excellent candidate for delivery by contact lenses because of the potential for reduced side ef fects. Our specific focus in Chapter 2,3 is to design a contact lens that will not release drug during packaging but will be triggered to release dr ug after insertion in the eye. For this study, o ur approach is based on encapsulating timolol in nanoparticl es and dispersing the particles in contact lenses. Several techniques have been explored in details for preparing nanoparticles such as ionotropic gelation, solvent evaporation, microemulsion, and emulsion. Our approach is based on a modification of the e mulsion based method for encapsulation of drugs, which is commonly used in several applications [36 4 5 ]. Emulsion based methods typically produce microparticles unless a very large amount of surfactant is used to create nanosiozed drops. Microparticles a re not suitable for incorporation in contact lenses because of the critical requirement of optical clarity. Also, a significant amount of surfactant may not be desirable in lenses as it could cause toxicity. To address these requirements, we here propose a novel method for preparing nanoparticles, described in details Chapter 2 in pHEMA hydrogel

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20 To ensure a long release duration from the nanoparticles, we prepare particles by polymerizing monomers with mutivinyl functionalities to create highly crosslin ked nanoparticles. As illustr ated in Figure 1 1, i f drug laden nanoparticles loaded contact lenses are placed on the eye, the d rug is expected to diffuse through the lens matrix, and enter the tear film, which is the thin tear film trapped between the cor nea and the lens. Furthermore, in Chapter 3 our aim is to develop a particle loaded silicone hydrogel materials that can be used as extended wear contact lenses, while also providing extended drug release at therapeutic rates. To accomplish this objective we dispersed timolol loaded PGT nanoparticles in silicone hydrogel contact lenses. The nanoparticle loaded silicone hydrogel lenses are characterized to explore drug release profiles and all properties relevant to extended wear contact lenses including t ransparency, modulus, and ion and oxygen permeabilities. In addition in vivo animal studies are conducted with Beagle dogs to establish safety and efficacy of glaucoma therapy by extended wear of nanoparticle loaded contact lenses. To our knowledge, the system developed here is the longest releasing extended wear contact lens for glaucoma therapy, and its pharmacodynamics efficacy is also proven through preliminary in vivo studies in Beagle dogs. Furthermore, s everal researchers have developed fornix ins erts [46 50] and Ocusert and Lacrisert have been commercialized to treat glaucoma and dry eyes, respectively. Lacrisert is a cylindrical insert 3.5 mm long and 1.27 mm in diameter made of hydroxypropyl cellulose [51] The insert dissolves over a period of a day after insertion leading to increased tear viscosity and lubrication. Thus, t he inserts proposed

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21 Chapter 4 were chosen to be geometrically similar to lacrisert and thus have a length ranging of about 7 mm and a length of about 1 mm. While retention of the fornix inserts is typoically a concern, cylindrical shaped inserts are considered best for retention in the conjunctival sac [52 54] The materials for designing the inset were chosen to be HEMA and PGT, which a re similar to the materials commonly used in ocular applications such as contact lenses Finally, in Chapter 5 we demonstrate the UV blocking nanoparticles loaded in contact lenses from silicone hydrogel. Currently, the U.S. Food and Drug administration ( FDA) has established the standards for UV blocking contact lenses based on American National Standards Institute (ANSI) standards. Based on these standards, the UV blocking contact lenses have been classified in two categories (Class 1 and Class 2) depend ing on the extent of the protection. Class 1 lenses must block more than 90% of UVA and 99% of UVB (280 315nm) radiation. The lenses in Class 2 category must block more than 70% of UVA and 95% of UVB radiation. The UVA radiation corresponds to the wavele ngth range of 316 400 nm. However, only wavelengths from 316 380 nm are considered in the ANSI standards for determining the classification of a contact lens. The health benefits of UV blocking in contact lenses has been well recognized, yet only three co mmercial contact lenses (ACUVUE Oasys ACUVUE Advnace and ACUVUE Advance for Astigmatism ) are categorized as Class 1 blocking lenses [55,56]. Exposure to the ultraviolet radiation from sun is known to have the potential to cause significant damage to the body including skin irritation and burning to serious diseases such as skin cancer. Environmental damage to the ozone layer has further

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22 exacerbated the damaging potential of exposure to sun. It has become a very common practice to use UV blocking lot ions and creams to minimize the damage from sun exposure. While the general population is well aware of the potential for damage to skin from UV radiation, there is less awareness of the possibility of damage to other organs, particularly eyes. The UV ra diation can cause mild irritation and a foreign body sensation in the eyes, and regular exposure can cause far more serious problems such as snow blindness, cataracts and rarely cancer of cornea or conjunctiva [5 7 60 ], and also photokeratitis, erythema of the eyelid, solar retinopathy, and retinal damage. The damage from the UV radiation is likely due to creation of free radicals that can cause protein modification and lipid peroxidation [61 ]. The intraocular lens in an adult eye filters out a majority of the UV light. The UV transmittance decreases with age and by the age of 25, the lens absorbs UV light almost completely. The accumulated exposure to UV light before the age of 25 could cause significant retinal damage [61 ]. The potential for retinal damage due to UV exposure is even higher in Aphakic patients, i.e., the patients that have lost the natural IOL. In addition to the retina damage, UV exposure is considered to be a potential risk factor for cataract, which is the leading cause of blindness in the world. Exposure of the ocular tissue to the UV radiation can be minimized by wearing glasses that block UV light. However, the extent of blocking depends on type of the lenses and a lso on the design of the sunglasses. Most sunglasses do not protect completely from the UV radiation because UV light can reach the eyes through the top, bottom and sides of the eyeglasses. The limitations of sun glasses could be overcome

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23 by wearing UV bl ocking contact lenses as contact lenses cover the entire cornea and thus can provide protection from light at all angles [61 ]. The UV blocking feature is typically incorporated into the contact lenses by adding a UV absorbing molecule to the lens composi tion. It is well known in the art that the absorption spectrum of a molecule depends on the molecular structure and a large number of molecules are known in literature to absorb UV radiation [62,63 ]. Also a number of patents are filed and issued for UV blocking contact lenses [6 4 69 ]. These patents focus on novel UV absorbers, or methods of producing a contact lens containing the absorbers. In all applications, the UV absorbing molecules are copolymerized with the polymer(s) used to manufacture the len s to eliminate any possible of leaching of the UV blocking agent either during processing steps after polymerization or during lens wear. The major challenges in preparing a contact lens loaded with the UV blocker are: (i) contact lenses are typically pol ymerized in molds by UV light, which gets absorbed by the UV blocking agent making the process of lens curing longer or requiring increased light intensity. Even with increased duration, or light intensity the properties of the final lens might be signifi cantly compromised: (ii) the presence of the UV blocking agent could also impact the kinetics of polymerization particularly if the molecule is loaded in appreciable amounts. Furthermore, the requirement of copolymerization of the UV blocking agent with t he lens matrix limits the choice of molecules, and the molecules that have the copolymerization feature may not be the most efficient UV absorbers. In the Chapter 5, we have developed a unique approach for incorporating UV blocking feature into the conta ct lenses which minimizes or eliminates some of the

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24 problems with the use of the current state of the art approaches. Our approach is based on incorporating UV blocking molecules in nanoparticles and dispersing the nanoparticles in the contact lenses. Th e nanoparticles can be incorporated into the contact lenses by addition to the polymerization mixture, followed by curing with UV light at suitable intensity and time to ensure curing. Alternatively, if the nanoparticles are prepared of a size sufficientl y small, the particles can be loaded into preformed contact lenses. The particles can be loaded into the lenses in a medium that significantly swells the lens to increase the pore size, allowing the particles to diffuse into the lenses. After particle lo ading, the lenses can be soaked in PBS to extract the loading medium and collapsing the pores in the lens matrix trapping the particles. This approach of loading the UV blockers post lens curing eliminates the problems associated with curing lens composit ions containing the UV blockers.

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25 Figure 1 1 Schematic illustration of the nanoparticle laden contact lens inserted in the eye

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26 CHAPTER 2 TIMOLOL LADEN NANOPA RTICLES LOADED IN HEMA HYDROGEL In this chapter we indicated to dispersing highly crosslinked triggered nano particles in polymeric contact lenses to increase the duration of drug release from 1 2 hours to about 1 2 months. The highly crosslinked particles were prepared by emulsion polymerization of a monomer w ith multivinyl functionalities of EGDMA (ethylene glycol dimethacrylate) and PGT (propoxylated glyceryl triacylate). A glaucoma drug, which is the base form of timolol was added to the monomer and was trapped in the na no particles. The timolol loaded nano particles were then dispersed in various polymers suitable for contact lenses including hydroxy methyl methacrylate (HEMA), which are common contact lens materials. This nanoparticle loaded in HEMA hydrogel proposed the mechanism of drug transport of hydrolysis of ester bonds that link timolol to the particle matrix which form during particle polymerization process. Additionally, the proposed mechanism of drug transport is a short time burst due to diffusion of drug from the bulk matrix, followed by an extended release drug from inside the particles. The mechanism was established by measuring transport in particle laden gels of various thicknesses. The duration of the burst phase is proportional to the square of the thic kness and the slow release is independent of the thickness. Most experiments in this chapter were done by loading nanoparticles into HEMA hydrogel to evaluate the properties resulted from the additional nanoparticles. Also, it had developed a hydrogel plat form for controlling bulk physical and transport properties of contact lenses and optimized the drug and particle loading amount into HEMA hydrogel matrix system.

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27 2 .1 Materials and Methods 2.1.1 Materials Hydroxyethyl methacrylate (HEMA) monomer, timolol maleate p luronic F127, ethylene glycol dimethacrylate (EGDMA) (PBS) w ere purchased from Aldrich Chemicals (St Louis, MO) P ropoxylated glyceryl triacrylate (PGT) was a gift from Sartomer Benzoyl peroxide (BP) (97% ) and methyl methacrylate (MMA) were purchased from Aldrich Chemicals (Milwaukee, WI) The initiator 2,4,6 Trimethylbenzoyl diphenyl phophineoxide ( Darocur TPO ) was kindly provided by Ciba (Tarrytown, NY). 2.1 2 Preparation of Highly Crosslinked PGT and E GDMA Nanoparticles The drug loaded highly crosslinked nanoparticles are prepared by emulsion polymerization of a mixture of the timolol base and the mutivinyl functionality monomers, which are typically called crosslinkers. The molecular structure of timo lol and the two monomers explored in this study are shown in Figure 2 1 The base form of timolol was prepared by a dd ing timolol maleate to 1.04 M NaOH (purged with nitrogen ). At such high pH, timolol maleate converts to the base form that is an oily liquid, which phase separates from the mixture The timolol base was pipetted from the bottom and 120 mg of the base was added to a mixture of 1.0 g of the crosslinker ( PGT or EGDMA ) and 7.5 mg of BP. The amount of base was increased to 360, 600, 800 and 1000 mg to prepare particles with a higher drug loading. The mixture was then added to 5 ml of DI water and then 1.65 ml of 2.08M NaOH w as added to the mixture Due to the high pH, timolol was converted almost completely to the hydrophobic base form, wh ich has a high solubility in the crosslinkers phase. The mixture was purged with nitrogen for 15 min and then heated in an 80 C hot water bath under stirring a t 1100 rpm for 8 hours The

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28 thermal polymerization results in formation of drug loaded nanoparti cles, which were separated from the suspension by centrifugation for 15 min 2.1 3 Preparation of Nanoparticle Laden pHEMA Gels The particle laden p HEMA gels were synthesized by free radical polymerization of the HEMA monomer mixed with the timolol loaded nanoparticles. Briefly, 1.45 ml of the HEMA monomer, 0.5 ml of DI water, 5 l of crosslinker (PGT or EGDMA ), and 0.4 g of the particle suspension were mixed in a vial, and purged with nitrogen for 15 minutes. Next, 3 mg of the photoinitiator ( Darocur TPO ) was added to the mixture under stirring for 5 min, and then the solution was poured into a mold that comprised of two glass plates separated with either 100 or 200 m thick plastic spacers. The molds were then placed on Ultraviolet transilluminiator UVB 10 (Ultra Lum, Inc.) and irradiated with UVB light (305 nm) for 40 min. This procedure led to formation of gels with a particle fraction of about 21.6% on dry basis. The amount of the particle suspension was reduced to 0.14 g and 0.07g to prepare gels with particle loadings of about 8.8 and 4.6%, respectively. The cured gels were cut into circular pieces of 1.65 cm diameter with a cork borer and dried in air overnight before further use. Control gels in which timolol was incorporated by direct addition to the polymerization mixture were prepared by mixing timolol base to a mixture of 1.45 ml of monomer HEMA, 5 l of crosslinker, and 0.5 ml DI water at a loading of 40m g /g The solution was purged by bubbling nitrogen for 15min and then polymerized by following the procedures described above. 2.1.4 Preparation of Highly Crosslinked PGT with Surfactant Nanoparticles The stability and particle sizes of emulsions can be contro lled by addition of emulsifiers, thereby impacting the size of the polymerized particles. To explore this

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29 issue, p luronic F127 was added to the emulsion of PGT in water Specifically, 0.1 g of p luronic F127 surfactant was added to the mixture of 120 mg of timolol base, 1 ml PGT and 7.5 mg of BP. The methods for particle preparation and incorporation of particles into the gels a re same as those described Section 2.1.2 and 2.1.3 2.1.5 Particle Size Distribution The particles were dispersed in ethanol at 5 m g/ml for the particle size measurement by dynamic light using Nanotrac Particle Size Analyzers ( Microtrac Inc. ) The data were analyzed using the microtrac FLEX application software program. Additionally, the particles were dispersed in 10% (w/w) of pluro nic F 127 surfactant solutions and particle sizes were measured. 2.1.6 Equilibrium Water Content (EWC) The dried lenses were weighed (W dry ) and then hydrated by soak ing in 3.5 ml of PBS for 24 hours The hydrated gels were weighed (W wet ) and the equilibriu m water content (EWC) was calculated by the following formula ( 2 1 ) 2.1.7 Transmittance The transmittance of nano particle laden p HEMA hydrogels was measured using UV Vis spectrophotometer (Thermospectronic Genesys 10 UV). The lenses were hydrated by soaking in PBS overnight, and then mounted on the outer surface of a quartz cuvette. The cuvette was placed in a spectrophotometer and the transmittance was measured at wavelengths ranging from 500 nm to 900 nm.

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30 2.1.8 Mechanical Properties Storage modulus of the gels w as measured using a dynamic mechanical analyzer (DMA Q800, TA instruments). H ydrated rectangular gel s 400 micron in thickness were mounted in the submersion tension clamp at room temperature. A prelo ad force applied 0.01N and force track was 115% used. Strain sweep tests were recorded to confirm the linear range at room temperature at 1 Hz Subsequently, frequency dependent storage modulus was obtained by fixing the strain within the linear range 2.1.9 Drug Release from the Nanoparticles The rates of drug release from the highly crosslinked nanoparticles were measured in diffusion cells. The donor compartment was filled with 18 ml of the nanoparticles dispersed in 10% w/w pluronic F 127 in PBS solu tion. The receiver compartment was filled with 32 ml of PBS and a 100 m thick HEMA gel was mounted in between the compartments. As a control experiment, drug without nanoparticles was dissolved in PBS solution in donor compartment. The drug concentrati on in the receiver was monitored periodically with a UV VIS spectrophotometer by measuring the absorbance at wavelength ranging from 252 nm to 321 nm 2.1.10 Drug Release from the Nanoparticles Laden pHEMA Gels The dried circular gels 1.65 cm in diameter w ere submerged in 200 ml DI water under minimal stirring of 140 rpm and at room temperature for 24 hours to extract the unreacted monomer. Next, the gels were soaked in a PBS solution and the dynamic drug concentration in the PBS was determined by measuring the absorbance at wavelengths ranging from 252 nm to 321 nm. The PBS volume was chosen to be 1.75 ml or 3.5 ml for the 100 m and 200 m thick gels, respectively. The temperature in the

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31 release medium was controlled through a temperature bath to vary fr om 25 to 100 0 C to explore the effect of temperature on drug release profiles. 2.1.10 Drug Release Rates from the Nanoparticles Laden pHEMA after Packaging Typi cally commercial contact lenses are packaged in blister packs that contain about 1 1.5ml solutio n such as PBS, which is tear mimic To simulate the packaging, gels were subjected to the monomer extraction as described above, and then stored in 1 ml of packaging solution (PBS) for durations ranging from 1 5 months at either room temperature or in a re frigerator at 4 C. After the storage in PBS, the gels were withdrawn, and the drug release profiles were measured by following the same procedures as described above. 2.2 Results and Discussion 2.2 1 Particle Size Distribution Since the nanoparticles are hydrophobic, these have to be either dispersed in liquids like ethanol or in water through addition of surfactants for the measurements of particle size distribution. Both aqueous solutions and ethanol solutions were transparen t suggesting that the particle size is less than the wavelength of visible light. Figure 2 2 shows the particle size distributions for timolol laden highly crosslinked nanoparticles dispersed in ethanol and dispersed in 10% (w/w) of pluronic F127 solution The mean particle size is about 3.5 nm in both mediums. Further experiments with varying surfactant concentrations in aqueous solutions showed no effect of the surfactant concentration on the measured particle size (data not shown). The size measuremen ts clearly prove that the size of the particles is uncorrelated to the size of the emulsion drops, and is instead controlled by the effect of the diluent on the polymerization dynamics. To further prove this hypothesis, experiments were

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32 conducted by remov ing the diluent (timolol base) from the polymerization mixture. The end product in this case was an agglomeration of large particles. Furthermore, some experiments were conducted in which surfactant was added to the emulsion to control the particle size. Typically, surfactant addition reduces the emulsion particle size, and consequently the size of the particles produced by emulsion polymerization. However, addition of surfactants of various types and in various loadings had no impact on the particle si ze, further proving that the particle size is controlled by the presence of the timolol base as the diluent. Section 2.2.2 we propose a hypothesis for the mechanisms that leads to formation of diluent encapsulated nanoparticles when a mixture of hydrophob ic oily monomer (crosslinkers in this case) and an oily diluent (timolol base) is polymerized. 2 2.2 Proposed Mechanism for Particles Formation In free radical polymerization, gelation occurs through formation and growth of nanogels that eventually grow an d merge to form a contiguous gel, or particles depending on the monomer concentration [ 70 72 ]. In the case of emulsion polymerization with a diluent, several nanogels form in each emulsion drop, and each encapsulate some diluent. The diluent timolol base has a higher solubility in the unpolymerized mixture compared to the nanogels, and thus the concentration of the diluents in the unpolymerized fraction increases with time. Eventually the diluent concentration is sufficiently high to prevent further growt h and merging of the nanogels, and thus each microdrop of the emulsion contains nanogels. Since the monomers (EGDMA or PGT) contains multivinyl groups, the nanogels are expected to form a highly crosslinked structure. These mechanisms are illustrated in a schematic in Figure 2 3

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33 This new preparation method requires a careful optimization of the type of diluents and its concentration. If the diluents concentration falls below a critical value, the nanogels in each drop merge and thus each microdrop poly merizes into a single drop, which is undesirable for our application. While a number of oils can be used as diluents we have successfully utilized vitamin E which is a well known nutraceutical [7 3 80 ], olive oil, and timolol base which is an important gla ucoma drug. When timolol base is used as the dilute, it serves the dual purpose of the diluent and the active drug which diffuses over a long period of time. As evident from the image, the Vitamin E loaded lenses are transparent irrespective of the Vitam in E loading, but attain a slightly yellowish color at high Vitamin E loadings. 2.2 3 Drug Release from Particles The nanoparticles were dispersed in surfactant solution and added to the donor compartment of the diffusion cell. The concentration of drug was then monitored in the receiving cell to determine the rate of drug release from the particles. Control tests w ere conducted to ensure that the particles did not cross the diffusion cell membrane. It was also established that the time scale for free drug to equilibrate between the donor and the receiving cells was substantially smaller than the time scale of relea se from the particles (data not shown). The % cumulative drug release profiles from the particles dispersed in the donor compartment are shown in Figure 2 4 The Figure 2 4 explains that t he amount of drug released was calculated by multiplying the conce ntration in the donor cell with the total fluid volume in the diffusion cell. The total amount of drug loaded into the particles was determined by multiplying the concentration of drug in the crosslinker (120mg/1g of PGT) by the mass of PGT dispersed in t he donor

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34 compartment. The data in Figure 2 4 shows that t he nanoparticles were dispersed in surfactant solutions in the donor compartment and the dynamic drug concentration in the receiving compartment was monitored. T he particles release drug for about a month with only 51.75% of cumulative release. A fraction of the remaining 48.25% was partitioned in the aqueous phase during the particle preparation or irreversibly trapped in the particles. These issues will be discussed later. 2.2.4 Optical Clartity of PGT Nanoparticle L oaded pHEMA Gels The clarity of the particle laden lenses was characterized by measuring the transmittance spectra of the 100 micron thick gels in the visible range from 500 to 900 nm. The transmittance of the particle laden gels and control HEMA gels is almost 100% in the entire range (data not shown). A photographic image of the particle laden gel also proves that the presence of particles does not reduce the visual clarity of the gel (data not shown). 2.2 5 Mechnical Properties of PGT Nanoparticle L oaded pHEMA Gels The frequency dependent storage modulus (G ) and loss modulus (G ) were obtained for nanoparticle laden p HEMA hydrogels. The effect of particle loading on the frequency dependent storage moduli for p HEMA gels is shown in Figure 2 5 and the values of the zero frequency storage modulus are listed in Table 2 1 The results clearly show that particle loading increases the storage modul us for the gels. The p HEMA gels with 21.6% particle loading have a zero frequency modul us of 4.01 MPa, which is about 4 times that of the gels without particles. The i ncreas e in modulus could be due to the presence of the nanoparticles and also partially to a high degree of crosslinking of the bulk due to incomplete reaction of the PGT duri ng the particle formation step. To explore the effect of particle loading on modulus, the storage

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35 modulus was also measured for a particle loading of 4.6% ( Figure 2 5 ). At 4.6% particle loading, the zero frequency storage modulus is 0.95 MPa, which is approximately equal to the value for p HEMA gels and also within the range of typical commercial lenses. 2.2 6 Equilibrium Water Content of PGT Nanoparticle L oaded pHEMA Gels The effect of particle addition on swelling of p HEMA hydro gels in PBS is shown in the Table 2 2 Since the particles are expected to not absorb water, the degree of swelling in PBS is expected to decrease in proportion to the amount of particles added. The p HEMA control gel swells 73 % in PBS. Thus, p HEMA gel with 21.6% particles is e xpected to swell by about 73 x (1 0.216) = 57.23%, which is about 23.96% higher than the measured value. This suggests that the presence of particles likely leads to increased crosslinking in the gels likely in the vicinity of the particles perhaps due to the dangling vinyl groups on the surface of the particles. The water content for a particle loading of 4.6% was determined to be 63.82 3.03%, which is within the range for commercial lenses. 2.2 7 Timolol Release of PGT Nanoparticle L oaded pHEMA Gels Th e cumulative drug release profiles from the particle loaded gels of two different thicknesses are shown by Figure 2 6 The release of drug from the particle loaded hydrogels could potentially be controlled either by the resistance in the particles or that in the gel. The rate controlling mechanism can easily be identified by conducting release experiments from gels with two different values of gel thickness ( h ). If the drug release is controlled by diffusion through the gel, the release time scales as h 2 and if the release is controlled by transport through the particles, the release time should be independent of h The release profiles in Figure 2 6 show that the drug rel ease duration is about the same for both 100 and 200 m thick gels proving that the rate controlling

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36 mechanism is transport of the drug molecules from inside the particles to the gel. The total amount of drug released from the thick gels is about twice t hat from the thin gels, which is expected due to the weight ratio of the two gels. To further prove that the release dynamics are controlled by particles, drug release experiments were conducted with HEMA gels in which drug was loaded by addition to the p olymerization mixture. The release profiles from these gels show that the release time is about 1 and 4 hours for the 100 and 200 m gels, respectively, which is substantially less than the drug release durations from the particle loaded lenses (Figure 2 7). It is noted that the particle loaded gels were subjected to an initial extraction step in which the gels were soaked in water for 24 hours. The gels loaded with highly crosslinked PGT nano particles lost about 19.97 % of the initially loaded drug in th e extraction stage. This fraction is presumably the drug that was present outside the nanoparticles. The drug release profiles for the subsequent release experiments are shown in Figure 2 6 in which the mass of drug released amount is plotted as a functi on of time for both the thin and the thick gels. The data in Figure 2 6 clearly shows that there is an extended drug release from the gel which lasts for about 1 month for both the 100 and 200 m gels, which is significantly longer than 1 or 2 hour drug re lease duration from the p HEMA gels without particles. The usual dose of timolol is one drop of 0.25% timolol maleate twice per a day, which amounts to a daily dosage of 125 g. It is reported that about 1 2% of timolol applied as eye drops reaches corne a, which suggests that the therapeutic requirement of timolol is about 2.5 g/day Both mathematical models and some clinical data suggest that about 50% of the drug delivered through contact lenses reaches cornea

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37 Furthermore, p HEMA based contact lenses can only be worn during the day due to the insufficient oxygen transport. Based on 50% bioavailability and 12 hour daily wear time, p HEMA lenses require a release of about 10 g/day from a contact lens. The 100 micron thickness particle laden HEMA gels rel ease timolol at a rate of about 7 g/day for 3 0 days, which is slightly smaller than the required release rate. 2 2.8 Effect of Timolol/PGT Ratio on Drug Release Profiles of PGT Nanoparticle L oaded pHEMA Gels Due to the increase in modulus on particle inc orporation, it is important to minimize the particle loading in the lenses while still loading adequate drug amounts for therapeutic effects. For fixed particle loading, the drug loading in the lenses can presumably be increased by increasing the drug loa ding in the particles. This hypothesis was tested by increasing the mass of timolol in the mixture of the drug and the crosslinker PGT, which was polymerized to prepare the particles. Specifically, the mass of timolol was increased from 120 mg to 360, 60 0, 800, and 1000 mg to obtain particles with increased drug loading. The mass of the crosslinker was kept fixed at 1 g in each case, and thus the ratio of the drug to PGT (T:PGT) was 0.12, 0.6, 0.8 and 1 for the four cases considered. It is noted that th e ratio of the drug and PGT in the polymerized particles could be slightly different than the ratios listed above because a fraction of timolol remains unencapsulated during particle formation. HEMA gels were then prepared with 4.6 and 8.8% particle loadi ng, which is about one fifth and two fifth of the particle loadings in the results described above. The drug release profiles from the gels with 4.6 and 8.8% particles but various drug s to PGT ratio are shown in Figure 2 8 A D for both the 100 m and 200 m gels. All gels release drug for about a month and the amount of drug release d from the thicker gels is about double of that from the

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38 thinner gels proving that but the time scales are unaffected by timolol and PGT ratios Furthermore, the mass of drug released from the gels with 4.6% loading is comparable to that from the gels with 21.6% particle loading containing particles that were prepared with a timolol to PGT ratio of 0.12. This proves that the cumulative drug release could be increased by increa sing the drug loading in the timolol PGT mixture. The gels with 4.6% particle loading have a zero frequency storage modulus of 0.95MPa, which is much less than that for the gels with 21.6% particle loading, and within the range of the modulus for commerci al contact lenses. The water content of the gels with 4.6% is 63.82 3.03 %, which is much larger than the water content of 33.27% for the gels with 21.6% particle loading. Thus the gels loaded with 4.6% particles are suitable for use as contact lenses. 2.2 9 Effect of Temperatures on Drug Release Profiles of PGT Nanoparticle L oaded pHEMA Gels The temperature dependence of the release rates is useful in understanding the rate limiting transport mechanisms, and so drug release experiments were performed for several temperatures ( 25, 40, 60, 80, and 100 C ) The effect of temperature on the drug release profiles shown in Figure 2 9 clearly proves that the drug release from the particle loaded lenses is highly temperature sensitive. An increase in temperat ure reduces the drug release duration and also increases the total amount of drug released from the lenses. 2.2 10 Release Mechanisms of PGT Nanoparticle L oaded pHEMA Gels The release profiles from the particle loaded lenses cannot be fitted to a diffusi on model (fits not shown) suggesting that diffusion through the particles is not the rate

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39 limiting step. Instead, the drug release data fits with the model for a first order process, i.e., (2 2 ) where M is the amount of drug released at a time t and M is the amount of drug released after infinite time, and k is the rate constant. The values of M and k can be determined by fitting the experimental data for drug release rates to the above equation. The best fit curves at each temperature are shown as the solid lines in Figure 2 9 A E, and the best fit values of M and k are shown in Table 2 2 for each te mperature, gel thickness and ratio of timolol to PGT in the initial mixture. The temperature dependency of the rate constant can be described by the Arrhenius equation (2 3 ) where A is the preexponential constant, E a is activ ation energy, R is the ideal gas constant, and T is the temperature in Kelvin. Based on the above equation, a plot of log(k) with 1/T, should be a straight line of slope E a /R The temperature dependence of the values of k obtained by fitting th e drug rel ease data for 200 m thick gels satisfy the above relationship as shown in Figure 2 10 The four sets of data correspond to particle loading, T:PGT ratios and 21.6% loading, T:PGT = 0.12, R 2 =0.9949, T:PGT ratios and 4.6% loading, T:PGT = 0.6, R 2 =0.9686, T:PGT = 0.8, R 2 =0.9677, T:PGT = 1.0, R 2 =0.9635 in Figure 2 10. The value of the activation energy obtained from the slopes of the graphs is 24.42 2.47 kJ/g mol. The excellent fit between the release rates and the first order model, along wit h the Arrhenius dependency of the rate constant suggests that the rate limiting step could be hydrolysis of an ester linkage that formed during the polymerization process. The

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40 potential mechanism for the ester formation is shown in Figure 2 11 The OH gr oup in timolol molecule carries a partial negative charge on oxygen in high pH condition, which can react with the polymeric network through nucleophilic reaction to form an ester, which then gets hydrolyzed during the drug release experiments. Several est ers of timolol have been synthesized for use as prodrugs and their rates of hydrolysis have been measured under various conditions [81 82 ]. The rate constant of the hydrolysis of timolol esters was reported to be 3.7 x 10 2 /min at 37 0 C in PBS buffer [81 ] The rate constant for the drug release from the highly crosslinked particles is 0.5747/day = 4.0 x 10 4 /min at 40 0 C in PBS, which is an order of magnitude lower than the reported rate constants for hydrolysis in bulk aqueous solutions. This large diff erence in the rates of hydrolysis could be due to the stearic constraints imposed by the small pore size and also due to the very small amounts of water available inside the hydrophobic particles. We note that it is plausible that the drug transport in th e particles is an activated process, i.e., the drug molecules are trapped in the pores whose sizes fluctuate with time. The rate of drug molecules jumping from out of the local tight pores determines the net rate of drug transport from the particles. In this case, the rate constant will characterize the time scale for the fluctuations in the pore structure of the particles. To distinguish between these two plausible mechanisms, we encapsulated timolol and lidocaine base, which is similar in size to timol ol but it does not possess the hydroxyl group that could potentially result in the ester linkage. The release experiments showed that lidocaine was released rapidly from the particles proving that the slow hydrolysis of timolol due to the small water cont ent of the particles is the rate controlling mechanism. To further prove that the release is controlled by hydrolysis, we soaked two groups of

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41 gels in ethanol one for 2 days and the other for 7 days. It was hypothesized that the hydrolysis will not occu r in ethanol but if physical entrapment is the rate limiting mechanisms, drug release in ethanol must be faster due to increased swelling. The lenses were withdrawn from ethanol after the 2 days or 7 days and then subjected to the drug release experiments in PBS. The release profiles in PBS from both sets of gels soaked in ethanol were similar, and also similar to the release profiles from gels that were not soaked in ethanol proving that hydrolysis is the rate limiting step (data not shown). While the t ime scale of drug release is independent of timolol to PGT ratio, the total mass of drug released depends on the ratio. The ratio of the total released amount from gels and the weight of gels (M /M gels ) is plotted as function of 1/T in Figure 2 12 for gel s loaded with particles prepared with various timolol to PGT ratios. The drug release amounts for each gel increase with an increasing temperature eventually reaching a plateau. The increase in M with increasing temperature is likely due to the decreasin g equilibrium partition coefficient between timolol bound as the ester in the particles and the free timolol. The plateau in M corresponds to the total drug loaded in the particles, and this value can be divided by the mass of drug added to PGT to obtain the encapsulation efficiencies for various ratios of timolol to PGT (Table 2 4). Interestingly, the encapsulation efficiency initially increases with increasing timolol to PGT ratio and then eventually begins to decrease beyond a ratio of 0.8. The effect of temperature on the drug release profiles shown in Figure 2 9 clearly proves that the drug release from the particle loaded lenses is highly temperature

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42 sensitive. An increase in temperature reduces the drug release duration and also increase s the total amount of drug released from the lenses. 2.2 11 Effect of Packaging of PGT Nanoparticle L oaded pHEMA Gels Typi cally commercial contact lenses are packaged in blister packs that contain about 1 1.5ml solution such as PBS, which is tear mimic s olution Contact lenses loaded with drugs could release into the packaging solution, thereby reducing the mass of drug loaded into the lens. Furthermore, a fraction of the drug released into the packaging solution could be delivered unintentionally in th e eye if the lenses are not rinsed before insertion into the eye. It is thus important to minimize the drug loss in the packaging solution. To explore the impact of packaging on the particle loaded contact lenses, we packaged gels in 1 ml of PBS for dura tions of 1 and 2 months. Several ophthalmic formulations are stored in refrigerated conditions. So we explored packaging at room temperature and also in refrigerator at 4 0 C. After 1 or 2 months of storage in the packaging solution, the gels were withdr awn, rinsed and then soaked into 1.75 ml or 3.5 ml fresh PBS for the 100 m or 200 m thick gel s, respectively. Also, the total amount of drug released in the packaging solution was measured. The profiles of drug released after the packaging stage are sh own in Figure 2 13 A D. Figure 2 13 A and B show the data for room temperature storage for 1 and 2 months packaging, respectively. Figure 2 13C and D show the release data for storage for 1 and 5 months, respectively, in refrigerator at 4 0 C. Each F igure shows release data for both 100 and 200 micron thick gels. The release profiles in Figure 2 1 3 A and B exhibit an initial burst followed by a slow release for about a month. The mass of drug released in the initial burst after 1 month packaging is 33 and 127 g for the 100 and 200 micron thick gels, respectively.

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43 This initial burst is presumably the release of the drug that diffused out of the particles into the gel during packaging. The initial burst for the 200 micron thick lens might be expected to be double that for the 100 micron. However, data shows that the amount released in the initial burst for the 200 m gels is more than twice that for 100 m gels likely because the volume of the packaging liquid was 1 ml for both lenses, and thus the ratio o f the gel to fluid volume is larger for the thicker lenses, resulting in smaller percentage release into the packaging liquid. The duration of the initial burst for the 200 micron thick gels is about 4 times that for the 100 micron thick gels, proving tha t the initial burst is arising from the drug that diffused into the gel during packaging. As mentioned above the /day. After the initial burst, the 100 owed by an release rate. Furthermore, lenses subjected to the 2 month packaging exhibited higher burst followed by gradual releas e at lower rates compared to those pack aged for 1 month because the equilibration duration for the particle loaded lenses is longer than 1 month. The sub therapeutic dose along with the potentially toxic initial burst reduces the viability of using the particle loaded gels for delivering ophth almic drugs such as timolol. In some cases where a high initial dose may be required such as for treating infections, the release profiles from the lenses packaged at room temperature might be desirable. Figure 2 13C shows the drug release profiles after packaging under refrigerated conditions. The release profiles after 1 month packaging do not exhibit any initial burst. Furthermore, the release profiles after 1 and 5 months packaging are almost

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44 indistinguishable from each other and also from the releas e profiles prior to packaging. These results along with the immeasurable drug release in the packaging solution show that the particles do not release any drug under refrigerated conditions. To further understand this issue, we utilize the Arrhenius rela tionship to determine the expected release duration from the particles during the storage in refrigerator. The predicted duration is at least 1 year, which along with the reduced partition coefficient between bound timolol as ester and the free timolol ex plains the negligible release from the particle loaded lenses in the refrigerator. Based on this data, we propose that the particle loaded lenses should be stored in refrigerated conditions prior to insertion in the eyes. Figure 2 13D shows that even aft er 5 month packaging in refrigerator, the gels exhibit sustained therapeutic release. 2.2 12 Effect of Surfactant Incorporation on Nanoparticle Laden Gels As described previously, we also prepared PGT nanoparticles through emulsions polymerization in pres ence of surfactants at various concentrations. The presence of surfactant did not impact the size distribution because the particle size is uncorrelated to the size of the emulsion drops. The particles prepared in presence of the surfactants were also in corporated in the HEMA gels, followed by characterization of the gels through measurements of transmittance, water content, modulus, and drug release dynamics. The gels loaded with PGT particles and surfactant were also transparent with almost 100% transm ittance in the visible range, and the modulus and water content of the gels was similar to those loaded only with particles ( Table 2 1 and Table 2 2 ). Furthermore, the drug release profiles from the gels loaded with surfactant and PGT particles ( Figure 2 14 ) are also very similar to those in which particles were prepared without surfactant. These results again show that the presence of surfactant does not

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45 impact any of the properties of the nanoparticles and the nanoparticle loaded gels proving that our a pproach of producing the nanoparticles is substantially different from the typical emulsion polymerization. Furthermore, these results also show that addition of surface active polymers or surfactants to the polymerization mixture that contains the nanopa rticles does not impact the release behavior from the lenses. This is an important result because typical lens manufacturing contains several long chain polymers such as poly vinyl pyrridone and poly vinyl alcohol that are added as comfort enhancers. 2.2 13 Comparison of Lenses Loaded with EGDMA and PGT Particles PGT was chosen as the monomer for preparing the nanoparticles due to its tri vinyl functionality. Other crosslinkers such as the divinyl ethylene glycol dimethacrylate (EGDMA) could also be used for preparing highly crosslinked nanoparticles. EDGMA nanoparticles were prepared with the same methods as those for preparing PGT nanoparticles, and EGDMA loaded gels were then prepared and characterized. The EGDMA nanoparticles laden pHEMA are also tra nsparent in the entire visible spectrum (data not shown ). The water content of the gels loaded with EGDMA particles is comparable to those loaded with the PGT nanoparticles ( Table 2 2 ) However, the EGDMA nanoparticle laden p HEMA gels with 21.6% particle loading have a zero frequency modulus of 100 MPa ( Table 2 1 ) which is 2 orders of magnitude larger than the modulus of commercial contact lenses. The drug release profiles from the EGDMA loaded lenses ( Figure 2 15 ) show a n extended release duration of about a month for both 100 and 200 micron thick lenses. However, the EG D MA nanoparticle loaded gels release only about a third of the drug amounts released by the PGT nanopa r ticle l oaded

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46 gels. Based on the increased modulus and the reduced drug release amounts, EGDMA nanoparticle load ed pHEMA gels are not suitable for contact lens use. 2.3 Conclusion In this study we showed that incorporation of highly crosslinked nanoparticles can increase drug release duration from a fe w hours to about 1 2 months. While several monomers with mutivinyl functionality can be used to prepare the nanoparticles, propoxylated glyceryl triacrylate (PGT) is the most suitable monomer. The particle preparation does not require surfactant but ins tead utilizes the oily drug timolol as the diluent that modifies the polymerization dynamics resulting in the nanoparticle formation. The encapsulation efficiency of timolol first increases and then decreases with increasing ratio of timolol to PGT and a ratio of 0.8 is optimal for producing the particles. At this ratio, 66.34 3.20% of the drug is encapsulated in the particles. The base form of timolol reacts with PGT during polymerization due to creation of nucleophiles to form as ester bond with the polymeric network. The slow hydrolysis of the ester bond results in an extended timolol release from the particles. The drug loaded nanoparticles can be loaded in p HEMA hydrogels by adding to the polymerization mixture. The nanoparticle loaded p HEMA gels are transparent and release drugs for about a month due to the slow hydrolysis of the timolol bound to the particles through the ester bond. The cumulative drug release profiles fit the model for a first order reaction with temperature dependent rat e constant. The rate constant for hydrolysis at 40 C is 4.0x10 4 /min which is much smaller than the previously reported rate constants for hydrolysis of other timolol esters, potentially due to stearic effects or the low water content of the highly cros slinked particles. The Arrhenius dependence of the rate constant yields an activation energy of 24.42

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47 2.47kJ/g mol. The temperature dependency of the rate constant is critical to our desired application of using contact lenses loaded with the nanoparti cles for ophthalmic drug delivery. The low kinetic rate constant under refrigerated conditions along with the low equilibrium partition coefficient for the free timolol ensures that a majority of the drug loaded in the lenses is retained in long periods o f storage under refrigerated conditions. The pHEMA hydrogels loaded with 4.6% nanoparticles are transparent with modulus comparable to commercial lenses but have a reduced water content compared to control pHEMA gels. The pHEMA based lenses are required t o be removed and submerged in PBS or the cleaning solutions overnight because pHEMA based lenses are not suitable for extended wear. The lenses will continue to release drug overnight at the same rate as measured in the in vitro studies. The drug release d overnight will be wasted and based on a 12 hour daily wear, a release of about 10 g/day of timolol is expected to be therapeutically effective. Based on all the results, we propose that a contact lens loaded with 4.6% PGT particles prepared with a rat io of drug to PGT ratio of 0.8 are most suitable for use as contact lenses. These lenses are transparent and maintain a modulus and water content comparable to pure p HEMA gels, while also releasing timolol at a rate of about 15 g/day for about 2 weeks, which is expected to be sufficient for therapeutic effects. The contact lenses will need to be packaged in refrigerated conditions to minimize the drug loss during packaging. When placed on the eyes, the increased temperature will trigger the drug releas e at the therapeutically desired rate. It is noted that commercial contact lenses are sterilized by autoclaving which is not viable for the system developed

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48 here because of the rapid drug release at high temperatures. Thus alternative approaches such as radiation, soaking in ethanol, etc will need to be explored for sterilization. While there are several systems capable of extended ophthalmic drug release through contact lenses, the system developed here is the first to release minimal drug amounts in packaging and be triggered to release drugs at elevated rates after lens insertion in the eye. Another advantage of the system proposed here is that the release profiles are non Fickian and decay at a slower rate compared to systems based on diffusion. Thus, the release rates are closer to zero order for the first 15 days compared to Fickian release profiles. This study focuses only on incorporating timolol into the particles. Several other ophthalmic drugs that have a reactive OH group which can form as ester linkage such as tilisilol, latanoprost could be encapsulated in the nanoparticles. Therefore, t he approach presented here could be very useful for extended delivery of several ophthalmic drugs. Additionally, this approach could be useful in a nu mber of applications such as transdermal drug delivery patches, wound healing patches, gold coating nanoparticle for attacking cancer cells, etc.

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49 A Timolol B Propoxylated glyceryl triacrylate (PGT) C Ethylene glycol dimethacrylate (EGDMA) Figure 2 1 Molecular structures of T imolol, PGT, and EGDMA

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50 Figure 2 2 Particle size distribution of PGT nanoparticles dispersed in ethanol and aqueous solutions with 10% pluronic F127 0 20 40 0.00 2.00 4.00 6.00 8.00 10.00 % Channel size (nm) F127 10 w% in PGT particle PGT particles Mean = 3.42 0.42

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51 Figure 2 3 Schematic illustration of the mechanism of formation of highly crosslinked nanoparticles du e to the presence of diluents. A ) an emulsion drop at very ea rly stages of polymerization B ) an emulsion drop containing several nanogels and high concentration of diluent in the unpolymerized liquid. and C ) solution with several emulsion drops, with several nanoparticles in each drop A B C

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52 Figure 2 4 Cumulative drug release from the na noparticles in a diffusion cell Data is shown as mean std (n = 3) 0.0 10.0 20.0 30.0 40.0 50.0 60.0 0 5 10 15 20 25 30 35 % Drug release Time (days) PGT particles

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53 Figure 2 5 Storage modulus of gels loaded with PGT nanoparticle s prepared with or without surfactant Data is shown as mean std (n = 3) 0 2 4 6 8 10 12 14 16 18 20 0 10 20 30 40 50 60 Storage Modulus (MPa) Frequency (Hz) 21.6% particle with surfactant 21.6% particle 8.8% particle 4.6% particle

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54 F igure 2 6 Cumulative drug release from 100 and 200 m thick PGT nanoparticle laden pHEMA gels. Data is shown as mean std (n = 3) 0 100 200 300 400 500 600 0 5 10 15 20 25 30 Drug release ( g) Time (days) 100 m 200 m

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55 Figure 2 7 Cumulative drug release from 100 and 200 m thick pHEMA gels without nanoparticles 0 500 1000 1500 2000 2500 0 0.5 1 1.5 2 2.5 3 3.5 Drug released (g) Time (days) m m

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56 Figure 2 8 Cumulative drug release from 100 and 200 m thick gels at room temperature Data is shown as mean std (n = 3) A) 8.8% particle loading in gels; particles prepared with timolol to PGT (T:PGT) ratio of 0.36. B) 4.6% particle loading in gels; particles prepared with T:PGT ratio of 0.6. C) 4.6% particle loading in gels; particles prepared with T:P GT ratio of 0.8. D) 4.6% particle loading in gels; particles prepared with T:PGT ratio of 1.0 0 100 200 300 400 500 600 700 800 900 0 5 10 15 20 25 30 Drug release ( g) Time (days) 100 m 200 m 0 100 200 300 400 500 600 0 5 10 15 20 25 30 Drug release ( g) Time (days) 100 m 200 m A B

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57 F igure 2 8 Continued 0 100 200 300 400 500 0 5 10 15 20 25 30 Drug release ( g) Time (days) 100 m 200 m 0 100 200 300 400 500 600 0 5 10 15 20 25 30 Drug release ( g) Time (days) 100 m 200 m C D

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58 Figure 2 9 Effect of temperature on drug release from the nanoparticle loaded gels. The solid lines are fits to the first order reaction model. A ) 100 m thick gels with 21.6% particle loading; particles prepared with T:PGT ratio of 0.12. B ) 200 m thick gels with 2 1.6% particle loading; particles prepared with T:PGT ratio of 0.12. C ) 200 m thick gels with 4.6% particle loading; particles prepared with T:PGT ratio of 0.6. D ) 200 m thick gels with 4.6% particle loading; particles prepared with T:PGT ratio of 0.8. E ) 200 m thick gel s with 4.6% particle loading; particles prepared with T:PGT ratio of 1.0. Data is shown as mean std (n = 3) 0 50 100 150 200 250 300 350 400 450 500 0 5 10 15 20 25 30 Drug release ( g) Time (days) 0 100 200 300 400 500 600 700 800 900 1000 0 5 10 15 20 25 30 Drug release ( g) Time (days) model fit 100C 80C 60C 40C 25C A B

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59 Figure 2 9 Continued 0 100 200 300 400 500 600 700 0 2 4 6 8 10 Drug release ( g) Time (days) 0 100 200 300 400 500 600 700 800 900 0 2 4 6 8 10 Drug release ( g) Time (days) model fit 100C 80C 60C 40C 25C C D

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60 Figure 2 9 Continued 0 100 200 300 400 500 600 700 800 900 0 2 4 6 8 10 Drug release ( g) Time (days) model fit 100C 80C 60C 40C 25C E

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61 Figure 2 10 Arrhenius fit between the rate constant and temperature. The gel thickness for each case is 200 m. Data is shown as mean std (n = 3) -1.5 -1 -0.5 0 0.5 1 1.5 2 0.002 0.0025 0.003 0.0035 0.004 log k 1/T T:PGT=0.12 T:PGT=0.6 T:PGT=0.8 T:PGT=1.0

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62 Figure 2 11 Potential mechanism for the formation of the ester link between timolol and PGT polymer

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63 Figure 2 12 Dependence of the total mass released from the gel per unit weight of the gel on 1/T. The y intercept represents the mass of dru g loaded into the gels 0 2 4 6 8 10 12 14 16 18 0.0025 0.0027 0.0029 0.0031 0.0033 0.0035 1/T T:PGT=0.12 T:PGT=0.6 T:PGT=0.8 T:PGT=1.0

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64 F igure 2 13 Amount of timolol release from 100 and 200 m thickness after packaging Data is shown as mean std (n = 3) A ) 1 month p ackaging at room temperature B ) 2 month pac kaging at room temperature C ) 1 month packaging in refrigerator at 4 0 C D ) 5 month packaging in refrigerator at 4 0 C 0 50 100 150 200 250 300 350 0 20 40 60 80 100 Drug release ( g) Time (days) 100 m 200 m 0 50 100 150 200 250 0 10 20 30 40 50 60 Drug release ( g) Time (days) 100 m 200 m A B

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65 Figure 2 13 Continued 0 100 200 300 400 500 600 700 0 5 10 15 20 25 30 Drug release ( g) Time (days) 100 m 200 m 0 100 200 300 400 500 600 0 5 10 15 20 25 30 35 Drug release ( g) Time (days) 100 m 200 m C D

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66 Figure 2 14 Cumulative release profiles from gels loaded with particles prepared in presence of F127 surfactant. Data is shown as mean std (n = 3) 0 100 200 300 400 500 600 700 800 0 10 20 30 40 Drug released (g) Time (days) 100 m 200 m

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67 Figure 2 15 Cumulative release profiles from pHEMA gels with 21.6% EGDMA particles. Data is shown as mean std (n = 3) 0 20 40 60 80 100 120 0 5 10 15 20 25 Drug released (g) Time (days) 100 m 200 m

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68 Table 2 1 Zero frequency s torage modulus of p HEMA hydrogels loaded with nanoparticles Hydrogel name Storage modulus (MPa) HEMA 1 0.30 HEMA loaded with 21.6% PGT particles 4.1 0.02 HEMA loaded with 21.6% particles prepared in presence of pluronic F127 surfactant 4.6 0.21 HEMA loaded with 21.6% EGDMA particles 100 HEMA loaded with 4.6% PGT particles 0.95 0.09 HEMA loaded with 8.8% PGT particles 1.35 0.005 Table 2 2 E quilibrium water content (EWC) of HEMA hydrogels loaded with nanoparticles Hydrogel name EWC ( in PBS ) HEMA 73.00 1.27 HEMA with 21.6% PGT particles 33.27 1.27 HEMA with 21.6% PGT particles prepared in presence of pluronic F127 surfactant 33.31 1.13 HEMA with 21.6% EGDMA particles 34.15 1.60 HEMA with 4.6% PGT particles 63.82 3.03 HEMA with 8.8% PGT particles 49.04 3.11

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69 Table 2 3 Dependency of total mass of drug released and the rate constant on temperature 100m gels T (C) T(K) M (g) k 25 298.15 234.72 7.52 0.0998 0.013 40 313.15 247.42 13.73 0.5747 0.055 60 333.15 318.21 6.85 2.0322 0.106 80 353.15 393.56 39.27 8.7882 1.394 100 373.15 372.77 51.32 28.4987 3.012 Table 2 4 Dependency of total mass of drug released and the rate constant on temperature 200 m gels T (C) T(K) M (g) k 25 298.15 534.89 15.82 0.1002 0.002 40 313.15 730.68 46.99 0.4084 0.035 60 333.15 799.91 13.93 1.9166 0.018 80 353.15 822.37 31.13 8.5919 1.154 100 373.15 823.98 45.55 22.5633 4.643

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70 Table 2 5 Effect of the composition of the particles on encapsulation efficiency Ratio of timolol and PGT (T:PGT) Encapsulation efficiency (%) 0.12 48.12 3.66 0.6 55.97 2.32 0.8 66.34 3.20 1.0 54.44 2.07

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71 CHAPTER 3 EXTENDED RELEASE OF NANOPARTICLE LOADED SILICONE HYDROGEL CONTACT LENSES In Chapter 2 we successfully developed extended wear contact lenses that can also provide extended glaucoma therapy, we disperse nanoparticles of PGT (propoxylated glyceryl triacylate) that contain a glaucoma drug timolol. In this Chapter, extended wear contact lenses that deliver glaucoma dru gs for extended periods could increase patient compliance, while also increasing the bioavailability. The particles can also be loaded into prefabricated lenses by soaking the lenses in a solution of particles in ethanol and by dispersing in polymerizatio n mixture. Nanoparticle incorporation in the silicone hydrogels were explored in here, including timolol release dutation, EWC, UV transmittace, modulus, ion and oxygen permeabilities, and the impact on each of these properties is proportional to the part icle loading. A gel with 5% particle loading can deliver timolol at therapeutic doses for about a month at room temperature, with a minimal impact on critical lens properties. Preliminary animal studies in Beagle dogs conducted with lenses in which parti cles are loaded by soaking the lenses in ethanol show a reduction in IOP. 3.1 Materials and Methods 3.1.1 Materials N,N Dimethylacrylamide (DMA), 1 vinyl 2 pyrrolidone (NVP), timolol maleate and ere purchased from Aldrich Chemicals (St Louis, MO) P ropoxylated glyceryl triacrylate (PGT) was purchased from Sartomer; Benzoyl peroxide (BP) (97%) was purchased from Aldrich Chemicals (Milwaukee, WI) The macromer bis alpha, mega (methacryloxypropyl) polyd imethylsiloxane (Macromer) was supplied by Clariant. 3 m ethacryloxypropyl tris(trimethylsiloxy)silane (TRIS) was

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72 gifted by Silar l aboratories (Scotia, NY) Methyacrylic acid (MAA) was purchased by Polysciences, Inc (Wattingyon, PA). 2,4,6 Trimethylbenzoyl diphenyl phophineoxide ( Darocur TPO ) was kindly provided by Ciba (Tarrytown, NY). Vitamin E (D alpha tocopherol, Covitol F1370) was girfted by Cogins Corporation. 3.1.2 Preparation of Nanoparticles and Silicone Hydrogels 3.1.2.1 Preparation of drug containing PGT Nanoparticles The drug loaded nanoparticles were prepared by thermal polymerization of a mixture of the timolol base and the PGT. The details of the process are available elsewhere. Briefly, timolol maleate was converted to the oily base f orm by increasing the pH of the aqueous solution. The timolol base was added to the crosslinker ( PGT ) and the initiator BP. The ratio of timol base and PGT was varied to prepare particles with various drug loadings. The mixture was then added to 5 ml of DI water and then 1.65 ml of 2.08M NaOH w as added to the mixture The mixture was purged with nitrogen for 15 min and then heated in an 80 C hot water bath under stirring a t 1100 rpm for 8 hours The thermal polymerization results in formation of drug l oaded nanoparticles, which were separated from the suspension by centrifugation for 15 mi n. 3.1.2.2 Preparation of Nanoparticle laden silicone hydrogels by adding particles to the polymerization mixture The particle laden silicone gels were synthesized by free radical polymerization. To prepare the silicone hydrogel, 0.8 ml of macromer bis alpha,omega (methacryloxypropyl) polydimethylsiloxane and 0.504 g of the drug laden nanoparticle suspension were added to 0.56 ml of N,N dimethylacrylamide ( DMA ) 0.24 ml of methylacrylic acid ( MAA ) 0.8 ml of 3 m ethacryloxypropyl tris(trimethylsiloxy)silane ( Tris ) and 0.12 ml of 1 vinyl 2 pyrrolidone (NVP) This composition results in formation

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73 of a gel with about 16.7 % (w/w) particle loading The mixture is purged by b ubbling nitrogen for 15 min. After adding 0.012g of the initiator Darocur TPO with stirring for 5 min, the mixture was poured in between two glass plates separated by a 100 or 200 m thick plastic spacer. The mold was then placed on Ultraviolet transilluminiator UVB 10 (UltraLum, Inc.) and irradiated with UVB light (305 nm) for 5 0 min. The molded gel was cut into circular pieces (about 1.65 cm in diameter) with a cork borer and dried in air overnight befor e further use. A dditionally pure silicone gels used for control s were prepared by same procedure as described above except that particle suspension was not added to the mixture 3.1.2.3 L oading Timolol PGT Nanoparticles into commercial contact lenses The timolol PGT particles can be loaded into polymerized commercial contact lenses by soaking the lenses in a solution of particles in ethanol. Due to the small size of the particles and the increased pore size in the lens matrix, particles diffuse into the lenses. After equilibration, the lenses are withdrawn from the ethanol solution and soaked in PBS to extract ethanol. The particles are retained in the lenses due to the hydrophobicity and the larger size compared to ethanol. Specifically, contact lense s (Acuvue Oasys of 3.5 D power and 14 mm diameter) were soaked in the 3% (w/w) solution of the particles in ethanol for a period of 24 hours. The lenses were then withdrawn and soaked in 100 ml PBS for extraction of ethanol for 24 hours. The particle lo aded commercial lenses were then utilized in in vitro drug release studies and in vivo pharmacodynamics studies in Beagle dogs.

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74 3.1 3 Particle Size Distribution The particle size distribution was determined by dynamic light scattering (DLS) The diameters of the investigated particle suspensions were analyzed by Nanotrac Particle Size Analyzers ( Microtrac Inc. ) 3.1 4 Transmittance The transmittance of nano particle laden silicone hydrogels was measured using UV Vis spectrophotometer (Thermospectronic Genesys 10 UV). The lenses were hydrated by soaking in PBS overnight, and then mounted on the outer surface of a quartz cuvette. The cuvette was placed in a spectrophotometer and the transmittance was measured at wavelengths ranging from 200 nm to 1000 nm. 3.1 5 Equilibrium Water Content (EWC) The dried lenses were weighed (W dry ) and then hydrated by soaking in 3.5 ml of PBS for 24 hours. The hydrated gels were weighed (W wet ) and the equilibrium water content (EWC) was calculated by the following formula (3 1) 3.1 6 Mechanical Properties Storage modulus of the gels w as measured using a dynamic mechanical analyzer (DMA Q800, TA instruments). H ydrated rectangular gel s 400 micron in thickness were mounted in the s ubmersion tension clamp at room temperature. A preload force applied 0.01N and force track was 115% used. Strain sweep tests were recorded to confirm the linear range at room temperature at 1 Hz Subsequently, frequency dependent storage modulus was obta ined by fixing the strain within the linear range

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75 3.1 7 Ion Permeability Extended wear contact lenses must be permeable to i on s to ensure homeostasis of ion concentration in the post lens tear film, which is necessary for lens motion. It has been claime d that extended wear contact lenses must possess ion permeability greater than 1.5 x 10 6 mm 2 /min. Details of the procedures for measuring ion permeability of gels and contact lenses are available elsewhere [ 8 2]. Briefly, gels (1.65 cm diameter, 100 m thick) were loaded with salt by soaking in 0.75 M sodium chloride solution until equilibrium was achieved. The gels were then soaked in 27.5 ml DI water and the dynamic sodium chloride concentration in the release medium was determined by measuring the co nductivity (Con 110 series sensor, OAKTON ). The salt release from the gels into the release medium can be fitted to a diffusion model for release in perfect sink to determine the effective ion diffusivity. 3.1 8 Oxygen Permeability Cornea is an avasular tissue and thus it obtains the oxygen required for metabolism directly from air. Details of the procedures for measuring oxygen permeability of gels and contact lenses are available elsewhere [83]. Briefly, gels ( 1.36 cm diameter, 100 m thick) were moun ted in a diffusion cell containing oxygenated water in the donor compartment. The 18 ml donor compartment was filled with DI water that was equilibrated with air. The 32 ml receiver compartment was filled with water that was deoxygenated by bubbling nitr ogen for an extended period of time. The oxygen concentration in the receiver was measured with OX TRAN Model 2/21 from MOCON and the data was fitted to a diffusion model to determine the oxygen permeability through the gel.

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76 3.1 9 Drug Release from the N anoparticles Laden Silicone Hydrogels The nanoparticle loaded gels prepared by UV polymerization were cut into circular pieces of 1.65 cm diameter, dried and weighted. The sol component i.e., the unreacted soluble fraction was extracted by soaking the gel in 200 ml DI water under stirring at room temperature for 24 hours. Next, the gels were transferred to fresh PBS for the drug release experiments. The PBS volume in the release medium was chosen to be ectively, and the temperature was varied from 25 to 100 0 C to explore the effect of temperature on drug release profiles. The drug concentration was determined periodically by measuring the UV VIS absorbance for wavelength ranging from 252 nm to 321 nm an d fitting the data to timolol spectra. 3.1 10 Drug Release Rates from the Nanoparticles Laden Silicone Hydrogels after Packaging Typi cally commercial contact lenses are packaged in blister packs that contain about 1 1.5 ml solution such as PBS, which is tear mimic solution To simulate the packaging, gels were subjected to the monomer extraction as described above, and then stored in 1 ml of packaging solution (PBS) for durations ranging from 1 5 months at either room temperature or in a refrigerator at 4 C. After the storage in PBS, the gels were withdrawn, and the drug release profiles were measured by following the same procedures as described above in Section 3.1.9 3.1.11 In vitro and In vivo studies from Nanoparticles Loaded Commercial Lesnes 3.1.11.1 In vitro Timolol PGT Nanoparticles into commercial contact lenses The drug release studies described above were also conducted with Acuvue Oasys lenses loaded with the timolol PGT nanoparticles. The same procedure as

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77 described above was utilized but the commercial lenses were packaged in refrigerator only for 2 weeks and release profiles were measured only at room temperature and 40 0 C. 3.1.11.2 In vivo pharmacodynamics studies in Beagle dogs The in vivo studies were d one in a 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 [85 ]. Beagle dogs have been used in several prior studies as animal models for glaucoma in cluding in some of our prior studies with contact lenses [86 91 ]. The cornea shape and size of these dogs are similar to that of human beings, and therefore the commercially available contact lenses for human can be used in this study without further modi fication. The pharmacodynamics experiments focused on measuring the intraocular pressure(IOP) after insertion of the particle loaded contact lenses. Timolol reduces the IOP and so dynamic measurements of the IOP after lens insertion could be utilized to determine the pharmacodynamic efficacy of the particle loaded contact lenses. The study utilized 10 adult Beagle dogs with inherited open angle glaucoma. Prior to the experiments with lenses, the baseline IOP was measured via applanation tonometry (Tono Pen XL (Mentor O and O, Norwell, MA)) in both eyes (OU) 2 times daily at the same times of day for 5 days. A topical anesthetic (proparacaine hydrochloride 0.5%) was applied to each eye prior to the measurement of IOP OU. After one week of washout, partic le loaded Acuvue Oasys lenses were placed in the right eyes of the 10 dogs and IOP OU were measured 2 times daily. In addition to measuring the IOP, the eyes were observed for ocular irritancy and the response was quantified according to the McDonald and S hadduck scoring system [92 ].

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78 All animals in this study were housed and cared for according to the guidelines from the Association for Research in Vision and Ophthalmology (ARVO) and the Institutional Animal Care and Use Committee (IACUC) at University of F lorida (UF) prior, during and after the experiments. All in vivo experiments procedures were approved by the IACUC at UF and were performed in compliance with the ARVO Statement for the Use of Animal in Ophthalmic and Vision Research. 3.2 Results and Discu ssion 3. 2 1 Equilibrium Water Content (EWC) The water content of silicone hydrogel lenses is expected to decrease on incorporation of PGT particles due to the negligible water uptake by the particles. The silicone control gel swells 47 % (on dry basis) in PBS, and thus the silicone gel with 16.7 % particles is expected to swell by about 47 x (1 0.167) = 39.5%. The gels with 16.7% particles however swell only by 26%, which is about 13.5% lower than the expected value. This suggests that the presence of part icles leads to increased crosslinking in the gels likely due to the PGT that remained unreacted during the particle preparation step. 3.2 2 Transmittance The clarity of the particle laden lenses was characterized by measuring the transmittance spectra of the 100 micron thick gels in the range from 200 to 1000 nm (Figure 3 1 A). Both control silicone hydrogel and the nanoparticle loaded gels exhibit near 100% tr ansmittance in the visible range. Additionally the nanoparticle loaded silicone gels show partial UV blocking in 240 300 nm range which is a desirable property for contact lenses. The UV blocking can be attributed to the increased amount of PGT fraction, which absorbs UV in 240 300 nm range. A photographic image of the

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79 particle laden gel shown in Figure 3 1B along with the transmittance data in Figure 3 1 A prove that the presence of particles does not reduce the visual clarity of the gel. 3.2.3 Mechanica l Properties The frequency dependent storage modulus (G ) and loss modulus (G ) were obtained for nanoparticle laden silicone hydrogels. The effect of particle loading on the frequency dependent storage and loss moduli for silicone hydrogels is shown in Fi gure 3 2. The results clearly show that particle loading loads to an increase in both the storage and the loss modul i for the gels. The i ncreas e in modulus could be due to the presence of the nanoparticles and also partially to a high degree of crosslinki ng of the bulk due to incomplete reaction of the PGT during the particle formation step. The increase in modulus is undesirable and thus we explored use of vitamin E as a diluent for reducing the gel modulus. Vitamin E is an antioxidant and so its incorp oration into contact lenses could have other beneficial effects in addition to the reduced modulus. Silicone hydrogels were prepared after adding 20% (w/w) vitamin E and 16.7% (w/w) nanoparticles to the silicone hydrogel formulation. The zero frequency storage modulus for the gels with 16.7% particles and 20% vitamin E is 1.07 MPa which is lower than both the silicone hydrogel with 16.7% particles (13.3 MPa) as well as the control silicone hydrogels (5.95 MPa) showing that vitamin E could be used as a di luent to control modulus of contact lenses. To determine whether vitamin E incorporation impacts drug release, timolol release profiles were measured from gels with 16.7% particles and 20% vitamin E. The data shows that vitamin E incorporation does not imp act the release profiles significantly (Figure 3 3).

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80 3. 2 4 D rug Release from the Nanoparticles Laden Silicone Hydrogels The rate mechanism can easily be identified by conducting equilibrium release experiments from gels with two different values of gel th ickness ( h ). If the drug release is controlled by diffusion through the gel, the release time scales as h 2 and if the release is controlled by particles, the release time should be independent of h. The drug release profiles from control silicone hydroge ls (without nanoparticles) do exhibit the h 2 dependency (data not shown). However the release profiles from the particle loaded gels ( Figure 3 4 ) show that the drug release duration is about the same for both 100 and 200 m thick gels proving that the presence of particles controls the drug release rates. The total amount of drug released from the thick gels is about twice that from the thin gels, which is expected due to the weight ratio of the two gels. The data in Figu re 3 4 clearly shows that there is an extended drug release from the gel which lasts for over 1 month for both the 100 and 200 m gels, which is significantly longer than drug release duration from the silicone gels without particles. 3. 2 5 Effect of Tempe ratures on Drug Release Profiles The rate mechanism can easily be identified by conducting equilibrium release experiments from gels with two different values of gel thickness ( h ). The mechanism of extended drug release from p HEMA gels loaded with the tim olol PGT particles was shown to be hydrolysis of the ester bond between timolol and the particle network [Section 2.2.10 ]. Accordingly, the drug release data should fit the first order reaction model, i.e., ( 3 2)

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81 where M is the amount of drug released at a time t and M is the amount of drug released after infinite time, and k is the rate constant. The rate constant k depends on temperature through the Arrhenius relationship, i.e., (3 3) where A is the pre exponential constant, E a is activation energy, R is the ideal gas constant, and T is the temperature in Kelvin. To explore temperature dependency, the drug release experiments were performed at 25, 40, 60, 80, and 100 C and the results are shown in Figure 3 5 The solid lines in the figures represent the fit of the release profiles to the first order reaction model (Eq. 3 2). Several esters of timolol have been synthesized for use as prodrugs and their rates of hydrolysis have been measured under various conditions [81,82]. The rate constant of the hydrolysis of timolol esters was reported to be 3.7 x 10 2 /min at 37 0 C in PBS buffer [81]. The rate constant for the drug release fr om the highly crosslinked particles dispersed in the silicone hydrogels is 0.2477/day = 1.72 x 10 4 /min at 40 0 C in PBS. The rate constant was determined to be 4.0 x 10 4 /min at 40 0 C for the particles dispersed in pHEMA gels [ Section 2.2.10 ]. The smaller rate constant for the dispersion in silicone hydrogels could potentially be due to smaller water content in the silicone hydrogels, which slows down the hydrolysis. The rate constants for the hydrolysis reaction in the highly crosslinked P GT particles is an order of lower than the reported rate constants for hydrolysis in bulk aqueous solutions for timolol esters [82]. This large difference in the rates of hydrolysis could be due to the stearic constraints imposed by the small pore size an d also due to the very small amounts of water available inside the hydrophobic particles.

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82 Based on the Arrhenius relationship, a plot of log(k) with 1/T, should be a straight line of slope E a /R The temperature dependence of the values of k obtained by fi tting the drug release data for 100 m thick gels satisfy the above relationship as shown in the inset in Figure 3 5 The value of the activation energy obtained from the slopes of the graphs is 23.40 kJ/gmol. The activation energy was the case of parti cles dispersed in the pHEMA gels was determined to be 24.42 kJ/gmol, which is similar to the activation energy for the case of silicone hydrogels. 3. 2 6 Effect of Packaging Contact lenses are typically packaged in blister packs that contain about 1 1.5 ml solution such as PBS, which is tear mimic solution Ophthalmic drugs loaded into contact lenses could diffuse out into the packaging solution during the extended shelf life of a few months, thereby reducing the mass of drug loaded into the lens. To explo re the impact of packaging on the particle loaded contact lenses, we packaged gels in 1 ml of PBS for durations of 1 and 2 months at room temperature and in a refrigerator at 4 0 C. After the storage in the packaging solution, the gels were withdrawn, rins ed and then soaked into 1.75 ml or 3.5 ml fresh PBS for the 100 m or 200 m thick gel s, respectively. Also, the total amount of drug released in the packaging solution was measured. The profiles of drug released after the packaging stage are shown in Fi gure 3 6 A D. Figure 3 6 A and B show the data for room temperature storage for 1 and 2 months packaging, respectively. Figure 3 6 C and D show the release data for storage for 1 and 5 months, respectively, in refrigerator at 4 0 C. Each figure shows rel ease data for both 100 and 200 micron thick gels. The release profiles in Figure 3 6 A and B exhibit an initial burst followed by a slow release for about a month. The mass of drug released in the initial burst after 1 month

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83 packaging is 150 and 350 g fo r the 100 and 200 micron thick gels for 60days, respectively. Th e initial burst is presumably due to the release of the drug that diffused out of the particles into the gel during packaging. The initial burst for the 200 micron thick lens might be expect ed to be double that for the 100 micron. The release profiles show that the amount released in the initial burst for the 200 m gels is slightly more than twice that for 100 m gels likely because the volume of the packaging liquid was 1 ml for both lense s, and thus the ratio of the gel to fluid volume is larger for the thicker lenses, resulting in smaller percentage release into the packaging liquid. The duration of the initial burst for the 200 micron thick gels is about 4 times that for the 100 micron thick gels, proving that the initial burst is arising from the drug that diffused into the gel during packaging. As mentioned above the desired daily release from the lenses is /day. After 30 days packaging at room temperature, the 100 m gel releases Furthermore, lenses subjected to the 2 month packaging exhibited higher burst compared to those packaged for 1 month because the equilibration duration for the particle loaded lenses is longer than 1 month. The release amount of 100 burst. The sub therapeutic dose along wi th the potentially toxic initial burst reduces the viability of using the particle loaded gels for delivering ophthalmic drugs such as timolol. Figure 3 6 C shows the drug release profiles after packaging under refrigerated conditions. The release profile s after 1 month packaging do not exhibit any initial burst. Furthermore, the release profiles after 1 and 5 months packaging are almost indistinguishable from each other and also from the release profiles prior to packaging. These results along with the immeasurable drug release in the packaging solution show

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84 that the particles do not release any drug under refrigerated conditions. Based on this data, we propose that the particle loaded lenses should be stored in refrigerated conditions prior to insertio n in the eyes. Figure 3 6 D shows that even after 5 month which is potentially in the therapeutic range. 3. 2 7 Maximizing Drug Loading in Nanoparticles While nanoparticl e incorporation leads to extended drug release, it also increases lens modulus, and reduces ion and oxygen permeabilities. To minimize the impact on modulus, it is beneficial to minimize the particle loading, while maximizing the drug loading in the parti cles to insure adequate drug loading in the lens. Particles with about 50% drug loading were prepared by following the same procedures as described in the Methods section but with a polymerization mixture composed of 1 g of timolol, 1 g of PGT and 7.5 mg of the initiator. Gels were then prepared by adding the particles with high drug loading to the polymerization mixture at 5% particle loading. The drug release profiles from the gels with 5% high drug loading particles are shown in Figure 3 7 The results shown that the 100 micron thick gels with 5% particle release about 4.8g/day which is comparable to the release from the gels with 16.7% particle loading but with a lower drug loading in particles. To determine the suitability of the gels loaded with 5% particles as contact lenses, modulus, water content and ion and oxygen permeabilities were also measured. The frequency dependent storage modulus is included in Figure 3 2. The zero frequency storage modulus of the gels is 7.01 MPa, which is only slight ly larger than that of the control gel, suggesting minor impact of particle incorporation on modulus. The modulus of contact lenses is typically about 1 MPa and so the modulus will be required to be reduced. This can be accomplished either by including a diluent

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85 such as vitamin E in the polymerization mixture or by choosing the control formulations utilized in manufacturing commercial contact lenses. The ion permeability (product of diffusivity and partition coefficient) of the 5% particle gels is 1. 8 x 10 3 0.000655 mm 2 /min which is comparable to that for the control, and larger than the critical value of 1.5 10 6 mm 2 /min required for on eye lens motion. Finally, the oxygen permeability (Dk) of the 5% particle gels is 48.50 12.23 [(10 11 (cm 2 /s)( mlO 2 /(ml mmHg)] (also known as Barrers) which is slightly less than that for the control. The oxygen permeability of contact lenses needs to be sufficiently high for extended wear to avoid corneal edema [92 94]. Oxygen transmissivity (Dk/thickness) values of about 20 and 75 [(10 9 (cm 2 /s)(mlO 2 /(ml mmHg)] are recommended to avoid corneal edema for the open and closed eyes, respectively. Based on the measured Dk value and a thickness of 80 micron for a contact lens, we obtain a value of Dk/t of 70.41 for co ntact lenses made from the gels with 5% high timolol loading particles. 3. 2 8 Drug Release from Nanoparticles Loaded Commercial Lens Timolol PGT particles with high drug loading (1:1 ratio of timol to PGT) were loaded into the Acuvue Oasys lenses by soakin g the lenses in 3% (w/w) solution of particles in ethanol. The drug release profiles from the particle loaded Acuv u e O asys lenses are shown in Figure 3 8 (at room temperature before packaging) and in Figure 3 9 (at room temperature after packaging in refrigerator for 2 weeks). The release profiles prior to packaging are similar to those from the lenses in which the particles were incorporated prior to polymerization. The particle loaded Acuvue Oasys lenses rele ase timolol for about 2 weeks with an average release rate of 6 g/day. The lenses packaged in refrigerator for 2 weeks release drug at a significantly reduced rates, but

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86 without any initial burst. The absence of the initial burst suggests that timolol w as not released from the PGT particle matrix and thus the reduction in the drug loading during packaging is due to diffusion of the timolol PGT particles into the packaging medium. Thus, particle loading by soaking of prefabricated lenses is not a viable approach for extended delivery of timolol. However the Acuvue Oasys lenses loaded freshly with the particles could be utilized in the animal studies to explore safety and the pharmacodynamic efficacy because the diffusion of the particles through the lens es is slow compared to the release at physiological temperatures. 3. 2 9 Safety and Pharmacodynamic Efficacy of the Nanoparticles Loaded Commercial Lenses The IOP is impacted by several factors including food uptake, time of the day, stress, etc and so typ ically the IOP of the untreated eye is utilized as a control, and thus the difference in the IOP between the treated and the untreated eyes is considered as an indicator of the pharmacodynamic effect of the drug delivery. No adverse event was recorded in this study suggesting that the nanoparticle loaded contact lenses are safe. The IOP difference between the untreated (OS) and the treated (OD) eye is plotted as a function of time in Figure 3 10 Also data for the IOP difference from the control study i s included. The in vitro drug release data at about 37 0 C, which is close to the ocular temperature, is included in the inset. The error bars in each set of data indicate the standard error with n = 10 for the control study and n = 10, 6, 6, 6, and 4 for days 1 5 for the lens study. The value of n in the lens study decreases with elapsed days because several animals lost the lenses at various time points in the study. The data show that lens insertion leads to a decrease in IOP but the effect is insign ificant after day five. Also there is no significant effect of lens

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87 insertion on the first day. It is noted that this animal model for glaucoma is treated by 2 eye drops each day of 0.5% timolol. Thus the mass of drug delivered each day by eye drops for therapeutic effect is about 300 g. Based on the estimated 10 fold increase in bioavailability for contact lenses, the expected therapeutic dose is about 30 g, which is larger than the drug release rate from the lenses after day four. 3.3 Conclusion Ti molol PGT particles release drug for an extended period of time likely due to the slow hydrolysis of the ester bond. The gels prepared by polymerizing the mixture of the silicone hydrogel monomers and the timolol PGT particles are transparent, and release for extended period of about a month at room temperature. The release duration increases to more than a year under refrigerated conditions thus providing an easy approach for preventing loss of drug in the packaging by storage in the refrigerator. The cu mulative drug release profiles fit the model for a first order reaction with temperature dependent rate constant. The rate constant for hydrolysis at 40 C is 1.7x10 4 /min which is much smaller than the previously reported rate constants for hydrolysis of other timolol esters, and also slightly smaller than the rate constant for particles dispersed in the pHEMA gels. The Arrhenius dependence of the rate constant yields activation energy of 23.40kJ/g mol, which is comparable to the activation energy for th e case of particles dispersed in the pHEMA gels. Dispersing the timolol PGT particles in the silicone hydrogels does not reduce transparency but it does have an undesirable effect on some other critical contact lens properties such as water content, modul us and ion and oxygen permeabilities. The impact on the properties is proportional to the percentage loading of the particles. It is thus important to maximize the drug encapsulation in the particles so that therapeutic doses can be delivered with minima l

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88 particle loading. The lenses loaded with 5 % of particles with a timolol:PGT ratio of 1:1 release therapeutic drug dose with minimal impact on other critical properties. Preliminary in vivo animal studies demonstrate safety and efficacy of the particle l oaded lenses at treating glaucoma. However the lenses utilized in the animal studies were loaded with the particles by soaking lenses in ethanol. This approach of loading drug PGT particles is not suitable for packaging due to the slow diffusion of the p articles into the packaging medium. The in vivo tests are however very useful at showing that a slow and extended release of timolol from contact lenses can have the desired pharmacodynamics effect of reduction in the IOP. While the results presented here are very encouraging, further in vivo safety and pharmacokinetics and pharmacodynamics tests particularly with lenses in which particles are loaded prior to lens fabrication are necessary to evaluate the potential of the proposed approach for glaucoma the rapy.

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89 Figure 3 1 Transmittace spectra and photo graphic image. A ) Transmittance spectra from 2 00 nm to 10 00 nm of silicone control and PGT nanoparticle laden silicone hydro gel. B ) A photo graphic image of the PGT nanoparticle laden silicone hydro gel 0 20 40 60 80 100 200 400 600 800 1000 Transmittance % Wavelength (nm) nanoparticle laden silicone gel silicone control gel A B

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90 Figure 3 2 Mechanical properties of gels loaded with PGT nanoparticles prepared and PGT nanoparticles including Vitamin E A) Storage moduli. B) loss modul i Data is shown as mean std (n = 3) 0 5 10 15 20 25 30 0 10 20 30 40 Storage Modulus (MPa) Frequency (Hz) silicone control gel 16.7% nanoparticle loading 16.7% nanoparticle with 20% vitamine E loading 5% nanoparticle loading 0 1 2 3 4 5 6 7 8 0 10 20 30 40 Loss Modulus (MPa) Frequency (Hz) silicone control gel 16.7% nanoparticle loading 16.7% nanoparticle with 20% vitamin E loading 5% nanopartlce loading A B

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91 Figure 3 3 Cumulative drug release from 100 and 200 m with 16.7% PGT nanoparticle and 20% Vitamin E laden silicone hydrogel 0 50 100 150 200 250 300 350 400 450 500 0 20 40 60 80 Drug release ( g) Time (day) 100 m 200 m

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92 Figure 3 4 Cumulative drug release from 100 and 200 m thick PGT nanoparticle laden silicone hydrog el. Data is shown as mean std (n = 3) 0 50 100 150 200 250 300 350 400 450 0 10 20 30 40 50 60 Drug released (g) Time (days) 100 m 200 m

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93 Figure 3 5 Effect of temperature on drug release from the nanoparticle loaded gels. The solid lines are fits to the first order reaction model. A) The gel thic kness for each case is 100 m. B) shows the Arrhenius fit between the rate constant and temperature. Data i s shown as mean std (n = 3) 0 50 100 150 200 250 300 350 400 0 5 10 15 20 25 30 Drug release ( g) Time (day) model fit 25C 40C 60C 80C 100C y = 2886.4x + 8.5066 R = 0.9797 -1.6 -1.2 -0.8 -0.4 0 0.4 0.8 1.2 0.002 0.0025 0.003 0.0035 log k 1/T A B

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94 A B C D Figure 3 6 Amount of timolol release from 100 and 200 m thickness after packaging A ) 1 month p ackaging at room temperature B ) 2 month pa ckaging at room temperature C ) 1 month packaging in refrigerator at 4 0 C. D ) 5 month packaging in refrigerator at 4 0 C. Data is shown as mean std (n = 3) 0 50 100 150 200 250 300 350 400 450 0 10 20 30 40 50 60 70 Drug release ( g) Time (days) 100 m 200 m 0 50 100 150 200 250 300 0 10 20 30 40 50 60 Drug release ( g) Time (days) 100 m 200 m 0 50 100 150 200 250 300 350 400 450 500 0 10 20 30 40 Drug release ( g) Time (day) 100 m 200 m 0 50 100 150 200 250 300 0 5 10 15 20 25 30 Drug release ( g) Time (day) 100 m 200 m

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95 Figure 3 7 Cumulative drug release from 5% PGT nanoparticle with timolol 1000 mg laden silicone hydrogel. Data is shown as mean std (n = 3) 0 100 200 300 400 500 600 700 0 10 20 30 40 50 60 70 Drug release ( g) Time (day) m m

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96 Figure 3 8 Drug release at room temperature from Acuve Oasys lenses loaded with particles by soaking the lenses in a 3% solution of particles in ethanol 0 20 40 60 80 100 120 140 0 5 10 15 20 25 30 Drug release ( g) Time (day)

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97 Figure 3 9 Drug release at room temperature after packing particle loaded Acuve Oasys lenses in refrigerator for 2 weeks. Data is shown as mean std (n = 3) 0 5 10 15 20 25 0 5 10 15 20 25 Drug release ( g) Time (day)

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98 Figure 3 10 Pharmacodynamic effect of extended wear of the particle loaded Acuve Oasys lenses. The in vitro drug release data at about 37 0 C -3 -2 -1 0 1 2 3 4 5 6 7 8 0 1 2 3 4 5 6 IOP (OS OD) T ime (days) Baseline Contact lens

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99 CHAPTER 4 EXTENDED RELEASE OF TIMOLOL FROM NANOPAR TICLE LOADED FORNIX INSERT FOR GLAUCOMA THERAPY Our goal in this Chapter 4 is to develop an ophthalmic insert that releases a glaucoma drug for an extended period of time, and can be safety and easily inserted in the fornix. Timolol, a beta adrenergic receptor antagonist, is used as the drug because it 1979 [96] and also because of the potential of significant cardiac side effects from systemic exposure to timolol [97] Although prostaglandins such as latanoprost are now more commonly prescribed than timolol, a combination of both of these drugs is frequently used as well. Several researchers have developed fornix inserts and Ocusert and Lacrisert have been commercialized to treat glaucoma and dry eyes, respectively. Lacrisert is a cylindrical insert 3.5 mm long and 1.27 mm in diameter made of hydroxypropyl cellulose. The insert dissolves over a period of a day after insertion leading to increased tear viscosity and lubrication. The inserts proposed here were chosen to be geometrically similar to lacrisert and thus have a length ranging of about 7 mm and a length of about 1 mm. While retention of the fornix inserts is typoically a concern, cylindrical shaped inserts are considered best for retention in the conjunctival sac. The materials for designing the inset were chosen to be HEMA and PGT, which are similar to the materials commonly used in ocular applications such as contact lenses. 4.1 Materials and Methods 4.1.1 Materials Hydroxyethylmethacrylate ( HEMA ) monomer t imolol m aleate Azobisisobutylonitrile (AIBN) saline (PBS) w ere

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100 purchased from A ldrich Chemicals (St Louis, MO) P ropoxylated glyceryl triacrylate (PGT) was kindly provided by Sartomer; Benzoyl peroxide (BP) (97%) was purchased from Aldrich Chemicals (Milwaukee, WI) 4.1.2 Preparation of Highly Crosslinked PGT Nanoparticles The drug loaded nanoparticles were prepared by thermal polymerization of a mixture of the timolol base and the PGT. The details of the process are available elsewhere. Briefly, timolol maleate was converted to the oily base form by increasing the pH of the aqueous solution. The timolol base (240 mg) was added to 1 g of the crosslinker ( PGT ) and 7.5 mg of the initiator BP. The ratio of timolol base and PGT was varied to prepare particles with various drug loadings. The mixture was then added to 5 ml of DI water and then 1.65 ml of 2.08M NaOH w as added to the mixture The mixture was purged with nitrogen for 15 min and then heated in an 80 C hot water bath under stirring a t 1100 rpm for 8 hours The thermal polymerization results in formation of drug loaded nanoparticles, which were separated from the suspension by centrifugation for 15 mi n. 4.1 3 Preparation of Nanoparticles Laden Inserts The timolol PGT particles were added to the HEMA monomer in various ratios ranging from 25:75 to 100:0 to create the polymerization mixture for fabricating t he inserts. The mixture was purged with nitrogen for 15 minutes to reduce the amount of dissolved oxygen. Next, 30 mg o f a thermal initiator a zobisisobutylonitrile (AIBN) was added to the mixture and stirred for 15 min. The mixture was then poured into 1 .02 mm inner diameter Silastic tubing which served as the molds for the polymerization. The silicone molds were sealed at both ends and then for 40 minutes for polymerization. After overnight drying, the cylindrical inserts were

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101 gently pulled out of the silicone molds. The cured inserts were 1 mm in diameter and were cut into 7.5 mm long sections (Figure 4 1) Control p HEMA inserts were also prepared by following the same procedure as described above except that the polymerization mixture contained a mixture of HEMA monomer with timolol base. 4.1 4 Drug Release Experiments The 7.5 mm long 1 mm diameter inserts were first submerged in 30 ml DI water under minimal stirring (12 0 rpm) and at room temperature for 24 hour s to extract the unreacted monomer. Next the inserts were transferred to 3ml of fresh PBS for the drug release experiments. During the release experiments, the drug concentration was measured once daily, with two additional measurements at 1 hour and 4 hou rs after the soaking in PBS. The time dependent concentrations of timolol in PBS were determined by measuring the absorbance as a function of time by UV Vis spectrophotometer in the 2 52 3 12 nm wavelength range. The spectra were compared with the timolol s pectra to ensure that the absorbance from components other than the drug was minimal. 4. 1 5 Packaging Tests The inserts could be packaged either in dry or in hydrated state. Hydrated inserts would likely be more comfortable compared to the dry inserts, but packaging in hydrated state could lead to loss of drug during packaging. To explore the impact of wet pac kaging, the 7.5 mm long, 1 mm diameter inserts were packaged in 1 mm of PBS in refrigerator at 4 0 C for 3 months. The inserts were subjected to the initial extraction in 30 ml DI water for 24 hours prior to the packaging. After 3 months of storage in the refrigerator, the inserts were submerged in 3 ml of fresh PBS for the drug release experiments.

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102 4.2 Results and Discussion 4. 2 1 Drug Release from Control p HEMA Inserts An image of the dried nanoparticle loaded insert is shown in Figure 4 1. The dried i nserts are rigid and clear, but become soft and lose transparency on hydration. Figure 4 2 shows the drug release profiles from the control p HEMA insert in which timolol was added directly to the polymerization mixture prior to thermal curing. The drug d iffuses out in about 12 hours, which is inadequate for the desired goal of extended release fornix inserts. 4. 2 2 Effect of Timolol PGT particles Loading in Inserts on Drug Release Figure 4 3 A shows the cumulative amount of drug released as a function of time from the particle loaded inserts. The four curves correspond to the various particle loadings, which are indicated in the legend. The solid lines are fits of a model described to the experimental data. The data in Figure 4 3 A clearly shows that th ere is an extended drug release from the inserts which lasts for about 1 month which is significantly longer than 6 hour drug release duration from the p HEMA inserts without particles (Figure 4 2 ). The increase in particle loading increases the total am ount of drug release, but it is accompanied by an increase in the release duration as well. The increase in total amount of drug released is expected because an increased particle loading results in an increased total drug loading in the insert. The incr ease in the drug release duration is interesting, particularly since the drug release rates are controlled by release from the particles and so an increase in particle loading is not expected to impact release durations. The increased duration could poten tially be caused by an increased crosslinking in the gel due to the presence of the particles and also due to the PGT fraction that remained unreacted during the particle preparation step. To explore

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103 whether diffusion limitations due to increased crosslin king are leading to the increased release duration, the cumulative release profiles are plotted as a function of t 1/2 in Figure 4 3B. If diffusion is the rate limiting step, the cumulative drug release should scale as t 1/2 at very short times, and thus th e plot of mass release as a function of t 1/2 should be linear for about 15 20% of the cumulative release from a cylindrical device [99] The curves in Figure 4 3B are clearly not linear for the first 15 20% release showing that diffusion through the gel is not rate limiting for any of the four cases shown in Figure 4 3. The likely rate controlling mechanism and the potential reasons for incre ase in release duration with an increase in particle loadings are described Section 4.2.4 4. 2.3 Comparison of Release Rates from Inserts to Therapeutic Doses The usual dose of timolol is one drop of 0.25% timolol maleate in the affected eye(s) twice per a The bioavailability of timolol delivered through eye drops is only about 1 2%, which implies that the therapeutic requirement of timolol is about 2.5 he bioavailability of drug released through the fornex inserts could be higher than that for eye drops based on the fact that the total dosage of pilocarpine administered by one Ocusert system over 7 days was about one eighth of the amount provided by the 28 applications (4 each day) of the 2% eyedrops [99]. Thus, a therapeutically desired release rate for the timolol insert could be about 15 insert with 25% particles release about 15 be suitable for use as a 1 week release system. The inserts with higher particle loadings release timolol over longer durations and also at rates closer to zero order. However the release rates decrease with increased particle loadings, and the insert with 100% pa rticles release only about 3 day. Although the release rates could

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104 potentially be increased by either increasing the length of the insert or by increasing drug loadings in the particles, wear duration of longer than a week may be difficult to achieve d ue to retention issues. 4. 2 4 Effect of Timolol Loading in the Particles on Drug Release To explore whether the release rates from the inserts could be increased by increasing the ratio of timolol to PGT in the particles, we prepared nanoparticles by follo wing the same procedure as described in the Methods section except that 360 mg of timolol base was added to 1 g of PGT. The drug release profiles from 7.5 mm insert with 25% particle loadings and increased drug percentage in the particles are shown in Fig ure 4 4, along with the profiles for the similar insert with the lower drug fraction in the particles. The data shows that increased drug loading in the particles leads to an increase in the release rates without impacting the total release duration. How ever the increase in the release rates for the first week is only about 33% even though the drug loadings in the particles increased by about 50%. This suggests that the encapsulation efficiency decreases with increasing drug loadings in the particles. 4 2 5 Mechanism of Release The mechanism of extended drug release from p HEMA contact lenses loaded with the timolol PGT particles was shown previously to be hydrolysis of the ester bond between timolol and the particle network (Section 2.2.10). Accordingly the drug release data should fit the first order reaction model, i.e., ( 4 1) where M is the amount of drug released at a time t and M is the amount of drug released after infinite time, and k is the rat e constant. In the previous studies with contact lenses, the particle loadings were substantially smaller. To test whether the

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105 same model can describe the drug release profiles even at very high particle loadings in the inserts, the release data in Figur e 4 3A was fitted to Eq. 4 1. The best fit curves are included in the figure and the parameters for the fits are included in Table 4 1. The model fits the data well proving that the release is controlled by the hydrolysis reaction. The rate constant k d ecreases with an increase in the particle loading because of the reduction in the water content, which is required for the hydrolysis reaction. The decrease in k leads to an increase in the total release duration. The value of M increases linearly with increasing particle loadings but then levels off suggesting that there is an increase in irreversibly trapped drug fraction with an increase in particle loading. 4. 2 6 Effect of Packaging To explore the effect of packaging, the particle laden inserts were soaked in 1 ml of packaging solution (PBS) for a period of 3 months at refrigerator and then subjected to drug release studies with protocols described in the Methods section. The results of these studies are presented in Figure 4 5 The release pr ofiles exhibit a slight burst release, followed by an extended release for about 20 days. The burst release is due to diffusion of the drug that was released from the particles into the gel due to hydrolysis during the packaging duration. Although hydrol ysis is slowed down in refrigerator, the burst release after packaging is undesirable, and thus it would be preferable to package the inserts in the dry state and hydrate them just prior to insertion. 4.3 Conclusion Glaucoma therapy through eye drops has several drawbacks including low bioavailability and low compliance. Extended release of glaucoma drugs through ophthalmic inserts could potentially increase both bioavailability and compliance.

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106 We have developed a particle loaded insert for extended delivery of beta blocker timolol. The release is controlled by hydrolysis of the ester bond that links timolol to the particle matrix. The release durations could be adjusted from about 10 days to a month by changing t he fraction of particles in the insert. The release rates could also be adjusted by changing the length and/or the drug loading in the particles. Based on approximate estimation a 7.5 mm and 1 mm diameter insert with 25% particles could be suitable for e xtended delivery for about a week. The inserts should be stored in dry state and hydrated prior to insertion. The approach of incorporation of timolol encapsulated nano particles into the conjunctival inserts and puncta plugs to prolong the drug release c ould potentially find application in several other drug delivery application such as such as ophtha coils, retinal implants, transdermal patches, wound healing patches, etc. While the results of this in vitro design study are encouraging, these need to be supplemented with animal studies to explore retention and therapeutic efficacy.

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107 Figure 4 1 Image of the nanopaticle laden fornix insert

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108 Figure 4 2 Cumulative drug release profiles from control pHEMA (without n anoparticles) inserts. Data is shown as mean std (n = 3) 0 100 200 300 400 500 600 700 0 0.5 1 1.5 2 2.5 Drug release ( g) Time (day)

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109 Figure 4 3 Cumulative drug release profiles fro m nanoparticle lo aded inserts. A) The four curves correspond to different particle loadings. B) The release is plotted as a function of square root of time to determine the transport mechanism. Data is shown as mean std (n = 3) 0 100 200 300 400 500 600 700 0 20 40 60 80 100 120 Drug release ( g) Time (day) model fit Insert (25:75) Insert (50:50) Insert (75:25) Insert (100:0) 0 100 200 300 400 500 600 0 2 4 6 8 10 12 Drug release ( g) Time^0.5 (days^0.5) model fit Insert (25:75) Insert (50:50) Insert (75:25) Insert (100:0) B A

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110 Figure 4 4 Cumulative drug release profiles from nanoparticle loaded inserts. The two curves correspond to different drug loadings in the particle 0 50 100 150 200 250 300 350 0 10 20 30 40 50 Drug release ( g) Time (day) T360 T240

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111 Figure 4 5 Cumulative drug release p rofiles from nanoparticle loaded inserts with 25% particle loading after 3 month packaging at 4 C 0 50 100 150 200 250 0 20 40 60 80 Drug release ( g) Time (day)

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112 Table 4 1 Dependency of total mass of drug released (M ) and the rate constant at different particl e loading Particle loading (%) M (g) k M total (g) 25 273.89 38.69 0. 1159 0.0 18 420 50 554.27 74.63 0.0360 0.0 07 840 75 682.82 13.62 0. 0184 0. 004 1260 100 696.36 59.81 0.0062 0. 0005 1680

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113 CHAPTER 5 UV BLOCKING NANOPARTICLE LOADED IN SILICONE HYDROGLE CONTACT LENSES The effect of UV radiation on eyes is now a big issue, because it may result in contracting diseases of ocular surface. Several researchers are currently exploring UV blocking contact lenses that prov ide significant defense against UV radiation. The focus of our research is to develop contact lenses that retain the loaded UV blocking materials, entrapped in highly crosslinked nanoparticles during polymerization and measure the transmittance for UVA and UVB. A new approach has been established to design the UV blocking particles by preparing highly crosslinked nanoparticles by polymerization of an emulsion of a monomer with multi vinyl functionalities of PGT (propoxylated glyceryl triacylate) in presenc e of oily diluents. Hydrophobic oily UV blocking materials can be loaded into the particles by adding into the polymerization mixture, which leads to trapping of the UV blocking materials. The UV protecting nanoparticles were soaked in ethanol, and then si licone hydrogels uptake the UV protecting nanoparticles. Those materials were made of several particle concentrations in ethanol and various UV blocking materials concentration in particles. Our studies indicate that the silicone hydrogel lenses containin g UV blocking materials loaded nanoparticles are transparent and can block UVA (315 400nm) and UVB (280 315 nm). The results of the various concentration show that some silicone hydrogel can reach the class 1 UV blocking. Currently, only three commercial c ontact lenses are the class 1 UV blocker. The classification of UV blocker will be described in details later.

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114 5.1 Materials and Methods 5.1.1 Materials N,N Dimethylacrylamide (DMA), 1 vinyl 2 phosphate buffered saline (PBS) were purchased from Aldrich Chemicals (St Louis, MO). Propoxylated glyceryl triacrylate (PGT) was purchased from Sartomer; Benzoyl peroxide (BP) (97%) and 1,3 diphenyl 1,3 propanedione (DP) were purchased from Aldrich Chemicals (Milwaukee, WI). The macrom er bis alpha, mega (methacryloxypropyl) polydimethylsiloxane (Macromer) was supplied by Clariant. 3 methacryloxypropyl tris(trimethylsiloxy)silane (TRIS) was gifted by Silar laboratories (Scotia, NY). Methyacrylic acid (MAA) was purchased by Polysciences, Inc (Wattingyon, PA). 2,4,6 Trimethylbenzoyl diphenyl phophineoxide (Darocur TPO) was kindly provided by Ciba (Tarrytown, NY). Vitamin E (D alpha tocopherol, Covitol F1370) was gifted by Cogins Corporation. 5.1.2 Preparation of Highly Crosslinked PGT Nanop articles with DP The UV blocker 1,3 diphenyl 1,3 propanedione (DP) is incorporated into highly crosslinked particles of propoxylated glyceryl triacrylate (PGT). The DP loaded PGT particles were prepared by a novel approach of using a diluent as a polymerization modifier. Vitamin E was chosen as the diluent due to its hydrophobicity a nd also its biocompatibility. To prepare the nanoparticles DP, PGT and Vitamin E are mixed in the desired ratios and added to 7.5 mg of polymerization initiator BP. T he mixture is added to DI water, purged with nitrogen for 15 min to remove the dissolved oxygen, and then heated in an 80 C hot water bath under stirring a t 1100 rpm for 8 hours The thermal polymerization results in formation of UV blocking nanoparticles The particles are separated from the suspension by centrifugation for 15 min

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115 The compositions chosen for preparing the nanoparticles are listed in Tables 1 and 2. In each composition varying amounts of DP were added to 120 mg of Vitamin E, 1 g of PGT and 7.5 mg of BP. The amount of DP was varied from 10 to 210 mg to create particles with DP fractions ranging from about 1% to 15.8 %. To further increase the DP fraction, it was necessary to increase the Vitamin E fraction as well. Accordingly 280 or 4 50 mg of DP were added to 150 mg Vitamin E, 1 g of PGT and 7.5 mg BP to produce particles with DP loadings of about 19.6 and 28.1%, respectively. 5.1 3 Preparation of Silicone gels Extended wear contact lenses are prepared from silicone hydrogels to obtain high oxygen and ion diffusion. The silicone gels are synthesized by free radical polymerization of a mixture of a silicone monomer with a hydrophilic monomer. Additionally a macromer is added to ensure solubilization of the silicone and the hydrophilic m onomers. Also some other components such as NVP are added to increase the water content. To prepare the silicone hydrogel, 0.8 ml of macromer bis alpha,omega (methacryloxypropyl) polydimethylsiloxane is added to 0.56 ml of N,N dimethylacrylamide ( DMA ) 0 .24 ml of methylacrylic acid ( MAA ) 0.8 ml of 3 m ethacryloxypropyl tris(trimethylsiloxy)silane ( Tris ) 0.12 ml of 1 vinyl 2 pyrrolidone (NVP) and 10 l of Propoxylated glyceryl triacrylate (PGT) The mixture is purged by bubbling nitrogen for 15 min. After adding 0.012g of the initiator Darocur TPO with stirring for 5 min, the mixture is poured in between two glass plates separated by a 100 m thick plastic spacer. The mold is then placed on Ultraviolet transilluminiator UVB 10 (UltraLum, Inc.) and irrad iated with UVB light (305 nm) for 5 0 min. The molded gel was cut into circular pieces (about 1.65 cm in diameter) with a cork borer and dried in air overnight before further use.

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116 5.1 .4 Preparation of Nanoparticle Loaded in Silicone Gels The UV blocking nan oparticles were loaded into the silicone hydrogels either by direct addition to the polymerization mixture or by soaking the preformed gel in a solution of nanoparticles in ethanol. 5. 1 5 Particles Added to Polymerization Mixture To prepare the silicone hy drogel with nanoparticles, varying amounts of DP nanoparticles were added to the composition described above. Specifically, 0.13, 0.25, or 0.504g of DP loaded nanoparticles were added to a mixture of 0.8 ml of macromer bis alpha,omega (methacryloxypropyl) polydimethylsiloxane, 0.56 ml of N,N dimethylacrylamide ( DMA ) 0.24 ml of methylacrylic acid ( MAA ) 0.8 ml of 3 m ethacryloxypropyl tris(trimethylsiloxy)silane ( Tris ) 0.12 ml of 1 vinyl 2 pyrrolidone (NVP) and 100 l of Propoxylated glyceryl triacrylate (PGT) The mixture is purged by bubbling nitrogen for 15 min. After adding 0.012g of the initiator Darocur TPO with stirring for 5 min, the mixture is poured in between two glass plates separated by a 100 ck plastic spacer. The molds were then placed on Ultraviolet (Ultra Lum, Inc.) and irradiated with UVB light (305 nm) for 2 hours. The duration of polymerization was chosen to be longer due to the attenuation in UV intensity because of the absorption by the particles. The details of the experimental designs regarding the DP loading in the nanoparticles and the nanoparticle loading in the polymerization mixture are presented in Table 5 1. Since the presence of the UV blocking particles inhibit UV initiat ed polymerization, thermally initiated polymerization was utilized in some cases to cure the gel. The mold was placed in an oven at 80C for 24 hrs to polymerize the gel. The thermal

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117 polymerization was utilized only for loading 5% of UV blocking particle s with 5.9% DP. Also in this case the gel thickness was increased to 200 microns. 5. 1 6 Parti cles Loaded by Soaking the Contact Lens in a Solution of Particles in Ethanol The silicone gels prepared with the procedures described above were soaked in 95% et hanol to remove the unreacted monomers To load the nanoparticles into the silicone gel, each gel was soaked in a 3 ml of a solution of nanoparticles in ethanol for 3 hours. After 3 hours of loading, the gels w ere withdrawn and submerged in 200 ml of DI w ater for 2 hours to extract ethanol from the gels. The gels were then withdrawn, gently wiped and then soaked in 5ml of DI water for further use. The concentration of nanoparticles in the ethanol solution was varied from 1 to 10%. The details of the expe rimental designs regarding the DP loading in the nanoparticles and the nanoparticle concentration in ethanol are presented in Table 5 2. 5. 1 7 Transmittance Measurements The transmittance of nano particle laden silicone hydrogels was measured using UV Vis spectrophotometer (Thermospectronic Genesys 10 UV). The lenses were hydrated by soaking in DI water overnight, and then mounted on the outer surface of a quartz cuvette. The cuvette was placed in a spectrophotometer and the transmittance was measured at wa velengths ranging from 200 nm to 480 nm. The transmittance data is utilized to calculate the average blocking in the range of UVC, UVB and UVA radiation. For the UVA range, average blocking is determined for the range 315 380 which is the range considere d for ANSI classification and also for 315 400 which represents the entire UVA spectrum.

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118 5.2 Results and Discussion 5. 2 1 Incorporation of Particles into Lenses by Adding Particles to the Polymerization Mixture 5.2.1.1 Effect of Particle Loading on UV blocking The transmission spectra in the 200 nm to 480 nm range from the control silicone hydrogel and those from gels loaded with particles by addition of particles to the polymerization mixture are shown in Figures 5 1. Each figure corresponds to partic les with a fixed loading of the UV blocker and the various curves in each figure correspond to multiple loading of the particles in the gels. Also the average absorbance was calculated in the ranges of UVA, UVB and UVC lights and these averages are listed in Tables 5 3 ,4,5,6 For the UVA range, average blocking is determined for the range 316 380 (Table 5 5 ) which is the range considered for ANSI classification and also for 316 400 (Table 5 6 ) which represents the entire UVA spectrum. Figure 5 1 A shows th e absorption spectra (% absorption) for wavelengths from 200 nm to 480 nm from gels loaded with varying percentage of particles with about 1% DP in the particles. The data clearly shows significant reduction in transmittance in the UV range due to loading DP containing nanoparticles in the gels. Also, the transmittance decreases with an increasing in the particle loading but the decrease is non linear suggesting that the UV attenuation is beyond the linear regime of the Beer sorption for the UVA, UVB and UVC ranges listed in Tables 5 3 ,4,5,6 show that gels loaded with 20% particles achieve class 1 UV blocking classification. The Figure also shows the absorption spectra from gels loaded with 20% particles that did not contain any DP. These gels also exhibit significant reduction in UV transmission due to the UV blocking from vitamin E which is also encapsulated in the

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1 19 particles. While the gels with 20% loading achieved class 1 blocking, the high particle loading could impact other lens properties such as modulus, oxygen and ion permeability, etc and thus it is important to achieve class 1 blocking with a much lower particle loading. To improve the UV blocking from the lenses, particles with higher DP loading were prepared and dispersed in the lenses by direct addition to the polymerization mixture. Figure 1b shows the % transmittance of from gels loaded with particles containing 2.6% DP. The particles with 2.6% DP were loaded at 5, 10 and 20% w/w in the gels (Figure 5 1 B ). Th e gels with 5% particles blocked UVC almost completely, but blocked only 91% UVB and thus the lenses do not qualify for the class 2 UV blocking classification. However, on increasing the loading of the particles to 10 and 20% w/w, the transmittance decrea sed to 0.8% and ~0%, respectively for the UVB light. Furthermore, these lenses blocked 84% and 87% in UVA ranges (316 400nm) respectively, and 98 and 99% in the 316 380 nm range which places the gels with 5% or 10% of the particles in the category of class 1 lenses. Figures 5 1C and D show the transmittance for gels loaded with particles with 5.9% and 9.7% DP, respectively. Gels were prepared with 5 and 10% particles for the 5.9% DP and with only 5% loading for the 9.7% DP particles. In all three cases th e gels block sufficient UV radiation to be classified as class 1 UV blocking gels. Thus, increasing DP amount in particles and increasing particle loading amount in silicone gels continuously improved the UV blocking. However, the silicone gels also beco me progressively softer due to the decrease in the polymerization rates due to the blocking of the UV light during polymerization. This effect could be overcome by thermal

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120 polymerization and so gels with 5% loading of the 5.9% DP particles were also cured by placing the mold in an oven at 80 0 C for 24 hours. These gels were prepared to be 200 micron thick. The UV blocking from these gels was adequate for the class 1 classification (Figure 5 1 E ). These gels blocked more UV light compared to the 100 micr on gels with the same particle loading because of the increased thickness. The intensity of light decreases exponentially with thickness (assuming uniform properties), and thus 200 micron thick gels block more radiation compared to a 100 micron thick gel. Thus, it is important to note that the class 1 or 2 classification also depends on the thickness of the contact lenses. The 200 thermally polymerized micron thick gels with 5% of 5.9% DP particles were also heated at 100 0 C for 2 hours to explore the st ability of the UV blocking effect after exposure to high temperatures. The gels retained class 1 classification and the UV blocking was only slightly reduced. Incorporation of particles into the lenses by direct addition to the polymerization solution can be used to prepare UV blocking lenses but it would require modifications to the curing procedures currently used by contact lens manufacturers. It was thus decided to explore the possibility of loading the particles into preformed gels/lenses by soaking the gels/lenses in a solution of DP loaded particles in ethanol. These results are described Section 5.2.2 5. 2 2 Incorporation of Particles into Lenses by Soaking the Lenses in Solution of Particles in Ethanol 5.2.2.1 Mass of Particles Loaded To prove that soaking gels in a solution of nanoparticles in ethanol can lead to diffusion of particles into the gels, we soaked the gels in a solution of 10.2% DP nanoparticles in ethanol and measured the weight gain of the gels after drying. The

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121 particle loading in the gels was varied from 2 to 10%. The dry weight of the gels both before and after particle loading and the % increase in weights are listed in Table 5 7 As shown Table 5 1, the increase in gel weight was proportional to the concentration of particles in ethanol. The particles diffused into the lenses during soaking because the pore size of the gels increases significantly in ethanol compared to that in PBS. DP containi ng nanoparticles were also loaded into commercial silicone hydrogels by soaking in a solution of 2% nanoparticles with 5.9% DP in ethanol. The dry mass of contact lenses both before and after particle loading, along with the fractional weight gain are lis ted in Table 5 8 The commercial contact lenses load a larger mass of nanoparticles compared to the gels prepared in our lab. Since contact lenses are typically autoclaved for sterilization, we also autoclaved the lenses after particle loading. The dry weight of the lenses after autoclaving is also included in Table 5 8 The data shows that the particles that diffuse into the lenses during soaking in ethanol are retained in the contact lenses even during autoclaving in PBS due to the significant decreas e in the pore size and the hydophobicity of the particles. Furthermore soaking in PBS at room temperature for extended periods also did not lead to any significant leaching out of the particles from the gels or the contact lenses. 5.2.2.2 Effect of UV Blo cking Particles on Transmittance Spectra of the Lenses The transmission spectra in the 200 nm to 480 nm range from the control silicone hydrogel and those from gels loaded with particles by soaking the control gel in a solution of particles in ethanol are shown in Figures 5 2. Each figure corresponds to particles with a fixed loading of the UV blocker and the various curves in each figure correspond to multiple loading of the particles in the ethanol. Also the average

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122 absorbance was calculated in the rang es of UVA (316 380nm and 316 400nm), UVB and UVC lights and these averages are listed in Tables 5 9,10,11,12 Figure 5 2 A shows the absorption spectra from gels soaked in solutions of particles with 10.2% DP in ethanol. The absorbance from the control gel is compared with those from gels soaked in 2%, 5% and 10% particle solutions in ethanol. The data clearly shows signific ant reduction in transmittance in the UV range due to partitioning of DP containing nanoparticles in the gels during soaking in ethanol. The nanoparticle loading in the gels is proportional to that in ethanol but the transmission spectra is only slightly different the three cases suggesting that the UV attenuation is beyond the linear ranges listed in Tables 5 9,10,11,12 show that gels soaked in the solution of particles wi th 5% DP block UVB and UVC radiation almost entirely. The gels block about 84, 87 and 89% of the UVA (316 400 nm) radiation for the cases of soaking in 2, 5 and 10% solutions in ethanol, and block more than 95% of the radiation in 316 380nm range and thu s can be classified as class 1 UV blockers. To achieve class 1 UV blocking with a lower particle loading and also to improve the UV blocking from the lenses beyond the class 1 requirements, particles with higher DP loading were prepared and loaded into the gels by soaking in solutions of particles in ethanol. Figure 5 2 B shows the % transmittance of from gels soaked in ethanol solutions of particles containing 15.8% DP. Lenses were soaked in ethanol solutions with of different particle concentrations (1,2 ,5,10,and 20%) to prepare gels with five different particle loadings. The curves in Figure 5 2 B and the average absorba nce values in Tables 5 9,10,11,12 show that all five gels almost completely blocked UVC,

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123 UVB and UVA (316 380 nm) radiation. The gels s oaked in 1 or 2% solutions block less than 90% UVA (316 400 nm) but the gels soaked in 5, 10 and 20% solutions blocked 90.31, 92.95 and 94.36% UVA (316 400 nm) radiation, respectively and these lenses can be considered as superior to class 1 UV blockers. P article loading in the gels clearly reduced UV transmittance but could potentially also impact other properties such as ion and oxygen permeability and so it is desirable to reach the class 1 blocking with minimum particle loading. Gels were loaded with n anoparticles containing 19.6% DP and the transmittance profiles are shown in Figure 5 2 C for several different concentrations of particles in the ethanol solution. The gels loaded with the 19.6% DP particles also almost completely block UVC, UVB and UVA ( 316 380 nm). The gels soaked in 7 and 10% solutions also reduce transmittance in the visible range due to a partial loss in transparency and are thus not suitable for use as contact lenses. The gels soaked in 5% solutions block 92.10% UVA (316 400 nm) an d thus can be considered to have UV blocking superior to class 1 blockers. In Figure 5 2D transmittance spectra are shown for particles soaked in ethanol solutions of particles containing 28.1% DP. In this case lenses soaked in 10% solutions are partial ly opaque. The gels soaked in 5% are class 1 blockers but even these block some visible light due to a slight loss of transparency. Gels soaked in 1 and 2% solutions are completely transparent and block sufficient UV radiation for class 1 classification. 5.2.2.3 Effect of Autoclaving Contact lenses are sterilized by autoclaving so it is important that the particles retain the UV blocking efficiency after autoclaving. Particles with 15.8 % DP were autoclaved at 121 0 C for a total cycle time of 1 hour and t hen loaded into the silicone hydrogel by soaking the gel in solution of the autoclaved particles in ethanol. The

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124 concentration of particles in ethanol was chosen to be 1, 5 and 10% (Figure 5 2 E ). The transmission spectra from the lenses loaded with autoc laved particles is compared with those loaded with the un autoclaved particles but with the same 15.8% DP loading in particles and the same 1, 5 and 10% particle concentration in solution (Figure 5 3). The spectra in Figure 5 3 are relatively unaffected b y autoclaving proving that the DP loaded highly crosslinked particles maintain the UV absorption efficiency after autoclaving. 5. 2 3 UV Absorption in Commercial Contact Lenses Loaded with DP Containing Nanoparticles Commercial contact lenses (Night & Day and Pure Vision) were soaked in 3 m of ethanol solution with 2 and 4% particles that were loaded with 19.6% DP. The procedures for loading the particles in the commercial lenses were identical to those described in the Methods section for loading par ticles in pre polymerized silicone hydrogels. The particle loaded lenses were autoclaved at 121 o C for 1 hour cycle time. The transmission spectra from the particle loaded contact lenses, both before and after autoclaving are shown in Figure 5 4 A and B T he specific % of UV blocking are shown Table 5 13 In all cases the particle loaded contact lenses almost completely blocked UVA and UVB radiation. The contact lenses soaked in 2 and 4 % solutions blocked 99.85 and 99.9% (Night and Day) and 99.9 and 99.9% (Pure Vision), respectively and thus can be classified as class 1 UV blockers. The % absorbance in 316 400 nm was 88.57 and 93.82 % (Night and Day) and 91.57 and 96.63 % (Pure Vision). 5.3 Conclusion The UV radiation can potentially cause significant dam age to ocular tissues leading to moderate to severe problems. Contact lenses can be more effective than glasses in reducing the UV exposure due to better protection against peripheral

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125 radiation. It is thus important to impact UV blocking feature to all c ontact lenses. While the importance of UV blocking in contact lenses has been recognized, there are only a few commercial contact lenses that provide Class 1 UV protection. We have developed a novel approach of incorporating UV blocking into contact lens es by dispersing nanoparticles that contain a UV blocking compound. The extent of UV blocking can be tailored by either manipulating the concentration of the UV blocking compound in the particles or by manipulating the particle loading in the gels. The n anoparticles are designed such that the UV blocking compound does not diffuse out of the particles due to the high croslinking in the particles. Furthermore the particles are prepared through a novel approach that leads to a particle size of about 4 nm wi thout using any surfactant. The particles can be loaded into the gel by addition to the polymerization mixture. However curing gels that contain UV blockers is difficult due to the attenuation of the generation of free radicals because of the absorption of the UV light typically used in polymerization. This problem could be overcome by either increasing the intensity or duration of UV curing or by using thyermal curing. Additionally, the ultrasmall particle size allows loading of particles into the lens es after polymerization by soaking the lenses in a solution of particles in ethanol. The pore size of the lenses is much larger in ethanol compared to PBS so the particles that diffuse into the lenses during ethanol soaking stay trapped in the lenses afte r ethanol is extracted and the lenses are hydrated in PBS. The particles do not leach out of contact lenses even during autoclaving. Also the UV blocking efficiency of the particles is retained during autoclaving. Several of the gels loaded with the pa rticles achieve class 2 and a few also achieve the class 1 UV block classification. The approach developed here can be

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126 integrated into any commercial contact lens, and also possibly into other biomedical devices.

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127 Figure 5 1 Transmittance (%) spectra for 100 m thick gels loaded with particles by direct addition to the polymerization mixture. A) 1% DP in the particles B) 2.6% DP in the particles C) 5.9% DP in the particles D) 9.7% DP in the particles E) 5% loading of particles with 5.9% DP after a thermal processing 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 20% of vitamin E 5% 20% 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 5% 10% 20% A B

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128 Figure 5 1 Continued 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 5% 10% 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 5% C D

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129 Figure 5 1 Continued 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 5% (before thermal treatment) 5% (after thermal treatment) E

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130 Figure 5 2 Transmittance (%) spectra for 100 m thick gels loaded with particles by soaking in a solution of ethanol with particles. A)10.2 % DP in ethanol. B) 15.8 % DP in ethanol. C) 19.6% DP in ethanol. D ) 28.1% DP in ethanol. E) 15.8 % DP that were autoclaved pr ior to addition to the ethanol 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 2% 5% 10% 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 1% 2% 5% 10% A B

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131 Figure 5 2 Continued 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 1% 2% 5% 7% 10% 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 1% 2% 5% 10% C D

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132 F igure 5 2 Continued 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 1% 5% 10% E E

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133 Figure 5 3 Effect of autoclaving on transmittance of gels loaded with particles contained 15.8% DP by soaking in a solution of particles in ethanol 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) control 1% 5% 10% 1% autoclave 5% autoclave 10% autoclave

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134 Figure 5 4 Transmittance spectra from commercial lenses of t he particles contained 19.6% DP A ) Night and Day B ) PureVision 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) N&D N&D_2% N&D_4% N&D_2% _autoclave N&D_4%_ autoclave 0 10 20 30 40 50 60 70 80 90 100 200 250 300 350 400 450 500 % Transmittance Wavelength (nm) PureVision PV_2% PV_4% PV_2% _autoclave PV_4% autoclave A B

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135 Table 5 1 Experimental design of particles in polymerization mixture % of UV blocking dye in particle Amount (mg) of UV blocking dye used Amount (mg) of Vitamin E % of particles in silicone polymerization mixture 1 10 120 5,10 2.6 30 120 5,10,20 5.9 70 120 5,10 9.7 120 120 5 Table 5 2 Experimental design of soaking the lenses in solution of particles in ethanol % of UV blocking dye in particle Amount (mg) of UV blocking dye used Amount (mg) of Vitamin E % of particles in ethanol 10.2 126.8 120 2,5,10 15.8 210 120 1,2,5,10,20 19.6 280 150 1,2,5,7,10 28.1 450 150 1,2,5,10

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136 Table 5 3 The percentage of UVC (below 280 nm) blocking range in nanoparticle laden silicone gels % of UV blocking dye in particle % of particle loading in polymerization 5 10 20 1 98.6 99.9 2.6 99.0 99.9 99.9 5.9 99.6, 99.9(thermo) 99.9 9.7 99.9 Table 5 4 The percentage of UVB (280 315 nm) blocking range in nanoparticle laden silicone gels % of UV blocking dye in particle % of particle loading in polymerization 5 10 20 1 7 9.6 99. 2 2.6 83.6 99.2 99.9 5.9 99.5 99.9(thermo) 99.9 9.7 99.9

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137 Table 5 5 The percentage of UVA (315 380 nm) blocking range in nanoparticle laden silicone gels % of UV blocking dye in particle % of particle loading in polymerization 5 10 20 1 61.0 95 6 2.6 90.6 98.2 99. 4 5.9 98.0, 99.8 (thermo) 99.5 9.7 99.7 Table 5 6 The percentage of UVA (315 400 nm) blocking range in nanoparticle laden silicone gels % of UV blocking dye in particle % of particle loading in polymerization 5 10 20 1 50.9 81.8 2.6 67.7 84.0 86.7 5.9 83.1, 91.9 (thermo) 87.2 9.7 94.2

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138 Table 5 7 The weight changes of silicone lenses after soaked in 10.2% DP in nanoparticles with ethanol solutions % of particle loading in EtOH Dry weight of silicone gel (g) Dry weight change (g) After soaked in nanoparticle and EtOH % of weight changes 2 0.0235 0.0237 0.85 5 0.0232 0.0241 3.89 10 0.0240 0.0254 5.83 Table 5 8 The weight changes of commercial contact lenses after autoclaving: 5.9 % of UV blocking dye loading in particles and then 2% particles in EtOH Dry weight (g) Control lenses UV blocking laden lenses % of weight changes After autoclaving of UV blocking laden lenses Acuvue Oasys 0.0219 0.0228 4.11 0.0225 O 2 Optix 0.0196 0.0202 3.06 0.0200 Pure Vision 0.0218 0.0226 3.67 0.0226 Night & Day 0.0223 0.0231 3.59 0.0230 Acuvue Advance 0.0200 0.0211 5.5 0.0211

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139 Table 5 9 The percentage of UVC (below 280 nm) blocking range in nanoparticle laden silicone gels % of UV blocking dye in particle % of particle loading in EtOH 1 2 5 7 10 20 10.2 99.5 99.9 99.9 15.8 99.7 99.9 99.9 99.9 99.9 19.6 99.43 99.9 99.9 99.9 99.9 28.1 99.89 99.9 99.9 99.9 Table 5 10 The percentage of UVB (280 315 nm) blocking range in nanoparticle laden silicone gels % of UV blocking dye in particle % of particle loading in EtOH 1 2 5 7 10 20 10.2 99.0 99.9 99.9 15.8 99.4 99.9 99.9 99.9 99.9 19.6 99.7 99.9 99.9 99.9 99.9 28.1 99.9 99.9 99.9 99.9

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140 Table 5 11 The percentage of UVA (315 380 nm) blocking range in nanoparticle laden silicone gels % of UV blocking dye in particle % of particle loading in EtOH 1 2 5 7 10 20 10.2 99.36 99.82 99.89 15.8 99.45 99.76 99.89 99.9 99.9 19.6 99.27 99.77 99.9 99.9 99.9 28.1 99.81 99.88 99.9 99.9 Table 5 12 The percentage of UVA (315 400 nm) blocking range in nanoparticle laden silicone gels % of UV blocking dye in particle % of particle loading in EtOH 1 2 5 7 10 20 10.2 84.38 87.05 89.38 15.8 85.25 86.21 90.31 92.95 94.36 19.6 83.63 86.17 92.10 92.39 93.86 28.1 86.73 88.48 93.54 99.61

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141 Table 5 13 The percentage of UV blocking range in 19.6% DP nanoparticle in Nigh&Day and PureVision Contact lenses Particle loading UVC UVB UVA (315 380 nm) UVA (315 400nm) Night & Day Control 82.27 25.17 36.92 29.58 2% particle in EtOH 99.9 99.9 99.85 88.57 After autocalving 99.9 99.9 99.75 88.00 4% particle in EtOH 99.9 99.9 99.9 93.82 After autocalving 99.9 99.9 99.9 90.34 Pure Vision Control 94.38 62.66 32.54 28.34 2% particle in EtOH 99.9 99.9 99.9 91.57 After autocalving 99.9 99.9 99.9 92.37 4% particle in EtOH 99.9 99.9 99.9 96.63 After autocalving 99.9 99.9 99.9 94.08

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142 CHAPTER 6 CONCLUSION M any ocular diseases such as glaucoma, dry eye syndrome, and infections are being treated by eye drops. However, o phthalmic drug delivered by eye drops is inefficient because of its low bioavailability. Only about 5% or less of the drug is absorbed into the cornea for several reasons including rapid tear turnover and nonproductive absorption. T hese problems may be ove rcome by the use of contact lenses that increase bioavailability to about 50% Such higher bioavailability decreases drug wastage and the side effects and increases residence time in cornea However it is important to note that commercial contact lenses are not suitable for delivering ophthalmic drugs because these release drugs only ab o ut 1 2 hours, which is a very short duration. Our study has conclusively shown the extended ophthalmic drug delivery by pHEMA ,silicone hydrogel contact lens and insert containing nanoparticles. In Chapter 2 and 3, we showed that incorp oration of highly crosslinked nanoparticles can increase drug release duration from a few hours to about 1 2 months. Nanoparticles are shown to control the release duration of timolol. Com parison between PGT highly crosslinked nanoparticles and surfactant in nanoparticles are exhibit ed by similar drug release profile of the timolol and the same particle size. Moreover, EGDMA having two vinyl groups and PGT having three vinyl groups have sig nificantly different drug release rates Thus, we focused on highly crosslinker PGT nanoparticles which were synthesized of emulsion polymerization. Furthermore, in Chapter 2 several properties including particle size distribution, transmittance, modulus, EWC, ion permeability and oxygen permeability are

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143 characterized to determine the pros and cons of loading timolol laden nanoparticles into buffered saline (PBS) for an extended period of time, 3~4 weeks at therapeutic dose. Nanoparticles are shown to control the release duration of timolol. Comparison between PGT highly crosslinked nanoparticles and surfactant in nanoparticles are exhibit ed by similar drug release profi le of the timolol and the same particle size. Moreover, EGDMA having two vinyl groups and PGT having three vinyl groups have significantly different drug release rates Thus, we focused on highly crosslinker PGT nanoparticles which were synthesized of emulsion polymerization. As shown in Chapter 2, t he release from the pHEMA gels after packaging is affected by the temperature, which is close to the desired release rates at refrig erator. In addition at various temperatures PGT nanoparticle laden pHEMA gels were shown significantly different timolol release amount and release duration. It could be use to control drug release rate at different temperatures. While particle addition increases the release duration, it has the undesirable effect of increase in modulus and decrease in water content. A major deficiency of these systems is a modulus of pHEMA gels. Due to highly crosslinked nanoparticles, these lenses may have to be used d iluents such as Vitamin E for minimized modu lus during preparing the gels. Chapter 3 is also shown to develop extended wear contact lenses that can also provide extended glaucoma therapy. We disperse nano particles of PGT (propoxylated glyceryl triacylate) that contain a glaucoma drug timolol. The particles can also be loaded into prefabricated lenses by soaking the lenses in a solution of particles in ethanol. T he release profiles fit a first order reaction with temperature dependent rate

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144 constant. Nano particle incorporation in the silicone hydrogels results in reduction in ion and oxygen permeabilities, and an increase in modulus, and the impact on each of these properties is proportional to the particle loading. A gel with 5% particle loading can deli ver timolol at therapeutic doses for about a month at room temperature, with a minimal impac t on critical lens properties. Preliminary animal studies in Beagle dogs conducted with lenses in which particles are loaded by soaking the lenses in ethanol show a reduction in IOP. Chapter 4 focuses on developing timolol laden nanoparticles into fornix insert for an extended period of time. PGT nanoparticles including timolol is used in this Chapter studies. The 1 mm diameter 7.5 mm long insert with 25% particles c an release timolol for about 10 days at an average rate of about 15 g/day. Longer release durations could also be achieved but the maximum wear duration would likely be limited by retention duration. The mechanism of release is hydrolysis of the ester bon d that links timolol to the PGT matrix, and thus the release profiles fit a first order reaction. The rate constant for the hydrolysis decreases with an increase in particle loading in the insert most likely due to the reduction in the water content. The inserts can be packaged in wet conditions and stored in refrigerator, but the inserts will exhibit a slight burst release due to release of drug from the particles into the p HEMA matrix during the shelf life. The burst effect could be avoided by packagi ng the inserts in dry state, with hydration prior to insertion. Chapters 5 demonstrated a novel approach of incorporating UV blocking feature into contact lenses by dispersing nanoparticles in contact lenses. The nanoparticles encapsulate a UV blocking co mpound which is trapped in the particles due to the very

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145 high crosslinking. The highly crosslinked nanoparticles prepared by using vitamin E as a diluents are about 4 nm in size. The nanoparticles were loaded with concentration of UV blocking compound DP ranging from 0 to 30% w/w. The DP loaded nanoparticles were dispersed in gels and lenses in varying loadings ranging from 0 to 20%. The particle loaded gels showed effective UV blocking and a number of compositions fall in Class 1 blocking. Autoclaving did not lead to particle release from contact lenses and also did not degrade UV blocking efficiency of the nanoparticles. In conclusion, t hese systems are promising for extended drug delivery. In the future, we will aim to find drugs that are more suitab le for delivery using nanoparticle systems. Furthermore these systems will further expand biomedical avenues to nano scale systems that could be helpful for a number of applications such as transdermal drug delivery patches, wound healing patches, gold coating nanopatricle for attacking cancer cells, etc. In chapter 5, DP loaded highly crosslinked particles can be loaded into contact lenses either before or after polymerization to prepare lenses with excellent UV blocking characteristic.

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154 BIOGRAPHICAL SKETCH Hyun Jung Jung was born in Pohang, South Korea in 1982. After graduating from High School in 2000, she began her undergraduate studies in Seoul, South Korea and graduated in 2004. And then, two yea r later she received her Master of Engineering degree at Korea University After receiving her Master of Science degree, she worked for Korea Institute of Science and Technology in Seoul and for Pharmaceutical Company in Ansan, S outh Korea as a researcher She then joined the department of Chemical Engineering at University of Florida in August 200 7 under the guidance of Dr. Anuj Chauhan on extended ophthalmic drug delivery using nanoparticles. She received a Doctor of Philosophy degree in August 2012