vitamin e loaded silicone hydrogel contact lenses...
TRANSCRIPT
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VITAMIN E LOADED SILICONE HYDROGEL CONTACT LENSES FOR EXTENDED OPHTHALMIC DRUG DELIVERY
By
CHENG-CHUN PENG
A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT
OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY
UNIVERSITY OF FLORIDA
2011
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© 2011 Cheng-Chun Peng
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To my dearest Mom and Dad, and all my family
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ACKNOWLEDGMENTS
It would have been next to impossible to finish this dissertation without the support,
patience and guidance of the following people through these years. It is to them that I
owe my deepest gratitude.
First and foremost, I would like to thank my doctoral advisor, Dr. Anuj Chauhan, for
not only his guidance on my research but also all his encouragement on my life in
Gainesville. He taught me the importance of keeping motivated and efficient on my
work, giving me enough freedom to explore, fail and improve. Most of all he has been a
good friend of mine, for which I will cherish. I also wish to extend many thanks to my
other doctoral committee members, Dr. Tanmay Lele, Dr. Peng Jiang, and Dr. Gregory
S. Schultz for their insightful viewpoints and willingness to participate in my doctoral
review process. In addition, I would like to thank Dr. Tanmay Lele and Dr. Spyros A.
Svoronos for the opportunity to serve as a teaching assistant under their guidance. I
would also express my gratitude to Dr. Caryn Plummer for her kindly help in providing
me the opportunity to conduct animal studies for my research. I am extremely grateful to
Dr. Yiider Tseng for his gracious support as both a scholar and a friend through my
years in Gainesville. I would also like to thank Dr. Nae-Lih Wu in National Taiwan
University for encouraging me to continue doctoral research in United States.
It is my privilege of working with many outstanding colleagues in Dr. Chauhan‟s
group. First of all, I would like to extend my special thanks to Dr. Jinah Kim, who not
only helped me establish the foundation of this dissertation, but also served as my role
model by showing me how to cherish and enjoy your life. I thank Dr. Yash Kapoor, Dr.
Brett Howell and Dr. Chhavi Gupta for their precious help in my research. The
demonstration of leadership, discipline, dedication and commitment toward work and life
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they showed me are one of the invaluable treasures I have obtained in my research
career in Gainesville. Hyun-Jung Jung has been a great friend and shared my best and
worst time in my research, for which I am extremely grateful. I also would like to thank
Michael Burke for his help in the lidocaine delivery experiments. Finally I thank the
current group members, Loki, Han and Ming, for sharing in my last days in Gainesville.
Many current and former staff members of the Department of Chemical
Engineering were also very accommodating during my time in Gainesville. I would like
to gratefully acknowledge the technical support for this research by Sean Poole, James
Hinnant and Dennis Vince. I would also like to thank Shirley Kelly, Deborah Sandoval,
Melissa Fox, and Carolyn Miller for their assistance during these years.
I am truly grateful to have so many wonderful friends in my life to stand by me
through all the good times and bad on the way to what I am now. Thanks my friends
and my family in Gainesville: Rob, Akhil, Poom, Jun, Can, Derek, Wei-Chiang, Tzung-
Hua, Jack and Hungta, and everyone whom I have had the privilege to share my life
with all these years. It is impossible to accomplish my work without all of your kindly
support.
Finally, words cannot describe my everlasting gratefulness to my big loving family
in Taiwan, especially to my parents, Yueh-Chin Tai and Kuo-Yuan Peng for their
unconditionally trust on every single decision I made in my whole life, and I hope I have
made you proud.
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TABLE OF CONTENTS
page
ACKNOWLEDGMENTS .................................................................................................. 4
LIST OF TABLES ............................................................................................................ 9
LIST OF FIGURES ........................................................................................................ 10
LIST OF ABBREVIATIONS ........................................................................................... 14
1 INTRODUCTION .................................................................................................... 18
2 CHARACTERIZATION OF VITAMIN E LOADED SILICONE HYDROGEL ............ 30
2.1 Materials and Methods ...................................................................................... 30 2.1.1 Materials .................................................................................................. 30 2.1.2 Vitamin E Loading into Pure Lenses........................................................ 31 2.1.3 Ion Permeability Measurements .............................................................. 31 2.1.4 Oxygen Permeability Measurements ....................................................... 32 2.1.5 Transmittance Measurement of Vitamin E Loaded Contact Lens ............ 33 2.1.6 Preparation of Silicone Hydrogel ............................................................. 33 2.1.7 Mechanical Properties Measurements .................................................... 34
2.2 Results and Discussion ..................................................................................... 34 2.2.1 Vitamin Loadings in the Lenses ............................................................... 34 2.2.2 Transparency of the Vitamin E-laden Contact Lenses ............................ 35 2.2.3 Water Content of Pure and Vitamin E Loaded Lenses ............................ 35 2.2.4 Size Change Due to Vitamin E Loading .................................................. 36 2.2.5 Ion Permeability of Vitamin E Loaded Lenses ......................................... 37 2.2.6 Oxygen Permeability of Vitamin E Loaded Lenses .................................. 38 2.2.7 Transmittance of Vitamin E Loaded Lenses ............................................ 41 2.2.8 Mechanical Properties of Vitamin E Loaded Silicone Hydrogel ............... 42
3 HYDROPHILIC DRUG DELIVERY BY VITAMIN E LOADED SILICONE HYDROGEL ............................................................................................................ 55
3.1 Materials and Methods ...................................................................................... 55 3.1.1 Materials .................................................................................................. 55 3.1.2 Drug Loading into Pure Lenses ............................................................... 56 3.1.3 Vitamin E Loading into Pure Lenses........................................................ 56 3.1.4 Drug Loading into Vitamin E Loaded Lenses .......................................... 57 3.1.5 Drug Release Experiments ...................................................................... 57
3.2 Results and Discussion ..................................................................................... 57 3.2.1 Dynamics of Drug Transport from Contact Lenses without Vitamin E ..... 57 3.2.2 Dynamics of Drug Transport from Vitamin E Loaded Lenses .................. 59
3.2.2.1 Timolol-Vitamin E loaded lenses .................................................... 59
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3.2.2.2 DXP-Vitamin E loaded lenses ........................................................ 62 3.2.2.3 Fluconazole-Vitamin E loaded lenses ............................................ 63
3.2.3 Model for Hydrophilic Drugs .................................................................... 64 3.2.4 Diffusivities of Drugs in Vitamin E Loaded Lenses .................................. 66
4 HYDROPHOBIC DRUG DELIVERY BY VITAMIN E LOADED SILICONE HYDROGEL ............................................................................................................ 82
4.1 Materials and Methods ...................................................................................... 83 4.1.1 Materials .................................................................................................. 83 4.1.2 Drug Loading into Pure Lenses ............................................................... 83 4.1.3 Vitamin E Loading into Pure Lenses........................................................ 83 4.1.4 Drug Loading into Vitamin E Loaded Lenses .......................................... 84 4.1.5 Drug Release Experiments ...................................................................... 84 4.1.6 Viscoelastic Measurement ....................................................................... 85
4.2 Results and Discussion ..................................................................................... 85 4.2.1 Dynamics of Drug Transport from Contact Lenses without Vitamin E ..... 85 4.2.2 Dynamics of Drug Transport from Vitamin E Loaded Lenses .................. 86 4.2.3 Diffusivities of Drugs in Vitamin E Loaded Lenses .................................. 90 4.2.4 Scaling Model for Effect of Vitamin E Loading on Extended DX
Delivery ......................................................................................................... 92
5 ANESTHETICS DELIVERY BY VITAMIN E LOADED SILICONE HYDROGEL ... 105
5.1 Materials and Methods .................................................................................... 107 5.1.1 Materials ................................................................................................ 107 5.1.2 Drug Loading into Pure Lenses ............................................................. 107 5.1.3 Vitamin E Loading into Pure Lenses...................................................... 108 5.1.4 Drug Release Experiments .................................................................... 109 5.1.5 Silicone Hydrogel Preparation ............................................................... 109 5.1.6 Partition Coefficient ............................................................................... 110 5.1.7 Determination of Critical Micelle Concentration (CMC) of Lidocaine ..... 111
5.2 Results and Discussion ................................................................................... 111 5.2.1 Dynamics of Drug Release from Contact Lenses .................................. 111
5.2.1.1 Drug uptake through drug-PBS solution....................................... 111 5.2.1.2 Drug uptake through drug-ethanol solution .................................. 114
5.2.2 Lidocaine Release Study (Surfactant Behavior) .................................... 115 5.2.3 Mechanisms of Extended Drug Release by Vitamin E Loaded Contact
Lens ............................................................................................................ 117
6 ION TRANSPORT OF SILICONE HYDROGEL .................................................... 138
6.1 Materials and Methods .................................................................................... 143 6.1.1 Materials ................................................................................................ 143 6.1.2 Preparation of Silicone Hydrogel ........................................................... 143 6.1.3 Water Fraction Measurements .............................................................. 144 6.1.4 Ion Permeability Measurements ............................................................ 145
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6.1.4.1 Salt release in perfect sink (kinetic desorption). ........................... 145 6.1.4.2 Ion transport in diffusion cell (direct permeation) ......................... 145
6.2 Results and Discussion ................................................................................... 146 6.2.1 Comparison of Transport Measurements from the Kinetic and
Permeation Approaches .............................................................................. 146 6.2.1.1 Kinetics of salt release in perfect sink .......................................... 146 6.2.1.2 Ion transport through permeation in a diffusion cell ..................... 150
6.2.2 Effect of Composition of Silicone Hydrogel on Ion Permeability ............ 154
7 CYCLOSPORINE DELIEVERY BY SILICONE HYDROGEL FOR CRONIC DRY EYE SYMDROME ................................................................................................ 185
7.1 Materials and Methods .................................................................................... 187 7.1.1 Materials ................................................................................................ 187 7.1.2 Drug Loading into Contact Lens ............................................................ 187 7.1.3 Drug Release Experiments from Lenses Loaded with CyA ................... 188 7.1.4 Vitamin E Loading into Contact Lens..................................................... 188
7.2 Results ............................................................................................................ 189 7.2.1 Drug Uptake by Pure Contact Lens ....................................................... 189 7.2.2 Drug Release by Pure Contact Lens ..................................................... 191 7.2.3 Drug Uptake by Vitamin E Loaded Contact Lens .................................. 191 7.2.4 Drug Release by Vitamin E Loaded Contact Lens ................................ 192
7.3 Discussion ...................................................................................................... 193 7.3.1 Release Mechanism and Model Fitting .................................................. 194 7.3.2 Therapeutic Release Rates ................................................................... 198
8 DRUG DELIVERY BY CONTACT LENS IN GLAUCOMATOUS DOGS ............... 211
8.1 Materials and Methods .................................................................................... 213 8.1.1 Materials ................................................................................................ 213 8.1.2 Drug and Vitamin E loading into Contact Lens ...................................... 213 8.1.3 Animal Model ......................................................................................... 214 8.1.4 Data Analysis ........................................................................................ 215
8.2 Results ............................................................................................................ 216 8.2.1 Contact Lens without Vitamin E ............................................................. 216 8.2.2 Contact Lens with Vitamin E .................................................................. 216 8.2.3 Eye Drop ............................................................................................... 217 8.2.4 Drug Administration Methods Comparison ............................................ 217
8.3 Discussion ...................................................................................................... 218
9 CONCLUSIONS ................................................................................................... 225
LIST OF REFERENCES ............................................................................................. 227
BIOGRAPHICAL SKETCH .......................................................................................... 240
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LIST OF TABLES
Table page 2-1 List of silicone hydrogel extended wear commercial contact lens (dipoter -
6.50) explored in this study (n = 6). .................................................................... 54
3-1 Model parameters obtained by fitting experimental data to the model ................ 81
4-1 Partition coefficient (K) of DX in lenses soaked in DX-PBS solution. ................ 104
6-1 Composition of silicone hydrogel. ..................................................................... 179
6-2 Parameters of different silicone hydrogels. ....................................................... 180
6-3 Parameters for GEL A3 with various sodium chloride concentrations for salt loading. ............................................................................................................. 181
6-4 Fitting results of ion transport by diffusion cell for silicone hydrogels. All samples are 0.13 mm thick and contain no preloaded salt. .............................. 182
6-5 Fitting results of ion transport by diffusion cell for Gel A3 with various sodium chloride concentrations in the donor compartment. .......................................... 183
6-6 Fitting results of ion transport by diffusion cell for silicone hydrogels which pre-soaked in sodium chloride solution with various concentrations. ............... 184
7-1 Results of CyA uptake by silicone contact lens. ............................................... 209
7-2 Results of CyA uptake by Vitamin E loaded ACUVUE® OASYSTM lenses. ....... 210
8-1 Summary of various drug delivery methods considered in this study. .............. 224
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LIST OF FIGURES
Figure page 1-1 Schematic illustration of ophthalmic drug delivery through eye drops. ............... 28
1-2 Schematic illustration of the microemulsion laden contact lens inserted in the eye. ..................................................................................................................... 29
2-1 Correlation of Vitamin E loading and concentration of soaking solution for different lenses. .................................................................................................. 44
2-2 Images of Commercial NIGHT&DAYTM contact lens and NIGHT&DAYTM lens with 30%Vitamin E loading. ................................................................................ 45
2-3 Plot of water content (Q) and EW of Vitamin E loaded lenses versus Vitamin E loading. ........................................................................................................... 46
2-4 Percent increase in diameter of dry lenses and wet lenses before and after loading Vitamin E. ............................................................................................... 47
2-5 Effect of Vitamin E loading on ion permeability of lenses. .................................. 48
2-6 Effect of Vitamin E loading on oxygen permeability (Dk). ................................... 49
2-7 Transmittance spectrum for commercial contact lenses. .................................... 50
2-8 Transmittance spectrum for NIGHT&DAYTM and ACUVUE® OASYSTM with different Vitamin E loading. ................................................................................. 51
2-9 Transmittance spectrum of NIGHT&DAYTM with Vitamin E loading .................... 52
2-10 Dependence of storage module of Vitamin E loaded silicone hydrogel on frequency.. .......................................................................................................... 53
3-1 Effect of timolol loading method on profile of timolol release by commercial contact lenses.. ................................................................................................... 70
3-2 Profiles of timolol release by Vitamin E loaded contact lenses. .......................... 71
3-3 Profiles of repeated timolol releases by Vitamin E loaded contact lenses. ......... 73
3-4 Profiles of DXP release by Vitamin E loaded contact lenses ............................. 74
3-5 Profiles of fluconazole release by Vitamin E loaded contact lenses ................... 76
3-6 Drug release duration increase by Vitamin E loaded contact lenses. ................. 78
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3-7 Plot of % timolol release by Vitamin E loaded NIGHT&DAYTM versus square root of time.......................................................................................................... 80
4-1 Effect of DX loading method on profile of DX release by contact lenses... ......... 96
4-2 Profiles of experimental and model fitted DX uptake and release by Vitamin E loaded contact lenses.. ....................................................................................... 97
4-3 Plot of % drug release by Vitamin E loaded lenses against square root of time. .................................................................................................................. 99
4-4 Fitted DX diffusivity and partition coefficient for contact lenses with different Vitamin E volume fraction. ................................................................................ 101
4-5 Effect of Vitamin E volume fraction on increase in drug uptake times. ............. 102
4-6 Dependence of the loss modulus G" on frequency for pure Vitamin E ............. 103
5-1 Molecular structures of model drugs. ................................................................ 123
5-2 Lidocaine release in PBS by O2OPTIXTM with various Vitamin E loading. ........ 124
5-3 Vitamin E release from O2OPTIXTM during lidocaine release in PBS. .............. 125
5-4 Bupivacaine release in PBS by O2OPTIXTM with various Vitamin E loading. .... 126
5-5 Short time and long time tetracaine release in PBS by O2OPTIXTM with various Vitamin E loadings. .............................................................................. 127
5-6 Lidocaine, bupivacaine and tetracaine release in PBS by O2OPTIXTM with various Vitamin E loading. ................................................................................ 128
5-7 Lidocaine release by O2OPTIXTM with 0.36g Vitamin E/g pure lens. ................ 130
5-8 The relationship between surface tension and lidocaine concentration in PBS. ......................................................................................................................... 131
5-9 The calculated partition coefficient (K) of Vitamin E loaded silicone hydrogel at various lidocaine hydrochloride concentrations. ........................................... 132
5-10 The relationship between the lidocaine partition coefficient in Vitamin E (KVE) and the bulk drug concentration. ...................................................................... 133
5-11 Model fitting for anesthetic drug release increase ratio on Vitamin E loading fraction in the silicone hydrogel. ....................................................................... 134
5-12 Lidocaine release from pure 0.2 mm-thick silicone hydrogel or gel with Vitamin E loading. ............................................................................................. 135
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5-13 Lidocaine release by silicone hydrogel with or without Vitamin E loading with various thickness.. ............................................................................................ 136
5-14 Lidocaine release by silicone hydrogel (with or without Vitamin E loading with various thicknesses. ......................................................................................... 137
6-1 NaCl release profile and model prediction (solid line) of different silicone hydrogel in perfect sink. .................................................................................... 161
6-2 NaCl release profile of 0.13 mm-thick Gel A3 with different pre-soaking NaCl concentration in perfect sink. ............................................................................ 162
6-3 NaCl release profile of Gel A3 with different thickness in perfect sink. ............. 163
6-4 Ion permeability test for silicone hydrogels by diffusion cell.. ........................... 164
6-5 Ion permeability test of Gel A3 by diffusion cell with various NaCl concentration in the donor compartment.. ........................................................ 165
6-6 Ion permeability test by diffusion cell of silicone hydrogel that were pre-soaked in sodium chloride solution with different concentration. . ................... 166
6-7 Ion permeability test of Gel A3 with different thickness by diffusion cell. .......... 168
6-8 NaCl partition coefficient (K), diffusivity(D), KD and Water content (Q) for silicone hydrogel with different TRIS/Macromer compositions.......................... 170
6-9 NaCl partition coefficient (K), diffusivity(D), KD and Water content (Q) for silicone hydrogel with different EGDMA compositions. ..................................... 172
6-10 NaCl partition coefficient (K), diffusivity(D), KD and Water content (Q) for silicone hydrogel with different DMA compositions. .......................................... 174
6-11 The relationship between salt partition coefficient (K) and water fraction (Q) of silicone hydrogel. .......................................................................................... 176
6-12 The relationship between salt diffusivity (D) and the reciprocal of water fraction (1/Q) of silicone hydrogels. .................................................................. 177
6-13 The relationship between salt permeability (KD) and reciprocal water fraction (1/Q) of silicone hydrogels. ............................................................................... 178
7-1 Cumulative drug uptake by ACUVUE® OASYSTM lens ..................................... 201
7-2 Cumulative CyA release by 1-DAY ACUVUE® ................................................. 202
7-3 Cumulative CyA release from silicone contact lens. ......................................... 203
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7-4 Cumulative drug release from Vitamin E loaded ACUVUE® OASYSTM lenses. ......................................................................................................................... 204
7-5 Daily average CyA release rate from Vitamin E loaded ACUVUE® OASYSTM lenses. .............................................................................................................. 205
7-6 Plot of % CyA release by silicone contact lenses versus square root of time. .. 206
7-7 Plot of % CyA release by Vitamin E loaded ACUVUE® OASYSTM versus square root of time. ........................................................................................... 207
7-8 Comparison of CyA and dexamethasone delivery by Vitamin E loaded ACUVUE® OASYSTM. ....................................................................................... 208
8-1 Effect of insertion of drug loaded contact lenses on the intraocular pressure. .. 220
8-2 Effect of drug administration through eye drops on the intraocular pressure.. .. 222
8-3 Comparison of the effect of various drug delivery methods on the differences in the IOP between the treated and the untreated eyes. .................................. 223
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LIST OF ABBREVIATIONS
BCL Bandage contact lens
CMC Critical micelle concentration
CyA Cyclosporine A
DI Deionized
DMA N, N-Dimethylacrylamide
DX Dexamethasone
DXP Dexamethasone 21-disodium phosphate
EGDMA Ethylene glycol dimethacrylate
GVHD Graft-versus-host disease
HEMA 2-hydroxyethyl methacrylate
IOP Intraocular pressure
LASIK Laser in situ keratomileusis
MAA Methacrylic acid
NVP 1-vinyl-2 pyrrolidone
OU Both eyes
PBS Phosphate buffered saline
PLTF Pre lens tear film
POLTF Post lens tear film
PRK Photorefractive keratectomy
TRIS 3-Methacryloxypropyl tris(trimethylsiloxy)silane
UV Ultraviolet
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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
VITAMIN E LOADED SILICONE HYDROGEL CONTACT LENSES FOR EXTENDED
OPHTHALMIC DRUG DELIVERY
By
Cheng-Chun Peng
August 2011
Chair: Anuj Chauhan Major: Chemical Engineering
Ophthalmic drug delivery via eye drops is inefficient as only 1-5% of the applied
drug enters the cornea and the rest is absorbed into the bloodstream. This absorbed
drug then enters other organs where it can cause side effects. Furthermore, drug
administration through eye drops results in a rapid variation in drug delivery rates to the
cornea that limits the efficacy of therapeutic systems and limit compliance. The purpose
of our study is to develop a novel soft contact lens system for long term and controlled
drug delivery to eliminate these problems. The aims are to characterize the lens system,
to establish the drug transport mechanism models, and to evaluate the potential of
practical applications for ocular disease treatments.
Our approach focuses on creating transport barriers to increase release duration
from commercial contact lenses. In the absence of diffusion barriers, drug molecules
diffuse out of the lens in about a few hours. In contrast, if diffusion barriers are present,
molecules have to diffuse around these, resulting in an increase in the path length and
the release duration. In this study the hydrophobic Vitamin E is loaded into silicone
hydrogel contact lens by being dissolved in ethanol that swells the lens and
subsequently forming aggregates inside the lens after solvent evaporation.
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Several properties including geometry, ion permeability, oxygen permeability and
UV transmittance are characterized to determine the pros and cons of loading Vitamin E
into the lenses. The results indicate the property changes caused by Vitamin E loading
do not disqualify these silicone hydrogels as extended-wear contact lens. In addition,
Vitamin E loading has a beneficial effect of blocking UV radiation which will reduce the
corneal damage due to UV light. Among all the lens property changes, ion permeability
demonstrated the strongest dependency on the Vitamin E loading in the lens, and thus
a further securitization of the ion transport of silicone hydrogels with various
compositions is discussed.
In vitro drugs release results show that the increase in release duration is
significantly dependent on the interaction between Vitamin E and the drug of interest.
For hydrophilic drugs (timolol, fluconazole, dexamethasone phosphate), the drug
release duration increases quadratically in Vitamin E loading. For example, for
NIGHT&DAYTM lens loadings of 10 and 40% Vitamin E increase release durations to 0.5
and 16 days, respectively. For hydrophobic drugs dexamethasone and cyclosporine A,
the effect of the Vitamin E inclusion is smaller but still significant for release. On the
other hands, for some amphiphilic anesthetic drugs, including lidocaine, bupivacaine
and tetracaine, the interfacial interaction between drug and Vitamin E aggregation plays
the determinative role in drug transport through the Vitamin E/silicone hydrogel matrix.
Ocular drugs delivery by contact lens can be viewed as a one-dimensional transport by
a flat thin film, and subsequent mathematical models based on the proposed
mechanisms are established.
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In addition, a case study was conducted to evaluate the use of Vitamin E loaded
silicone hydrogel contact lenses for chronic dry eye treatment. In vitro release
experiments in perfect sink condition demonstrate that, through incorporation of Vitamin
E, the loaded cyclosporine A in the lens can be released from 2 weeks to the total wear
time of the lens, which is about a month. Vitamin E incorporation also renders the
release profile closer to „zero order‟ such that the release rates are within the estimated
therapeutic window. The long release duration along with the higher bioavailability
compared to commercial eye drops (RESTASIS® ) suggested that these lenses could
potentially be useful for treatment for chronic dry eye and also for reducing the
symptoms of contact lens mediated dry eyes.
Finally, an in vivo study was conducted to investigate the feasibility and
effectiveness of timolol delivery to glaucomatous dogs via drug-impregnated contact
lenses. By utilizing contact lens to deliver timolol to the eye, the intraocular pressure in
the treated eye decreased effectively to similar degree compared to that by eye drop
treatment, while it significantly reduced the risk of systemic drug exposure.
In conclusion, silicone hydrogel contact lenses with Vitamin E are promising
candidates for extended ophthalmic drug delivery. These Vitamin loadings can
significantly attenuate the drug delivery rate, reduce wastage and provide safer
treatment route.
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CHAPTER 1 INTRODUCTION
While easy accessibility is clearly an advantage for ophthalmic drug delivery,
several other features of the ocular physiology, anatomy and biochemistry render the
ocular environment impervious to foreign substances, thus posing unique challenges for
delivery drugs to eyes [1-4]. The most common ophthalmic dosage forms are solutions,
ointments and suspensions, which together account for nearly 90% of currently
available formulations in the United States [5]. Amongst all the dosage forms for
ophthalmic drug delivery, eye drop solutions are the preferred choice since they are
relatively simple to prepare, filter, and sterilized.
Ophthalmic drugs are commonly delivered to the front of the eye through
instillation of drug-laden eye drops into the tear film, as shown in Figure 1-1. The human
tear film is about 10 L in volume and 3-10 m in thickness [6, 7]. The tear film is
typically considered to comprise of three layers: a thin mucin layer lying on the corneal
surface, which renders the corneal surface hydrophilic; the aqueous layer which
comprises the bulk of the tear film; and a very thin oily layer that minimizes evaporation
of tears into air [8]. The mucin layer is produced mainly by secretions from conjunctival
goblet cells [9]; the aqueous layer is secreted by the lacrimal glands and supplemented
by conjunctival secretion, and the lipid layer is produced by the tarsal meibomian glands
[7, 8]. The tear film is lined by the cornea and the conjunctiva, which is further classified
as the bulbar or the palpebral conjunctiva. The volume and composition of the tears are
maintained through a dynamic balance between the various secretion and elimination
pathways for the fluid and the ions. The tear fluid is produced by the lacrimal glands,
and from secretion from the conjunctiva, and eliminated through evaporation and
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drainage of tears through the canaliculi into the lacrimal sac [7]. The transport of ions in
the tear film includes inflow through the lacrimal secretions and outflow through the tear
drainage. Additionally, active ion transporters, pumps, and osmotic and electro-osmotic
transport across the cornea and the conjunctiva contribute to the ion transport in the
tear film [10].
The transport pathways for ions and tear fluid also contribute to the transport of
drug delivered to the tear film through instillation of an eye drop. A typical eye drop is
about 30 L in volume, which is too large to be accommodated in the tear film, and thus
a fraction of the instilled drop typically rolls off after instillation [7]. The remaining
amount rapidly mixes with the tear volume due to blinking. Subsequently, a fraction of
the drug penetrates the corneal and the conjunctival epithelium, and the remaining
amount drains into the lacrimal sac and the nasal cavity, where it can get absorbed into
the blood stream through the mucus membrane. A large fraction of the drug that is
absorbed in the conjunctiva also enters systemic circulation because of the high blood
perfusion in this tissue. The amount of drug absorbed by the conjunctiva is significantly
more than that absorbed by the cornea because of the larger area and permeability of
the conjunctiva in comparison to those for the cornea [11]. The drug absorbed by the
cornea can diffuse across the three layers of the cornea, i.e., the epithelium, stroma,
and the endothelium to reach the anterior chamber. A fraction of the drug is then
cleared through the drainage of the aqueous humor, and a fraction can bind to ocular
tissues such as the lens, ciliary bodies, etc. Due to the limited residence time of the
drug in the tear film, the drug concentration at the target tissue could exhibit a short
duration burst above the toxicity threshold, followed by a long duration below the
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therapeutic threshold. A few factors that aid in increasing the duration of action of the
drug at the target tissues include a strong drug binding to the target tissue and/or
adsorption-desorption to other ocular tissue.
The clearance mechanisms in eyes render ophthalmic drug delivery via eye drops
relatively inefficient. Most ophthalmic drugs have a short residence time of
approximately 2 minutes [12] due to the rapid tear turnover. With such short residence
time, only about 2-5 % of the applied drug penetrates the cornea to reach the
intraocular tissue [12, 13]. The remaining drug is either absorbed in the conjunctiva or it
drains with tears into the lacrimal sac, leading to drug wastage and sometimes adverse
side effects [5]. Additionally, the poor drug bioavailability and short residence time in
tears results in the need for several daily administrations, reducing the patient
compliance.
To overcome the drawbacks of eye drops, several ophthalmic drug delivery
systems aim for sustained drug release have been proposed such as suspension of
nanoparticles, nanocapsules, liposomes and noisome, ocular inserts like collagen
shields and Ocusert® , and therapeutic contact lenses. Among these, contact lenses
have been widely studied due to the high degree of comfort and biocompatibility. As
illustrated in Figure 1-2, if drug loaded contact lenses are placed on the eye, the drug is
expected to diffuse through the lens matrix, and enter the post-lens tear film (POLTF),
which is the thin tear film trapped between the cornea and the lens. In the presence of a
lens, drug molecules will have a residence time of about 30 minutes in the post-lens
tear film, compared to about 2 minutes in the case of topical application as drops [12, 14,
15]. The longer residence time will result in a higher drug flux through the cornea and
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reduce the drug inflow into the nasolacrimal sac, thus reducing the drug absorption into
the blood stream. Drug loaded contact lenses also have the potential to provide
continuous drug release, because of the slow diffusion of the drug molecules through
the lens matrix compared to that in aqueous solution. However, the maximum drug
loading is limited by the solubility of the drugs in the gel matrix. Also the only additional
resistance to drug transport is diffusion through the gel matrix, which is significantly
determined by the materials and microstructures of the contact lens, and for most ocular
drugs the release time is less than a few hours [16].
The development of hydrogel based contact lens started in 1936 when the first
hydrogel lens was introduced by Wichterle by using polymethylmethacrylate (PMMA), a
resin that has greater clarity than glass [17, 18]. Since then, researches of new hydrogel
polymers were continuously motivated by the obvious pursuit for a competitive position
in the new growing vision-correction soft contact lens market [19]. The evolution of new
soft contact lens materials have also driven by an increased understanding of the
physiological needs of the cornea, since the ocular environment places high demands
on the performance of contact lenses as biomaterials. The contact lens on the eye must
maintain a stable, continuous tear film for clear vision; it needs to be resistant to
deposition of tear film components and has proper mechanic strength to be nonirritating
and comfortable for the wearer during blinking. In addition, permeability to oxygen and
ions are two key performance characteristics for contact lenses. Not like other tissues in
human, cornea obtains its oxygen directly from the air to maintain its clarity, structure
and function, and hence a contact lens must be sufficiently permeable to oxygen to
maintain normal corneal metabolism. Effective ion permeation in the contact lens allows
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the movement of the lens on the eye. Therefore, to satisfy all the criteria, the lens must
have excellent surface characteristics being neither hydrophobic nor lipophilic, and have
the appropriate bulk polymer composition and morphology at the same time.
Design for new extended wearable contact lens has been widely studied since the
negative physiological effect of cornea oxygen deprivation began to be realized [20, 21].
Traditional hydrophilic hydrogel contact lens has been proved fails to provide the
minimum oxygen transmissibility of 125 barrer/mm to completely avoid low oxygen-
related effects and to be safely worn overnight [22, 23]. To overcome the performance
hurdle, silicone hydrogel contact lens was developed due to its high oxygen
permeability since the first commercial one was produced in 1980 by Dow Corning.
However, polysiloxane has negligible water content and manifest typically rubbery
behavior. To offer the softness, wettability and on-eye comfort of a conventional
hydrogel, while at the same time provide the higher oxygen transmission required by the
cornea, new silicone hydrogels contact lens for extended wear were developed based
on the combination of silioxane, hydrophilic co-monomers, and bi-functional macromer
to find the optimized performance [24, 25].
In our study we aim to develop an “extended-wear extended-release” silicone
hydrogel contact lens which can safely and continuously provide ophthalmic drug
delivery. In previous studies on ocular drug delivery by contact lenses, drug can be
loaded in the contact lenses by either soaking the gels in drug solution [26-31] or by
dissolving drug in the monomer solution before polymerization [32-34]. The major
problem of loading drugs by uptake from drug solutions is that in most cases the loading
capacity of the soaked contact lenses is inadequate. An additional problem that can
23
occur when absorbing drugs in hydrophilic contact lens is that the preservatives
included in the drug are often preferentially absorbed in the lens to a level that is toxic
[35]. For directly loading drug in the monomer solution, although it can allow higher
loadings of the drugs in the lens, it can result in an activity loss during polymerization,
and a majority of the drug diffuses from the lenses into the packaging medium.
Moreover, even though providing higher bioavailability, commercial lenses release
ophthalmic drugs rapidly in a time of only a few hours and therefore limited their
application for extended drug delivery. Recently Karlgard et al. measured the in vitro
uptake and release of a number of ophthalmic drugs by commercially available pHEMA
based and silicone contact lenses [16]. Drugs, including cromolyn sodium, ketotifen
fumarate, ketorolac tromethamine and dexamethasone sodium phosphate, were loaded
into these lenses by soaking these in drug solutions for a limited period of time. The
release studies showed that a majority of the drug taken up by the gels was released in
a short period of time for less than a few hours, which is not suitable for extended drug
delivery.
To increase drug release durations, several different drug-impregnated contact
lens systems have been developed, including nanoparticle-laden lenses [36-40],
biomimetic and „imprinted‟ contact lenses [41-48], and contact lens with layer structures
[49]. While these approaches are effective at extending the drug release duration from
contact lenses, all studies cited above focused on lenses based on hydrophilic hydrogel
materials, which are not suitable for extended wear due to insufficient oxygen
permeability. In addition, while the approaches listed above can increase the loading
24
capacity and the release duration from contact lenses, the lenses are still not suitable
for extended release lasting a week or longer.
The aim of my study is to develop a contact lens system containing diffusion
barriers in the hydrogel matrix that can provide sufficient drug transport resistance, and
this system should be still satisfy all the criteria as extended wear contact lens material.
This extended-wear extended release contact lens could serve as the drug delivery
vehicle for long term ocular disease treatment, such as glaucoma and chronic dry eye,
and/or alleviate common untoward syndromes for extended contact lens wearers.
Glaucoma, which is the second largest cause of blindness in the world after cataract,
affects about 60.5 million people, leaving 8.2 million with bilateral blindness [50]. Most
current glaucoma therapies are based on drug delivery via eye drops that are inefficient
and have lower than 50% compliance [51], with further lower compliance for multiple
drops and/or multiple drugs therapies [52]. For chronic dry eye, In 2007, there were
about 35 million contact lens wearers in North America and about half of those reported
some symptoms of dryness and discomfort, more commonly experienced at the end of
the day [53]. Our designed contact lenses can also be used to release topical
anesthetics for pain control after photorefractive keratectomy (PRK) surgery. The case
studies of these potential applications are further discussed in Chapters 5, 7 and 8.
The release of a molecule from a contact lens can be considered as controlled by
diffusion within the lens material. For a one dimensional diffusion controlled process, the
duration of release can be approximately calculated by l2/D, where l is the path length
that a compound needs to traverse and D is the molecular diffusivity. For a typical
contact lens, l is the thickness of the lens, which varies in the radial direction but is on
25
average approximately 80 to 100 m for a typical lens. The period of time over which a
drug is released from a contact lens can be increased by either increasing l or by
decreasing D. In most diffusion controlled systems, augmentation of diffusivity has been
performed by changing the bulk material to one of a different diffusivity. However,
because of the strict requirements of a contact lens where many material properties
cannot be compromised, there are practical limits to the selection of the bulk material.
Furthermore, an effective strategy to modifying the diffusion process must be applicable
to a wide range of bioactive agents with a similar bulk material.
Our proposed concept is directed towards controlling the diffusion of a bioactive
agent in a contact lens matrix by the creation of diffusion barriers within the lens, such
that an included bioactive agent is forced to take a long tortuous path to diffuse from the
lens, resulting in extended release. The concept of using transport barriers has been
explored extensively for designing membranes that retard gas transport [54-57]. To our
knowledge, this concept has not been applied to retard drug transport from a biomedical
device. The diffusion barrier can be any solid or liquid material that is able to be
dispersed within the lens material in a manner that keeps the lens transparent. The
most important requirement for an effective barrier material is that the barrier should be
relatively impermeable to the molecules whose diffusion needs to be attenuated. A
number of ophthalmic drugs are charged at physiological pH and so hydrophobic
molecules will likely form effective barriers. It is also important to ensure that the barrier
material is biocompatible so that diffusion of the compound forming the barrier into the
tear film does not cause toxicity.
26
The diffusion barrier material of interest in our study is Vitamin E, which is a highly
hydrophobic liquid. It is a powerful antioxidant and has been shown in some animal
studies that the topical application of Vitamin E inhibits a number of eye diseases
including keratocyte apoptosis after surgery, ethanol-induced apoptosis in the corneal
epithelium, etc. [58, 59]. Also, there have been a number of in vivo studies suggesting
Vitamin E retards cataract development [60-64]. Because of the potential benefits of
delivering Vitamin E to the eye, there have been several attempts to develop ophthalmic
solutions containing Vitamin E [65, 66]. Once Vitamin E is trapped inside the gel matrix,
it should be stable in the hydrated contact lens due to its high hydrophobicity. In
Chapter 2 we explored the approaches to introducing Vitamin E into the silicone
hydrogels, and characterized the properties of these Vitamin E loaded silicone
hydrogels as extended contact lens, including ion permeability, oxygen permeability,
geometry changes, elastic modules, water transport, and light transmittance, etc.
Among these properties, we especially focused the ion transport through the silicone
hydrogels since first it can be correlated to some charged drugs transport in the gel
matrix, which is very common at physiological pH, and the results are discussed in
Chapter 6.
The transport mechanisms of extended ocular drugs by Vitamin E loaded silicone
hydrogel are further discussed throughout Chapters 3 to 5. Chapter 3 discussed the
release of three different hydrophilic ophthalmic drugs by Vitamin E loaded silicone
hydrogel contact lens, including timolol (beta blocker used for treating glaucoma),
dexamethasone 21-disodium phosphate (anti-inflammatory corticosteroid), and
fluconazole (anti-fungal). These drugs were chosen because they are hydrophilic at the
27
physiological pH, which should have negligible affinity to the desired Vitamin E barriers.
In addition, we also explored the drug delivery these Vitamin E loaded silicone
hydrogels for hydrophobic drug such as dexamethasone (anti-inflammatory) and
cyclosporine A (chronic dry eye treatment), and amphiphilic anesthetics drugs, including
lidocaine, bupivacaine and tetracaine. The drug transport results and mechanisms by
silicone hydrogel for these hydrophobic and amphiphilic drugs are discussed in
Chapters 4, 5 and 7.
Finally, in Chapter 8 we demonstrate the in vivo evaluation of the safety and
efficacy of glaucoma therapy through release of timolol from silicone hydrogel contact
lenses. Timolol is a beta-blocker that is widely used to treat glaucoma by reducing IOP
through decreasing the production of aqueous humor [67]. We focus on this drug
because of the large number of glaucoma patients in the world [50], and also because
of the potential of serious side effects from systemic exposure to timolol [68]. We
choose the colony of beagle dogs who are affected by or carriers of a hereditary form of
primary open angle glaucoma, the most common form of glaucoma in human beings
[69]. Beagle dogs have been used in several prior studies on glaucoma therapy [70-76].
Another benefit of using the Beagle dogs is that the cornea shape and size of these
dogs are similar to that of human beings, and therefore the commercially available
contact lens for human can be used in this study without further modification.
28
Figure 1-1. Schematic illustration of ophthalmic drug delivery through eye drops.
Tear film
Cornea
Conjunctiva
29
Figure 1-2. Schematic illustration of the microemulsion laden contact lens inserted in the eye.
Post-lens tear film
Cornea Drug loaded
contact lens
Pre-lens tear film
30
CHAPTER 2 CHARACTERIZATION OF VITAMIN E LOADED SILICONE HYDROGEL
In this chapter we explored the possibility to introducing Vitamin E into the silicone
hydrogel polymer matrix as contact lenses materials. Hydrophobic Vitamin E is loaded
into silicone hydrogel contact lens by being dissolved in ethanol that swells the lens and
subsequently forming aggregates inside the lens after solvent evaporation. This “in situ”
approach to create diffusion barriers is particularly suited for biomedical applications in
which the polymer processing steps could potentially damage some materials that are
used as transport barriers.
Most experiments in this chapter were done by loading Vitamin E into commercial
silicone hydrogel contact lens to evaluate the properties chance resulted from the
additional Vitamin E inside the gel matrix system. Important properties for the bulk
material, including geometry, equilibrium water content, ion permeability, oxygen
permeability, and light transmittance. In addition, to understand the effect of Vitamin E
on the elastic modulus of silicone hydrogel, a lab synthesized silicone hydrogel was
prepared as a substation of commercial silicone hydrogel contact lenses.
2.1 Materials and Methods
2.1.1 Materials
Five commercial silicone contact lenses (diopter -6.50) that are used in this study
are described in Table 2-1. 2-hydroxyethyl methacrylate (97%), sodium hydroxide
pellets (97+%), ethanol (99.5%), and Dulbecco‟s phosphate buffered saline
(PBS) were purchased from Sigma-Aldrich Chemicals (St. Louis, MO) and ethylene
glycol dimethacrylate (EGDMA) from Sigma-Aldrich Chemicals (Milwaukee, WI).
31
Sodium chloride (99.9+ %) were purchased from Fisher Chemical (Fairlawn, NJ).
Vitamin E (D-alpha tocopherol, Covitol® F1370) was gifted by Cognis Corporation.
For preparation of silicone hydrogel, ethylene glycol dimethacrylate (EGDMA,
98%), N, N-Dimethylacrylamide (DMA, 99%) and 1-vinyl-2 pyrrolidone (NVP, 99+ %)
were purchased from Sigma-Aldrich Chemicals (Milwaukee, WI). The macromer
acryloxy(polyethyleneoxy)-propylether terminated poly(dimethylsiloxane) (DBE-U12,
95+%) were purchased from Gelest Inc. (Morrisville, PA). 3-Methacryloxypropyl
tris(trimethylsiloxy)silane (TRIS) was supplied by Silar laboratories (Scotia, NY), and 2,
4, 6-trimethylbenzoyl-diphenyl-phosphineoxide (Darocur® TPO) were kindly provided by
Ciba Specialty Chemicals (Tarrytown, NY).All chemicals were used as received without
further purification if not specifically mentioned.
2.1.2 Vitamin E Loading into Pure Lenses
Vitamin E was loaded into lenses by soaking the lens in 3 mL of a Vitamin E-
ethanol solution for 24 hours. Vitamin E-ethanol solutions of various concentrations
were prepared by simply mixing Vitamin E and ethanol with vortexing for a few seconds
followed by moderate magnetic stirring for several minutes. After the loading step, the
lens was taken out and excess Vitamin E-ethanol solution on the lens surface was
blotted, and the lens was then dried in air overnight.
2.1.3 Ion Permeability Measurements
Ion permeability of lenses was measured by using a homemade horizontal
diffusion cell that consists of a donor and a receiving compartment, which were both
fabricated from Plexiglas. The ion permeability of the lens was determined by
measuring the rate of transport of ions across the lens. To mount the lens in the
diffusion cell, the circular edge of the dried lens was glued to the outer edge of a 1 cm
32
hole cut into a plastic spacer. The spacer along with the lens was then soaked in
deionized (DI) water for longer than three hours to fully hydrate the lens. The excess
water on the spacer was wiped off and the spacer was subsequently placed in between
the two compartments of the diffusion cell, and clamped. Latex O-rings were also
inserted in between the spacer and each of the compartments to ensure sealing. The
latex O-rings were boiled in DI water for 40 minutes for three times before placing in the
diffusion cell to leach out impurities. After assembling the diffusion chamber, the
receiving chamber was filled with 30 mL of DI water and the donor chamber was filled
with 18 mL of 0.1 M NaCl solution. The ion conductivity of the fluid in the receiving
chamber was measured as a function of time with a conductivity meter with temperature
sensor (Con 110 series, OAKTON® ), and linear regression was applied to the data after
reaching pseudo-steady state (after 70 minutes) to obtain the best fit slopes. The rate of
conductivity change can be converted to the rate of ion transport, which can then be
related to the ion permeability of the lenses by using Fick‟s law.
2.1.4 Oxygen Permeability Measurements
To measure the oxygen permeability, lenses were mounted in a horizontal
diffusion cell by following the same procedure as described in Section 2.2.3. To create
oxygen gradients in the cell, the donor compartment was filled with 18 mL of DI water
that was equilibrated with air, and the receiving chamber was filled with 32 mL of DI
water that was degassed by bubbling nitrogen for 10 minutes. Both compartments were
kept well-stirred with minimal boundary layer thicknesses adjacent to the lens by stirring
at 900 rpm. The dissolved oxygen concentration in the receiving reservoir was
measured every 12 seconds by an oxygen sensor (DO-BTA, Vernier® ) for a total
33
duration of 2 hours. The measured data was fitted to a mathematical model described
later to determine the oxygen permeability of the lens.
2.1.5 Transmittance Measurement of Vitamin E Loaded Contact Lens
The transmittance of Vitamin E laden lenses was measured using UV-VIS
spectrophotometer (Thermospectronic Genesys 10 UV). The lenses were hydrated by
soaking in PBS overnight, then cut into stripes and mounted on the outer surface of a
quartz cuvette. The cuvette was placed in the spectrophotometer and the transmittance
values were measured at wavelengths ranging from 200 nm to 500 nm.
2.1.6 Preparation of Silicone Hydrogel
To prepare the silicone hydrogel, hydrophilic monomers with high ion permeability
are copolymerized along with the hydrophobic silicone monomer with high oxygen
permeability, and a macromer is needed in the monomer mixture to ensure
solubilization of all monomers. In this study, TRIS was used as the hydrophobic
monomer, DMA was the hydrophilic monomers, and DBE-U12 was selected as the
macromer. Highly hydrophilic NVP monomer was also added to increase water content
of the hydrogel and EGDMA was introduced in the monomer mixture for controlled
crosslinking. To prepare the polymerizing mixture, 2.4 mL of a mixture that comprises
0.8 mL TRIS and 0.8 mL macromer and 0.8 mL of the hydrophilic DMA/MAA mixture
was combined with 0.12 mL of NVP and 0.1 mL of EGDMA. After well mixed with
vortexing for few second, the mixture was purged with bubbling nitrogen for 15 minutes
to reduce the dissolved oxygen. To each monomer mixture, 12 mg of photoinitiator
Darocur® TPO was added with stirring for 5 minutes and the final mixture was
immediately injected into a mold which is composed of two 5 mm thick glass plates. The
plates were separated by a plastic spacer with various thicknesses. The mold was then
34
placed on ultraviolet transilluminator UVB-10 (UltraLum Inc.) and the gel mixture was
cured by irradiating with UVB light (305 nm) for 50 minutes. The synthesized hydrogel
was either cut into circular pieces (about 1.65 cm diameter) with a cork borer for
subsequent experiments. Prior to conducting further tests, the prepared hydrogel was
soaked in ethanol for 3 hours then dried at ambient temperature overnight to remove
the unreacted monomer within.
2.1.7 Mechanical Properties Measurements
The mechanical properties of gels are analyzed in tensile mode by using a
dynamic mechanical analyzer (DMA Q800, TA instruments). A 0.4 mm thick rectangular
hydrated gel was mounted on the tension clamp while submerged in water at room
temperature. A periodic tensile force was applied in the longitudinal direction with varied
frequency and the response (storage modulus and loss modulus) of gel was determined.
A static preload force of 0.01 N was applied and a 115% of force track was used to
keep the sample taut on the tension clamps. Strain sweep tests were conducted at
room temperature at 1 Hz to determine linear viscoelastic range, and proper strain
which was therefore confirmed within linear range by the strain sweep test will be
chosen for subsequent strain controlled frequency sweep experiments.
2.2 Results and Discussion
2.2.1 Vitamin Loadings in the Lenses
Vitamin E loadings into each lens for different initial concentration of Vitamin E
loading solutions are shown in Figure 2-1. Vitamin E loading has a linear dependency
on the concentration of Vitamin E loading solutions. In addition, ACUVUE® OASYS™
and NIGHT&DAY™ have the highest and the lowest affinity for Vitamin E, respectively.
35
2.2.2 Transparency of the Vitamin E-laden Contact Lenses
An image of a Vitamin E loaded contact lens is shown in Figure 2-2. As evident
from the image, the Vitamin E loaded lenses are transparent irrespective of the Vitamin
E loading, but attain a slightly yellowish color at high Vitamin E loadings.
2.2.3 Water Content of Pure and Vitamin E Loaded Lenses
Water contents (Q) of lenses are listed on each lens package and were also
measured as
100
eq
veleq
W
WWWQ (2-1)
where Weq, Wl, and Wve are mass of hydrated lens at equilibrium, mass of dry pure
lens, and mass of Vitamin E loaded in the lens, respectively. Both the listed and
measured Q‟s are shown in Table 2-1. Additionally, the values of the equilibrium water
content (EW) which is defined as mass of water absorbed by unit mass of pure lens, i.e.,
100
l
veleq
W
WWWEW (2-2)
are also listed in Table 2-1. Results show that ACUVUE® ADVANCE™ has the highest
EW (86.0 ± 2.3) and NIGHT&DAY™ has a relatively low EW (31.1 ± 5.5). The effect of
Vitamin E loading on Q and EW are clearly seen in Figure 2-3. In Figure 2-3A, water
content of Vitamin E loaded lenses tends to decrease relatively linearly as Vitamin E
loading increases. However, Weq of Vitamin E loaded lenses increases as Vitamin E
loading, which may be causing the decrease in the Q values. To observe the effect of
Vitamin E loading on water amount absorbed in lens polymers, EW was plotted verses
Vitamin E loading in Figure 2-3B. The EW for Vitamin E loaded lenses is also less than
that for the pure lenses for each type of lens but the trends are different. The EW‟s of
36
ACUVUE® OASYS™ and PureVision™ lenses linearly decrease and the values of EW
are 46% and 44% respectively for about 20% Vitamin E loading. The EW‟s of
NIGHT&DAY™ and O2OPTIX™ lenses decrease by about 10% for Vitamin E loadings
of about 10% but there is negligible decrease in EW‟s with further increase in Vitamin E
loadings. The latter behavior for the NIGHT&DAY™ and O2OPTIX™ lenses suggests
that at low loadings, the Vitamin E is solubilized in the lens and so it reduces the water
content of the gel because of its hydrophobicity but beyond a critical weight fraction the
extra Vitamin E simply phase separates, and thus it has no further effect on the EW.
The critical Vitamin E loading which can be solubilized by the NIGHT&DAY™ and
O2OPTIX™ appears to be less than 10%, which is consistent with the values obtained
in the later sections based on drug transport data (6.2% for NIGHT&DAY™ and 9.7%
for O2OPTIX™). The continuous linear decrease in EW for ACUVUE® OASYS™ and
PureVision™ lenses suggests that these lenses can either solubilize large amounts of
Vitamin E or the Vitamin E that phase separates coats the polymer and thus continues
to reduce the EW.
2.2.4 Size Change Due to Vitamin E Loading
The sizes of the contact lenses are expected to increase due to Vitamin E uptake.
The diameters of the lenses both with and without Vitamin E were measured both in dry
and hydrated states, and the size changes of lenses after loading the Vitamin E are
shown in Figure 2-4. The % dry and hydrated diameter increase are the increase in the
dry and hydrated diameter divided by the dry and hydrated diameter of the lens without
Vitamin E, respectively. The solid lines in the figure are the best fit straight lines. Figure
2-4A shows that the dry diameter change of lenses is about 30 % of the Vitamin E
loading. For example, about 30 % Vitamin E loaded lens shows increase of about 10%
37
in diameter in dry state, which suggests that the expansion of lens by Vitamin E loading
is isotropic. In Figure 2-4B, wet diameter change is much less than dry diameter change,
which is expected because Vitamin E does not absorb water. For example, lenses with
about 30% Vitamin E loaded lens expand about only 6.5 % in diameter. From
application perspective, changes in wet diameter should be small to preserve the power
of the contact lens, and all the lenses show less than 8 % increase in wet diameter for
about 40% of Vitamin E, which can likely be tolerated by eyes. There may be further
changes to the corrective power due to refractive index changes in the lens. In any case,
if there is a significant change in the power of the lens, the listed power for a lens can
be modified from the original value.
2.2.5 Ion Permeability of Vitamin E Loaded Lenses
Ion permeability of contact lenses is a critical variable for lens motion on the eye
according to Domscheke et al [77]. The thickness of the lens varies in the radial
direction and the exact profiles are not available in literature. To obtain the permeability,
each lens was treated as a section of a sphere with radius equal to the known base
curve of the lens and 80 m in thickness. The calculated values of ion permeability are
plotted in Figure 2-5A as a function of the Vitamin E loading for O2OPTIX™,
NIGHT&DAY™ and ACUVUE® OASYS™ lenses. The results show that the ion
permeability of pure O2OPTIX™ is highest among three lenses and is about 3.4 fold
and 2.5 fold that of the pure NIGHT&DAY™ and ACUVUE® OASYS™, respectively.
Also it is clearly seen that the ion permeability decreases as Vitamin E loading
increases for all the lenses.
38
The decrease ion permeability for Vitamin E loaded lens can be seen more clearly
in Figure 2-5B in which the ratio of ion permeability of lens with and without Vitamin E is
plotted as a function of the Vitamin E loading. Interestingly, the graphs are almost the
same for O2OPTIX™ and NIGHT&DAY™ and the decrease in ion permeability by
Vitamin E is much larger for ACUVUE® OASYS™ for the same Vitamin E loadings
compared to the other two lenses. Dion should be larger than 6.0 10-6 mm2/min for
sufficient on-eye movement of lens according to Nicolson et al [25]. Figure 2-5 indicates
that all Vitamin E loaded lenses in our study have adequate ion permeability to maintain
on-eye motion.
2.2.6 Oxygen Permeability of Vitamin E Loaded Lenses
The oxygen permeability of extended-wear contact lenses must be sufficiently high
to avoid deprivation of oxygen to cornea, which could cause adverse responses [20, 21].
The lens permeability (Dk) is the product of the diffusivity D and the oxygen partition
coefficient k, and it is typically expressed in units of 10-11 (cm2/sec)·(mLO2/(mL·mmHg))
or 10-11 mLO2·cm/(sec·cm2·mmHg), which is also referred as a barrer or a Fatt. The
oxygen permeability is an intrinsic property of a material to transport oxygen through its
bulk and is independent of thickness. The oxygen transmissibility, Dk/t, refers to the
oxygen transport capacity of a specific contact lens with thickness t, and it generally
expressed in units of 10-9 cm·mLO2/(sec·mL·mmHg) or 10-9 mLO2/(sec·cm2·mmHg). To
avoid hypoxia, an extended wearable contact lens must provide at least a minimum
oxygen transmissibility (Dk/t) of 87, which cannot be achieved by traditional hydrophilic
contact lens [22]. Recently, the suggested minimum value of Dk/t to avoid hypoxia has
been proposed to increase to 125 [23]. The reported values of Dk values of various
39
commercial contact lenses are 140 for NIGHT&DAY™, 110 for O2OPTIX™, 103 for
ACUVUE® OASYS™ and 91 for PureVision™. With an approximate average thickness
of 80 m, these commercial silicone hydrogel contact lenses can provide sufficient
oxygen transmissibility to be used for extend wear.
The influence of Vitamin E loading on oxygen transport through the contact lenses
was determined by mounting the lenses is a diffusion cell with gradient in the dissolved
oxygen concentration across the lens, and then measuring the oxygen concentration in
the receiver chamber. Below a model is presented to fit the measured oxygen
concentration data to determine the oxygen diffusivity through the lens.
Overall mass balance of dissolved oxygen in the closed diffusion cell is given by
ddrrddrr CVCVCVCV 00 (2-3)
where Vr and Vd are the DI water volumes of the receiving and donor compartments,
respectively, and Cr and Cd indicate the dissolved oxygen concentrations with initial
concentration of Cr0 and Cd0 in the receiving and donor chambers, respectively. Since
the lens volume is substantially less than the fluid volume, the system reaches a
pseudo-steady state very rapidly and thus the oxygen flux through the lens can be
expressed as
)( rd
r
r CCh
ADk
dt
dCV (2-4)
where A and h are the surface area and the average thickness of hydrated lens
respectively; D is the oxygen diffusion coefficient of the lens material and k is the
oxygen partition coefficient between lens and DI water. The above equation implicitly
assumes negligible mass transfer resistance in the boundary layers in the receiver and
donor compartments. This assumption was verified by showing that the measured
40
oxygen concentration profiles were not sensitive to stirring at stirring speeds of 900 rpm.
Equations 2-3 and 2-4 can be combined to give:
)( 00
r
dr
dr
d
r
r
dr CVV
VV
V
C
V
C
h
ADk
dt
dC (2-5)
The solution to the above equation with the initial condition Cr (t = 0) =Cr0 is
)exp()exp(1)( 0
00 th
ADk
VV
VVCt
h
ADk
VV
VV
V
C
V
C
VV
VVC
dr
dr
r
dr
dr
d
r
r
d
dr
dr
r
(2-6)
The parameter DkA/h can be obtained by fitting the experiment data to the above
equation using the function „fminsearch‟ in MATLAB® . The exact value of D through
various lenses could not be directly obtained because the detailed shapes of the lenses
were not available in literature, but could be calculated by using the approximate
surface area of these lenses described in Chapter 2.3.5. The validity of this approach
was established by measuring oxygen diffusivity through pHEMA hydrogels prepared of
two different thicknesses from the procedures reported in earlier study [38]. The
measured value of 14.6±1.3 for the synthesized hydrogel with water content 41.1% was
in good agreement with reported value of 12.9 for conventional hydrogel materials of
which oxygen permeability is primarily determined by its water content [78].
The effect of Vitamin E loading on Dk of silicone contact lenses is shown in Figure
2-6. The calculated Dk values were 148, 118, and 111 for NIGHT&DAYTM, ACUVUE®
OASYSTM and O2OPTIXTM, respectively, which were also in good agreement with the
reported Dk values from the manufacturers and other research groups, providing the
accuracy of the measurement methods [79]. The results show Vitamin E loading in
NIGHT&DAYTM slightly reduces the oxygen permeability when the Vitamin E amount
goes up to about 75%. On the other hand, no significant change was observed for
41
ACUVUE® OASYSTM and O2OPTIXTM up to about 35% of Vitamin E loading in the lens.
While it is not feasible to quantitatively evaluate the effect of Vitamin E on Dk values
due to the relatively large standard deviations in the measured values, it is clear that the
Dk value of these Vitamin E loaded lenses with average thickness 80 m are still
sufficiently high to meet the minimum requirements to avoid hypoxia. The results that
Vitamin E loading in the lens has much higher influence on ion transport than on oxygen
transport suggests that most Vitamin E aggregates exist in the hydrophilic polymer
region in the gel matrix. This is plausible because Vitamin E likely has a much lower
solubility in the hydrophilic regions than in the hydrophobic silicone-rich region in the gel
matrix. Since ion transport occurs primarily through the hydrophilic channels, the
presence of Vitamin E aggregates significant reduces the ion permeability. On the other
hand, oxygen transport occurs mainly through the silicone-rich channels, which may not
contain Vitamin E aggregates resulting in a minimal attenuation in oxygen permeability.
2.2.7 Transmittance of Vitamin E Loaded Lenses
In addition to correcting vision, a contact lens could potentially also prevent or
minimize exposure of the corneal tissue to damaging effects of UV light. Currently
ACUVUE® is the only brand that claims the benefit of protection against UV radiation
[80]. Figure 2-7 shows the measured transmittance spectrum for three commercial
contact lenses used in our study. NIGHT&DAYTM and O2OPTIXTM have no significant
protection against UVB (280-315 nm) and UVA (315-400 nm), while ACUVUE®
OASYSTM completely blocks UVB and UVA radiation. These results match the reported
UV transmittance characteristics of silicone hydrogel contact lenses reported by the
manufacturers and other independent research group [80, 81].
42
The effect of Vitamin E loading on the transmittance spectrum of NIGHT&DAYTM
and ACUVUE® OASYSTM is shown in Figure 2-8. The results clearly show that Vitamin
E loaded NIGHT&DAYTM lenses completely block UVB radiation and also partially block
UVA radiation proportionally to the Vitamin E loading. Since ACUVUE® OASYSTM block
UV radiation, Vitamin E loading only marginally increases the UV protection for these
lenses. The UV radiation is known to induce photo-oxidation of Vitamin E transforming
Vitamin E into various photoproducts [82, 83]. To explore the effect of photo-oxidation
on protection again UV radiation, the transmittance spectra of lenses was measured as
a function of time while exposing the lenses to natural light. While the UVB blocking
effect of Vitamin E was retained, the ability to absorb UVA radiation decreased for a few
days and then reached equilibrium, as shown in Figure 2-9.
2.2.8 Mechanical Properties of Vitamin E Loaded Silicone Hydrogel
Kim et al. had found that the frequency storage moduli (E‟) and the loss moduli (E‟‟)
have significant dependency on the compositions of silicone hydrogels [84]. It was
shown that as the content of TRIS in the gel increases, both E‟ and E‟‟ increase. These
viscoelastic properties also depend on the frequency. The modulus of a lens has
important consequences on vision correction and safety. The modulus of silicone
hydrogels is larger than that of hydrogels and the most suitable lens is the one that
balances the advantages of silicone hydrogels while not significantly influencing
modulus. The modulus of commercial silicone hydrogel lenses lies in the range of 1-1.5
MPa [44].
The dynamic mechanical analyze results for hydrogel with different Vitamin E
loading are shown in Figure 2-10, which indicates that the storage modulus of the
silicone hydrogel decreased as the amount of Vitamin E in the hydrogel increased. In
43
addition, with similar amount of Vitamin E loading, the storage modules of the gel by
directly adding Vitamin E into the monomer mixture before photo-curing is smaller than
that by soaking the polymerized silicon hydrogel in Vitamin E/ethanol solution.
In this chapter we have shown the approach of in situ creation of transport barriers
of Vitamin E, and the material properties as extended wear contact lens were
characterized. The results suggested that with proper amount of Vitamin E loading the
Vitamin E loaded commercial silicone hydrogel lens can still be qualified for continuous
wear. For example, Vitamin E loading in the NIGHT&DAYTM lens leads to slight increase
in lens sizes (6.5% increase for 30% loading), a slight reduction in oxygen diffusion
(about 40% reduction for 75% loading), and a more significant reduction in the ion
permeability (50% reduction for 10% loading). However, these changes are not
sufficient to preclude use of the Vitamin E laden commercial lenses for extended wear.
Additionally, Vitamin E loading has a beneficial effect of blocking UV radiation which will
reduce the corneal damage due to UV light.
44
0.0
0.2
0.4
0.6
0.8
1.0
1.2
0.00 0.10 0.20 0.30 0.40
Concentraition of soaking solution
(g vitamin E/ml ethanol)
Vit
am
in E
lo
ad
ing
(g
vit
am
in E
/g p
ure
le
ns
)
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
PureVision™
Figure 2-1. Correlation of Vitamin E loading and concentration of soaking solution for
different lenses. The lines are the best fit straight line to data. The slope and R2 of the line are 5.26, 0.9692 (ACUVUE® OASYS™), 2.53, 0.9860 (NIGHT&DAY™), 3.30, 0.9918 (O2OPTIX™), 4.35, 0.9997 (PureVision™), respectively.
45
Figure 2-2. Images of Commercial NIGHT&DAYTM contact lens (left in panel A) and
NIGHT&DAYTM lens with 30%Vitamin E loading (right image in panel A and panel B). Photos courtesy of Cheng-Chun Peng.
B A
46
0
5
10
15
20
25
30
35
40
0.0 0.2 0.4 0.6 0.8 1.0 1.2
Vitamin E loading (g vitamin E/g pure lens)
Q (
%)
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
PureVision™
0
10
20
30
40
50
60
70
0.0 0.2 0.4 0.6 0.8 1.0 1.2
Vitamin E loading (g vitamin E/g pure lens)
EW
(%
)
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
PureVision™
Figure 2-3. Plot of A) water content (Q) B) EW of Vitamin E loaded lenses versus
Vitamin E loading.
B
A
47
0
5
10
15
20
25
30
35
0.0 0.2 0.4 0.6 0.8 1.0
Vitamin E loading (g vitaming E/g pure lens)
% D
iam
ete
r in
cre
as
e (
dry
)
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
PureVision™
0
5
10
15
20
25
30
35
0.0 0.2 0.4 0.6 0.8 1.0
Vitamin E loading (g vitamin E/g pure lens)
% D
iam
ete
r in
cre
as
e (
we
t)
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
PureVision™
Figure 2-4. Percent increase in diameter of A) dry lenses B) wet lenses before and
after loading Vitamin E. Lines are best fit straight lines passing zero to the data.
A
B
48
1.0E-06
1.0E-05
1.0E-04
1.0E-03
1.0E-02
0 0.2 0.4 0.6 0.8
Vitamin E loading (g Vitamin E/g pure lens)
Ion
ofl
ux
dif
fus
ion
co
eff
icie
nt,
Dio
n
(mm
2/m
in)
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
0
0.2
0.4
0.6
0.8
1
1.2
0 0.2 0.4 0.6 0.8
Vitamin E loading (g Vitamin E/g pure lens)
Dio
n / D
ion
(p
ure
len
s)
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
Figure 2-5. Effect of Vitamin E loading on ion permeability of lenses. The error bars
denote 95% confidence intervals. The solid dash line in A) indicates the minimum requirement for sufficient on-eye movement [25].
A
B
49
0
20
40
60
80
100
120
140
160
180
0 0.2 0.4 0.6 0.8
Vitamin E loading (g Vitamin E/g pure lens)
Dk
(B
arr
er)
NIGHT&DAY
ACUVUE® OASYS
O2OPTIX
pHEMA gel (0.1 mm-thick)
pHEMA gel (0.2 mm-thick)
Figure 2-6. Effect of Vitamin E loading on oxygen permeability (Dk). Data are presented
as mean ± S.D. with n 3. The reported values from manufacturers are shown in hollow marker. The solid dash line indicates the minimum requirement to avoid deprivation of oxygen to cornea [22].
50
0
20
40
60
80
100
200 250 300 350 400 450 500
Wavelength (nm)
Tra
ns
mit
tan
ce
(%
)
Night & Day
Acuvue Oasys
O2 Optix
Figure 2-7. Transmittance spectrum for commercial contact lenses. All measurements
were conducted within 24 hours after sample preparation, and data are presented as mean ± S.D. with n = 3.
51
0
20
40
60
80
100
200 250 300 350 400 450 500
Wavelength (nm)
Tra
ns
mit
tan
ce
(%
)
0 g Vitamin E/g pure gel0.15 g Vitamin E/g pure gel0.28 g Vitamin E/g pure gel0.42 g Vitamin E/g pure gel
0
20
40
60
80
100
200 250 300 350 400 450 500
Wavelength (nm)
Tra
ns
mit
tan
ce
(%
)
0 g Vitamin E/g pure gel
0.21 g Vitamin E/g pure gel
0.42 g Vitamin E/g pure gel
Figure 2-8. Transmittance spectrum for a) NIGHT&DAYTM and b) ACUVUE® OASYSTM
with different Vitamin E loading. All measurements were conducted within 24 hours after sample preparation, and data are presented as mean ± S.D. with n = 3.
A
B
52
0
20
40
60
80
100
200 250 300 350 400 450 500
Wavelength (nm)
Tra
ns
mit
tan
ce
(%
)
1 day
10 days
21 days
0
20
40
60
80
100
200 250 300 350 400 450 500
Wavelength (nm)
Tra
ns
mit
tan
ce
(%
)
1 day
10 days
21 days
Figure 2-9. Transmittance spectrum of NIGHT&DAYTM with Vitamin E loading A) 0.15 g
Vitamin E/g pure lens and B) 0.28 g Vitamin E/g pure lens, and data are presented as mean ± S.D. with n = 3.
B
A
53
0
2
4
6
8
10
12
14
16
0 10 20 30 40 50 60
Frequency (Hz)
Sto
rag
e M
od
ulu
s (
MP
a)
Pure gel
0.11 g VE/g gel
(soaking)0.20 g VE/g gel
(soaking)0.30 g VE/g gel
(soaking)0.10 g VE/g gel
(direct loading)0.20 g VE/g gel
(direct loading)
Figure 2-10. Dependence of storage module of Vitamin E loaded silicone hydrogel on
frequency. Vitamin E was loaded by either soak the silicone hydrogel into Vitamin E-ethanol solution or by directly adding Vitamin E into monomer mixture before polymerization.
54
Table 2-1. List of silicone hydrogel extended wear commercial contact lens (dipoter -6.50) explored in this study (n = 6).
Commercial name a
(manufacturer) Material a
Dry weight measured
[mg]
Water content, Q
measured (listed a)
[%]
EW measured
[%]
Diameter [mm]
Wet measured (listed a)
Dry measured
ACUVUE® ADVANCE™ (Johnson&Johnson Vision Care, Inc., Jacksonville, FL) ACUVUE® OASYS™ (Johnson&Johnson Vision Care, Inc., Jacksonville, FL) NIGHT&DAY™ (Ciba Vision Corp., Duluth, GA) O2OPTIX™ (Ciba Vision Corp., Duluth, GA) PureVision™ (Bausch&Lomb, Inc., Rochester, NY)
Galyfilcon A
Senofilcon A
Lotrafilcon A
Lotrafilcon B
Balafilcon A
19.7 ± 0.3
21.7 ± 0.1
22.2 ± 0.3
25.9 ± 0.2
21.0 ± 0.2
46.2 ± 0.7 (47)
36.9 ± 0.9 (38)
23.6 ± 0.3 (24)
31.5 ± 1.3 (33)
35.0 ± 0.7 (36)
86.1 ± 2.3
58.4 ± 1.5
27.3 ± 0.6
46.0 ± 2.7
53.9 ± 1.7
14.40 ± 0.31 (14.0)
14.12 ± 0.26 (14.0)
13.92 ± 0.07 (13.8)
14.43 ± 0.23 (14.2)
14.18 ± 0.15 (14.0)
11.46 ± 0.34
12.18 ± 0.29
12.85 ± 0.15
12.78 ± 0.12 12.49 ± 0.17
a Referred from product packages
55
CHAPTER 3 HYDROPHILIC DRUG DELIVERY BY VITAMIN E LOADED SILICONE
HYDROGEL
In Chapter 2 we successfully established the approach to prepare silicone
hydrogel contact lens that containing Vitamin E and evaluated its potential as
continuous wear contact lens. In this chapter we focused on the investigation of
the efficacy of Vitamin E aggregates in the gel matrix as diffusion barriers to
extend the drug release duration through the gel matrix. Three different
ophthalmic drugs were explored in this chapter, including timolol (beta blocker
used for treating glaucoma), dexamethasone 21-disodium phosphate (DXP, anti-
inflammatory corticosteroid), and fluconazole (anti-fungal). These drugs were
chosen because they are hydrophilic at the physiological pH, which should have
negligible affinity to the desired Vitamin E barriers.
3.1 Materials and Methods
3.1.1 Materials
Five commercial silicone contact lenses (diopter -6.50) that are used in this
study are described earlier in Table 2-1. Dexamethasone 21-disodium phosphate
(DXP, 99%), timolol maleate (98%), fluconazole (≥98%), 2-hydroxyethyl
methacrylate (97%), sodium hydroxide pellets (97+%), ethanol (99.5%), and
Dulbecco‟s phosphate buffered saline (PBS) were purchased from Sigma-Aldrich
Chemicals (St. Louis, MO) and ethylene glycol dimethacrylate (EGDMA) from
Sigma-Aldrich Chemicals (Milwaukee, WI). Sodium chloride (99.9+ %) were
purchased from Fisher Chemical (Fairlawn, NJ). Darocur® TPO was kindly
provided by Ciba Specialty Chemicals (Tarrytown, NY) and Vitamin E (D-alpha
56
tocopherol, Covitol® F1370) was gifted by Cognis Corporation. All chemicals
were used as received without further purification if not specifically mentioned.
3.1.2 Drug Loading into Pure Lenses
The commercial silicone contact lenses were rinsed with DI water and then
dried in air before further use. The drug timolol maleate was converted to timolol
base for further use by increasing the pH of timolol maleate solution, and then
separating out the precipitated timolol base. All other drugs were used as
supplied. The drug (timolol, DXP, fluconazole) was loaded into the lenses by
soaking the lens either in 2 mL of a drug-PBS solution for 1 or 7 days or in the
same volume of a drug-ethanol solution for 3 hours. During soaking the lens in
either solution, the dynamic concentration in the solution was not monitored since
the absorbance of these drugs in this concentration range was beyond the
measurement limit of the UV-VIS spectrometer. At the end of the loading stage
the lens was taken out and excess drug solution was blotted from the surface.
The lens was then dried in air overnight, and used for later release experiments.
3.1.3 Vitamin E Loading into Pure Lenses
Vitamin E was loaded into lenses by soaking the lens in 3 mL of a Vitamin
E-ethanol solution for 24 hours. Vitamin E-ethanol solutions of various
concentrations were prepared by simply mixing Vitamin E and ethanol with
vortexing for a few seconds followed by moderate magnetic stirring for several
minutes. After the loading step, the lens was taken out and excess Vitamin E-
ethanol solution on the lens surface was blotted, and the lens was then dried in
air overnight.
57
3.1.4 Drug Loading into Vitamin E Loaded Lenses
The drug was loaded in Vitamin E loaded lenses either by directly adding
drug in the Vitamin E-ethanol solution before soaking the pure lens in the solution
or by soaking the Vitamin E loaded lens in a drug/PBS solution. For the case of
adding drug in a Vitamin E/ethanol solution, the drug was dissolved in 3 mL of a
Vitamin E/ethanol solution and then the pure lens was soaked in this
drug/Vitamin E/ethanol solution for 24 hours. For the case of soaking in
drug/PBS solution, the Vitamin E loaded lens was soaked in 2 mL of a drug/PBS
solution until equilibrium.
3.1.5 Drug Release Experiments
The drug release experiments were carried out by soaking a drug loaded
lens in 2 mL of PBS. During the release experiments, the dynamic drug
concentration in the PBS was analyzed by measuring the absorbance of solution
with a UV-VIS spectrophotometer (Thermospectronic Genesys 10 UV). The
absorbance of solution was measured at wavelength of 241 nm for DXP, 294 nm
for timolol, and 210 nm for fluconazole. Control experiments were conducted to
ensure that diffusion of Vitamin E from the lenses was negligible and so it did not
interfere with the drug detection.
3.2 Results and Discussion
3.2.1 Dynamics of Drug Transport from Contact Lenses without Vitamin E
Figure 3-1 shows the dynamics of timolol release by each of five contact
lenses soaked in 0.8 mg/mL of timolol-PBS solution or timolol-ethanol solution.
The soaking duration was either 24 hours or 7 days in PBS and 3 hours in
ethanol, but the release profiles for 24 hours in PBS were not drawn in Figure 3-1
58
since they were identical to those for 7 days soaking in PBS. To observe the
effect of different loading methods on timolol release dynamics, mass of drug
released divided by total drug released is plotted as a function of time. All the
lenses release 90% of timolol in less than 1.5 hours. In addition, timolol release
profiles for different loading methods overlap for each lens except for
PureVision™ lens, which shows a slightly faster release from the lens soaked in
timolol-ethanol solution than that soaked in PBS medium. ACUVUE® OASYS™
lens releases 90% of timolol relatively slowly for 1.2 hours compared to the other
lenses. ACUVUE® ADVANCE™ lens exhibits rapid timolol release lasting less
than 0.5 hour and the other three lenses show comparable release durations. It is
observed that the release durations of timolol are not correlated to the water
content of the lenses. The total amount of drug released is the highest by
PureVision™ (about 57 g), lowest by NIGHT&DAY™ (about 22 g), and those
of the other lenses are similar ranging 26-30 g based on PBS medium soaking
method. The amounts of timolol uptake and release are also uncorrelated to the
water content, likely due to differences in the hydrophilic components of the
lenses, which lead to differences in drug binding to the hydrophilic component
rich phases in the lenses. It is interesting that all the lenses soaked in ethanol
solution for 3 hours release substantially high total amount of timolol; about 2.5 -
3 times more than those soaked in PBS solution. For example, ACUVUE®
OASYS™ lens soaked in PBS solution for 7 days releases 28 g of timolol, but
that soaked in ethanol solution for 3 hours release about 95.7 g. The increased
uptake of timolol from ethanol soaking is likely due to the fact that timolol does
59
not ionize in ethanol and so it preferentially binds to the polymer. In PBS, the
drug is almost entirely ionized, which leads to a very large solubility in water, and
consequently to small binding to the gel.
The drug release from control lenses, i.e., without Vitamin E, were also
conducted with the other two drugs (DXP and fluconazole) but these are not
presented here because the major conclusions are the same as those mentioned
above in the context of timolol. The % release profiles were independent of the
method of loading and the total release durations were all about 1-10 hours.
These control data are presented in later sections while comparing the results
with the release from the Vitamin E loaded lenses.
3.2.2 Dynamics of Drug Transport from Vitamin E Loaded Lenses
3.2.2.1 Timolol-Vitamin E loaded lenses
Figure 3-2 shows timolol release dynamics by Vitamin E loaded lenses for
different loadings of Vitamin E. Timolol and Vitamin E were loaded into lenses
simultaneously by soaking the lens in 0.8 mg/mL of timolol-Vitamin E-ethanol
solution for 24 hours. For pure lenses (no Vitamin E loading), timolol was loaded
by soaking in timolol-ethanol solution of 0.8 mg/mL for 3 hours. It is clearly seen
in the figure that the rate of timolol release by all the lenses except PureVisionTM
decreases as Vitamin E loading increases, while the total drug release amount
does not change significantly. Specifically, NIGHT&DAY™ shows 9.8-fold
release time for 16% Vitamin E loading corresponding to release time of about
5.5 hours, 76-fold for 27% corresponding to 43 hours release, and 341-fold for
74% corresponding to 192 hours release. The total amount of timolol released by
NIGHT&DAY™ is lowest at about 50 g. The drug transport data for
60
PureVision™ lenses suggests that the Vitamin E simply dissolves in the matrix
leading to negligible barrier effect. However the drug transport data for ACUVUE®
OASYS™ lenses shows a significant barrier effect, which in combination with the
EW data suggests that the barrier effect in these lenses likely arises due to
Vitamin E that coats polymer fibers rather than forming larger aggregates, which
appears to be the mechanism for NIGHT&DAY™ and O2OPTIX™ lenses.
To explore the effect of the loading method, timolol was also loaded into
Vitamin E containing lenses by soaking the lenses in timolol-PBS solution for 7
days. Timolol release profiles of the NIGHT&DAY™, ACUVUE® OASYS™ and
O2OPTIX™ lenses for sequential loading of Vitamin E and timolol are also shown
in Figure 3-2. It can be clearly seen that this method also increases timolol
release duration compared to the control lenses without Vitamin E. Additionally,
there is an increase in the total amount of drug released for the higher Vitamin E
loading (74% for NIGHT&DAY™, and 97% for O2OPTIX™). Therefore, loading
timolol and Vitamin E at the same time through ethanol medium is much more
efficient way for preparation of timolol-Vitamin E loaded lenses. For O2OPTIX™,
with same amount of Vitamin E loading, the release profiles from the lenses
where timolol and Vitamin E were loaded sequentially are almost the same as for
the case where timolol and Vitamin E were loaded simultaneously. However, for
ACUVUE® OASYS™ and NIGHT&DAY™ even though the release durations
from different loading methods are similar to each other, the release profiles are
slightly different. The difference is likely to be resulted from the
non-homogeneous distribution of timolol inside the lens. Timolol loaded by
61
drug-PBS solution goes into the gel matrix by diffusion for longer time, leading to
a well distribution in the lens. On the other hand, timolol uptake in drug-ethanol
solution might result in high drug concentration in the center region of lens after
ethanol evaporation.
The morphology of the Vitamin E laden lens could potentially change over
time, which could impact the drug transport. To investigate this issue,
NIGHT&DAY™ lenses with various Vitamin E loadings that were utilized in the
drug release experiments were soaked in 2 mL PBS solution after the release
experiment were over. The lenses were subsequently stored for 6 months and
then further soaked in 250 mL DI water with moderate stirring for 48 hours to
remove the residual timolol prior to be used in second release experiment. The
cleaned Vitamin E lenses were dried and weighed to ensure that the Vitamin E
loading was kept the same as the initial loading. The dry weight of the lens was
within 1% difference of that measured immediately after the initial Vitamin E
loading, which proves that Vitamin E does not diffuse out into PBS during the
storage. The lenses were then soaked in 0.8 mg/mL timolol-PBS solution for 7
days to load the drug. After the drug loading, the drug release profiles were
measured in 2 mL PBS (Figure 3-3). The release profiles in this case were
almost identical to the first release profiles; this proves that the morphology of the
Vitamin E laden lenses is stable even when soaked in PBS for 6 months, and
thus the drug release behavior of these lenses will not degrade during packaging
and shelf storage. The morphology of the Vitamin E laden lenses does not
62
change during PBS soaking likely because of the negligible solubility of Vitamin E
in PBS.
3.2.2.2 DXP-Vitamin E loaded lenses
DXP release profiles for various Vitamin E loaded commercial lenses are
shown in Figure 3-4. The dry Vitamin E loaded lenses were soaked in 0.7 mg/mL
DXP-PBS solution for sufficient time to reach equilibrium. In all experiments of
DXP-Vitamin E loaded lenses explored here, the uptake periods were longer
than the release equilibrium time, suggesting that equilibrium was achieved
during loading. Figure 3-4 indicates that, similar to the release rates for timolol,
the DXP release rates from all lenses decrease as the Vitamin E loading
increases, while the total drug release amount is relatively independent of the
Vitamin E loading. With similar Vitamin E loading, ACUVUE® OASYS™ has the
longest drug release time, followed by NIGHT&DAY™ and O2OPTIX™, while
PureVision™ shows negligible increase. For example, ACUVUE® OASYS™ lens
releases about 27 g of DXP in 7 days for 10% Vitamin E loading and in 3 weeks
for 23% Vitamin E loading, while 40 g of DXP in PureVision™ released in only
8 hours even with 36% Vitamin E loaded inside. In addition, even though the
drug release duration is much longer, the duration release time increase ratio by
Vitamin E loaded lens for DXP is similar to that for timolol with similar Vitamin E
loading amount. This suggests that the attenuation in drug release rates is similar
for all hydrophilic drugs even though the diffusivities of the drugs in the pure
lenses may be vastly different, which will be further discussed later.
63
3.2.2.3 Fluconazole-Vitamin E loaded lenses
To further validate the hypothesis that the attenuation in drug release rates
is similar for all hydrophilic drugs, we explored transport of an antifungal drug
fluconazole in Vitamin E laden lenses. Figure 3-5 shows the fluconazole release
dynamics from Vitamin E loaded NIGHT&DAY™, ACUVUE® OASYS™ and
O2OPTIX™ lenses. PureVision™ was not tested because of the marginal impact
of Vitamin E loading on transport rates of timolol and DXP from this lens.
To load drugs into lenses, the Vitamin E loaded lenses were soaked in 0.7
mg/mL fluconazole-PBS solution for sufficient time to reach equilibrium. The
results clearly show a significant reduction in release rates due to Vitamin E
loading in the lenses. For example, NIGHT&DAY™ lenses release about 60 g
of fluconazole in 10 hours for 17% Vitamin E loading, in 24 hours for 26%, 88
hours for 39 % and 227 hour for 66% Vitamin E loading, which is a 6.2, 14, 55
and 142-fold release duration increase, respectively. The total amount of
fluconazole released by different lens is similar, with the exception of O2OPTIX™,
which has a slightly higher drug release of about 80 g. With similar Vitamin E
loading, ACUVUE® OASYS™ shows longer fluconazole release period than
NIGHT&DAY™ and O2OPTIX™.
The effect of Vitamin E loading on hydrophilic drug transport is summarized
in Figure 3-6. The increase in the release times from Vitamin E loaded lenses
relative to release times from the control lenses without Vitamin E is relatively
similar for the three hydrophilic drugs particularly for ACUVUE® OASYS™ and
O2OPTIX™ lenses. There are some differences from NIGHT&DAY™ lens;
64
fluconazole released by NIGHT&DAY™ lens exhibits a smaller time increase
compares to timolol and DXP. The data also clearly shows that for each drug,
the release time is quadratic to the Vitamin E loading. These issues are
discussed below in the model development section.
3.2.3 Model for Hydrophilic Drugs
The hydrophilic drugs have a negligible partitioning in Vitamin E. The
increase in release times for charged drugs is likely due to the presence of
Vitamin E aggregates inside the gel that act as diffusion barriers. These barriers
lead to an increase in the length of the path that molecules take to diffuse from
inside the gel to the fluid reservoir. The path length of the tortuous path l should
scale as ))(1( * h , where h is the half thickness of lens, and depends on
the microstructure, including particle size and aspect ratio, of the Vitamin E
aggregates distribution in the gel; is the volume ratio of Vitamin E in the dry gel,
and )( * is the fraction that is present as the Vitamin E particles. The fraction
* is assumed to be either existing as bound to the polymer gel or as particles but
in regions of the gel that do not contribute to drug transport. For a diffusion-
controlled release, the time for release can be scales as l2/D. The gel thickness
increases due to Vitamin E uptake, and by assuming isotropic expansion and
small Vitamin E loading, it can be written as )3/1(0 hh , where h0 is a half
thickness of pure lens. The time of release thus scales as
2*
22
0 13
1~
D
h (3-1)
65
The term (1+/ 3)2 does not make a significant contribution to increase in
release time as for as large as 1, this term is less than 2. By neglecting this
term we get
22*22*
0
21~
(3-2)
where time is the duration in which 90% of release is completed and 0 is
the corresponding duration for the lens without Vitamin E. It is noted that
Equation 3-2 is only valid for * . The parameters and * can be obtained by
fitting the data shown in Figure 3-6 to the above model. The error between the
experimental data and model prediction was defined as
exex )/(/})/()/{( 0
2
00 , where )/( 0 and ex)/( 0 are the predicted
release time ratio by model and the experimental release time ratio, respectively.
The parameters and * for timolol, fluconazole and DXP were obtained using
the function „fminsearch‟ in MATLAB® minimizing the error and are listed in Table
3-1. For a given lens, the same value of was imposed in all fits since this
parameter should be the same for all the drugs as it only depends on the
interaction of Vitamin E with the lens matrix. Also the values of should be
similar for all drugs since this is a geometric parameter that only depends on the
microstructure of the Vitamin E laden lenses. The good fits between the model
and the data with identical and similar for each drug further substantiate the
mechanisms and the model presented above.
66
3.2.4 Diffusivities of Drugs in Vitamin E Loaded Lenses
Contact lenses have a complex geometry including curvature with variable
thicknesses from center to edge depending on power. However, a diameter of a
lens (about 14 mm) is much larger than its thickness (about 80 to 100 m) and so
we can simplify the geometry of lens as thin flat film with variable thickness.
Under this assumption, the mass transfer problem for transport in the contact
lens can be described by the following equations:
2
2
y
CD
t
C
(3-3)
where C is the drug concentration in the gel, D is the effective diffusivity and y
and t denote the transverse coordinate and time, respectively. The boundary
conditions for the drug release experiment are
wKCxhytC
yty
C
))(,(
0)0,( (3-4)
where h is the half-thickness of the gel, which depends on the curved lateral
coordinate x, Cw is the drug concentration in the release medium. The first
boundary condition assumes symmetry at the center of the gel and the second
boundary condition assumes equilibrium between the drug concentration in the
gel and that in the PBS phase. A mass balance on the PBS in the beaker yields
dxy
CxPD
dt
dCV hy
S
w
w
)(2
0
(3-5)
67
where Vw is the PBS volume, P(x) is the perimeter of the lens at the coordinate x,
and S is a half of maximum arc length. Finally the initial conditions for the drug
release experiments are
0)0(
)0,(
tC
CtyC
w
i (3-6)
The fluid volume is much larger than lens volume and the solubility of
timolol, fluconazole and DXP is very high in PBS of this pH 7.4, which satisfies
perfect sink condition. Under perfect sink conditions, the set of equations listed
above can be solved analytically to give the following solution for the
concentration profile in the lens:
0
)(4
)12(2
22
))(2
)12(cos(
)12(
4)1(
n
Dtxh
n
i
n
eyxh
n
n
CC
(3-7)
In short time limit, the concentration profile can also be expressed as
Dt
yh
i deCC
4
0
22
(3-8)
This result is only valid for times shorter than the Dtxh 4/)( . By using
Equation 3-5 and Equation 3-8, we obtain the following equation:
surface
i
S
iw
w ADt
CDdxxP
Dt
CD
dt
dCV
4
2)(
4
22
0 (3-9)
where Asurface is the total surface area of the lens. Equation 3-9 can be
integrated to give,
w
surface
iwV
AC
DtC
2 (3-10)
68
The fractional release igel
ww
CV
CVf can thus be expressed as
2
22
h
Dt
V
ADtf
gel
surface
(3-11)
where h is the mean thickness of the gel defined as surface
gel
A
Vh . The above
equation is only valid for times shorter than Dth 4/min , where hmin is the
minimum gel thickness, which typically equals the center thickness for negative
power contact lenses.
Figure 3-7 plots % drugs release by Vitamin E loaded NIGHT&DAYTM
lenses as a function of square root of time for timolol. The lines in the figure are
the best fit straight line to short time release data. The fits are all good with R2
values larger than 0.98 showing that the drug transport in these lenses is
diffusion controlled. The short-time data in the drug release profiles from Vitamin
E laden lenses is linear for all drugs and all lenses (data only shown for timolol
release from NIGHT&DAYTM) proving that the transport is diffusion limited for all
cases.
The results reported in this chapter conclusively show that Vitamin E
loading in commercial silicone contact lens can substantially increase the release
duration of hydrophilic drugs. The mechanism of increase in duration is due to
the barrier effect of Vitamin E. While it is reasonable to assume that the effect is
caused by the presence of particles of Vitamin E, it is also possible that Vitamin
E does not form macroscopic aggregates and is simply adsorbed on the polymer
gel. The surface adsorption could impede surface diffusion of the drug along the
69
polymer leading to a reduction in effective diffusion rates. Also, the release
profiles from the Vitamin E laden contact lenses are not zero-order and that may
have significant clinician implications.
70
0
20
40
60
80
100
120
0.0 1.0 2.0 3.0 4.0 5.0 6.0
Time (hr)
Dru
g r
ele
as
e,
M/M
f (%
)
NIGHT&DAY, PBS for 7 days (17.7 ug)
NIGHT&DAY, ethanol for 3 hours (46.4 ug)
O2 OPTIX, PBS for 7 days (27.6 ug)
O2 OPTIX, ethanol for 3 hours (71.8 ug)
Pure Vision, PBS for 7 days (56.2 ug)
Pure Vision, ethanol for 3hours (141.3 ug)
ACUVUE OASYS, PBS for 7 days (28.0 ug)
ACUVUE OASYS, ethanol for 3 hours (95.7 ug)
ACUVUE ADVANCE, PBS for 7 days (29.1 ug)
ACUVUE ADVANCEl-ethanol for 3 hours (91.5 ug)
Figure 3-1. Effect of timolol loading method on profile of timolol release by
commercial contact lenses. Drug release (M) divided by total amount released (Mf) are plotted as a function of time. Timolol was loaded by soaking the lens in 0.8 mg/mL of indicated medium for indicated duration of time. Total amount of drug released for each lens is marked in parenthesis on the legends.
71
0
20
40
60
80
100
120
0 100 200 300 400 500
Time (hr)
Dru
g r
ele
as
e (
g)
0 g vitamin E/g pure lens *
0.11 g vitamin E/g pure lens
0.24 g vitamin E/g pure lens
0.42 g vitamin E/g pure lens
0.68 g vitamin E/g pure lens
0 g vitamin E/g pure lens
0.23 g vitamin E/g pure lens
0
10
20
30
40
50
60
0 100 200 300 400 500
Time (hr)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/g pure lens *0.09 g Vitamin E/g pure lens0.16 g Vitamin E/g pure lens0.27 g Vitamin E/g pure lens0.75 g Vitamin E/g pure lens0 g Vitamin E/g pure lens0.16 g Vitamin E/g pure lens0.74 g Vitamin E/g pure lens
Figure 3-2. Profiles of timolol release by Vitamin E loaded contact lenses.
Timolol and Vitamin E were loaded together by soaking A) ACUVUE® OASYS™ B) NIGHT&DAY™ C) O2OPTIX™ D) PureVision™ contact lens in either timolol/Vitamin E-ethanol solution (shown as solid markers), or in timolol-PBS solution (shown as hollow markers). Vitamin E loadings are indicated. Some of data are presented as mean± S.D. with n = 3.
0
10
20
30
40
50
60
0 2 4 6 8
A
B
0
20
40
60
80
100
120
0 2 4 6 8
72
0
10
20
30
40
50
60
70
80
0 100 200 300 400 500
Time (hr)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/g pure lens *0.10 g Vitamin E/g pure lens0.19 g Vitamin E/g pure lens0.34 g Vitamin E/g pure lens0.97 g Vitamin E/g pure lens0 g vitamin E/g pure lens0.19 g Vitamin E/g pure lens0.99 g Vitamin E/g pure lens
0
20
40
60
80
100
120
140
160
180
0 50 100 150
Time (hr)
Dru
g r
ele
as
e (
g)
0 g vitamin E/g pure lens *
0.11 g vitamin E/g pure lens
0.22 g vitamin E/g pure lens
0.39 g vitamin E/g pure lens
Figure 3-2. Continued.
0
50
100
150
200
0 2 4 6 8
0
20
40
60
80
100
0 2 4 6 8
D
C
73
0
5
10
15
20
25
30
0 100 200 300 400
Time (hr)
Dru
g r
ele
as
e (
g)
0 g vitamin E/g pure lens (1st release)
0 g vitamin E/g pure lens (2nd release)
0.16 g vitamin E/g pure lens (1st release)
0.16 g vitamin E/g pure lens (2nd release)
0.74 g vitamin E/g pure lens (1st release)
0.74 g vitamin E/g pure lens (2nd release)
Figure 3-3. Profiles of repeated timolol releases by Vitamin E loaded contact lenses. For the second releases timolol was loaded by used soaking Vitamin E loaded lens in timolol-PBS solution (0.8 mg/mL) for 7 days. Vitamin E loadings are indicated. Some of data are presented as mean ± S.D. with n = 3.
0
5
10
15
20
0 2 4 6 8
74
0
5
10
15
20
25
30
0 500 1000 1500
Time (hr)
Dru
g R
ele
as
e (
g)
0 g vitamin E/g pure lens
0.10 g vitamin E/g pure lens
0.23 g vitamin E/g pure lens
0.32 g vitamin E/g pure lens
0
5
10
15
20
25
0 200 400 600 800
Time (hr)
Dru
g R
ele
as
e (
g)
0 g vitamin E/g pure lens
0.09 g vitamin E/ g pure lens
0.16 g vitamin E/g pure lens
0.24 g vitamin E/g pure lens
Figure 3-4. Profiles of DXP release by Vitamin E loaded contact lenses A)
ACUVUE® OASYS™ B) NIGHT&DAY™ C) O2OPTIX™ D) PureVision™. Vitamin E was loaded first by soaking pure contact lens in Vitamin E-ethanol solution and the lens was dried. And then DXP was loaded by soaking Vitamin E loaded lens in DXP-PBS solution (0.7 mg/mL). Vitamin E loadings are indicated.
0
5
10
15
20
25
30
0 20 40 60 80 100
0
5
10
15
20
0 12 24 36 48 60 72
A
B
75
0
5
10
15
20
25
30
0 100 200 300 400 500 600 700
Time (hr)
Dru
g R
ele
as
e (
g)
0 g vitamin E/g pure lens
0.09 g vitamin E/g pure lens0.20 g vitamin E/ g pure lens
0.35 g vitamin E/g pure lens
0
5
10
15
20
25
30
35
40
45
0 10 20 30 40 50 60
Time (hr)
Dru
g R
ele
as
e (
g)
0 g vitamin E/g pure lens
0.21 g vitamin E/g pure lens
0.36 g vitamin E/g pure lens
Figure 3.4. Continued
0
5
10
15
20
25
30
0 12 24 36 48 60 72
0
10
20
30
40
50
0 2 4 6 8
C
D
76
0
10
20
30
40
50
60
70
80
0 100 200 300 400 500 600
Time (hr)
Dru
g R
ele
as
e (
g)
0 g vitamin E/g pure gel
0.11 g vitamin E/g pure lens0.21 g vitamin E/g pure lens
0.40 g vitamin E/g pure lens
0
10
20
30
40
50
60
70
0 200 400 600 800Time (hr)
Dru
g R
ele
as
e (
g)
0 g vitamin E/g pure lens0.09 g vitamin E/g pure lens0.17 g vitamin E/g pure lens0.26 g vitamin E/g pure lens0.39 g vitamin E/g pure lens0.66 g vitamin E/g pure lens
Figure 3-5. Profiles of fluconazole release by Vitamin E loaded contact lenses A)
ACUVUE® OASYS™ B) NIGHT&DAY™ C) O2OPTIX™. Vitamin E was loaded first by soaking pure contact lens in Vitamin E-ethanol solution and the lens was dried. And then fluconazole was loaded by soaking Vitamin E loaded lens in fluconazole-PBS solution (0.7 mg/mL). Vitamin E loadings are indicated.
0
10
20
30
40
50
60
70
80
0 20 40 60 80
0
10
20
30
40
50
60
70
0 20 40 60 80 100
A
B
77
0
10
20
30
40
50
60
70
80
90
100
0 100 200 300 400
Time (hr)
Dru
g R
ele
as
e (
g)
0 g vitamin E/g pure lens
0.09 g vitamin E/g pure lens
0.18 g vitamin E/g pure lens
0.29 g vitamin E/g pure lens
0.47 g vitamin E/g pure lens
0.68 g vitamin E/g pure lens
Figure 3-5. Continued
0
20
40
60
80
100
0 10 20 30 40 50
C
78
0
20
40
60
80
100
120
140
160
180
0.0 0.1 0.2 0.3 0.4 0.5
Dru
g r
ele
as
e t
ime
in
cre
as
e, /
0
Timolol
DXP
Fluconazole
0
100
200
300
400
500
600
0.0 0.1 0.2 0.3 0.4 0.5
Dru
g r
ela
se
tim
e in
cre
as
e, /
0
Timolol
DXP
Fluconazole
Figure 3-6. Drug release duration increase by Vitamin E loaded contact lenses.
A) ACUVUE® OASYS™ B) NIGHT&DAY™ C) O2OPTIX™ D) ) PureVision™. The lines are best fit 2nd order polynomial curves to data of each lens. Drug release time is the duration in release of 90 % of total drug released.
A
B
79
0
50
100
150
200
250
300
350
0.0 0.1 0.2 0.3 0.4 0.5
Dru
g r
ele
as
e t
ime
in
cre
as
e, /
0
Timolol
DXP
Fluconazole
0
5
10
15
20
25
0.0 0.1 0.2 0.3 0.4 0.5
Dru
g r
ele
as
e t
ime
in
cre
as
e, /
0 Timolol DXP
Figure 3-6. Continued.
C
D
80
0
20
40
60
80
100
120
0 5 10 15 20 25
Time0.5
(hr0.5
)
% D
rug
re
lea
se
0 g vitamin E/g pure lens
0.09 g vitamin E/g pue lens
0.16 g vitamin E/g pure lens
0.27 g vitamin E/g pure lens
0.75 g vitamin E/g pure lens
Figure 3-7. Plot of % timolol release by Vitamin E loaded NIGHT&DAYTM versus
square root of time. The lines are the best fit straight for short time data. All R2‟s are larger than 0.98. Some of data are presented as mean ± S.D. with n = 3.
81
Table 3-1. Model parameters obtained by fitting experimental data to the model
Contact lenses
Timolol Fluconazole DXP
ACUVUE® OASYS™
0.0117 24.2 22.0 28.8
NIGHT&DAY™ 0.0621 47.6 31.5 42.8
O2OPTIX™ 0.0973 35.2 35.9 42.1
PureVision™` 0.1019 1.06 _ 10.95
82
CHAPTER 4 HYDROPHOBIC DRUG DELIVERY BY VITAMIN E LOADED SILICONE HYDROGEL
In Chapter 3, we have showed that the release duration of hydrophilic drugs from
commercial silicone-hydrogel contact lenses can be significantly increased by
incorporating Vitamin E into the lenses. The Vitamin E loaded lenses exhibit slower
release of hydrophilic drugs because Vitamin E is a hydrophobic solute and so
hydrophilic molecules need to diffuse around the Vitamin E barriers, leading to an
effective increase in release times. On the other hand, for hydrophobic drugs the
transport mechanism should be significantly different since now the drug can partition
into the Vitamin E aggregates. We propose that hydrophobic molecules could partition
and diffuse through Vitamin E and the relative high viscosity of Vitamin E compared to
the gel matrix will lead to reduced drug diffusivity. To test this hypothesis, we explore
the transport of dexamethasone (DX), a hydrophobic corticosteroid through Vitamin E
laden silicone hydrogel contact lenses.
DX is a glucocorticoid steroid that relieves eye inflammation and swelling, heat,
redness, and pain caused by chemicals, infection, and/or severe allergies. Prolonged
systemic administration of steroid can cause serious side effects such as diabetes,
hemorrhagic ulcers, skin atrophy, myopathies, osteoporosis and psychosis [85]. In view
of the potential for side effects, controlled release of DX from contact lenses could be
clinically useful. Furthermore, there are several other ophthalmic drugs that are
hydrophobic and have size similar to DX, and thus it can be considered as a test drug to
explore transport of small, hydrophobic molecules through Vitamin E-laden silicone
hydrogel contact lenses. This study will lead to an understanding of the effect of Vitamin
83
E loading on extended drug delivery for hydrophobic drugs, which in turn will allow for
rational design for extended release of other drugs from the lenses.
4.1 Materials and Methods
4.1.1 Materials
Five commercial silicone contact lenses (diopter -6.50) are used in this study,
including ACUVUE® ADVANCETM and ACUVUE® OASYS™ from Johnson&Johnson
Vision Care, Inc. (Jacksonville, FL), NIGHT&DAY™ and O2OPTIX™ from Ciba Vision
Corp. (Duluth, GA) and PureVision™ from Bausch&Lomb, Inc. (Rochester, NY).
Dexamethasone (DX, 98%), ethanol (99.5%), and Dulbecco‟s phosphate buffered
saline (PBS) were purchased from Sigma-Aldrich Chemicals (St. Louis, MO). Vitamin E
(D-alpha tocopherol, Covitol® F1370) was kindly provided by Cognis Corporation. All
chemicals were used as supplied without further purification.
4.1.2 Drug Loading into Pure Lenses
The commercial silicone contact lenses were rinsed with DI water and then air-
dried before further use. To evaluate the effect of different loading approaches, DX was
loaded into the lenses by either soaking the lens in either 2 mL of a drug-PBS solution
for 1 or 7 days or in the same volume of a drug-ethanol solution for 3 hours. While
soaking the lens in either solution, the dynamic concentration in the solution was not
monitored. At the end of the loading stage the lens was taken out and excess drug
solution was blotted from the surface of the lens. The lens was then air-dried and
subsequently used for release experiments.
4.1.3 Vitamin E Loading into Pure Lenses
Vitamin E was loaded into lenses by soaking the lens in 3 mL of a Vitamin E-
ethanol solution for 24 hours. Vitamin E-ethanol solutions of various concentrations
84
were prepared as reported in Chapter 3. After the loading step, the lens was taken out
and excess Vitamin E-ethanol solution on the lens surface was blotted out, and the lens
was then air-dried overnight. The Vitamin E loading amount was determined by
measuring the weight of dry lens before and after loading Vitamin E into the lens.
4.1.4 Drug Loading into Vitamin E Loaded Lenses
The drug was loaded in Vitamin E loaded lenses either by directly adding drug in
the Vitamin E-ethanol solution before soaking the pure lens in the solution or by soaking
the Vitamin E loaded lens in a drug-PBS solution. For the case of adding drug in a
Vitamin E-ethanol solution, the drug was dissolved in 3 mL of a Vitamin E-ethanol
solution and then the pure lens was soaked in the drug/Vitamin E-ethanol solution for 24
hours. For the case of soaking in drug-PBS solution, the Vitamin E loaded lens was
soaked in 2 mL of a drug-PBS solution until equilibrium. While loading DX into lenses,
changes in drug concentration of soaking solution were monitored. The total amount of
drug loaded into the gel was determined by finding the total amount of drug-loss from
the aqueous solution by measuring the absorbance of final solution after soaking at 241
nm for DX with a UV-VIS spectrophotometer (Thermospectronic Genesys 10 UV).
4.1.5 Drug Release Experiments
The drug release experiments were carried out by soaking a drug loaded lens in 2
mL of PBS. During the release experiments, the dynamic drug concentration in PBS
was analyzed in the same manner as described above for the drug loading experiments.
Control experiments were conducted to ensure that diffusion of Vitamin E from the
lenses was negligible and thus did not interfere with the drug detection.
85
4.1.6 Viscoelastic Measurement
The viscoelastic response of pure Vitamin E was measured in a as a function of
frequency with 0.1% strain in a cone and plate rheometer (AR-G2, TA Instruments, New
Castle, DE) with 1000 m gap at 25 oC.
4.2 Results and Discussion
4.2.1 Dynamics of Drug Transport from Contact Lenses without Vitamin E
The DX release profiles from five different contact lenses for three different loading
methods are shown in Figure 4-1. Since DX is a hydrophobic drug and has limited
solubility in PBS, DX-PBS solution of 0.08 mg/mL, which is close to the maximum
solubility of DX in PBS at room temperature, was used for DX loading into lenses. The
concentration of DX-ethanol was the same, i.e., 0.08 mg/mL, as that of DX-PBS solution
for comparison, though the solubility of DX in ethanol is about 1 mg/mL. For ACUVUE®
ADVANCETM, ACUVUE® OASYSTM and O2OPTIXTM, the DX release profiles of three
different loading methods are identical. However, the DX release behaviors by
NIGHT&DAYTM and PureVisionTM lenses exhibit a slight dependency on loading
methods. For these lenses, there is not much difference in the total release amount of
DX from the lenses soaked in DX-PBS solution for two different soaking times, but
slower DX release is observed from lenses that were soaked for 7 days than that for 24
hours. This suggests that equilibrium time for DX loading for these two lenses could be
longer than 24 hours. Among five lenses, NIGHT&DAY™ lens shows the longest
release time (16 hours for 90% of total release) followed by ACUVUE® OASYSTM (10.5
hours), O2OPTIXTM (9.5 hours), and PureVisionTM (8.5 hours), and then ACUVUE®
ADVANCETM has the shortest release time (4.5 hours) by loading the drug with DX-PBS
solution for 7 days. There is a good correlation between the water content of the lenses
86
reported by the manufacturers and the duration of release as shown in Figure 4-1F, with
increasing water content resulting in shorter release durations. For total release amount
of DX, PureVisionTM and ACUVUE® OASYSTM lenses release relatively smaller amounts
(c.a. 28 g and 35 g, respectively) compared to the other three lenses (c.a. 38 to 41
g). There is no correlation between amount of drugs released and the water content,
which is likely because the hydrophobic drugs are expected to partition in the silicone
rich phases, and so the partition coefficients in the gels will be mainly influenced by the
silicone composition of the gels. All the lenses soaked in DX-ethanol solution release
substantially low amount of DX (2 to 8 g). The solubility of DX in ethanol is very high
and the partition coefficient of DX between lens and ethanol is very low in the drug
loading step, which results in low loading of DX.
4.2.2 Dynamics of Drug Transport from Vitamin E Loaded Lenses
The dynamics of DX uptake and release by Vitamin E loaded lenses for four
different Vitamin E loadings are shown in Figure 4-2. The insets in the figure show the
magnified views of the plots for drug release during the initial hours. In these
experiments, Vitamin E was loaded in the lens first then air-dried, and subsequently DX
was loaded by soaking the lens in the DX-PBS solutions. The method of loading by
direct addition of DX in Vitamin E-ethanol was not used since DX loading through PBS
medium was much more efficient as shown earlier. In the figure, all the lenses exhibit
increase in loading or release time as Vitamin E loading increases. With similar Vitamin
E loadings in the lenses, DX loading time is longest for ACUVUE® OASYSTM, followed
by NIGHT&DAYTM, O2OPTIXTM, and shortest for PureVisionTM. For DX loading, the
effect of Vitamin E loading is similar for NIGHT&DAY™ and O2OPTIX™ with about 2-
87
fold loading time increase for about 10% Vitamin E loading, and about 10-fold for about
30% loading. However, the effect of about 10% Vitamin E loading for PureVision™ lens
on loading duration is negligible and even of about 40% loading shows only 6-fold
increase. These behaviors are similar for DX release time increase, even though the
changes in release duration are slightly less than in loading duration. For example,
NIGHT&DAYTM lenses with 27% Vitamin E loading shows 6.5-fold increase in release
duration compared to 9-fold increase in loading duration with the same Vitamin E
loading. The difference between the measured DX delivery time of uptake and release
is likely caused by the accumulation error of drug loss during the measurement process.
The release experiment is conducted in a lower drug concentration range than the
uptake experiment, and therefore contains larger relative error. The comparison for the
DX uptake and release delivery to an estimated model in perfect condition will be
discussed later.
It is noted that the effect of Vitamin E on uptake or release duration increase for
hydrophilic drugs as shown in Chapter 3 is much larger than that for DX with
comparable Vitamin E loading. For example, by comparing the hydrophilic drug release
and DX uptake experiment results, NIGHT&DAYTM with 27% Vitamin E loading has 76
times increase in timolol delivery time while it has only 8.8 times increase in DX even
though actual delivery time is longer for DX (142 hours) than for timolol (43 hours).
O2OPTIX™ with 34% Vitamin E loading also shows larger increase with 34.3-fold for
timolol while 15.5-fold for DX. Furthermore, while there is no significant difference in
drug delivery time for DX and dexamethasone 21-disodium phosphate (DXP) by pure
lens (For example, 10.5 hours and 14 hours by ACUVUE® OASYS™ , respectively), the
88
Vitamin E loaded lenses deliver DXP for longer duration compared to DX. With about
27% Vitamin E loading, NIGHT&DAY™ shows 40-fold increase in release time for DXP
which is about 12 days, and only 8.8-fold delivery time (4.5 days) for DX. These results
also support the theory for Vitamin E aggregates inside lens serving as diffusion barriers.
Since timolol and DXP are hydrophilic ionic drugs, it cannot diffuse through the highly
hydrophobic Vitamin E particles while the hydrophobic DX can partition and diffuse
through Vitamin E. The reduction in release rates for hydrophilic drugs is thus likely due
to presence of Vitamin E particles that act as diffusion barriers which create an
extended tortuous diffusion path. For DX, while it can diffuse through the Vitamin E
barrier, the diffusivity may be reduced because of increased viscosity and/or altered
adsorption to the polymer, and this reduction in diffusivity of DX through the Vitamin E
barrier could lead to the reduction in drug uptake and release rates. This will be
discussed further in the section on model development.
To understand the mechanism of transport of hydrophobic drugs through the
Vitamin E laden lenses, it is instructive to determine the partition coefficient of the drugs
both in the pure gel without Vitamin E and in the lenses with various Vitamin E loadings.
These data can then be used to obtain the partition coefficient of the drug in the Vitamin
E aggregates later.
For loading experiment, the partition coefficient of drug in the Vitamin E loaded
lens (K) was defined as
w,fl
w,fw,iw
w,f
fl
CV
)C(CV
C
CK
, (4-1)
where Vw and Vl are the volumes of the drug-PBS solution and the dry lens (either with
or without Vitamin E loading), respectively, and Cl,f, Cw,i and Cw,f are the equilibrium
89
concentrations of the drug in the lens phase, and the initial and equilibrium
concentrations in the aqueous phase, respectively, in the loading experiment. Partition
coefficient of drug in the pure lens (Kpl) can be also written as
w,fpl
w,fw,iw
w,f
fpl
plCV
)C(CV
C
CK
, (4-2)
where Vpl and Cpl,f are the volume of the dry pure lens and the equilibrium concentration
of the drug in the pure lens phase, respectively. The mass balance of drug in the vial
yields
wfwvefwveplfwplwfwvefveplfpli VCVCKVCKVCVCVCM ,,,,,, (4-3)
where Mi is total mass of drug in the vial and Cve,f is the equilibrium concentration of the
drug in the Vitamin E aggregates. Vve is the volume of Vitamin E aggregates in the lens
and is calculated by )( * lV , where is the volume ratio of Vitamin E in the dry lens
and is the Vitamin E loading the could either existing in the form that bounds to the
polymer gel or as particles but in regions of the gel that do not contribute to drug
transport, which we have obtained previously in Chapter 3. Partition coefficient of drug
in Vitamin E phase (Kve) can be obtained as
)(
/*
,,
l
wplplfwi
w,f
fve
veV
VVKCM
C
CK (4-4)
The values of K and Kve are listed in Table 4-1. K and Kve are comparable for DX,
which is due to the hydrophobic nature of the drug and Vitamin E. These partition
coefficient values will be utilized in the model presented below.
90
4.2.3 Diffusivities of Drugs in Vitamin E Loaded Lenses
The thickness of each commercial contact lens varies in the radial direction and
depends on the base curve, but the average thickness is about 80-100 m, which is
much smaller compared to the diameter of lens (about 14 mm). Thus, the drug delivery
by contact lens can be considered as a one-dimensional diffusion transport. To confirm
whether the DX uptake and release by Vitamin E loaded lenses are controlled by one-
dimensional diffusion as expected, the drug release profiles can be plotted as
percentage of drug release versus square root of time. For diffusion-controlled transport,
the percentage of drug release will be linear to the square root of time, and the results
are shown in Figure 4-3. The lines in the figure are the best fit straight line to short time
release data. The fits are all good with R2 values larger than 0.99 showing that the drug
transport in these lenses is diffusion controlled.
Below we develop a model based on the one-dimensional diffusion equation to fit
the experiment results and obtain the diffusion coefficient of DX in the lenses. Due to
the large aspect ratio, we assume that the geometry of contact lens can be modeled as
a flat thin film with homogenous thickness 80 m, which is the typical average thickness
of commercial contact lens. The thickness variation in the radial direction can easily be
integrated into the model but is not presented here for simplicity. If the drug diffusivity
(D) and partition coefficient (K) are independent of the drug concentration, the drug
transport to the transverse y-direction can be described as
2
2
y
CD
t
C gg
(4-5)
where Cg is the drug concentration in the lens gel matrix. The boundary conditions for
the drug release experiment are
91
wg
g
KChytC
yty
C
),(
0)0,( (4-6)
where h is the half-thickness of the gel, which is about 40m for pure contact lens
without Vitamin E loading. The half-thickness is adjusted with Vitamin E loading amount
by isotropic expansion assumption. The first boundary condition assumes symmetry at
the center of the gel and the second describes equilibrium of DX concentration between
the gel and the aqueous phase. A mass balance on the aqueous reservoir in the beaker
yields
hy
g
gw
wy
CDA
dt
dCV
2 (4-7)
where Vw is the water volume in the beaker and Ag is the cross-sectional area of the
lens.
In addition, the initial conditions for the DX delivery are
iww
igg
CtC
CtyC
,
,
)0(
)0,(
(4-8)
For uptake, Cg,i is zero and Cw,i is the initial concentration of DX solution for
loading (0.08 mg/mL). For release, Cg,i is the final equilibrium DX concentration in the
lens after drug uptake process, and Cw,i is zero. The equations were solved by finite
difference method with MATLAB® , and the fitted D and K are determined by fitting the
model with experimental results by using the function of „fminsearch‟ in MATLAB® . The
fitting results were shown in Figure 4-2 as solid lines. The good fits between the
experiment and model results suggest the validity of our proposed model. It is noted
that the fits are better for the uptake profiles compared to the release profiles,
92
particularly in the long time period, where the observed DX release amount are less
than predicted value. This is very likely caused by the accumulated drug loss during the
experiments, which also explains the observation that the experimental partition
coefficients for all lenses explored in this study are larger for release than those for
uptake. Therefore, the diffusivity values fitted to the uptake data are expected to be
more reliable than those from release. Figure 4-4 shows the fitted D and K for
ACUVUE® OASYSTM, NIGHT&DAYTM and O2OPTIXTM with different Vitamin E loading.
For all lenses, while the diffusivity decreases significantly as the amount of Vitamin E in
the lens increases, the drug partition coefficient almost remains the same regardless of
the Vitamin E loading. The results suggest that while Vitamin E has a similar partition
coefficient to the lens gel matrix, the DX diffusivity for Vitamin E is much smaller than
that for the lens matrix, likely due to the high viscosity of Vitamin E.
4.2.4 Scaling Model for Effect of Vitamin E Loading on Extended DX Delivery
The scaling model proposed in Chapter 3 for hydrophilic drugs delivery by Vitamin
E loaded silicone hydrogel contact lens is likely not valid for hydrophobic drugs that can
partition into the Vitamin E phase. For these hydrophobic drugs the transport occurs
partially by diffusion around the Vitamin E aggregates and partially by dissolution and
diffusion through these aggregates. Accordingly, the increase in release time is much
larger for hydrophilic drugs such as timolol compared to hydrophobic drugs such as DX.
The hydrophobic drugs can partition into the Vitamin E aggregates, diffuse through
these, and then diffuse into the gel matrix. Thus the transport of hydrophobic drugs
through the Vitamin E-laden gels can be considered as diffusion through regions of the
gel matrix and regions of Vitamin E arranged in series. Since the diffusivity of DX is
much smaller for Vitamin E than that for the gel matrix, the drug transport time will be
93
determined mainly by the diffusion through the Vitamin E region when the Vitamin E
loading amount increases.
For one-dimensional drug diffusion in a pure lens without Vitamin E loading with
average thickness h, the drug transport duration 0 can be estimated as h2/DG, where
DG is the drug diffusivity in the gel matrix. For Vitamin E loaded contact lens, the time it
takes for the drug diffuse through the Vitamin E aggregates region can be scaled as
(h(-*))2/DV, where DV is drug diffusivity in the Vitamin E aggregates. Thus, the ratio of
the transport time increase by Vitamin E loaded lenses (/0) is given by the following
expression:
1)( 2*
0
V
G
D
D (4-9)
The values of* were obtained by fitting the drug transport data for the
hydrophilic drugs, which is 0.0117, 0.0621, and 0.0973 for ACUVUE® OASYSTM,
NIGHT&DAYTM and O2OPTIXTM, respectively. The only unknown parameter DG/DV can
then be obtained by fitting the experimental data to the above equation. The fitting
results for DX uptake duration increase by Vitamin E loaded commercial lenses are
shown in Figure 4-5 and the fitted DG/DV values are 330 for ACUVUE® OASYSTM, 395
for NIGHT&DAYTM and 405 for O2OPTIXTM, respectively. The fitted results are satisfied
with the assumption in our model that DG >>DV.
The reduced diffusivity of DX through the Vitamin E barrier is likely due to the high
viscosity of Vitamin E. The diffusivity is inversely related to the viscosity and thus the
ratio DG/DV may be related to the ratio of the viscosity of Vitamin E and water. To test
this speculation, the dynamic viscosity of Vitamin E was measured by cone and plate
94
rheometer. The slope of the log-log plot of loss modulus (G") versus the angular
frequency is one, as shown in Figure 4-6, suggesting that Vitamin E can be
characterized as a Newtonian fluid. The measured viscosity of Vitamin E is 1.918 Pa∙s,
which is about 2100-fold to water at 25 oC (0.89 mPa∙s). The ratio of diffusivity is about
20% of the viscosity ratio, which is encouraging. The differences between the diffusivity
and the viscosity ratios could perhaps be attributed to channeling of drug through
specific paths, viz. silicone rich hydrophobic channel, and thus a fraction of the Vitamin
E loaded in the gel may not function as a barrier. If one assumes that only about 50%
of the precipitated Vitamin E acts as barriers, the ratio of DG/DV obtained by fitting the
data will increase to 4 times the values reported above bringing it in reasonable
agreement with the viscosity ratio.
In this chapter we show that the drug delivery duration for DX from contact
lenses can be significantly increased to more than a week by incorporation of Vitamin E
into the contact lenses. The mechanism for the extended release is likely related to the
reduced diffusivity of DX through the Vitamin E barriers due to its high viscosity. A
mathematical model based on diffusion controlled transport fits the uptake and release
profiles from the Vitamin E loaded lenses well showing that the transport is diffusion
controlled, and a scaling model fits the dependence of effective diffusivity on the Vitamin
E loading.
While in vivo studies are necessary to explore the efficacy of Vitamin E loaded
lenses for ophthalmic drug delivery, the results of this study along with those from our
prior studies in Chapters 2 and 3 strongly suggest that Vitamin E loaded contact lenses
95
could be very useful vehicles for extended drug delivery of both hydrophobic and
hydrophilic drugs.
96
0
20
40
60
80
100
120
0 10 20 30 40 50
Time (hr)
Dru
g r
ele
ase,
M/M
f (%
)
DX-PBS for 24 hours (41.0 ug)
DX-PBS for 7 days (41.0 ug)
DX-ethanol for 3 hours (7.4 ug)
0
20
40
60
80
100
120
0 10 20 30 40 50
Time (hr)
Dru
g r
ele
ase,
M/M
f (%
)
DX-PBS for 24 hours (37.6 ug)
DX-PBS for 7 days (38.6 ug)
DX-ethanol for 3 hours (3.7 ug)
0
20
40
60
80
100
120
0 10 20 30 40 50
Time (hr)
Dru
g r
ele
ase,
M/M
f (%
)
DX-PBS for 24 hours (32.0 ug)
DX-PBS for 7 days (34.7 ug)
DX-ethanol for 3 hours (5.7 ug)
0
20
40
60
80
100
120
0 10 20 30 40 50
Time (hr)
Dru
g r
ele
ase,
M/M
f (%
)DX-PBS for 24 hours (29.2 ug)
DX-PBS for 7 days (27.7 ug)
DX-ethanol for 3 hours (3.2 ug)
0
20
40
60
80
100
120
0 10 20 30 40 50
Time (hr)
Dru
g r
ele
ase,
M/M
f (%
)
DX-PBS for 24 hours (37.4 ug)
DX-PBS for 7 days (41.3 ug)
DX-ethanol for 3 hours (2.6 ug)
0.00
0.05
0.10
0.15
0.20
0.25
0 10 20 30 40 50
Water content (%)
(DX
rele
ase t
ime)-1
(h
r-1)
ACUVUE® ADVANCE™ACUVUE® OASYS™NIGHT&DAY™O2OPTIX™PureVision™
Figure 4-1. Effect of DX loading method on profile of DX release by A) ACUVUE®
ADVANCETM B) ACUVUE® OASYSTM C) NIGHT&DAYTM D) O2OPTIXTM E) PureVision™ contact lenses. F) The plot of (DX release time)-1 versus water content of contact lenses. Drug release (M) divided by total amount released (Mf) are plotted as a function of time. DX was loaded by soaking the lens in 0.08 mg/mL of indicated medium for indicated duration of time. Total amount of drug released for each lens is marked in parenthesis on the legends.
A
B
C
E
F
D
97
0
20
40
60
80
100
120
0 200 400 600 800 1000
Time (hr)
Dru
g u
pta
ke
& r
ele
as
e (
g)
0 g vitamin E/g pure lens0.11 g vitamin E/g pure lens0.25 g vitamin E/g pure lens0.42 g vitamin E/g pure lens0.57 g vitamin E/g pure lens
0
20
40
60
80
100
120
0 200 400 600 800 1000
Time (hr)
Dru
g u
pta
ke
& r
ele
as
e (
g)
0 g vitamin E/g pure lens0.10 g vitamin E/g pure lens0.17 g vitamin E/g pure lens0.27 g vitamin E/g pure lens0.35 g vitamin E/g pure lens
Figure 4-2. Profiles of experimental and model fitted DX uptake and release by Vitamin
E loaded contact lenses A) ACUVUE® OASYS™ B) NIGHT&DAY™ C) O2OPTIX™ D) PureVision™. Experiment results are presented by solid and hollow markers for uptake and release, respectively, and model fitted results are presented in solid line. Vitamin E was loaded first by soaking pure contact lens in Vitamin E-ethanol solution and the lens was dried. And then DX was loaded by soaking the Vitamin E loaded lens in DX-PBS solution (0.08 mg/mL). The data are presented as mean ± S.D. with n = 3.
0
10
20
30
40
0 20 40 60 80 100
0
10
20
30
40
0 20 40 60 80 100
A
B
98
0
20
40
60
80
100
120
0 200 400 600 800
Time (hr)
Dru
g u
pta
ke
& r
ele
as
e (
g)
0 g vitamin E/g pure lens0.12 g vitamin E/g pure lens0.21 g vitamin E/g pure lens0.34 g vitamin E/g pure lens0.46 g vitamin E/g pure lens
0
20
40
60
80
100
120
140
0 50 100 150 200 250
Time (hr)
Dru
g u
pta
ke
& r
ele
as
e (
g)
0 g vitamin E/g pure lens0.13 g vitamin E/g pure lens0.39 g vitamin E/g pure lens
Figure 4-2. Continued.
0
10
20
30
40
0 20 40 60 80 100
D
C
0
10
20
30
40
0 24 48 72
99
0
10
20
30
40
50
60
0.0 5.0 10.0 15.0 20.0 25.0
Time0.5
(hr0.5
)
% D
rug
re
lea
se
0 g vitamin E/g pure lens0.11 g vitamin E/g pure lens0.25 g vitamin E/g pure lens0.42 g vitamin E/g pure lens0.57 g vitamin E/g pure lens
0
5
10
15
20
25
30
35
40
45
0.0 5.0 10.0 15.0 20.0 25.0
Time0.5
(hr0.5
)
% D
rug
re
lea
se
0 g vitamin E/g pure lens0.10 g vitamin E/g pure lens0.17 g vitamin E/g pure lens0.27 g vitamin E/g pure lens0.35 g vitamin E/g pure lens
Figure 4-3. Plot of % drug release by Vitamin E loaded lenses versus square root of time. The lines are the best fit straight for short time data of DX release by A) ACUVUE® OASYS™ , B) NIGHT&DAY™ , C) O2OPTIX™ and D) PureVision™. All R2‟s are larger than 0.99. Some of data are presented as mean ± S.D. with n = 3.
B
A
100
0
5
10
15
20
25
30
35
40
0.0 5.0 10.0 15.0 20.0
Time0.5
(hr0.5
)
% D
rug
re
lea
se
0 g vitamin E/g pure lens0.12 g vitamin E/g pure lens0.21g vitamin E/g pure lens0.34 g vitamin E/g pure lens0.46 g vitamin E/g pure lens
0
5
10
15
20
25
30
0.0 5.0 10.0 15.0
Time0.5
(hr0.5
)
% D
rug
re
lea
se
0 g vitamin E/g pure lens
0.13 g vitamin E/g pure lens
0.39 g vitamin E/g pure lens
Figure 4-3. Continued.
C
D
101
0.0E+00
1.0E-05
2.0E-05
3.0E-05
4.0E-05
5.0E-05
6.0E-05
7.0E-05
0.0 0.1 0.2 0.3 0.4 0.5
Fit
ted
DX
dif
fus
ivit
y, D
(m
m2/h
r)
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
0
20
40
60
80
100
120
140
160
0.0 0.1 0.2 0.3 0.4 0.5
Fit
ted
DX
pa
rtit
ion
co
eff
icie
nt,
K
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
Figure 4-4. Fitted DX diffusivity and partition coefficient for contact lenses with different
Vitamin E volume fraction ().
102
0
10
20
30
40
50
60
0.0 0.1 0.2 0.3 0.4 0.5
DX
up
tak
e t
ime
in
cre
as
e, /
0
ACUVUE® OASYS™
NIGHT&DAY™
O2OPTIX™
Figure 4-5. Effect of Vitamin E volume fraction () on increase in drug uptake times.
The solid lines are best fits to the data based on Equation 4-9.
103
Figure 4-6. Dependence of the loss modulus G" on frequency for pure Vitamin E (as supplied). The slope of the log-log plot of G” versus the angular frequency is one, suggesting that Vitamin E can be characterized as a Newtonian fluid.
The value of Viscosity () estimated from the linear fit of G" to frequency was 1.918 Pa.S.
104
Table 4-1. Partition coefficient (K) of DX in lenses soaked in DX-PBS solution.
Contact lenses
Vitamin E loading
[g Vitamin E /g pure lens]
K for loading
Kve for loading
K for release
Kve for release
ACUVUE® OASYS™
0 77.4 - 105.7 -
0.11 89.7 211.6 116.9 234.4
0.24 76.7 80.3 96.2 61.5
0.42 94.9 154.3 149.6 294.7
0.7 80.2 89.0 120.4 152.4
NIGHT&DAY™ 0 119.2 - 137.7 -
0.1 120.3 - 137.8 -
0.17 111.7 112.2 126.5 98.6
0.28 106.7 82.1 137.7 250.6
0.35 110.4 109.5 162.8 299.6
O2OPTIX™ 0 131.3 - 146.3 -
0.12 131 - 143.9 -
0.21 120.7 140.9 134.5 153.9
0.34 119.4 154.0 141.8 205.4
0.46 115.9 113.8 149.1 214.5
PureVision™ 0 290.7 - 326.5 -
0.13 159.6 - 241.7 -
0.39 181.3 - 203.8 -
105
CHAPTER 5 ANESTHETICS DELIVERY BY VITAMIN E LOADED SILICONE HYDROGEL
Excimer laser vision correction has been widely accepted in our daily life since the
first proceduce was approved by FDA in 1995, and now more than one million
procedures are performed annually in the United States [86]. Well-developed excimer
refractive surgery techniques for low to moderate refraction errors include myopia,
hyperopia and astigmatism, but currently laser in situ keratomileusis (LASIK) is the most
preferred surgery for the treatment, followed by photorefractive keratectomy (PRK)
which comprises a much smaller fraction [87-89]. LASIK is the procedure of choice for
most patients in the civilian community mainly because of the significantly less
postoperative discomfort, faster visual recovery, and maintenance of an intact
Bowman‟s membrane [90, 91]. However, the higher risk of several serious potential
complications associated with LASIK, including corneal flap loss, tear or striae, and
keratectasia limits its general applications [91-95]. For people with thin corneas, anterior
basement membrane dystrophy and significant dry eye [96, 97], PRK remains the
preferred procedure to LASIK. PRK is also preferred by active people who are subject
to trauma, such as those in the military or involved in contact sports, because the
potential problems with flap stability after LASIK could lead to flap dislocation with
trauma [98]. In the United States military healthcare system, PRK is the preferred
refractive surgical procedure, while LASIK has not been approved [99]. Thus, the
current focus on PRK research is to improve the postoperative pain control as well as
reduce the visual recovery time after surgery.
Patients who have PRK generally receive a bandage contact lens (BCL)
postoperatively. Several studies have shown that the BCL protects the deepithelialized
106
cornea, leads to a faster reepithelialization, and reduces pain [98-102]. Lenses are
generally worn for 4 to 5 days after surgery, though typically the corneal reepithelializes
in 2 to 4 days with BCL [98, 99]. Contact lens with higher oxygen permeability are
preferred because lens with low DK/t may lengthen postoperative healing time because
of the decreased oxygen exposure [98, 103]. After PRK is preformed, BCL is placed on
the treated eye, followed by post medication including antibiotics, anti-inflammatory,
lubricant eye drops, and oral and topical anesthetics. For example, the patient might
need to apply one drop every 2 hours as needed for as long as the first 72 hours of
topical nonpreserved 0.5% tetracaine hydrochloride to control the pain after PRK [96].
Reports indicate that pain starts approximately 3 hours after PRK and reaches its
maximum at about 7 hours, and usually is over about 24 hours following surgery [104,
105]. The frequent dosage requirements interfere with the patients daily activities and
can lead to the potential risk of drug overdose.
The aim of the study in this chapter is to in vitro investigate the potential of using
Vitamin E loaded contact lens for postoperative treatment after PRK to obtain better
pain control for patients. Three common topical anesthetic drugs are explored here,
including lidocaine, bupivacaine and tetracaine, and the molecular structures of these
drugs are shown in Figure 5-1. The pKa values are 7.4, 8.1 and 8.4 for lidocaine,
bupivacaine and tetracaine, respectively, and thus these three drugs present in ionized
forms at physiological pH. To clarify the drug transport mechanism inside these
composite hydrogel systems, we also prepared lab-synthesized silicone hydrogel to
obtain further understanding for future model prediction.
107
5.1 Materials and Methods
5.1.1 Materials
Commercial silicone hydrogel contact lenses O2OPTIX™ (Lotrafilcon B, diopter -
6.50) from Ciba Vision Corp. (Duluth, GA) were used in this study. Lidocaine
hydrochloride, bupivacaine hydrochloride, tetracaine hydrochloride, ethanol (>99.5%)
and Dulbecco‟s phosphate buffered saline (PBS) were purchased from Sigma-Aldrich
Chemicals (St. Louis, MO), and Vitamin E (D-alpha tocopherol, Covitol® F1370) were
kindly gifted by Cognis corporation (Kankakee, IL). For preparation of silicone hydrogel,
ethylene glycol dimethacrylate (EGDMA, 98%), N, N-Dimethylacrylamide (DMA, 99%)
and 1-vinyl-2 pyrrolidone (NVP, 99+ %) were purchased from Sigma-Aldrich Chemicals
(Milwaukee, WI). The macromer acryloxy(polyethyleneoxy)-propylether terminated
poly(dimethylsiloxane) (DBE-U12, 95+%) were purchased from Gelest Inc. (Morrisville,
PA). 3-Methacryloxypropyltris(trimethylsiloxy)silane (TRIS) was supplied by Silar
laboratories (Scotia, NY), and 2, 4, 6-trimethylbenzoyl-diphenyl-phosphineoxide
(Darocur® TPO) were kindly provided by Ciba Specialty Chemicals (Tarrytown, NY). All
chemicals in this study were reagent grade and used as supplied without further
purification.
5.1.2 Drug Loading into Pure Lenses
O2OPTIX™ lens was rinsed with DI water and then air-dried before further use.
Drugs were loaded into the lenses by soaking the lens in 3 mL of a drug-PBS solution
for at least 7 days until reaching equilibrium. The initial drug concentrations were 5, 2.5
and 1 mg/mL for lidocaine, bupivacaine and tetracaine, respectively. While soaking the
lens in either solution, the dynamic concentration in the solution was not monitored. At
the end of the loading stage the lens was taken out and excess drug solution was
108
blotted from the surface of the lens. The lens was then air-dried and subsequently used
for release experiments.
Since all the drugs explored in this study have high solubility in ethanol, these
drugs were also loaded into the contact lens by soaking the lens in a 3 mL of
drug/ethanol solution for 24 hours. The initial soaking drug concentrations were 10, 10
and 1 mg/mL for lidocaine, bupivacaine and tetracaine, respectively. At the end of the
loading stage the lens was taken out and excess drug solution was blotted from the
surface of the lens, and subsequently used for release experiments.
5.1.3 Vitamin E Loading into Pure Lenses
Vitamin E was loaded into contact lens by soaking a lens in 3 mL of a Vitamin E-
ethanol solution for 24 hours. After the loading step, the lens was withdrawn and blotted
to remove excess solution on the surface. The lens was then dried in air overnight. The
Vitamin E loading amount was determined by measuring the increase in lens weight.
The mass of Vitamin E loaded into a lens is directly proportional to the concentration of
Vitamin E loading in the ethanol-Vitamin E solution [106]. The concentrations of the
Vitamin E in the loading solution in this study were 0.05, 0.10 and 0.15g Vitamin E/g
ethanol, which leaded to about 0.18, 0.37 and 0.55 g Vitamin E/g pure lens, respectively.
Subsequently, these Vitamin E loaded contact lens were soaked into drug/PBS
solutions through the same drug loading procedure for pure lens, as described in
Section 5.2.2.
The drugs were also loaded into lens by directly adding drug into Vitamin
E/ethanol solution. Each drug was dissolved in a 3 mL of 0.05, 0.10 or 0.15g Vitamin
E/g ethanol solution for 24 hours, and the drug concentration was designed to be 10, 10
and 1 mg/mL for lidocaine, bupivacaine and tetracaine, respectively.
109
5.1.4 Drug Release Experiments
The drug release experiments were carried out by soaking drug-impregnated lens
in 2 mL of fresh PBS. Since all these three anesthetic drugs have high solubility in water
and the volume of aqueous medium is much larger than that of the hydrated contact
lens, the drug release can be thus viewed as drug transport at perfect sink condition.
Therefore, the amount of residue drug in the lens at final equilibrium is negligible, and
the initial drug loading is equal to the total amount of drug release. The dynamic drug
concentration in aqueous solution was determined by measuring the absorbance in the
wavelength range with a UV-VIS spectrophotometer (Thermospectronic Genesys 10
UV). UV–VIS absorbance was converted to the concentration of drug by following the
absorbance spectra deconvolution method reported previously to detect both drug and
potential Vitamin E release [107]. The absorbance was measured in from 231 to
291 nm for lidocaine and bupivacaine, and 195 to 255 nm for tetracaine. The
absorbance was measured in scanning range rather than at a single wavelength to
ensure that the experimental methods did not lead to drug degradation which will
manifest as changes in the absorption spectrum. In this study, the drug release
duration by the lens is defined as the time it takes to complete 90% of final drug release
amount at equilibrium.
5.1.5 Silicone Hydrogel Preparation
To prepare the silicone hydrogel, hydrophilic monomers with high ion permeability
are copolymerized along with the hydrophobic silicone monomer with high oxygen
permeability, and a macromer is needed in the monomer mixture to ensure
solubilization of all monomers. In this study, TRIS was used as the hydrophobic
monomer, DMA was the hydrophilic monomers, and DBE-U12 was selected as the
110
macromer. Highly hydrophilic NVP monomer was also added to increase water content
of the hydrogel and EGDMA was introduced in the monomer mixture for controlled
crosslinking. To prepare the polymerizing mixture, 2.4 mL of a mixture that comprises
0.8 mL TRIS and 0.8 mL macromer and 0.8 mL of the hydrophilic DMA/MAA mixture
was combined with 0.12 mL of NVP and 0.1 mL of EGDMA. After well mixed with
vortexing for few second, the mixture was purged with bubbling nitrogen for 15 minutes
to reduce the dissolved oxygen. To each monomer mixture, 12 mg of photoinitiator
Darocur® TPO was added with stirring for 5 minutes and the final mixture was
immediately injected into a mold which is composed of two 5-mm thick glass plates. The
plates were separated by a plastic spacer with various thicknesses. The mold was then
placed on ultraviolet transilluminator UVB-10 (UltraLum Inc.) and the gel mixture was
cured by irradiating with UVB light (305 nm) for 50 minutes. The synthesized hydrogel
was either cut into circular pieces (about 1.65 cm diameter) with a cork borer for
subsequent experiments. Prior to conducting further tests, the prepared hydrogel was
soaked in ethanol for 3 hours then dried at ambient temperature overnight to remove
the unreacted monomer within.
5.1.6 Partition Coefficient
The synthesized silicone hydrogel was used for investigation of the dependency of
lidocaine partition coefficient in silicone hydrogel on drug concentration. To load Vitamin
E into the hydrogel, each circular piece of hydrogel was soaked in a 5 mL of 0.35 g/mL
Vitamin E/ethanol for 24 hours, which resulted in a loading of 0.28 0.01 g Vitamin E/ g
pure gel. The hydrogel, with or without Vitamin E, was then soaked in 5 mL of
lidocaine/PBS solution with various concentration until reached equilibrium. The
111
equilibrium drug loading in the hydrogel was subsequently determined by conducting
drug release in fresh PBS. The volume of the release medium was adjusted with the
drug loading to assure the drug concentration in the aqueous medium maintained in
measurable range.
5.1.7 Determination of Critical Micelle Concentration (CMC) of Lidocaine
Surface tension isotherm of lidocaine was measured at room temperature by
creating a pendant drop of lidocaine/PBS solution against ambient atmosphere. The
drop shape was digitally imaged and then fitted to the Young-Laplace equation by using
the Drop Shape Analysis System DSA100 (KRÜ SS) to calculate the surface tension.
The concentration of lidocaine solution was varied from 0.01 mg/mL to 360 mg/mL.
5.2 Results and Discussion
5.2.1 Dynamics of Drug Release from Contact Lenses
5.2.1.1 Drug uptake through drug-PBS solution
The results of lidocaine release by O2OPTIXTM with various Vitamin E loadings
were shown in Figure 5-2. The drug release duration increases as the Vitamin E
loading increases. For instance, pure O2OPTIXTM released 90% of its initial drug loading
in 1.8 hours, while lenses with 27% and 36% of Vitamin E can extend the release
duration to 6.2 and 10.8 hours, respectively. In addition, Vitamin E inside the lens can
enhance the total lidocaine loading when soaked in drug/PBS solution. Since Vitamin E
is highly hydrophobic, when the Vitamin E loaded lens is equilibrium with drug/PBS
solution it generally does not affect hydrophilic drug loading, such as timolol,
dexamethasone phosphate, and fluconazole [106]. The pH of the lidocaine/PBS
solutions ranges from 6.0 to 7.4 based on the drug concentration in this study, which is
lower than the pKa of lidocaine. Thus, the majority of lidocaine should present in the
112
hydrophilic ionized form, which is highly unlikely to partition into the Vitamin E
aggregates. The increase of drug loading could be resulted from the surface adsorption
of lidocaine on the interface between Vitamin E and the gel matrix, and we will further
examine this assumption later.
Another support of the interaction between lidocaine and Vitamin E were found,
as we observed tiny amount of Vitamin E (less than 1% of the Vitamin loading inside the
lens) release into the aqueous reservoir during the drug release experiment, as shown
in Figure 5-3. Since no Vitamin E release were detected for other hydrophilic drugs by
the same system, it is reasonable to assume that the existence of lidocaine enhanced
the solubility of Vitamin E in PBS. In practical, the Vitamin release should not cause
significant difference on the performance of these Vitamin E loaded lens because first
the loss amount is negligible and second in practice use the drug concentration is much
lower than our experiment condition, which should lead to even lower Vitamin E loss.
Figure 5-4 exhibited the bupivacaine release by O2OPTIXTM. Similar to lidocaine,
the release duration of bupivacaine extended with the amount of Vitamin E loading in
the lens. Pure O2OPTIXTM released 90% of its initial drug loading in 3.2 hours, while
lenses with 27% and 36% of Vitamin E loading can release in 10.2 and 20.7 hours,
respectively. The drug uptake by the lens was also enhanced with Vitamin E, while
similar Vitamin E loss were detected during bupivacaine release as well (data not
shown). The results of tetracaine release by O2OPTIXTM with various Vitamin E loadings
were shown in Figure 5-5. Again, higher Vitamin E loading in the lens resulted in longer
drug release duration. Pure O2OPTIXTM released 90% of its initial drug loading in 2.4
hours, while lenses with 27% and 36% of Vitamin E loading can release the loaded
113
tetracaine in 13.9 and 22.6 hours, respectively. However, unlike lidocaine and
bupivacaine, the total drug uptake does not significantly increased as the Vitamin E
loading amount increased. This is possibly due to the higher light sensitivity of
tetracaine compared to lidocaine and bupivacaine, which leads to drug degradation
during release experiment. The degradation can be observed by the gradual drug loss
in long time, as shown in Figure 5-5B. It was also observed by the absorbance
spectrum change in the long time release (data not shown). Vitamin E release is not
discussed during tetracaine release since its absorbance spectrum is not overlapped
with that of tetracaine in our study.
Even though Vitamin E inside the lens can effectively extend the release duration
of these anesthetic drugs, the effect is not as significant as those on other model drugs.
For example, with ca. 0.36 g Vitamin E/g pure lens loading, the drug release time
increased to from 1.8 hour by pure lens to 6.2 hours, which is a 3.5-fold increase. With
similar amount of Vitamin E loading, the lens can release the hydrophilic drug timolol for
28 hours, a 40-fold increase compared to pure lens; it can also release the hydrophobic
drug dexamethasone for 150 hours, which is a 15-fold increase. For hydrophilic drugs,
such as timolol, the Vitamin E loading inside the lens act as diffusion barriers for drug
transport due to the negligible affinity between drug and Vitamin E aggregates; for
hydrophobic drugs, such as dexamethasone, the drug can freely partition into the highly
viscous Vitamin E aggregates. The fact that we observed both hydrophobic and
hydrophilic behavior of lidocaine in the Vitamin E loaded silicone hydrogel composite
system implies that lidocaine could act as a surfactant-like molecule in this system, and
thus the transport mechanism of these anesthetics cannot be simply considered as
114
either hydrophilic or hydrophobic drug alone, but have to be controlled by other
mechanism based on the unique interaction between Vitamin E and these anesthetics
drugs, which will be discussed later.
5.2.1.2 Drug uptake through drug-ethanol solution
The drug release results by lens which the drug uptake was through drug/ethanol
solution were shown in Figure 5-6. With same amount of Vitamin E in the lens, the drug
release duration significantly increased by drug/ethanol uptake compared to that by
drug/PBS uptake. For instance, for O2OPTIXTM with 27% of Vitamin E loading, the drug
release duration of lidocaine and bupivacaine are 70 and 32 hours, respectively, which
is much higher than 6.2 and 10.2 hours by drug/PBS uptake. In addition, the total drug
uptake amount by the lens is independent of Vitamin E loading, which is different than
what we observed in the drug/PBS uptake. On the other hand, similar characterizations
such as degradation of tetracaine (Figure 5-6C) and Vitamin E release from lidocaine
and bupivacaine (data not shown). The possible mechanisms that result in these
differences will be discussed later.
In this study we explore two different approaches to load drug into the contact lens.
Drug uptake through drug/PBS solutions requires longer time to reach equilibrium, and
it need two steps to load Vitamin E and drug separately. However, since the lens was
kept in the drug/PBS solution, this method eliminates the drug loss in the package
solution. Since the anesthetic drug release time can be increased to about 20 hours
with 35% Vitamin E through drug/PBS uptake, these lenses can be used to provide
sustained topical anesthetics release based on daily replacement, which still has its
practical potential as the postoperative pain usually felt by the patient within the first day
after PRK surgery. On the other hand, drug uptake through drug/ethanol requires much
115
less time and can directly load Vitamin E and drug in one step. However, these lens
need to be kept in a package condition for practical use, with could lead to drug loss
and reduce the effect of extension of drug release duration.
To explore the packaging effect on the lidocaine release by contact lenses,
O2OPTIXTM lens was first soaked in a 3 mL of 10 mg/mL lidocaine/ethanol-Vitamin E
solution for 24 hours, where the Vitamin E concentration is 0.1 g Vitamin E/mL ethanol,
and the drug loaded lens was subsequently removed into a 3 mL of 10 mg/mL
drug-PBS solution for 1 and 7 days prior to in vitro release. As shown in Figure 5-7,
while through drug/ethanol uptake the lens is able to release 90% of the included
lidocaine in 70 hours, the release duration decreases to 40 hours after soaking the lens
in drug/PBS solution for 1 day, and further reduced to about 7 hours after 7 days, which
is similar to the duration by lens through drug/PBS uptake.
5.2.2 Lidocaine Release Study (Surfactant Behavior)
The observed behavior of lidocaine transport in Vitamin E loaded contact lens
implied that the lidocaine molecular could act as a surfactant-like molecule between the
hydrophobic Vitamin E aggregation and hydrophilic regions inside the hydrogel. To
further verify this hypothesis, the surface tension isotherm of lidocaine hydrochloride in
PBS was measured, and the results were plotted against drug concentration in
logarithmic scale in Figure 5-8. The result clearly demonstrated that the lidocaine affects
the surface tension of the drop when the concentration is above 0.2 mg/mL, and the
CMC of lidocaine in PBS is above 360 mg/mL, which is much higher than the
concentrations explored in our loading-release studies.
The drug partition coefficient in a pure silicone hydrogel (Kgel) based on the mass
balance in the uptake experiment can be defined as:
116
)/(,,
,,
wgelfgeliw
fgel
w,f
fgel
gelVVCC
C
C
CK
(5-1)
Where Cgel,f and Cw,f is the final drug concentration in the hydrogel and in the aqueous
medium when the uptake experiment reached equilibrium, and the volume of hydrogel
and aqueous medium were notated as Vgel and Vw, respectively. The only unknown
parameter Cgel,f can be determined by the drug extract experiment at perfect sink.
Similarly, the overall apparent drug partition coefficient in a Vitamin E loaded silicone
hydrogel (Kve-gel) can be defined as:
)/(,,
,,
wgelvefgelveiw
fgelve
w,f
fgelve
gelveVVCC
C
C
CK
(5-2)
In addition, the drug uptake by the Vitamin E loaded silicone hydrogel is the
summation of the amount of drug uptake by pure hydrogel and by Vitamin E aggregates
in the gel, and thus the drug partition coefficient in Vitamin E aggregates can be
determined as:
ve
gelgelgelvegelve
w,f
fve
veV
VKVK
C
CK
, (5-3)
The results of Kgel, Kve and Kve-gel at different initial soaking lidocaine concentration
are shown in Figure 5-9. Since lidocaine is highly likely acting as a surfactant-like
molecule in our Vitamin E loaded silicone hydrogel system, it is reasonable to assume
that the majority of lidocaine uptake by Vitamin E aggregates inside the hydrogel is
through surface binding. According to general Langmuir adsorption model at equilibrium,
which relates the adsorbed surface concentration of the drug on the Vitamin E
aggregates (Γ) to the free drug concentration in the aqueous phase(C) by the following
equation
117
Ckk
Ck
add
ad
(5-4)
where Γ∞ is the surface concentration at the maximum packing on the surface, and
kad and kd is the rate constants for adsorption and desorption of the drug on the Vitamin
E surface, respectively. In addition, the previously obtained Kve can be related as:
S
VC
C
C
K ve
w
ve
1 (5-5)
where V and S are the volume and surface area of Vitamin E aggregates,
respectively. By Equation 5-4 and 5-5 we can derive the following relation:
S
VC
Sk
Vk
K ad
d
VE
1 (5-6)
Therefore, if the Langmuir adsorption model successful describes the interaction
between lidocaine and Vitamin E, the inverse of KVE should be linear to the bulk drug
concentration, and the parameter V/S∞ can be obtained as the value of the slope. The
good fitting of experimental results to this linear model, as shown in Figure 5-10, highly
supported the validity of our assumption, and the value of V/S∞ is determined as
0.0447 mL/mg.
5.2.3 Mechanisms of Extended Drug Release by Vitamin E Loaded Contact Lens
The anesthetic drugs are in charged form at or below physiological pH in our study,
and thus these drugs should have a negligible partitioning into the Vitamin E aggregates
inside the lens matrix. Since it is observed that these anesthetic drugs have strong
interfacial interaction at the surface of Vitamin E aggregates, it is reasonable to assume
that the Vitamin E aggregates inside the gel matrix act as diffusion barriers for the drug
transport. When encountering the Vitamin E aggregates, drugs in the gel matrix need to
118
take a detoured route around the surface of Vitamin E with specific surface diffusivity to
diffuse out of the lens. The increased path length of the tortuous path l should scale
as h)( * , where h is the half thickness of lens, and is the parameter that depends
on the microstructure, including particle size and aspect ratio, of the Vitamin E
aggregates distribution in the gel; is the volume ratio of Vitamin E in the dry gel, and
)( * is the fraction that is present as the Vitamin E particles. The fraction * is
assumed to be either existing as bound to the polymer gel or as particles but in regions
of the gel that do not contribute to drug transport.
For a diffusion-controlled release, the release duration from a pure lens can be
scaled as:
gel
oo
D
h2
(5-7)
where ho is the half thickness of pure lens, and Dgel is the diffusion coefficient of the
drug inside the pure gel matrix. Where Vitamin E was loaded into the lens, the gel
thickness h increases due to Vitamin E uptake, and by assuming isotropic expansion it
can be written as )3/1(0 hh . Thus, we can estimate the drug release duration from
Vitamin E loaded contact lens as
s
o
gel
o
D
h
D
h2*
2 )])(3
1([
(5-8)
where Ds is the surface diffusivity of drug on the interface between Vitamin E
aggregates and the gel matrix. Due to the surfactant-like behavior of these anesthetics,
we expected that Ds should be higher than Dgel. By combining Equation 5-7 and 5-8, we
can obtain:
119
2*22 )()3
1(1
s
gel
o D
D (5-9)
From our previous studies in Chapter 3, for O2OPTIXTM, is about 35 and * is
0.0937. Equation 5-9 were fitted to the experimental data by using the function
„fminsearch‟ in MATLAB® to obtain the parameter Dgel/Ds for each drug. The model
fitting results were shown in Figure 5-11, and the fitted Dgel/Ds are 0.0291, 0.0319 and
0.0460 for lidocaine, bupivacaine and tetracaine, respectively.
Earlier studies have shown that the release behavior of most drugs by the contact
lens, with or without Vitamin E, is independent on the drug loading approaches [106,
108]. However, the drug release duration and the drug loading capacity of these three
anesthetic drugs by Vitamin E loaded lens through drug/ethanol uptake are significantly
different than those through drug/PBS uptake. When lidocaine is loaded into the lens
through drug/ethanol solution, since lidocaine should be in uncharged form in ethanol, it
is reasonable to assume that the uncharged lidocaine is mostly contained in the
hydrophobic regions of the silicone hydrogel matrix after ethanol evaporation, and the
drug loading capacity is simply determined by the equilibrium ethanol uptake of the lens.
When Vitamin E is also included in the drug/ethanol solution, while the total loading
capacity is not affected, the uncharged lidocaine loading now should distribute in both
the hydrophobic silicone region of the gel matrix and the Vitamin E aggregates based
on the respective partition coefficient. Since the uncharged lidocaine will changed to
charged form once it encounter the PBS during drug release which can be explained by
Equation 5-8, the extra resistance of lidocaine transport observed form the lenses
through drug/ethanol uptake should be arose from the drug diffusion inside the
hydrophobic regions inside the Vitamin E/hydrogel matrix.
120
Therefore, the effect on lidocaine release extension by loading Vitamin E and
lidocaine simultaneously could results from two possible mechanisms. First, the majority
of drug could partition into the hydrophobic region of silicone hydrogel matrix, and the
Vitamin E aggregates on the boundary between hydrophobic and hydrophilic region
serves as diffusion barriers to hinder the drug release from hydrophobic region to
hydrophilic region. In this case, the additional drug release time can be scaled as
siliconeDh /))('( 2* , where and Dsilicone are the respective aspect ratio parameter
and diffusion coefficient for lidocaine diffusion around the Vitamin E aggregates from
silicone-rich hydrophobic region in the gel matrix. Another alternative hypothesis is that
the additional drug release duration is controlled by the drug partitioned in the Vitamin E
aggregates with relatively smaller diffusivity compared to that of hydrogel matrix. In
other words, Vitamin E works as a drug reservoir that provides sustained release in this
case, and the one-dimensional diffusion inside Vitamin E aggregate can be scaled as
EVitaDr min
2 / , where r is the effective radius of the Vitamin E aggregates. Thus, when
drug loaded gel is soaked in fresh PBS, the overall lidocaine transport is the
combination of the drug distribution between gel matrix and Vitamin E, the
thermodynamic equilibrium between charged and uncharged form of lidocaine at the
interface between Vitamin E and the hydrated hydrogel, and the drug diffusion inside
the Vitamin E aggregates.
To further examine the above assumptions on the mechanism of lidocaine
transport by Vitamin E loaded silicone hydrogel, the lab-synthesized silicone hydrogel
with different thickness were used to conduct the drug uptake/release experiments. As
shown in Figure 5-12, the synthesized silicone hydrogel (with and without Vitamin E)
121
demonstrated similar lidocaine transport behaviors as those by commercial O2OPTIXTM.
While the 0.2 mm-thick pure gel released 90% of the loaded lidocaine in about 5 hours,
the hydrogel with 0.25 g Vitamin E/g pure gel loading can extend the release duration to
15 and 48 hours through drug/PBS and drug/ethanol uptake, respectively. The loading
capacity increased with Vitamin E loading through drug/PBS uptake, but kept the same
through drug /ethanol uptake, which is the same as we observed from O2OPTIXTM.
Figure 5-13 presented the lidocaine release results by silicone hydrogel with
various thicknesses, of which the drugs were loaded through drug/PBS uptake. The
drug release time increased when the Vitamin E loading or the gel thickness increased.
If Equation 5-8 holds to explain the lidocaine release transport here, than the kinetic
release time should be proportion to the square of gel thickness; i.e., if we define a
scaled time as time/(gel thickness/0.1 mm)2, then the release results from gels with
different thickness should be overlapped when plotted against the scaled time if the
drug transport is controlled by one-dimensional diffusion, and the results were shown in
Figure 5-13B. For both the gels with or without Vitamin E loading, the % releases
overlapped with different thicknesses, which support the validity of our proposed
transport mechanisms.
For the lidocaine release by gel through drug/ethanol uptake, if the additional drug
release resistance comes from the detour around Vitamin E aggregates from
hydrophobic regions to hydrophilic regions of hydrogel matrix, the drug release time
should be still proportional to the square of thickness when the Vitamin E loading inside
the gel kept the same. However, if the drug release was mainly controlled by the drug
diffusion in the Vitamin E aggregates, it is only affected by the size of Vitamin E
122
aggregates in the gel matrix instead of overall gel thickness. As shown in Figure 5-14,
the significant difference between the scaled releases of Vitamin E loaded gel with
different thicknesses suggested that the latter should be the more appropriate
assumption to describe the lidocaine transport by Vitamin E loaded silicone hydrogel
through drug/ethanol uptake.
In this study we investigate the potential to provide extended anesthetics delivery
by Vitamin E loaded silicone hydrogel contact lenses for postoperative pain control,
especially for patients who accepted PRK for vision correction. The thermodynamic
properties of these anesthetic drugs, including the amphiphilic behaviors and the
dependency on the pH of environment, significantly affects the mechanisms of drug
transport by Vitamin E loaded silicone hydrogel contact lenses at different drug loading
conditions. The Vitamin E loaded silicone contact lens can provide continuous
anesthetics release for about 1 day through drug/PBS uptake. The release duration by
lenses through drug/Vitamin E-ethanol uptake can be further increased, while the
packaging effect needs to be overcome for future practical use. Future in vivo studies
are also needed to further evaluate the feasibility of sustained anesthetic release by
contact lenses. Furthermore, the potential complexity of drug interactions between the
anesthetics and other ophthalmic drugs during the involved in postoperative treatment
should be taken into consideration in the future work.
123
Bupivacaine
Lidocaine
Tetracaine
Figure 5-1. Molecular structures of model drugs.
124
0
100
200
300
400
500
600
0 20 40 60 80
Time (hour)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/g pure lens
0.18 g Vitamin E/g pure lens
0.37 g Vitamin E/g pure lens
0.55 g Vitamin E/g pure lens
Figure 5-2. Lidocaine release in PBS by O2OPTIXTM with various Vitamin E loading.
125
0
10
20
30
40
50
60
70
80
0 20 40 60 80 100 120 140
Time (hour)
Vit
am
in E
re
lea
se
( g
)
0.18 g Vitamin E/g pure lens
0.37 g Vitamin E/g pure lens
0.55 g Vitamin E/g pure lens
Figure 5-3. Vitamin E release from O2OPTIXTM during lidocaine release in PBS.
126
0
100
200
300
400
500
600
700
800
900
1000
0 20 40 60 80 100
Time (hour)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/ g pure lens
0.36 g Vitamin E/g pure lens
0.55 g Vitamin E/g pure lens
Figure 5-4. Bupivacaine release in PBS by O2OPTIXTM with various Vitamin E loading.
127
A
0
10
20
30
40
50
60
70
0 20 40 60
Time (hour)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/g pure lens
0.17 g Vitamin E/g pure lens
0.37 g Vitamin E/g pure lens
0.55 g Vitamin E/g pure lens
B
0
10
20
30
40
50
60
70
0 100 200 300 400
Time (hour)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/g pure lens
0.17 g Vitamin E/g pure lens
0.37 g Vitamin E/g pure lens
0.55 g Vitamin E/g pure lens
Figure 5-5. A) Short time and B) long time tetracaine release in PBS by O2OPTIXTM with various Vitamin E loadings.
128
0
100
200
300
400
500
600
700
800
0 200 400 600
Time (hour)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/g pure lens
0.39 g Vitamin E/g pure lens
0.58 g Vitamin E/g pure lens
0
100
200
300
400
500
600
700
800
0 200 400 600
Time (hour)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/g pure lens
0.36 g Vitamin E/g pure lens
0.55 g Vitamin E/g pure lens
Figure 5-6. A) Lidocaine B) bupivacaine and C) tetracaine release in PBS by O2OPTIXTM with various Vitamin E loading. Drugs were loaded by soaking in drug/ethanol-Vitamin E solution for 24 hours.
A
B
129
0
10
20
30
40
50
60
70
80
0 200 400 600
Time (hour)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/g pure lens0.17 g Vitamin E/g pure lens0.36 g Vitamin E/g pure lens0.55 g Vitamin E/g pure lens
Figure 5-6. Continued.
C
130
0
20
40
60
80
100
120
0 20 40 60 80 100
Time (hour)
% D
rug
re
lea
se
Ethanol
PBS
Ethanol + PBS (1 day)
Ethanol + PBS (7 days)
Figure 5-7. Lidocaine release by O2OPTIXTM with 0.36g Vitamin E/g pure lens. Lidocaine was loaded into the lens by soaking in 10 mg/mL drug/ethanol for 24 hours (Ethanol) and subsequently soaked in 10 mg/mL drug/PBS for 1 day (Ethanol +PBS (1 day)) or 7 days (Ethanol +PBS (7 days) prior to release in 2 mL PBS. Drug was also loaded by soaking the lens in 10 mg/mL drug/PBS solution for 7 days.
131
40
45
50
55
60
65
70
75
0.01 0.1 1 10 100 1000
Lidocaine/PBS (mg/ml)
Su
rfa
ce
Te
ns
ion
(m
N/m
)
Figure 5-8. The relationship between surface tension and lidocaine concentration in
PBS.
132
0
1
2
3
4
5
6
7
8
9
10
0.1 1 10 100 1000
Lidocaine HCl Concentration (mg/ml)
Pa
titi
on
Co
eff
icie
nt,
K
Pure hydrogel (Kgel)
Vitamin E loaded hydrogel (Kve-gel)
Vitamin E aggregates (Kve)
Figure 5-9. The calculated partition coefficient (K) of Vitamin E loaded silicone hydrogel
at various lidocaine hydrochloride concentrations.
133
y = 0.0447x + 0.1798
R2 = 0.9946
0
0.5
1
1.5
2
2.5
3
0 20 40 60
Lidocaine HCl Concentration (mg/ml)
1/K
ve
Figure 5-10. The relationship between the lidocaine partition coefficient in Vitamin E
(KVE) and the bulk drug concentration.
134
0
2
4
6
8
10
12
14
0 0.1 0.2 0.3 0.4 0.5
Vitamin E volume fraction,
Dru
g r
ele
as
e d
ura
tio
n i
nc
rea
se
( /
o)
Lidocaine
Bupivacaine
Tetracaine
Figure 5-11. Model fitting for anesthetic drug release increase ratio on Vitamin E loading
fraction in the silicone hydrogel.
135
0
5
10
15
20
25
0 50 100 150
Time (hour)
Dru
g r
ele
as
e (
mg
/g p
ure
ge
l)Pure-EtOH
VE-EtOH
Pure-PBS
VE-PBS
0
20
40
60
80
100
120
0 20 40 60 80 100
Time (hour)
% D
rug
re
lea
se
Pure-EtOH
VE-EtOH
Pure-PBS
VE-PBS
Figure 5-12. Lidocaine release from pure 0.2 mm-thick silicone hydrogel or gel with
Vitamin E loading (0.25 g Vitamin E/g pure gel). Lidocaine were loaded into hydrogels through drug/PBS solution (10 mg/mL) or drug/(ethanol+Vitamin E) solution (10 mg/mL).
136
0
5
10
15
20
25
0 20 40 60 80
Time (hour)
Dru
g r
ele
as
e (
mg
/g p
ure
ge
l)
Pure_0.2 mm
VE_0.2 mm
Pure_0.3 mm
Pure_0.4 mm
VE_0.3 mm
0
20
40
60
80
100
120
0 2 4 6 8
Time/(thickness/0.1 mm)2 (hour)
% D
rug
re
lea
se
Pure_0.2 mm
VE_0.2 mm
Pure_0.3 mmPure_0.4 mm
VE_0.3 mm
Figure 5-13. Lidocaine release by silicone hydrogel with or without Vitamin E loading
(0.25 g Vitamin E/g pure gel) with various thickness. Lidocaine was loaded into the sample through soaking the gel into 10 mg/mL drug/PBS solution.
137
0
0.5
1
1.5
2
2.5
3
3.5
4
4.5
5
0 100 200 300
Time (hour)
Dru
g r
ele
as
e (
mg
/g p
ure
ge
l)
Pure_0.2 mm
VE_0.2 mm
Pure_0.4 mm
VE_0.4 mm
0
20
40
60
80
100
120
0 4 8 12 16
Time/(thickness/0.1 mm)2 (hour)
% D
rug
re
lea
se
Pure_0.2 mm
VE_0.2 mm
Pure_0.4 mm
VE_0.4 mm
Figure 5-14. Lidocaine release by silicone hydrogel (with or without Vitamin E loading
(0.25 g Vitamin E/g pure gel)) with various thicknesses. Lidocaine was loaded into the sample through soaking the gel into 10 mg/mL drug/(ethanol + Vitamin E) solution.
138
CHAPTER 6 ION TRANSPORT OF SILICONE HYDROGEL
Contact lenses for correcting vision are available in several different wear-
modalities including extended continuous wear for a period of 1-4 weeks, depending on
the type of the lens. Extended wear lenses are required to allow rapid oxygen transfer
because cornea is an avascular organ, and so it gets its oxygen supply directly from air
[22, 23]. Silicone based materials were explored as potential candidates for achieving
the high oxygen transport due to the very high oxygen solubility in these materials.
However, silicone based contact lenses were not useful as contact lens materials
because the lenses made of such materials adhered to the cornea. It may be
speculated that the hydrophobic nature of the silicone lenses leads to the adherence. It
was however determined that surface treatment of a lens to render the surface
hydrophilic is not enough to prevent adherence to cornea. If a particular lens material
did not move on the eye, application to the surface of contact lens did not change the
outcome significantly [19]. It was later discovered that the extended wear contact
lenses are required to also allow sufficient ion transport to maintain on-eye movement
and not adhere to the cornea. The importance of ion permeability of contact lens
material for maintaining lens motion was first described by Domschke et al., in a 1997
ACS presentation [77]. The need of ion transport through the lens is attributed to the
requirement of a fluid hydrodynamic boundary layer between the lens and the cornea
[109]. In the absence of the fluid layer, the lens can adhere to the cornea. Nicolson et al.
reported in a US patent to claim the ionoflux diffusion coefficient (Dion), i.e. the ion
permeability of the silicone hydrogel material, should be at least larger than 1.5 10-6
mm2/min for sufficient on-eye movement of lens [25]. While ion permeability is critical to
139
lens motion, an increase in the permeability beyond a critical value does not lead to a
further increase in on eye movement of the lens [19].
To compensate the drawback of the negligible water content from pure silicone,
hydrophilic monomers are copolymerized along with the hydrophobic silioxane
monomer. In general, the polysiloxanes and the hydrophilic polymers are immiscible
and thus a proper macromer is needed in the monomer mixture to ensure solubilization
of all monomers. The final silicone hydrogel matrix can be best described as isotropic
structures, either as a dispersed system with only one continuous phase, or as an
interpenetrating polymer network that both the hydrophilic and hydrophobic phases are
continuous throughout the material. A number of the commercial extended wear contact
lenses are reported to possess a bicontinuous microstructure that facilitates rapid
exchange of oxygen through the silicone rich phase and ions through the hydrophilic
phase [19]. Thus, fine tuning of each composition in the hydrogel mixture with proper
microstructure is critical for ensure the balance of all the key properties of the extended
wear contact lenses, including oxygen permeability, ion permeability, water content,
elastic modulus, surface wettability, etc.
While the importance of ion permeability has been known for extended wear
contact lenses, surprisingly there are a very limited number of studies focusing on ion
transport through these silicone hydrogels. Most prior studies on ion transport in contact
lenses have merely reported the ion permeability, i.e., the product of the diffusivity and
the partition coefficient, by direct permeation approach in a diffusion cell, which is a
better mimic to the real physiological environment [19, 25, 41, 84, 106, 110]. Ion
permeability was established by measurement of the flux of ions from the donor
140
reservoir, across the lens and into the receiver. The ion permeability could then be
calculated by applying Fick‟s law after a pseudo steady state is achieved. While ion
permeability is important because it directly determines the net ion flux across the
lenses at pseudo steady state, independent measurements of the diffusivity and
partition coefficient are important as well because these two relate to different aspects
of the lenses. Since partition coefficient is a thermodynamic property, it likely depends
only on the total fraction of the two phases (silicone and hydrophilic) in the hydrogel
matrix, while the diffusivity depends strongly on the connectivity of these phases
(dispersed or bicontinuous morphology). Thus, a further scrutinized study on the salt
transport through these silicone hydrogel materials is needed.
The diffusion of solutes in conventional hydrogels has been widely studied due to
the interest of its wide application, including separation process such as
chromatography [111], water purification [112-114], and hemodialysis as an artificial
kidney operation [115, 116]. Salt transport in hydrogels is typically explored through
direct membrane permeation and kinetic sorption/desorption methods. In the direct
membrane permeation study, the permeability of a solute is calculated based on the
measured permeation flux [112, 117-123]. In the sorption/desorption experiments, the
overall sorption or desorption kinetics are analyzed to characterize the diffusion
coefficients of the solutes in the hydrogel [112, 113, 119-121, 123, 124]. Both models
rely on ideal Fick‟s law of diffusion and the solution-diffusion model to extract the
transport parameters, and thus is suitable for homogeneous hydrogel systems where
the transport of solute is merely controlled by self-diffusion [125]. In both simplified
model only the salt release in the initial short time period (kinetic sorption/desorption) or
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the pseudo steady state ion flux (direct permeation) were used to obtain the transport
parameters. Since silicone hydrogel could have more complicated microstructure then
these conventional hydrogels, the validity of these models on silicone hydrogel systems
needed to be verified.
There are a number of models describing the diffusion of solutes in hydrogels, and
an overview of these models and experimental data was earlier made by Amsden [126].
Since the solute transport occurs primarily within the water-filled regions in the space
bounded by the polymer chains, any factor which affects the size of these spaces will
have an effect on the transport of the solute through the hydrogel matrix, including the
relative size of solute and of the openings between polymer chains, polymer chain
mobility, and the existence of charged groups on the polymer which may have strong
interaction with the solute molecule [126]. One of the most popularly used model to
explain sodium chloride transport in hydrated homogenous hydrogels was first proposed
by Yasuda et al. [120]. The model is based on the free volume theory by Cohen and
Turnbull which explains the process of solute diffusion in a pure liquid composed of hard
molecular spheres [127]. Their derivation is based on the concept that molecular
transport occurs by the movement of molecules into voids, with a size greater than
some critical value, formed by redistribution of the “free volume” vf. In such a system the
solute molecules move with the gas kinetic velocity u but most of the time is confined to
a cage delineated by their immediate neighbors. Occasionally, there is a fluctuation in
density which opens up a hole within a cage large enough to permit a considerable
displacement of the molecule contained by it. Such a displacement leads to diffusive
motion only if another molecule jumps into the hole before the first can return to its
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original position. Therefore, by calculating the statistical redistribution of the free volume,
Cohen and Turnbull derived the relation between the diffusion constant D in a liquid of
hard spheres and the free volume as D=Aexp(-v*/vf), where v* is the minimum required
volume of the void; A is some constants that related to the average thermal velocity and
the solute diameter, and is a numerical factor used to correct for overlap of free
volume available to more than one molecule [126, 127]. Yasuda et al. incorporated the
free volume theory to hydrated hydrogel systems and derived that the salt diffusion
coefficient D in the hydrogel should be exponentially proportional to the reciprocal water
fraction of the hydrated hydrogel. While this free volume theory model has successfully
described the sodium chloride transport in a variety of homogeneous hydrogels [112,
113, 117-120], to our knowledge, currently the salt transport in the silicone hydrogels as
extended-wear contact lens materials has not been scrutinized with similar rationale.
In this study, we use both permeation measurement and kinetic
sorption/desorption approach to explore the sodium chloride transport in the silicone
hydrogel contact lens materials. To confirm the mechanism of salt transport in the
hydrogel matrix, in addition to direct measure ion permeability with ion flux at pseudo
steady state, we propose to model the early transients in the permeation measurement
data to determine both diffusivity and the partition coefficient of the lens as well.
Additionally, we propose to utilize the kinetics sorption/desorption approach to extract
the diffusivity and the partition coefficient by loading the lens with salt by soaking and
then measuring the release rates of the salt under perfect sink conditions. The salt
release data is then utilized in a one-dimensional diffusion controlled transport model
[84, 106, 107, 128] to determine the partition coefficient and diffusivity of the lens.
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Results are compared from both approaches (diffusion cell and release in perfect sink).
Also, fundamental issues related to mechanisms of transport are addressed and the
dependence of the partition coefficient and the diffusivity on composition and
microstructure are explored.
6.1 Materials and Methods
6.1.1 Materials
Ethylene glycol dimethacrylate (EGDMA, 98%), N, N-Dimethylacrylamide (DMA,
99%) and 1-vinyl-2 pyrrolidone (NVP, 99+ %) were purchased from Sigma-Aldrich
Chemicals (Milwaukee, WI). Timolol maleate (98%), ethanol (99.5+ %) and Dulbecco‟s
phosphate buffered saline (PBS) were obtained from Sigma-Aldrich Chemicals (St.
Louis, MO). Sodium chloride (NaCl, 99.9+ %) were purchased from Fisher Chemical
(Fairlawn, NJ). The macromer acryloxy(polyethyleneoxy)-propylether terminated
poly(dimethylsiloxane) (DBE-U12, 95+ %) were purchased from Gelest Inc. (Morrisville,
PA). Methacrylic acid (MAA, 99.5%) were obtained from Polysciences, Inc. (Warrington,
PA) and 3-Methacryloxypropyltris(trimethylsiloxy)silane (TRIS) was supplied by Silar
laboratories (Scotia, NY). 2, 4, 6-trimethylbenzoyl-diphenyl-phosphineoxide (Darocur®
TPO) were kindly gifted by Ciba Specialty Chemicals (Tarrytown, NY). All chemicals
were used without further purification.
6.1.2 Preparation of Silicone Hydrogel
TRIS was used as the hydrophobic monomer; DMA and MAA were the hydrophilic
monomers; and DBE-U12 was selected as the macromer. Highly hydrophilic NVP
monomer was also added to increase water content of the hydrogel and EGDMA was
introduced in the monomer mixture for controlling the crosslinking. These six
components were mixed in several different ratios listed in Table 6-1 to prepare the
144
polymerization mixture. Four different series of mixtures were designed in this study. In
Series A the ratio of DMA to MAA is varied while keeping the fraction of hydrophilic
monomers (DMA+MAA) fixed; in series B the ratio of TRIS to Macromer is varied while
keeping the fraction of silicone component (TRIS + Macromer) fixed; in series C the
amount of crosslinker is varied while keeping amount of all other components fixed, and
finally in Series D the composition is randomly chosen. As an example, to prepare Gel
A1, 2.4 mL of a mixture that comprised 0.8 mL TRIS and 0.8 mL macromer and 0.8 mL
of the hydrophilic DMA/MAA mixture was combined with 0.12 mL of NVP and 0.1 mL of
EGDMA. After vortexing for few second, the mixture was purged with bubbling nitrogen
for 15 minutes to reduce the dissolved oxygen. To each monomer mixture, 12 mg of
photoinitiator Darocur® TPO was added with stirring for 5 minutes and the final mixture
was immediately injected into a mold which is composed of two 5-mm thick glass plates.
The plates were separated by a plastic spacer of a desired thickness, which is 0.13 mm,
0.26 mm or 0.40 mm in this study. The mold was then placed on Ultraviolet
transilluminator UVB-10 (UltraLum Inc.) and the gel mixture was cured by irradiating
with UVB light (305 nm) for 50 minutes. The synthesized hydrogel was either cut into
circular pieces (about 1.65 cm diameter) with a cork borer or other desired size and
shape by scissors for subsequent experiments. Prior to conducting further tests, the
prepared hydrogel was soaked in ethanol for 3 hours then dried at ambient temperature
overnight to remove the unreacted monomer.
6.1.3 Water Fraction Measurements
To determine the weight fraction of water (Q) in the hydrated gel, a dry gel of mass
Wd is soaked in DI water overnight or longer to ensure equilibrium. The hydrated lens is
then weighted and the equilibrium water fraction in the lens is calculated as
145
100[%]
eq
deq
W
WWQ (6-1)
where Weq is the mass of hydrated gel at equilibrium.
6.1.4 Ion Permeability Measurements
6.1.4.1 Salt release in perfect sink (kinetic desorption).
The circular hydrogel with 1.65 cm diameter was soaked in 0.5 M, 0.75 M or 1.0 M
NaCl solution until equilibrium was achieved. The salt-loaded gel was then transferred
into a DI water sink with a constant stirring at 300 rpm. The NaCl concentration of the
aqueous medium is determined by measuring the conductivity by Con 110 series
sensor OAKTON® followed by calculation from pre-established calibration curve. In this
study, the NaCl concentration is within the range that is linear to the measured
conductivity, with a slope of 8.58×10-6 M/s. For salt loading process, the volume ratio
of the solution to the gel was maintained about 80. For example, a piece of 0.13 mm
thick gel was soaked in 3.5 mL sodium chloride solution; for salt release experiment, the
volume ratio of the DI water and the gel was kept about 600. For example, a piece of
0.13 mm-thick gel was soaked in 27.5 mL DI water, and the amount of aqueous medium
was proportional to the thickness of the gel to maintain the liquid/gel volume ratio.
Because the volume of aqueous medium is much larger than that of the gel, the total
amount of NaCl loaded in the gel can be viewed as equal to the total salt release in DI
water reservoir, and the partition coefficient of NaCl in the gel can be thus determined.
6.1.4.2 Ion transport in diffusion cell (direct permeation)
The ethanol-extracted sample gel was soaked in DI water or in NaCl solutions with
various concentrations overnight at room temperature. The fully hydrated gel was
subsequently mounted in a horizontal diffusion cell, and then 18 mL of NaCl solution
146
and 30 mL of DI water were placed into the donor and receiver compartments with a
constant stirring at 300 rpm, respectively. After the ion transport starts, the NaCl
concentration in the receiving compartment is determined by measuring the conductivity
of the solution for 3 hours. The conductivity increases linearly in time after pseudo-
steady state is attained, and the slope of the increase in conductivity with time is a
measure of the ion permeability. The value of the ion permeability can also be obtained
by solving the diffusion equations for ion transport in the hydrogel, which will be
discussed later.
6.2 Results and Discussion
Below we first compare the transport measurements from the two different
methods and then discuss the dependency of the transport parameters on composition.
The two methods are compared only for Gel A1–4, and then the composition
dependency of transport parameters measured through the kinetic approach is explored
for all the gels prepared.
6.2.1 Comparison of Transport Measurements from the Kinetic and Permeation Approaches
6.2.1.1 Kinetics of salt release in perfect sink
In this approach the lens is soaked in a salt solution till equilibrium. The salt
loaded lens is then soaked in DI water and the release dynamics are measured. The
salt partition coefficient K in silicone hydrogel can be determined by
w,fg
w,fw,iw
w,f
g,f
CV
)C(CV
C
CK
(6-2)
where Vw and Vg are the volumes of the aqueous phase and the fully hydrated gel,
respectively, and Cg,f, Cw,i and Cw,f are the equilibrium concentrations of NaCl in the gel,
147
and the initial and equilibrium concentrations in the aqueous phase in the loading step,
respectively.
The dynamic ion release mechanism of silicone hydrogels can be viewed as one-
dimensional diffusion problem since the diameter of the hydrogel sample is much larger
than its thickness. Therefore, the ion transport can be described by the diffusion
equation, i.e.,
2
2
y
CD
t
C
(6-3)
The boundary conditions for the ion release experiment are
wKChytC
yty
C
),(
0)0,( (6-4)
where h is the half-thickness of the gel. The first boundary condition assumes symmetry
at the center of the gel and the second describes equilibrium of salt concentration
between the gel and the aqueous phase. A mass balance on the aqueous reservoir in
the beaker yields
hygel
w
wy
CDA
dt
dCV
2 (6-5)
where Vw is the water volume in the beaker, Agel is the cross-sectional area of the gel,
and C is the sodium chloride concentration in the gel.
In addition, the initial conditions for the drug release experiments are
0)0(
)0,(
tC
CtyC
w
i
(6-6)
Since the aqueous reservoir volume in the beaker is much larger than the gel
volume, it is reasonable to assume that the concentration Cw is negligible in this perfect
148
sink condition to simplify our calculation. The above set of equations can thus be solved
analytically to yield
)1(
12
1)(8
)2
)12(cos(
)12(
4)1(
2
22
2
22
4
)12(
022
0
4
)12(
Dth
n
nf
geli
w
n
Dth
n
i
n
enV
VCC
eyh
n
n
CC
(6-7)
The error between the experiment data and model-predicted value was defined
as NYYex )/1( 2 , where N is the number of data points; Yex and Y are the ratio of
measured NaCl release amount to that at equilibrium obtained by experiment and
model calculation, respectively. In addition, a parameter was introduced in this model
to compensate the offset error of time zero in salt release experiments, i.e. t in the fitting
model is equal to tapp - where tapp is the apparent experiment time. This imaginary
time offset is a parameter that presents the effects of diffusion boundary layer on the
surface, interaction between polymer chains and charged ions, and the default errors
from experimental setup on the initial reading of salt release. The above model was
used to fit the experiment data to determine the D and of the silicone gel by using the
function “fminsearch” in MATLAB® , and the results are summarized in Table 6-2 along
with the values of the water content and the salt partition coefficient. The comparison of
the experiment and model-predicted release profile for different gels are plotted in
Figure 6-1. It is clear that the model prediction in Figure 6-1 is well-matched with the
experiment release profiles, with a small time offset of time zero less than 40 seconds.
The model we proposed above based on two assumptions: the partition coefficient and
the diffusivity of hydrogel are independent of ionic concentration, and the dominating
149
mechanism of ion transport here is Fickian diffusion. These two assumptions are
needed to be confirmed prior to further discussion.
The effect of ionic strength on the ion transport of silicone hydrogel was studied by
conducting salt release in perfect sink with gels that have different initial NaCl amounts.
The gels were soaked in NaCl solution with different concentration until equilibrium, and
the following release profiles and fitted parameters of these gels are shown in Figure
6-2 and Table 6-3. Within the concentration range we explored here, the final release
amount of a given silicone hydrogel is proportional to the initial soaking concentration,
i.e. the ion partition coefficient keeps constant. The normalized release profiles for
different loading are overlapping each other implied that the diffusivity is independent on
ion concentration, as shown in Figure 6-2B. These results also suggest that the
electrostatic interactions of the double layer on the gel surface do not cause significant
impact on ion transport within explored ion strength range.
Another assumption in our proposed mechanism is that the ion transport of
silicone hydrogel here is diffusion controlled. From Equation 6-7 we know that if this
process is diffusion controlled with homogeneous diffusivity, the release time will
proportional to h2. The NaCl release profiles from gels with different thickness were
conducted in perfect sink and the dynamic release profiles were obtained as the method
for 0.13 mm-thick samples. The volume of aqueous reservoir in both uptake and release
experiments varied with the gel thickness so that the gel-fluid volume ratios were
maintained at the same as those of 0.13 mm-thick samples, and the measured release
profiles from gel with various thickness were plotted versus scaled time, which is
defined as time/(gel thickness/0.1 mm)2, and the results were shown in Figure 6-3. It is
150
obvious that within the margins of experimental error, the release profiles of the tested
silicone hydrogels with different thickness are overlapped, proving that the dominating
mechanism of ion transport of silicone hydrogel in perfect sink is the ion diffusion
through the gel phase.
6.2.1.2 Ion transport through permeation in a diffusion cell
Ion permeability of silicone hydrogel was also measured by diffusion cell, and the
results were shown in Figure 6-4. To obtain the ion permeability of the hydrogel, the
slope of the increase in conductivity with time is calculated after pseudo-steady state is
attained. The ionoflux diffusion coefficient, Dion, is then determined by solving the
diffusion flux equation for ion transport in the lens as:
dxdcA
nDion
/
(6-8)
where Dion is ionoflux diffusion coefficient (mm2/min), n' is the rate of ion transport
(mol/min), A is the area of ion transport (mm2), dc is concentration difference between
donor and receiving compartment (mol/mm3), and dx is the thickness of lens (mm). The
values of Dion for the gels were listed in Table 6-4.
In additional to use the diffusion flux equation at pseudo-steady state, we attempt
to model the ion transport of the hydrogel in the diffusion cell by the same diffusion
equation as Equation 6-3, with the following boundary conditions:
*),(
0)0,(
KChytC
ytC
(6-9)
where h is the thickness of the gel, C* is NaCl concentration in the donor
compartment, and K is the salt partition coefficient of the gel. The first boundary
condition is based on the assumption that the salt concentration in the receiver
151
compartment is much smaller than that in the donor compartment, which can be
considered as perfect sink condition. The second boundary condition assumes
equilibrium between the salt concentration in the gel and that in the donor solution. A
mass balance on the receiving compartment yields
0
y
w
wy
CDA
dt
dCV (6-10)
where Vw is the DI water volume of the receiving compartment, A is the cross-
sectional area of the gel, and Cw is the NaCl concentration in the receiving medium.
Finally the initial conditions for the drug release experiments are
0)0(
)0,(
tC
KCtyC
w
i (6-11)
where Ci is the concentration of NaCl solution in which the gel was soaked
overnight prior to conducting the diffusion cell experiment, and thus KCi is the initial salt
concentration in the gel.
The above set of equations can be solved analytically to yield
h
tDCenCnC
n
h
V
AKtC
yh
KCey
h
nnCnC
n
KC
Dth
n
n
i
n
w
w
n
Dth
n
i
*)(
1
*
2
1
*
1
)(*
)1()1)(cos()cos(2)1(
)(
)sin()1)(cos()cos(2
2
2
(6-12)
When the pseudo-steady state of ion transport is attained, Equation 6-12 will
reduced to Equation 6-8, where KD=Dion. The comparison of the experiment and model-
predicted diffusion profile for different gels are plotted in Figure 6-4, and the error
between the experiment data and model-predicted value was defined as the same for
previous fitting of salt release in perfect sink. was also introduced to compensate the
offset error of time zero, and K of each gel was calculated from previous perfect sink
152
test results. D and of each gel were subsequently obtained by using the function
“fminsearch” in MATLAB® .
Table 6-4 summarized the diffusivity fitting results for ion transport of silicone
hydrogel in diffusion cell. For each sample, the ionoflux diffusivity obtained by
Equation 6-9 is consistent with KD, the product of salt partition coefficient and the fitted
diffusivity from our diffusion controlled model. By comparing the fitting results of the
identical sample gel from salt release in perfect sink and from transport in diffusion cell,
we can conclude that the fitted diffusivity matches in both approaches, which confirms
the validity of both proposed model above. However, we noticed that a significant offset
to time zero in the diffusion cell experiment for gels that have no pre-loaded NaCl within,
which will be discussed later.
Similar to the model for salt release in perfect sink, the effect of ion strength also
need to be taken into concerned for our proposed diffusion-limited transport mechanism
in diffusion cell. A clean Gel A3 sample that contains no pre-loaded salt in the sample
was mounted to the diffusion cell and the ion transport with various NaCl concentrations
in the donor compartment was explored, as shown in Figure 6-5. Figure 6-5b is the
transport profile of normalized concentration versus time, and the normalized
concentration is defined as the measured NaCl concentration in the receiving
compartment divided by the NaCl concentration in the donor compartment. The results
from various donor concentrations overlapped within the margins of experiment error,
suggesting that the ion transport of silicone hydrogel in diffusion cell is linear in the
explored concentration range. The fitted parameters were listed in Table 6-5, and the
fitted Dion, D, and KD from experiments of various donor concentrations are similar to
153
each other respectively, which satisfy the assumption in our model that the ion diffusivity
of silicone hydrogel is independent of salt concentration. A significant large time offset
for around 10 minutes was also observed here. The fact that is not apparently
dependent on the donor concentration suggests that the large offset of time zero of the
ion transport in diffusion cell might not be caused by the surface diffusion resistance
boundary layer at the interface of silicone hydrogel and the NaCl solution in donor
compartment, since the thickness of this boundary layer is significantly dependent on
the ionic strength in the bulk solution.
Silicone gels with various amounts of pre-loaded NaCl were prepared by soaking
in solution with different NaCl concentration overnight prior to release experiment in
diffusion cell, and the release results were shown in
Figure 6-6 with the fitted constants listed in Table 6-6. The results show that the
fitted Dion, D and K has no significant dependency on the amount of pre-loaded NaCl in
the gel, but only depends on the composition of silicone gel. However, the fitted for
each gel significantly decreases when the gel contains some NaCl initially prior to the
diffusion cell experiment. The fact that the large offset from time zero for the ion
transport is only observed when the sample gel has no pre-loaded NaCl implies that an
adsorption/desorption barriers of ions on the gel matrix could be involved. For pure
hydrated gel without pre-loaded salt, the initial ion transport could be limited by this
adsorption/desorption process, and subsequently controlled by the ion diffusion through
the gel matrix after the ion adsorption/desorption process on the gel matrix reached
equilibrium.
154
The hydrogels with different thickness were also used in the ion transport
experiment in diffusion cell to confirm the assumption in our proposed model. From
Equation 6-12 we can observe that for long period of time, if the ion transport of silicone
gel is diffusion controlled with homogeneous diffusivity, the release time will proportional
to gel thickness. Thus we can define the scaled time for ion transport in diffusion cell
here as apparent release time divided by the ratio of gel thickness to 0.1 mm, and the
measured conductivity change in the receiving compartment for Gel A3 with various
thickness were plotted versus scaled time in Figure 6-7. All the ion transport profiles of
Gel A3 with different thickness overlap within the margins of experiment error, as we
observed in the salt release experiments in perfect sink. The results again confirm that
the rate limiting step of ion transport of silicone hydrogel in diffusion cell is the ion
diffusion through the gel matrix, which is the basic assumption in our proposed model.
6.2.2 Effect of Composition of Silicone Hydrogel on Ion Permeability
We used the developed methods above to further investigate the effect of silicone
hydrogel composition on ion permeability. All the experiments results described below
were conducted by using kinetic sorption/desorption method at perfect sink. The
thickness of sample hydrogel is 0.13 mm and the soaking NaCl concentration is 0.75 M
if not further specified.
Figure 6-8 shows the ion release experiment results for hydrogels with different
TRIS and Macromer ratios. Because the amount of hydrophilic monomer (DMA and
NVP) is fixed, it is expected that the water content and NaCl partition coefficient have
only a slight dependency on TRIS/macromer composition. For example, when we
increase the TRIS/(TRIS+macromer) amount from 25% to 75% , it will only leads to less
than 30% decrease on water content and NaCl partition coefficient. The decrease is
155
caused by the higher hydrophobility of silicone than the macromer after polymerization.
However, the fitted D exhibited a more complicated dependency on
TRIS/(TRIS+macromer) composition, as shown in Figure 6-8b. For the gel with
TRIS/(TRIS+macromer) composition less than 50%, D decrease in a slower rate than
that observed for gel with TRIS/(TRIS+macromer) ratio higher than 50%. The difference
might be caused by the structure difference of these gels. For the gels with higher
macromer content, it is easier to form bicontinuous polymer structure, which could
provide more diffusion channels for ions to diffuse through the gel matrix. On the other
hand, without enough macromer in the mixture the synthesized hydrogel could contains
a lot of phase-separated hydrophilic or hydrophobic regions, and therefore retards the
ion transport.
The effect of the crosslinker, EGDMA, on ion transport properties of the
synthesized silicone hydrogel was also studied. Figure 6-9 indicates that both water
content and NaCl partition coefficient decreases as the amount of EGDMA increased.
This is expected since when the crosslinking inside the gel matrix increase, the mobility
of the synthesized polymer chain decreases, and thus leading to lower water content
and salt solubility. However, with about 2% of EGDMA in the monomer mixture will help
the synthesized silicone hydrogel to obtain the optimal value of D, as shown in Figure
6-9b. With proper amount of EGDMA in the mixture as crosslinker, it is possible for the
monomers to be polymerized in structures that enhance the ion permeability; when the
amount of EGDMA over the critical value, the extra EGDMA could form higher
crosslinked region inside the gel matrix that retards the ion transport.
156
Finally, the effect of hydrophilic DMA composition in the hydrogel on ion
permeability were studied by using gel with various DMA, TRIS and macromer
composition (Gel D1-D6), while the rest ingredient amount in the gel mixture (NVP,
EGDMA and initiator) were kept the same, and the results are shown in Figure 6-10. If
we neglect the effect of non-hydrophilic monomers in the composition and focused on
the content of DMA of the synthesized silicone gel, we can clearly observed the linear
dependency of water content to the DMA composition in the gel, as shown in Figure
6-10d. The NaCl partition coefficient shows the similar linear dependency to DMA
component; on the other hand, even the fitted D increased as the DMA amount increase,
the slight non-linearity suggests other factors were involved as well. As a result, the KD
of these silicone gels shows a much stronger dependency on the DMA amount.
The summary of the composition effect on ion transport properties, including salt
partition coefficient, salt diffusivity and salt permeability, can be discussed based on the
water fraction of the synthesized silicone hydrogels. It is noted that even though the
water fraction is better to be measured in salt solution rather than DI water for these
comparisons, the variation within the salt concentration range explored here is not
significant, as observed in both our studies and previous studies on conventional
hydrogels systems [112, 113]. Ideally, if the salt only partition in the water phase inside
the gel matrix, the salt partition coefficient should be equal to the water fraction inside
the hydrogel, as the case in pure water. However, as shown in Figure 6-11, even
though the salt partition coefficient is strongly correlated to the water content of these
silicone hydrogels, the correlation is under the line where partition coefficient and water
fraction are equal. Similar results were observed in different hydrogel system as well
157
[112, 113, 119, 120, 124]. This deviation is considered to be resulted from the
interaction between the polymer chain and the solute. Even though this interaction is
generally much smaller compare to that between the water molecule and sodium
chloride, this effect becomes more dominant as the water content of the hydrated
hydrogel decreases. When the water fraction increases, the partition of sodium chloride
in the water phase becomes more significant, and thus the salt partition coefficient
gradually approaches to the ideal assumption as pure water.
In the free volume theory model of hydrated polymer membranes proposed by
Yasuda et al., the volume fraction of diluent is expressed by the hydration H. In the
consideration of permeation of salt, it seems to be quite obvious that salt alone will not
permeate through most pure polymer matrix. Consequently, it is reasonable to assume
that salt can permeate only through the hydrogel system in which the diluent of the
polymer is also a good solvent of the salt considered. The free volume of a hydrated
hydrogel system can be expressed by
polymerfOHff vHHvV ,2, )1( (6-13)
Since salt alone will not permeate through the polymer matrix, the free volume in
the polymer chains is negligible, and thus the free volume available for salt permeation
in the hydrogel system is approximately equal to Hvf,H2O. Thus, combined with the
Cohen-Turnbull relation, the diffusion constant D of salt through the hydrated membrane
can be written as a function of H as,
)]11
(exp[0 H
KDD (6-14)
where D0 is the diffusion constant of sodium chloride in pure water at the temperature of
the experiments, and K is a proportionality constant related to the characteristic volume
158
v* and vf,H2O. Then, the plot of log D versus 1/H is expected to be linear starting from the
diffusion constant of NaCl in pure water (1/H=1.0), which is 1.5×10-5 cm2/s at 25oC [129].
Figure 6-12 plotted the ion diffusivity (D) on logarithmic scale versus the reciprocal
of weight water fraction (1/Q) of these silicone hydrogels. It is noted that here assume
that the difference between the polymer density of these silicone hydrogel matrix is
negligible, and thus the weight volume fraction can be viewed as the volume fraction.
Since the water content of commercial silicone hydrogel contact lenses is between 20 to
60% [106, 130], here we only applied linear regression analysis on the hydrogel of
which the water fraction is greater than 15%, and the results showed that with proper
water content, the diffusion of sodium chloride through these silicone hydrogels can be
predicted by the free volume theory. However, when the water fraction decreases, the
free volume theory failed to relate the ion diffusivity to the water fraction. This is
reasonable since the water fraction is low, it is highly possible that the hydrophilic phase
of the hydrogel cannot form a continuous phase throughout the matrix, but dispersed
inside the continuous silicone phase.
The ion permeability is the product of salt partition coefficient and salt diffusion
coefficient. Even though the salt partition coefficient is dependent on the water fraction,
as we discussed earlier, this dependence is much smaller compared to the exponential
relationship between the ion diffusivity and water fraction. Thus, in most case the
relationship between ion permeability and water fraction can be also described by the
free volume theory model. As shown in Figure 6-13, the free volume theory model
linearly fitted the data from these silicone hydrogels with sufficient water fraction. When
water fraction becomes lower, in addition to the tendency to form phase-separated
159
regions inside the hydrogel matrix, the interaction between polymer matrix and salt also
becomes more significant. Both conditions violate the assumptions of the free volume
theory, which requires the hydrogel to be homogeneous and the interaction between
polymer and solute should be negligible. Some previous data of other silicone hydrogel
contact lens materials were also compared to our study. It is noted that for the
commercial continuous wear contact lenses NIGHT&DAYTM (Lotrafilcon A) and
O2OPTIXTM (Lotrafilcon B), the relationship between its ion permeability and water
content can be predicted by our regression results. Since both of these contact lens
were made by TRIS and DMA as the main components [130], it is reasonable to
consider our silicone hydrogels should have similar microstructures as these
commercial contact lenses. Therefore, these silicone hydrogels can be used to as a
substitute to further investigate other import transport properties of extended wear
contact lens, such as oxygen permeability and drug diffusivity.
In brief summary, critical properties of silicone hydrogels can be manipulated by
changing composition. Improvements in one property frequently occur at expense of
other properties, and so a careful optimization is needed to design contact lenses for
extended drug release. In this study a simple ion permeability measurement for silicone
hydrogel is established. While water content and ion partition coefficient of the hydrogel
generally depend on the total composition only, the ion diffusivity will also significantly
depend on the structures of different phases of the hydrogel. Thus, water content and
salt partition coefficient can be predicted by the bulk composition, but the diffusivity and
thus ion permeability are much more complicated, since it depends on the structure.
When the silicone hydrogel has proper composition to form bi-continuous
160
microstructures in the matrix, the ion transport through the gel can be described by the
free volume theory model. These results provided a solid foundation to further explore
other solute transport properties of silicone hydrogel contact lens, especially some
ionized hydrophilic ophthalmic drugs. In the future studies, other properties such as
elastic modulus and oxygen permeability should also be included to evaluate the design
of new silicone hydrogel as extended wear-extended release contact lens.
161
0.0
1.0
2.0
3.0
4.0
5.0
6.0
0 50 100
Time (minute)
Na
Cl re
lea
se
(m
g/g
ge
l)
Gel A1
Gel A2
Gel A3
Gel A4
Figure 6-1. NaCl release profile and model prediction (solid line) of different silicone
hydrogel in perfect sink. All samples are 0.13mm thick and pre-soaked in 0.75 M NaCl(aq) prior to conducting release experiment. Data are presented as mean ± S. D. with n ≥ 3.
162
0.0
0.5
1.0
1.5
2.0
2.5
3.0
3.5
4.0
0 20 40 60 80
Time (min)
Na
Cl re
lea
se
(m
g/g
ge
l)
1 M
0.75 M
0.5 M
0.0
0.2
0.4
0.6
0.8
1.0
1.2
0 20 40 60 80
Time (min)
Na
Cl re
lea
se
(C
/Cf)
1 M
0.75 M
0.5 M
Figure 6-2. NaCl release profile of 0.13 mm-thick Gel A3 with different pre-soaking
NaCl concentration in perfect sink. The solid lines are the model prediction and all experiment data are presented as mean ± S. D. with n ≥ 3.
163
0.0
0.5
1.0
1.5
2.0
2.5
3.0
0 50 100 150 200
Time (min)
Na
Cl re
lea
se
(m
g/g
ge
l)
0.13 mm
0.26 mm
0.40 mm
0.0
0.2
0.4
0.6
0.8
1.0
1.2
0 2 4 6 8 10 12
Scaled time (min)
Na
Cl re
lea
se
(C
/Cf)
0.13 mm
0.26 mm
0.40 mm
Figure 6-3. NaCl release profile of Gel A3 with different thickness in perfect sink. The
pre-soaked concentration of NaCl(aq) is 0.75M. The solid lines are the model prediction and all experiment data are presented as mean ± S. D. with n ≥ 3.
164
0.0
20.0
40.0
60.0
80.0
100.0
120.0
140.0
0 50 100 150 200
Time (min)
Co
nd
uc
tiv
ity
(
s)
Gel A1
Gel A2
Gel A3
Gel A4
Figure 6-4. Ion permeability test for silicone hydrogels by diffusion cell. All samples are
0.13 mm thick with no preloaded salt. The NaCl concentration in the donor compartment is 0.5 M. The solid lines are the model prediction and all experiment data are presented as mean ± S. D. with n ≥ 3.
165
0
5
10
15
20
25
30
35
40
0 50 100 150 200
Time (min)
Co
nd
uc
tiv
ity
(
s)
0.75 M
0.5 M
0.25 M
0.1 M
0
50
100
150
200
250
300
350
400
450
500
0 50 100 150 200
Time (min)
No
rma
lize
d c
on
ce
ntr
ati
on
0.75 M
0.5 M
0.25 M
0.1 M
Figure 6-5. Ion permeability test of Gel A3 by diffusion cell with various NaCl
concentration in the donor compartment. All samples are 0.13 mm thick with no preloaded salt. The solid lines are the model prediction and all experiment data are presented as mean ± S. D. with n ≥ 3.
166
A
0
20
40
60
80
100
120
140
160
0 50 100 150 200
Time (min)
Co
nd
uc
tiv
ity
(
s)
100% donor conc.
DI water
B
0
10
20
30
40
50
60
70
80
90
0 50 100 150 200
Time (min)
Co
nd
uc
tiv
ity
(
s)
100% donor conc.
DI water
Figure 6-6. Ion permeability test by diffusion cell of silicone hydrogel that were pre-
soaked in sodium chloride solution with different concentration. The samples and the sodium chloride concentration in the donor compartment are A) Gel A1, 0.5M, B) Gel A2, 0.5M, C) Gel A3, 0.5M and D) Gel A3, 0.25 M, respectively. All the samples are 0.13 mm thick. The solid lines are the model prediction and all experiment data are presented as mean ± S. D. with n ≥ 3.
167
C
0
5
10
15
20
25
30
35
0 50 100 150 200
Time (min)
Co
nd
uc
tiv
ity
(
s)
100% donor conc.
25% donor conc.
DI water
D
0
2
4
6
8
10
12
14
16
18
0 50 100 150 200
Time (min)
Co
nd
uc
tiv
ity
(
s)
100% donor conc.
25% donor conc.
DI water
Figure 6-6. Continued.
168
A
0.0E+00
2.0E-05
4.0E-05
6.0E-05
8.0E-05
1.0E-04
1.2E-04
1.4E-04
0 50 100 150
Time (min)
Na
Cl c
on
ce
ntr
ati
on
(M
)
0.13 mm
0.26 mm
0.40 mm
B
0.0E+00
5.0E-05
1.0E-04
1.5E-04
2.0E-04
2.5E-04
0 50 100 150
Time (min)
Na
Cl c
on
ce
ntr
ati
on
(M
)
0.13 mm
0.26 mm
0.40 mm
Figure 6-7. Ion permeability test of Gel A3 with different thickness by diffusion cell. The
sodium chloride concentration in the donor compartment is A) 0.25 M, with no preload salt in the gel; B) 0.5 M, with no preloaded salt in the gel and C) 0.25 M, the gel was pre-soaked in 0.25 M NaCl(aq), respectively. All experiment data are presented as mean ± S. D. with n ≥ 3.
169
C
0.0E+00
2.0E-05
4.0E-05
6.0E-05
8.0E-05
1.0E-04
1.2E-04
1.4E-04
0 50 100 150
Time (min)
Na
Cl c
on
ce
ntr
ati
on
(M
)
0.13 mm
0.26 mm
Figure 6-7. Continued.
170
A
0.00
0.02
0.04
0.06
0.08
0.10
0.12
0.14
0 0.2 0.4 0.6 0.8 1
TRIS/(TRIS+macromer)
Pa
rtit
ion
co
eff
icie
nt,
K
B
0.0000
0.0005
0.0010
0.0015
0.0020
0.0025
0.0030
0 0.2 0.4 0.6 0.8 1
TRIS/(TRIS+macromer)
D (
mm
2/m
in)
Figure 6-8. A) NaCl partition coefficient (K) B) diffusivity (D) C) KD and D) Water content (Q) for silicone hydrogel with different TRIS/Macromer compositions.
171
C
0.0E+00
5.0E-05
1.0E-04
1.5E-04
2.0E-04
2.5E-04
3.0E-04
3.5E-04
0 0.2 0.4 0.6 0.8 1
TRIS/(TRIS+macromer)
KD
(m
m2/m
in)
D
0
5
10
15
20
25
30
35
0 0.2 0.4 0.6 0.8 1
TRIS/(TRIS+macromer)
EC
W (
%)
Figure 6-8. Continued.
172
A
0.00
0.02
0.04
0.06
0.08
0.10
0.12
0.14
0.16
0 5 10 15 20
EGDMA (%)
Pa
rtit
ion
co
eff
icie
nt,
K
B
0.0000
0.0005
0.0010
0.0015
0.0020
0.0025
0.0030
0 5 10 15 20
EGDMA (%)
D (
mm
2/m
in)
Figure 6-9. A) NaCl partition coefficient (K) B) diffusivity(D) C) KD and D) Water content
(Q) for silicone hydrogel with different EGDMA compositions.
173
C
0.0E+00
5.0E-05
1.0E-04
1.5E-04
2.0E-04
2.5E-04
3.0E-04
3.5E-04
4.0E-04
0 5 10 15 20
EGDMA (%)
KD
(m
m2/m
in)
D
0
5
10
15
20
25
30
35
0 5 10 15 20
EGDMA (%)
EC
W (
%)
Figure 6-9. Continued.
174
A
0.00
0.05
0.10
0.15
0.20
0.25
0.30
0.35
0.40
0 20 40 60 80
DMA (%)
Pa
rtit
ion
co
eff
icie
nt,
K Gel P1~P6
Gel A~D
B
0
0.001
0.002
0.003
0.004
0.005
0.006
0.007
0.008
0.009
0 20 40 60 80
DMA (%)
D (
mm
2/m
in)
Gel P1~P6
Gel A~D
Figure 6-10. A) NaCl partition coefficient (K) B) diffusivity (D) C) KD and D) Water
content (Q) for silicone hydrogel with different DMA compositions.
175
C
0.0E+00
5.0E-04
1.0E-03
1.5E-03
2.0E-03
2.5E-03
3.0E-03
0 20 40 60 80
DMA (%)
KD
(m
m2/m
in)
Gel P1~P6
Gel A~D
D
0
5
10
15
20
25
30
35
40
45
50
0 20 40 60 80
DMA (%)
EW
C (
%)
Gel P1~P6
Gel A~D
Figure 6-10. Continued.
176
0
0.05
0.1
0.15
0.2
0.25
0.3
0.35
0.4
0.45
0.5
0 10 20 30 40 50
Q (%)
Pa
rtit
ion
Co
eff
icie
nt,
K
Series A
Series B
Series C
Series D
Figure 6-11. The relationship between salt partition coefficient (K) and water fraction (Q) of silicone hydrogel. The dash line indicates K=Q, and the solid line is draw for visual guidance.
177
1.0E-05
1.0E-04
1.0E-03
1.0E-02
1.0E-01
0 5 10 15
1/Q
D (
mm
2/m
in)
Series A
Series B
Series C
Series D
Water
Figure 6-12. The relationship between salt diffusivity (D) and the reciprocal of water
fraction (1/Q) of silicone hydrogels. The solid line is the linear best fit for data with Q greater than 0.15.
178
1.0E-07
1.0E-06
1.0E-05
1.0E-04
1.0E-03
1.0E-02
1.0E-01
0 5 10 15
1/Q
KD
(m
m2/m
in)
Series ASeries BSeries CSeries DWaterCommerical Contact LensesKim et al.
ACUVUE OASYS
O2 OPTIX
NIGHT&DAY
Figure 6-13. The relationship between salt permeability (KD) and reciprocal water
fraction (1/Q) of silicone hydrogels. The solid line is the best linear fit for data with Q greater than 0.15. Data were compared with selected literature data for commercial silicone hydrogel contact lenses [106] and other silicone hydrogels as contact lens material [84].
179
Table 6-1. Composition of silicone hydrogel. Each monomer mixture was mixed with
additional 0.12 mL of NVP before polymerization.
Sample TRIS (mL) Macromer (mL) DMA (mL) MAA (mL) EGDMA (L)
A1 0.80 0.80 0.80 - 100
A2 0.80 0.80 0.68 0.12 100
A3 0.80 0.80 0.56 0.24 100
A4 0.80 0.80 0.40 0.40 100
B1 1.20 0.40 0.80 - 100
B2 1.00 0.60 0.80 - 100
B3 0.60 1.00 0.80 - 100
B4 0.40 1.20 0.80 - 100
C1 0.80 0.80 0.80 - 5
C2 0.80 0.80 0.80 - 10
C3 0.80 0.80 0.80 - 20
C4 0.80 0.80 0.80 - 50
C5 0.80 0.80 0.80 - 75
C6 0.80 0.80 0.80 - 300
C7 0.80 0.80 0.80 - 500
D1 1.37 0.34 0.69 - 100
D2 1.37 0.69 0.34 - 100
D3 0.60 0.60 1.20 - 100
D4 0.48 0.48 1.44 - 100
D5 0.96 0.48 0.96 - 100
D6 0.40 0.40 1.60 - 100
180
Table 6-2. Parameters of different silicone hydrogels.
Sample Equilibrium Water
Content, Q (%) Partition
coefficient, K Diffusivity, D
(10-6 ×mm2/min) (min)
A1 22.5 0.1024 1957.7 0.160
A2 19.7 0.0834 1012.1 0.007
A3 16.6 0.0571 533.7 0.283
A4 11.1 0.0338 36.1 0.018
181
Table 6-3. Parameters for GEL A3 with various sodium chloride concentrations for salt
loading.
Soaking NaCl concentration (M)
Partition coefficient, K
Diffusivity, D (10-6 ×mm2/min) (min)
1.0 0.0585 563.2 0.750
0.75 0.0531 534.7 0.283
0.5 0.0570 583.0 0.698
182
Table 6-4. Fitting results of ion transport by diffusion cell for silicone hydrogels. All
samples are 0.13 mm thick and contain no preloaded salt.
Sample
Ionoflux Diffusivity, Dion (10-
6×mm2/min) Diffusivity, D
(10-6×mm2/min) KD
(10-6×mm2/min)
(min)
A1 193.5 1673.0 171.3 9.246
A2 112.3 1163.2 97.1 7.251
A3 39.6 596.6 34.3 13.786
A4 1.5 43.8 1.5 9.750
183
Table 6-5. Fitting results of ion transport by diffusion cell for Gel A3 with various sodium
chloride concentrations in the donor compartment. All samples are 0.13 mm thick and contain no preloaded salt.
Donor NaCl concentration
(M) Ionoflux Diffusivity, Dion (10-6×mm2/min)
Diffusivity, D (10-6×mm2/min)
KD (10-6×mm2/min) (min)
0.1 38.3 588.9 33.9 8.827
0.25 39.7 631.8 36.3 10.757
0.5 39.6 582.9 33.5 13.787
0.75 42.2 596.6 34.3 10.580
184
Table 6-6. Fitting results of ion transport by diffusion cell for silicone hydrogels which
pre-soaked in sodium chloride solution with various concentrations. All samples are 0.13 mm.
Sample
Donor NaCl conc.
(M)
Soaking NaCl conc.
(M)
Ionoflux Diffusivity, Dion (10-6× mm2/min)
Diffusivity, D (10-6×
mm2/min)
KD (10-6×
mm2/min) (min)
A1 0.5 0 193.5 1673.0 171.3 9.246
0.5 0.5 199.0 1997.0 204.5 3.848
A2 0.5 0 112.3 1163.2 97.0 7.251
0.5 0.5 111.3 1290.5 107.6 3.703
A3 0.5 0 39.6 582.9 33.5 13.786
0.5 0.125 38.5 550.4 31.6 2.701
0.5 0.5 42.9 633.8 36.4 3.897
0.25 0 39.7 631.8 36.3 10.756
0.25 0.125 38.7 600.3 34.5 0.952
0.25 0.5 38.0 719.4 41.4 3.184
185
CHAPTER 7 CYCLOSPORINE DELIEVERY BY SILICONE HYDROGEL FOR CRONIC DRY EYE
SYMDROME
Cyclosporine A (CyA, also known as cyclosporin or cyclosporine) is an
immunosuppressant drug used for treatment of many ocular diseases including
keratoconjunctivites sicca (dry eye syndrome) [131], uveitis in children and adolescent
[132, 133], vernal keratoconjunctivitis [134] and peripheral ulcerative keratitis [135]. It
also can be used to prevent allograft rejection [136]. CyA likely mitigates dry eyes
symptoms by reducing inflammation in eyes and tear-producing glands, potentially
resulting in increased tear production. Due to its low aqueous solubility (27.67 μg/mL at
25°C) [137], CyA is delivered through eye drops of an oil-in-water emulsion containing
0.05% cyclosporine [138]. While emulsion based eye drops is FDA approved, this
delivery approach suffers from low bioavailability with more than 95% of drug reaching
systemic circulation through transnasal or conjunctival absorption. Also, CyA delivery
through eye drops is particularly complicated for moderate dry eye patients who are
also contact lens wearers. In 2007, there were about 35 million contact lens wearers in
North America and about half of those reported some symptoms of dryness and
discomfort, more commonly experienced at the end of the day [53]. CyA emulsion can
also be used to alleviate these dryness and discomfort symptoms, but the patients need
to remove the lens prior to applying the CyA eye drop and then reinsert the lens 15
minutes afterwards [138].
To address the disadvantages of CyA delivery through eye drops, researchers
have been focused on the development of sustained release devices. For example,
biodegradable silicone-based ocular implants have been developed to continuously
186
release CyA for several years [139, 140]. Preclinical evaluation and animal studies
suggest significant potential of these implants for life-long treatment of ocular graft-
versus-host disease (GVHD) such as severe keratoconjunctivites sicca [141]. The
device can also be used for long-term treatment of uveitis [142], and can prevent high-
risk penetration keratoplasty (PKP) rejection [143]. These devices have higher
bioavailability for deep ocular tissues compared to topical delivery [144, 145]. Insertion
and replacement of these CyA implants require surgery and thus these appear to be
more suitable for severe ocular diseases that need long-term treatment. For the
treatment of moderate dry eye syndrome, which is one of the most common ocular
ailment that affect more than 30 million people in United States [146, 147], other CyA
delivery devices could be useful. For example, Gupta et.al proposed a novel
combination therapy based on extended CyA delivery lasting about 3 months from
punctual plug made of pHEMA [148].
In this Chapter we aim to develop contact lenses for providing sustained delivery
of CyA. To our knowledge, this study is the first one on CyA drug delivery through
contact lenses. The specific goal of this study is to develop silicone hydrogel contact
lenses that can delivery CyA for the entire duration of wear time, which is about 1 month
for several lenses. The approach explored here focuses on first measuring the CyA
release profiles from commercial Silicone hydrogel contact lenses and then
incorporating Vitamin E into the lenses to extend the release time to about 1 month.
Such lenses will particularly benefit dry eye sufferers who also wear contact lenses as
they will not need to remove lenses to deliver CyA eye drops. CyA has also shown
promising results in treating contact lens mediated dry eyes and thus the CyA releasing
187
contact lenses could also potentially mitigate or minimize the discomfort and the
dryness symptoms that arise due to contact lens use [149].
7.1 Materials and Methods
7.1.1 Materials
Five commercial contact lenses (diopter -6.50) were used in this study, including
1-DAY ACUVUE® and ACUVUE® OASYS™ from Johnson&Johnson Vision Care, Inc.
(Jacksonville, FL), NIGHT&DAY™ and O2OPTIX™ from Ciba Vision Corp. (Duluth, GA)
and PureVision™ from Bausch&Lomb, Inc. (Rochester, NY). Cyclosporine A (CyA) was
purchased from LC Laboratories (Woburg, MA). Ethanol (>99.5%) and Dulbecco‟s
phosphate buffered saline (PBS) was purchased from Sigma-Aldrich Chemicals (St.
Louis, MO), and Vitamin E (D-alpha tocopherol, Covitol® F1370) were kindly gifted by
Cognis corporation (Kankakee, IL). All chemicals were reagent grade and used as
supplied without further purification.
7.1.2 Drug Loading into Contact Lens
The commercial contact lenses were rinsed with DI water, dried in air and then
weighed. The stock CyA-PBS solution was prepared by dissolving 2.5 mg CyA in 100
mL PBS with moderate stirring at about 5oC for 24 hours. The stock solution was later
diluted with PBS to the desired drug concentration. The lenses were loaded with CyA by
soaking in 10 mL of a 15 g/mL CyA-PBS solution for 7 days. At the end of the loading
stage, the lenses were withdrawn from the solution, and excess drug solution was
blotted from the surface of the lens. The dynamic changes of drug concentration of the
soaking solution were not monitored during the drug loading process. The total amount
of CyA loaded into a lens was calculated at the end of the loading phase by determining
the total amount of drug-loss from the aqueous solution. The drug concentration in
188
aqueous solution was determined by measuring the absorbance in the wavelength
range from 208 to 220 nm with a UV-VIS spectrophotometer (Thermospectronic
Genesys 10 UV) and then using a pre-established calibration curve. The absorbance
was measured in the 208 to 220 nm range rather than at a single wavelength to ensure
that the experimental methods did not lead to drug degradation which will manifest as
changes in the absorption spectrum.
As shown in the latter sections, the 7-day duration was insufficient for equilibration,
and thus the duration of the uptake phase was increased to achieve equilibrium. The
equilibrium studies were conducted only for ACUVUE® OASYSTM lenses (dry weight =
22.3 0.3 mg). A lens was soaked into 10 mL of 17 g/mL CyA/PBS solutions for a
sufficiently long period of time to reach equilibrium. The mass of drug loaded into the
lens was determined by measuring the reduction in the drug concentration of the
soaking solution.
7.1.3 Drug Release Experiments from Lenses Loaded with CyA
The drug release experiments were carried out by soaking a CyA loaded lens in
1.75 mL of PBS. During the release experiments, the dynamic drug concentration in the
PBS was analyzed with the method described above. The release medium was
replaced by fresh 1.75 mL PBS after each measurement to maintain perfect sink
conditions.
7.1.4 Vitamin E Loading into Contact Lens
Vitamin E was loaded into lenses by soaking a lens in 3 mL of a Vitamin E-ethanol
solution for 24 hours. The concentration of Vitamin E in ethanol varies from 0.05 to 0.4
g/mL in the loading solution. After the loading step, the lens was withdrawn and blotted
to remove excess solution on the surface. The lens was then dried in air overnight. The
189
Vitamin E loading amount was determined by measuring the increase in lens weight.
CyA was loaded into the Vitamin E containing contact lenses by soaking the lens in 10
mL of a 17 g/mL CyA-PBS solution. The loading duration was increased with
increasing Vitamin E loading to provide sufficient time for equilibration. Subsequently,
drug release profiles were measured from the lenses that containing CyA and Vitamin E
by following the procedures described earlier.
7.2 Results
7.2.1 Drug Uptake by Pure Contact Lens
The data from the loading experiments was utilized to calculate the fractional
amount of drug absorbed (f) by the lenses during the 7-day soaking by using the
following relationship:
w,i
w,fw,i
C
)C(Cf
(7-1)
where Cw,i and Cw,f are the initial and final concentrations in the aqueous phase,
respectively. The calculated f by the lenses during the 7-day loading is listed in Table
7-1. For 1-DAY ACUVUE® , about 94% of CyA in the initial drug solution remained in
PBS medium after soaking for 7 days, while Silicone hydrogel lenses absorbed a
majority of CyA (51.6% to 75.6%) from the PBS solution into the gel matrix.
The partition coefficient (K) of CyA contact in lenses is an important parameter as
it determines the loading capacity. The value of K can also be calculated from the data
collected in the drug loading experiments through the following relationship:
w,fpl
w,fw,iw
w,f
fpl
CV
)C(CV
C
CK
, (7-2)
where Vw and Vpl are the volumes of the drug-PBS solution and the dry pure lens,
190
respectively, and Cpl,,f, is the final concentrations of the drug in the lens phase. It is
noted that the above relation defines the partition coefficient only if the concentration in
the gel and the fluid is in equilibrium at the end of the loading experiment. To confirm
whether the 7-day duration of the drug loading experiments is sufficiently long to reach
equilibrium, drug uptake experiments were conducted with ACUVUE® OASYSTM for
longer times. In these the dynamic drug concentration in the loading solution was also
measured. Figure 7-1 indicates that CyA concentrations between PBS medium and
ACUVUE® OASYSTM reaches equilibrium after about 15 days, and the final equilibrium
partition coefficient is 677.5 48.9. The shape of absorbance spectrum for CyA at
various times remains unchanged (data not shown), suggesting that the drug is stable
during the duration of the experiment.
Since the equilibration time for CyA loading is larger than 7-days, the data from the
7-day drug loading experiments cannot be strictly used to determine the partition
coefficients. However the data can still be used to obtain approximate values of the
partition coefficients. By comparing the K value for ACUVUE® OASYSTM, it is evident
that the partition coefficients reported in Table 7-1 which are based on 7-day soaking
are slightly lower than the true values. The estimates for K for other Silicone hydrogel
lenses are still useful as these allow comparison of the drug uptake potential for various
lenses. The calculated K values for different commercial contact lenses after 7-days
loading process are listed in Table 7-1. The drug partition coefficients are large for each
silicone hydrogel lens and are at least one order of magnitude larger than that for the 1-
DAY ACUVUE® lens.
191
7.2.2 Drug Release by Pure Contact Lens
Figure 7-2 shows the results of CyA release by 1-DAY ACUVUE® under perfect
sink condition. The release duration from the 1-DAY ACUVUE® is around 24 hours,
which is sufficient for this lens because it is intended for 1-day use only. Also, the 1-day
release duration implies that 7 days is sufficiently long for establishing equilibrium in the
loading phase, and thus the partition coefficient value for this lens reported in Table 7-1
is accurate.
The cumulative mass of drug released under perfect sink conditions is plotted as a
function of time for various silicone hydrogel contact lenses in Figure 7-3A, and the %
Release profiles are plotted in Figure 7-3B. The data clearly shows that each of the four
commercial silicone hydrogel contact lenses release CyA for extended period lasting
more than 7 days, which is significantly longer than the release duration by p-HEMA
hydrogel lenses (1-DAY ACUVUE® ). The data shows that after 7 days, ACUVUE®
OASYSTM lens releases about 82% of the loaded CyA, and the other three types of
lenses release about 50% of the loaded drug. In first 3 days, the ACUVUE® OASYSTM
lens releases about 15 g of drug each day and the other three types of lenses release
about 10 g CyA each day.
7.2.3 Drug Uptake by Vitamin E Loaded Contact Lens
The results of drug uptake by ACUVUE® OASYSTM lenses with various Vitamin E
loading are listed in Table 7-2. It is clearly seen that Vitamin E incorporation into the
lens significantly increases the mass of CyA absorbed. For example, after 120 days,
95% of the drug in the initial soaking solution was absorbed into the contact lens with
about 40% Vitamin E loading compared to 60% for the lens without Vitamin E. The
192
shape of the CyA spectra again is unchanged with time suggesting that the drug is
stable during the experimental duration of 4-months (data not shown). In addition, the
concentration of Vitamin E in the release medium was undetectable. Also, there was
negligible weight loss from the Vitamin E loaded contact lens during the release
experiment. Both of these observations suggest that the Vitamin E is not released into
the solution due to its very small solubility in aqueous solutions. By using the previously
determined equilibrium CyA partition coefficient of pure ACUVUE® OASYSTM and using
a mass balance [108], we estimated the CyA partition coefficient of the Vitamin E phase
in the lens (Kve) (Table 7-2). The average value of Kve (1.27±0.18 ×104) is relatively
independent of the amount of Vitamin E loaded into the lenses validating the accuracy
of the measured value, and also suggesting that the duration of the loading phase for
the Vitamin-E containing lenses was sufficient to achieve equilibrium. The average Kve
for CyA is about 20-fold higher than the partition coefficient of pure ACUVUE®
OASYSTM, which implies that the hydrophobic CyA has much higher affinity to the
Vitamin E phases than to the silicone gel matrix.
7.2.4 Drug Release by Vitamin E Loaded Contact Lens
Figure 7-4 shows the results of CyA release from Vitamin E loaded ACUVUE®
OASYSTM lenses under perfect sink condition. As shown in Figure 7-4A, the initial drug
release rate decreases as the Vitamin E loading amount in the contact lenses increases,
even though the total drug uptake amount increases as we discussed earlier. For
example, after the first day of release, ACUVUE® OASYSTM with 0% and 20% of
Vitamin E loading released at an average rate of 10.2 and 6.1 g/day, respectively. It is
also noted that the drug release duration by these contact lenses are significantly
193
extended with Vitamin E loading as shown in Figure 7-4B. For example, for pure
ACUVUE® OASYSTM lens, 60% of the inclusive CyA in the lens was released in 7 days,
while it took 16 and 46 days for ACUVUE® OASYSTM with 10% and 20% Vitamin E
loading to release the same percentage of loaded drugs, respectively. Moreover,
Vitamin E loading reduces the variations in the drug release rates with time. As shown
in Figure 7-5, the daily CyA release rate from AUCUVE® OASYSTM lenses decrease
rapidly with time from 10.2 g/day in day 1 to 1.3 g/day in day 15, while AUCUVE®
OASYSTM with 20% of Vitamin E releases CyA 6.1 g in day 1 and 1.8 g in day 15.
The solid lines in the figure are fits to the diffusion model which will be described later.
7.3 Discussion
CyA is a highly hydrophobic drug with very limited solubility in PBS, so its partition
coefficient is high in polymeric contact lenses. The partition coefficients are higher in
silicone hydrogel contact lenses compared to the hydrophilic p-HEMA lens (1-DAY
ACUVUE® ) because CyA has significantly higher affinity for the hydrophobic silicone
rich phases compared to that for the hydrophilic p-HEMA phase. The release time of
CyA from p-HEMA based contact lenses is much shorter than that from the silicone
hydrogel lenses due to the smaller partition coefficient. The ratio of the release times is
roughly equal to the ratio of the partition coefficients. The release duration from the
1-DAY ACUVUE® is about 1 day, which is consistent with the studies of Kapoor et al.,
who showed that 100 m thick pHEMA hydrogel releases CyA for about 1 day [38]. The
release duration of CyA from extended wear silicone hydrogel contact lenses is about
15 days. The CyA release duration from contact lenses is considerably longer than the
release duration for other ophthalmic drugs such as timolol, dexamethasone,
194
dexamethasone phosphate, and fluconazole, as we observed in Chapter 3 and 4. The
reason for the long release time for CyA is its higher molecular weight and very high
partition coefficient in the Silicone hydrogel contact lens. The larger size and the higher
binding of the drug to the polymer in the lens reduce the effective diffusivity leading to
long release times. While 15 day release is longer than that for other drugs, it is not
adequate because extended wear lenses are prescribed for about 1-month use. Vitamin
E incorporation in silicone contact lens leads to an increase in release duration of a
number of ophthalmic drugs, without a significant impact on any critical lens property.
Vitamin E incorporation also increases the release duration of CyA from ACUVUE®
OASYSTM. For example, the release duration increases to more than a month on 10%
Vitamin E incorporation into the lens. Incorporation of Vitamin E also increases the
effective partition coefficient of CyA into the lens which will have the additional benefit of
reducing the drug loss from the lens into the packaging solution. The impact of Vitamin
E incorporation on other silicone hydrogel lenses is expected to be similar to that on
ACUVUE® OASYSTM.
7.3.1 Release Mechanism and Model Fitting
To understand the mechanism of CyA release from the contact lenses, it is
instructive to compare the release profiles with a one-dimensional diffusion controlled
model which yields the following equation for % Release at short times [106, 108],
1002
(%)2
h
tD
M
M t
(7-3)
where Mt is the accumulated mass of drug released at time t, M the accumulated
mass of drug release as time approaches infinity and for perfect sink condition M = M0
(initial drug loading). The above equation predicts that the plot of % Release with
195
square root of time should be linear at short times. Figure 7-6 plots % drug release by
silicone hydrogel contact lens as function of square root of time for CyA. The best fit
straight lines are also shown in the figure. The straight lines fit the data well with R2
values larger than 0.96, showing that the drug transport in these lenses is diffusion
controlled. The slope is clearly larger for ACUVUE® OASYSTM showing the drug
diffusivity is highest for these lenses amongst those explored here. It is again noted that
Equation 7-3 is not strictly valid because the system did not reach equilibrium during
loading however the short time data should still satisfy Equation 7-3 as the
concentration in the region of the lens close to the surface was near equilibrium
concentration.
The CyA % release from Vitamin E loaded ACUVUE® OASYSTM versus square
root of time and the best fit straight lines for short time release are also plotted in Figure
7-7. All the R2 values of the fitting results for these ACUVUE® OASYSTM lenses with
various Vitamin E loaded lens are about 0.99, suggesting that the CyA release from the
Vitamin E loaded Silicone hydrogel lenses are controlled by diffusion as well. The
increased release times are due to the partitioning of the drug and slow diffusion
through the Vitamin E aggregates in the lens. This effect is similar to the proposed
mechanism of extended release of hydrophobic dexamethasone by the same systems
in Chapter 4. The comparison of CyA and dexamethasone delivery duration increase by
Vitamin E loaded ACUVUE® OASYSTM are plotted in Figure 7-8. In this figure, the ratio
of the release times after inclusion of Vitamin E (0) and the release time from lenses
without Vitamin E () are plotted as a function of the Vitamin E loading in the lenses.
While the trends are similar, Vitamin E loading has a slightly larger impact on
196
dexamethasone transport compared to CyA. This could perhaps indicate that the
diffusivity of CyA through the Vitamin E regions is attenuated to a lesser extend
compared to dexamethasone.
It is also noted that the profile of % drug release vs. square root of time is slightly
curved at short times and then becomes linear. The flux of drug transport from the gel
to the surrounding fluid is determined by two resistances in series; resistance in the gel
and then that in the fluid. The resistance in fluid is typically very small and can
frequently be neglected. However in instances of low gel diffusivity, particularly at short
times when the boundary layer thickness in the gel is very small, the fluid resistance can
dominate. For such cases, the drug flux is independent of time because the fluid
boundary layer thickness is unchanged in time [148, 150]. A constant drug flux results
in a linearly increasing cumulative release with time, which implies that the % Release
varies as the square of √t, which is clearly evident in the plots in Figure 7-6 and Figure
7-7.
Since the dynamic CyA release mechanism of contact lens can be viewed as one-
dimensional diffusion, it can be described by the diffusion equation, i.e.,
2
2
y
CD
t
C
(7-4)
For simplicity, we assume the diffusivity of CyA in the lens is independent of drug
concentration, which is usually valid in low drug concentration range. The diffusivity is
also assumed to be independent of position, which may be not precisely correct
because the distribution of Vitamin E in the lens may not be uniform. Since the exact
distribution of Vitamin E in the lens is not measured, we adopt the simplifying
assumption of uniform and fixed diffusivity.
197
The boundary conditions for the CyA release experiment are
wKChytC
yty
C
),(
0)0,( (7-5)
where h is the half-thickness of the lens, which can be assumed as 40 m. The
thickness of the lens is position dependent due to the lens curvature, but this variation is
also neglected for simplification of the model, and the lens is treated as a flat film with
thickness equal to the average thickness of the curved lens. The first boundary
condition assumes symmetry at the center of the gel and the second describes
equilibrium of drug concentration between the gel and the aqueous phase. Since the
aqueous reservoir in the vial was replaced with fresh PBS during the release, it is
reasonable to assume that the release occurs under perfect sink conditions, i.e., Cw can
be assumed to be negligible.
In addition, the initial conditions for the drug release experiments are
iCtyC )0,( (7-6)
The above set of equations can thus be solved analytically to yield
0
4
)12(2
22
)2
)12(cos(
)12(
4)1(
n
Dth
n
i
n
eyh
n
n
CC
(7-7)
Therefore, at the surface of the lens, the flux of drug is:
0
4
)12(2
22
2
n
Dth
n
i
hy
eh
C
dy
dC
(7-8)
We can thus relate the accumulated drug release rate measured by experiments
to these equations by a mass balance in the vial:
198
hylens
t
y
CDA
dt
dM
2 (7-9)
where Alens is the cross-sectional area of the contact lens. The mass of drug release in
period of time from t1 to t2 can be determined by computing Mt(t2) – Mt(t1). The data for
daily amount released as a function of time is plotted in Figure 7-5 and the fitted data is
plotted as the solid lines. The best fit values of diffusivity are 3.34 and 0.92
10-6mm2/hour for ACUVUE® OASYSTM with 0% and 20% of Vitamin E loading,
respectively. It is noted that this model neglects the mass transfer resistance in the fluid
phase, which is a reasonable assumption except at very short times.
7.3.2 Therapeutic Release Rates
Currently, CyA is delivered through 2-drops per day of oil-in-water emulsion
(Restasis® , Allergan) that deliver about 28 g (assuming a drop volume of 28 L) of
drug to the eye for the treatment of moderate dry eye [151]. Gupta et al. recently
determined the bioavailability of CyA delivered through Restasis® to be 2.8%, which
indicating that about 0.78 g/day of CyA is delivered to cornea and conjunctiva through
this treatment route [148]. CyA delivery through contact lenses will likely have a much
higher bioavailability due to the increase in the residence time of the drug in the tears.
The bioavailability needs to be determined through animal experiments but initial
estimate based on mathematical models supported by clinical data is about 50% [128].
Based on the 50% bioavailability for contact lenses and 2.8% for Restasis® , a release
rate of 1.6 g/day of CyA by contact lens should be able to provide equivalent
therapeutic effects. According to the clinical studies of CyA emulsion eye drops (later
Restasis® ), the concentration of CyA in the emulsion can be increased to 0.4% (Phase
II) or 0.1% (Phase III) without significant adverse effect after 12 weeks treatment [152,
199
153]. While the higher concentrations are non-toxic, no additional benefits were
observed with the higher concentrations. Thus, it can be anticipated that the therapeutic
window for daily dose of CyA delivered via eye drops is between 28 to 224 g.
Accordingly the therapeutic window for CyA delivery through contact lenses can be
estimated to be between 1.6 g/day and 12.8 g/day. Based on these estimates, a
suitable contact lens needs to deliver about 12.8 g/day on the first day and 1.6 g/day
on the last day of the wear-duration. The amount of drug released from the contact
lenses in a day is plotted as a function of time (days) in Figure 7-5 for lenses with and
without Vitamin E. The data shows that the lens without Vitamin E can maintain
delivery rates within the therapeutic window for 14 days, while ACUVUE® OASYSTM
with 20% of Vitamin E loading can maintain CyA release within the safe region at the
rate above the equivalent Restasis® release rate for 20 days. The duration of
therapeutic release cannot be increased any further for the lens without Vitamin E
without causing toxicity in the early phase of the release. However, the duration of
therapeutic release from ACUVUE® OASYSTM with 20% of Vitamin E can be further
increased by increasing the drug loading in the lens. The drug loading in the lens can
be increased by soaking the lens in solution with higher drug concentration since the
lens was soaked in CyA solution at 60% of the solubility limit in this study.
To brief summarize the study in this chapter: CyA is loaded into commercial
silicone hydrogel contact lenses by soaking the lenses in CyA-PBS solution.
Subsequent in vitro release experiments in perfect sink condition demonstrate that the
loaded CyA in the lens can be release for about 2 weeks. The release duration can be
further increased to the total wear time of the lens which is about a month, through
200
incorporation of Vitamin E into the lens. In addition to increasing the release duration,
Vitamin E incorporation also renders the release profile closer to „zero order‟ such that
the release rates are within the estimated therapeutic window for the entire 1-month
period. The long release duration along with the higher bioavailability compared to eye
drops suggests that Vitamin E loaded silicone hydrogel lenses could be very useful for
extended and controlled release of CyA. These lenses could potentially be useful for
treatment for chronic dry eye and also for reducing the symptoms of contact lens
mediated dry eyes. It is noted though that in vivo release and toxicity studies are
needed to fully determine the benefits of CyA release from extended wear contact
lenses.
201
0
20
40
60
80
100
120
0 5 10 15 20 25 30 35
Time (Day)
Dru
g U
pta
ke
(
g)
Figure 7-1. Cumulative drug uptake by ACUVUE® OASYSTM lens soaked in 17 g/mL of 10 mL CyA/PBS solutions.
202
0
2
4
6
8
10
12
0 10 20 30 40 50
Time (hour)
Dru
g r
ele
as
e (
g)
Figure 7-2. Cumulative CyA release by 1-DAY ACUVUE® . Drug was loaded in the lens
by soaking in 15 g/mL of 10 mL CyA/PBS solutions for 7 days. The release profiles were measured in 1.75 mL fresh of PBS that was replaced after every
measurement. Data are plotted as meanSD (n=3).
203
Figure 7-3. Cumulative CyA release from silicone contact lens. The drug was loaded in
the lenses by soaking in 10 mL of 15 g/mL CyA/PBS solution for 7 days. The release profiles were measured in removed to 1.75 mL fresh of PBS that was
replaced after every measurement. Data are plotted as mean SD (n=3).
0
10
20
30
40
50
60
70
80
90
0 20 40 60 80 100 120 140 160 180
Time(hour)
% d
rug
re
lea
se
NIGHT&DAY
O2OPTIX
ACUVUE OASYS
Pure Vision
0
10
20
30
40
50
60
70
0 20 40 60 80 100 120 140 160 180
Time(hour)
Dru
g r
ele
as
e (
g)
NIGHT&DAY
O2OPTIX
ACUVUE OASYS
Pure Vision
204
0
10
20
30
40
50
60
70
80
90
100
0 10 20 30 40 50 60
Time (day)
% D
rug
re
lea
se
0 g Vitamin E/g pure lens
0.1 g Vitamin E/g pure lens
0.2 g Vitamin E/g pure lens
Figure 7-4. Cumulative drug release from Vitamin E loaded ACUVUE® OASYSTM
lenses. The drug was loaded in the lenses by soaking in 10 mL of 17 g/mL CyA/PBS solution for 30 to 60 days. The release profiles were measured in 1.75 mL fresh of PBS that was replaced after every measurement. Data were
plotted as mean SD (n=3).
0
20
40
60
80
100
120
0 10 20 30 40 50 60
Time (day)
Dru
g r
ele
as
e (
g)
0 g Vitamin E/g pure lens
0.1 g Vitamin E/g pure lens
0.2 g Vitamin E/g pure lens
205
Figure 7-5. Daily average CyA release rate from Vitamin E loaded ACUVUE®
OASYSTM lenses. The drug was loaded in the lenses by soaking in 10 mL of
17 g/mL CyA/PBS solution for 30 to 60 days. The experimental data is
plotted as solid markers (mean SD (n=3)). The solid lines are the model fits based on diffusion model (Equation 7-9). The dash lines represent the estimated therapeutic window on the basis of Phase II and Phase III studies for Restasis® along with estimated bioavailability from drops and lenses [152, 153].
0
2
4
6
8
10
12
14
0 10 20 30 40
Dru
g r
ele
as
e r
ate
(
g/d
ay)
Time (day)
0 g Vitamin E/g pure lens
0.2 g Vitamin E/g pure lens
Phase II
Phase III
206
0
10
20
30
40
50
60
70
80
90
100
0 5 10 15
Time0.5
(hour0.5
)
% d
rug
re
lea
se
Night and Day
O2OPTIX
ACUVUE OASYS
Pure Vision
Figure 7-6. Plot of % CyA release by silicone contact lenses versus square root of time. The lines are the best fit straight lines. The fitted slope and R2 are 6.7967 and 0.9908 for ACUVUE® OASYSTM, 3.7416 and 0.9823 for O2OPTIXTM, 3.6036 and 0.9940 for Pure VisionTM and 3.3205 and 0.9660 for NIGHT&DAYTM, respectively. Data are presented as mean ± S.D. with n = 3.
207
0
10
20
30
40
50
60
70
80
90
100
0 10 20 30 40
Time0.5
(hour0.5
)
% d
rug
re
lea
se
0 g Vitamin E/g pure lens
0.1 g Vitamin E/g pure lens
0.2 g Vitamin E/g pure lens
Figure 7-7. Plot of % CyA release by Vitamin E loaded ACUVUE® OASYSTM versus square root of time. The lines are the best fit straight line. The fitted slope and R2 are 4.3901 and 0.9894, 3.0685 and 0.9891, 1.9404 and 0.9882 for lens with 0%, 10% and 20% of Vitamin E loading, respectively. Data are presented as mean ± S.D. with n = 3.
208
0
5
10
15
20
25
30
35
40
45
50
0.0 0.1 0.2 0.3 0.4 0.5
Vitamin E volume fraction
Dru
g d
eli
ve
ry t
ime
in
cre
as
e,
/ 0
Dexamethasone
Cyclosporine A
Figure 7-8. Comparison of CyA and dexamethasone [108] delivery by Vitamin E loaded
ACUVUE® OASYSTM.
209
Table 7-1. Results of CyA uptake by silicone contact lens. Each lens was soaked in 10
mL of 15 g/mL CyA-PBS solution for 7 days. Data are shown as mean SD (n=3)
Vpl (mL) CyA Uptake
(g)
Fraction of CyA Uptake by
Contact Lens (f)
Partition Coefficient,
K
1-Day ACUVUE® 0.0224 ± 0.0004 9.0 ± 2.7 0.060 ± 0.018 31.6 ± 10.3
NIGHT&DAYTM 0.0224 ± 0.0005 98.7 ± 4.9 0.658 ± 0.033 910 ± 118
O2OPTIXTM 0.0249 ± 0.0003 100.4 ± 2.3 0.700 ± 0.015 858 ± 59
ACUVUE® OASYSTM 0.0227 ± 0.0002 77.4 ± 2.4 0.516 ± 0.016 486 ± 30
Pure VisionTM 0.0224 ± 0.0003 110.4 ± 2.5 0.736 ± 0.017 1330 ± 104
210
Table 7-2. Results of CyA uptake by Vitamin E loaded ACUVUE® OASYSTM lenses.
Each lens was soaked in 10 mL of 17 g/mL CyA-PBS solution for various
drug uptake durations. Data are shown as mean SD (n=3)
Vitamin E loading (g
Vitamin E/ g pure lens)
Drug uptake duration (days)
Drug uptake (g) K Kve
0 30 103.4 ± 3.1 6.78×102
0.106 45 134.2 ± 1.3 1.58×103 1.15×104
0.200 60 151.1 ± 5.4 2.62×103 1.33×104
0.426 120 160.4 ± 4.7 3.70×103 1.12×104
0.653 120 166.4 ± 0.7 6.10×103 1.48×104
211
CHAPTER 8 DRUG DELIVERY BY CONTACT LENS IN GLAUCOMATOUS DOGS
While there is now extensive literature on in vitro studies for release of ophthalmic
drugs through contact lenses, there are very few animal or human studies. A number of
human studies were reported in 1970s focusing on management of glaucoma with
hydrophilic lenses soaked in pilocarpine. The lenses used in these studies were mostly
afocal, 0.2 mm thick, 13.5 -14 mm of diameter and radius of curvature between 7.8 and
8.6 mm Sauflon lenses [27-29]. Hillman et al. compared clinical response of Sauflon
lenses soaked in 1% pilocarpine solution with that of intensive pilocarpine 4% therapy,
which comprises instilling 1 or 2 drops per minutes for 5 minutes, every 5 minutes for
half an hour and then every 15 minutes for 90 minutes [28]. Hillman et al. reported that
the clinical response to the contact lens soaked in 1% solution was better than that for
pilocarpine therapy [28]. In another study with the same type of lenses and the same
treatment methodologies, Hillman observed a 54.8% drop in intraocular pressure (IOP)
with the contact lenses and a 49.7% reduction with the 4% pilocarpine regimen [27].
The encouraging results of these studies clearly prove the potential of glaucoma
therapy through contact lenses. Also the amount of drug loaded in the contact lenses
was substantially less than that delivered through eye drops supporting the predictions
of Li and Chauhan regarding higher bioavailability of contact lenses compared to drops.
These studies were however conducted with lenses that release the drug in a short
burst much like the profiles from eye drops. In view of the current renewed interest in
ophthalmic drug delivery by extended wear contact lenses, in vivo animal and human
studies with extended wear contact lenses are much needed. Such studies are
212
necessary to demonstrate that extended and continuous release of drugs can achieve
the same or better therapeutic efficacy as eye drops.
The goals of this study are to demonstrate the efficacy of glaucoma therapy
through release of timolol from silicone hydrogel contact lenses. Timolol is a beta-
blocker that is widely used to treat glaucoma by reducing IOP through decreasing the
production of aqueous humor [67]. We focus on this drug because of the large number
of glaucoma patients in the world [50], and also because of the potential of serious side
effects from systemic exposure to timolol [68]. For this in vivo study, we choose the
colony of beagle dogs who are affected by or carriers of a hereditary form of primary
open angle glaucoma, the most common form of glaucoma in human beings [69].
Beagle dogs have been used in several prior studies on glaucoma therapy [70-76].
Another benefit of using the Beagle dogs is that the cornea shape and size of these
dogs are similar to that of human beings, and therefore the commercially available
contact lens for human can be used in this study without further modification.
In this chapter we compare the pharmacodynamics of IOP reduction for contact
lenses with that from eye drops. Two different drug loadings are considered for contact
lenses to explore the effect of drug loadings and also to compare bioavailability of
contact lenses with that for eye drops. Since silicone hydrogel contact lenses release
timolol relatively rapidly, studies are also conducted with contact lenses loaded with
Vitamin E, which have longer release times compared to control lenses without Vitamin
E. In fact the in vitro results have shown that the timolol release time can be extended
from 1 hour by pure NIGHT&DAYTM contact lens to 70 hours by NIGHT&DAYTM with ca.
25% of Vitamin E loading, as shown in Chapter 3.
213
8.1 Materials and Methods
8.1.1 Materials
NIGHT&DAYTM (Lotrafilcon A, Ciba Vision Corp., Duluth, GA) contact lenses
(diopter -6.50) are used in this study. Timolol maleate (98%), ethanol (99.5%), and
Dulbecco‟s phosphate buffered saline (PBS) were purchased from Sigma-Aldrich
Chemicals (St. Louis, MO). Vitamin E (D-alpha tocopherol, Covitol® F1370) was gifted
by Cognis Corporation. All other chemicals were of reagent grade and used without
further purification.
8.1.2 Drug and Vitamin E loading into Contact Lens
NIGHT&DAYTM lenses were rinsed with DI water before further use. To load
Vitamin E into the contact lens, each rinsed lens was soaked in a 3 mL of 0.1 g/mL
Vitamin E-ethanol solution for 24 hours. After the loading step, the lens was taken out
and excess Vitamin E-ethanol solution on the lens surface was gently blotted out, and
then the lens was removed into a 30 mL of DI water to extract the residual ethanol in the
lens. After 3 hours the lens was removed into fresh DI water to repeat the extraction
process. After extraction, the lens was taken out and the excess water on the surface
was blotted out, followed by quickly dipping the lens in ethanol for few seconds to
remove the Vitamin E on the lens surface. The lens was then transferred into PBS
solution before further use. The Vitamin E loading in the lens is ca. 25% w/w according
to previous studies, which will release ca. 75% of concluded timolol into tear film within
24 hours at perfect sink condition [106].
To load timolol into the lens, NIGHT&DAYTM lens with or without Vitamin E loading
was soaked in a 3.5 mL of timolol maleate-PBS solution for at least 5 days. The
concentrations of timolol maleate-PBS solution are 2.67 mg/mL and 8 mg/mL, which are
214
chosen according to our previous in vitro studies to provide the capacity of the lens with
ca. 25% Vitamin E loading to release 20 g and 60 g of timolol into the eye within 24
hours, respectively [106, 128]. The drug release capacity of each delivery method was
summarized in Table 8-1. During the whole preparation process all lenses are kept in
hydrated state to maintain their original shape.
In the remainder of this Chapter, the notation CON_H and CON_L denote control
lenses without Vitamin E that are loaded with the higher (200 g) and the lower (67 g)
amounts of drug, respectively. Similarly, the notation VIT_H and VIT_L refer to Vitamin
E loaded lenses (25% w/w) loaded with the higher (200 g) and the lower (67 g)
amounts of drug, respectively.
8.1.3 Animal Model
Before investigating the effect of these timolol loaded contact lens on
glaucomatous dogs, each enrolled study dog (12 adult Beagle dogs with inherited open
angle glaucoma) had their IOP estimated via applanation tonometry (Tono-Pen-XL
(Mentor O and O, Norwell, MA)) in both eyes (OU), determined 3 times daily at the
same times of day for 5 days to establish a baseline for each individual animal for each
of these parameters. A topical anesthetic (proparacaine hydrochloride 0.5%) was
applied to each eye prior to the measurement of IOP OU.
After one week period of drug washout, 10 dogs were used in the investigation of
using eye drop to delivery timolol to the eye. Each study animal received one drop of
timolol maleate 0.5% ophthalmic solution to one eye (randomly selected) daily for 5
days. IOP OU was measured at time zero and then three times daily through the
duration of drug administration. Control experiments were also done to establish that
215
Tono-Pen-XL can be used to accurately measure IOP even after insertion of a contact
lens.
After one week period for drug washout, contact lenses with or without Vitamin E
loading which designed to release timolol 20g or 60 g within 24 hours was placed in
one eye (randomly selected) in each of the study dogs and IOP OU were measured at
time zero and then three times daily through the duration of drug administration (3 days).
12 dogs were randomly separated into 2 groups with 6 dogs each, and in the first week
CON_L and CON_H were conducted in each groups, then followed by VIA_L and
VIA_H. During the experiment, freshly loaded contact lenses replaced the previous
day‟s contact lens on a daily basis so that each contact lens will remain on the eye for
24 hours. The designed doses in this study have been tested and believed safe in the
dog via application. All experiments involving animals in this study were approved by
the Institutional Animal Care and Use Committee at University of Florida and were
performed in compliance with the ARVO Statement for the Use of Animal in Ophthalmic
and Vision Research.
8.1.4 Data Analysis
Each of the measured parameters was compared between each administration
method and the untreated controls to determine if there was a difference in the effects
among these therapeutic methods. The drug delivery methods comparisons were
performed using SPSS programs utilizing one-way ANOVA tests for multi-comparison
and Games-Howell tests for Post Hoc test since the sample size are not equal. Within
each test week, the average measurements for IOP for each day were compared with
subsequent measurements to detect significant changes (P < 0.05) using the Games-
216
Howell tests and ANOVA for repeated measurements. Each measured parameter for
drug-treated eyes was compared both to baseline and to the values for untreated eye.
Measured parameters for treated eyes were compared between methods as well.
8.2 Results
8.2.1 Contact Lens without Vitamin E
The mean ± SEM changes in IOP for each of the tested timolol delivery methods
are summarized in Figure 8-1. Timolol delivered by CON_L significantly decreased IOP
from the baseline in the treated eye by 3.18 ± 0.71 mmHg (P=0.002) and insignificantly
increased IOP (P=0.167) in the untreated control eye by 2.27 ± 0.79 mmHg from the
baseline, as indicated in Figure 8-1A. The difference in IOP decline between the treated
and untreated eyes was significant (5.45 ± 0.98 mmHg, P<0.001).
As shown in Figure 8-1B, timolol delivered by CON_H significantly decreased IOP
in the treated eye by 5.02 ± 0.83 mmHg (P<0.001) from the baseline, but the IOP
difference in the untreated control eye was not significant from the baseline (P=1.000).
The difference in IOP decline between the treated and untreated eyes was significant
after one day (5.20 ± 1.26 mmHg, P=0.004). There was no significant difference
between CON_L and CON_H in treated eye IOP (1.84 ± 1.01 mmHg, P=0.761).
8.2.2 Contact Lens with Vitamin E
Timolol delivered by VIT_L decreased IOP in the treated eye by 4.80 ± 0.63
mmHg (P<0.001) from the baseline, but showed no significant change in the untreated
control eye (P=0.996), as shown in Figure 8-1C. The difference in IOP decline between
the treated and untreated eyes was significant (4.18 ± 0.80 mmHg, P<0.001).
As shown in Figure 8-1D, timolol delivered by VIT_H significantly decreased IOP
in the treated eye by 3.27 ± 0.71 mmHg from the baseline (P=0.002). The IOP decrease
217
in the untreated control eye from the baseline was not significant (P=0.998). The
difference in IOP decline between the treated and untreated eyes was significant (3.90
± 0.91 mmHg, P=0.003). There was no significant difference between VIT_L and VIT_H
in treated eye (1.53 ± 0.83 mmHg, P=0.777).
8.2.3 Eye Drop
As shown in Figure 8-2, timolol delivered by eye drops decreased IOP in the
treated eye by 4.64 ± 0.41 mmHg from the baseline and in the untreated control eye by
3.17 ± 0.41 mmHg. The IOP decrease from baseline was significant for both the treated
eye (P<0.001) and the untreated eye (P<0.001). The difference in IOP decline between
the treated and untreated eyes by eye drops was significant (1.47 ± 0.43 mmHg,
P=0.011).
8.2.4 Drug Administration Methods Comparison
Timolol delivery by different methods was compared here to evaluate the effects.
Values for IOP in treated eye were not significantly different among all 5 different tested
timolol treatment methods (P=0.075). The decline of IOP in treated eye from baseline
was significant for all these five methods (P<0.001).
Figure 8-3 compares the difference of IOP between treated and untreated eye
from different drug administration methods. The results indicated that there was no
significant difference among all 4 contact lens delivery methods (P=0.068), while eye
drops leaded to significantly smaller IOP difference between treated and untreated eye
than CON_L (-4.24 ± 0.73 mmHg, P<0.001), CON_H (-3.47 ± 0.90 mmHg, P=0.003)
and VIT_L (-2.46 ± 0.71 mmHg, P=0.010), but not significantly lower than VIT_H (-1.84
± 0.76 mmHg, P=0.134).
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8.3 Discussion
All the five different methods tested in this study effectively lower the IOP in
treated eyes from the baseline, and the effects are similar among these methods. The
IOP reduction with contact lenses was not improved on increasing the drug loading
likely due to saturation of the target sites with the drug. In fact, future studies are
warranted with further lowering of drug loadings in the lens to determine if IOP
reductions comparable to those with eye drops can be achieved with even smaller
loadings in the lenses. By comparing the IOP reduction in the lenses with the smaller
drug loading with those from the eye drops, we can deduce that contact lenses deliver a
larger fraction of the loaded drug into the cornea compared to eye drops.
The IOP reduction in the untreated eye is significantly larger for the eye drops
compared to the lenses. The reduction in IOP in the untreated eye is commonly
attributed to the drug transport into the other eye through systemic circulation. For the
drug delivered through contact lenses, the untreated eye‟s IOP did not decrease likely
because of the smaller loss of the drug to the systemic circulation. This data further
supports the hypothesis that contact lenses deliver a larger fraction of the loaded drug
to the cornea.
In this study we did not observe any significant difference in IOP reduction
between contact lenses with or without Vitamin E. Inclusion of Vitamin E in the contact
lens increases the release duration but the total amount of drug loaded into the lenses
with and without Vitamin E was equal. In this study lenses were replaced each day and
the lenses loaded with Vitamin E release only about 75% of the drug in the wear time of
a day. Thus the total release of drug was actually smaller for the Vitamin E loaded
lenses. Furthermore, due to the larger release duration, the rate of drug released from
219
the Vitamin E loaded lenses was substantially smaller that the lenses without Vitamin E.
It is thus encouraging that the lenses with Vitamin E can achieve the same therapeutic
effect as eye drops in spite of the significantly smaller release rates albeit over a longer
duration. This shows that continuous release of ophthalmic drugs from contact lenses
could be efficacious. To further evaluate the effect of extended release by Vitamin E
loaded contact lens, in the future we can increase the lens wear interval from daily
replacement to 3-5 day replacement, which will provide further understanding for the
effect of extended release from Vitamin E loaded contact lens in animal models.
This pilot in vivo study demonstrates the potential advantages of delivering
ophthalmic drugs through contact lenses. Contact lenses achieved same efficacy as
eye drops but with one-third of the drug loading, and also resulted in smaller IOP
reduction in the untreated eye signifying reduced drug loss to the systemic circulation.
Incorporation of Vitamin E into the lenses does not lead to toxicity but it also does not
improve efficacy in this case likely because t lenses were replaced every day. In future,
in vivo studies with continuous lens wear should be conducted to further investigate the
feasibility of extended drug release by contact lens both with and without Vitamin E.
220
A
10
15
20
25
30
35
0 1 2 3
Time (Day)
Intr
ao
cu
lar
pre
ss
ure
(m
mH
g)
Treated (CON_L)
Untreated (CON_L)
Baseline
B
10
15
20
25
30
35
0 1 2 3
Time (Day)
Intr
ao
cu
lar
pre
ss
ure
(m
mH
g)
Treated (CON_H)
Untreated (CON_H)
Baseline
Figure 8-1. Effect of insertion of drug loaded contact lenses on the intraocular pressure
(IOP), A) CON_L, B) CON_H, C) VIT_L and D) VIT_H. Data is presented as
(mean SEM). Lenses are inserted in the treated eye at initial time and replaced every 24 hours.
221
C
10
15
20
25
30
35
0 1 2 3
Time (Day)
Intr
ao
cu
lar
pre
ss
ure
(m
mH
g)
Treated (VIT_L)
Untreated (VIT_L)
Baseline
D
10
15
20
25
30
35
0 1 2 3
Time (Day)
Intr
ao
cu
lar
pre
ss
ure
(m
mH
g)
Treated (VIT_H)
Untreated (VIT_H)
Baseline
Figure 8-1. Continued.
222
10
15
20
25
30
35
0 1 2 3 4 5
Time (Day)
Intr
ao
cu
lar
pre
ss
ure
(mm
Hg
)
Treated
Untreated
Figure 8-2. Effect of drug administration through eye drops on the intraocular pressure.
Data is presented as (mean SEM). Drug loaded drops are instilled in the treated eye at the initial time and then every 24 hours.
223
-2
0
2
4
6
8
10
12
0 1 2 3
Time (Day)
IOP
dif
fere
nc
e (
mm
Hg
)
Eye DropCON_LCON_HVIT_LVIT_H
Figure 8-3. Comparison of the effect of various drug delivery methods on the
differences in the IOP between the treated and the untreated eyes. Data is
presented as (mean SEM).
224
Table 8-1. Summary of various drug delivery methods considered in this study.
Methods Description Drug Release
Capacity (g)
Estimate release within
24 hours (g)
Estimate uptake by
eye (g)
CON_L Pure NIGHT&DAYTM, soaked in 2.67 mg/mL timolol maleate solution
67 67 27
VIT_L NIGHT&DAYTM with 25% Vitamin E, soaked in 2.67 mg/mL timolol maleate solution
67 50 20
CON_H Pure NIGHT&DAYTM, soaked in 8 mg/mL timolol maleate solution
200 200 80
VIT_H NIGHT&DAYTM with 25% Vitamin E, soaked in 8 mg/mL timolol maleate solution
200 150 60
Eye drop 0.5% ophthalmic solution, one drop
150 150 7.5
225
CHAPTER 9 CONCLUSIONS
Our study has conclusively shown the feasibility of extended ophthalmic drug
delivery by silicone hydrogel contact lens containing Vitamin E as diffusion barriers. In
Chapter 2 several properties including geometry, ion permeability, oxygen permeability
and UV transmittance are characterized to determine the pros and cons of loading
Vitamin E into the lenses. The results indicate the property changes caused by Vitamin
E loading do not disqualify these silicone hydrogels as extended-wear contact lens. In
addition, Vitamin E loading has a beneficial effect of blocking UV radiation which will
reduce the corneal damage due to UV light. For extended wear, the most effected
critical property of silicone hydrogel contact lens with Vitamin E is ion permeability, and
thus a further securitization of the ion transport of silicone hydrogels with various
compositions is discussed in Chapter 6.
Chapters 3 to 5 demonstrated the In vitro drugs release by the Vitamin E loaded
silicone hydrogel contact lenses, and the results indicated that the increase in release
duration is significantly dependent on the interaction between Vitamin E and the drug of
interest. For hydrophilic drugs (timolol, fluconazole, dexamethasone phosphate), the
drug release duration increases quadratically in Vitamin E loading; for hydrophobic
drugs dexamethasone and cyclosporine A, the effect of the Vitamin E inclusion is
smaller but still significant for release. On the other hands, for some amphiphilic
anesthetic drugs, including lidocaine, bupivacaine and tetracaine, the interfacial
interaction between drug and Vitamin E aggregation plays the determinative role for
drug transport through the Vitamin E/silicone hydrogel matrix. Ocular drugs delivery by
contact lens can be viewed as a one-dimensional transport by a flat thin film, and
226
subsequent mathematical models based on the proposed mechanisms were
established. In addition, a case study in Chapter 7 explored the potential of silicone
hydrogel contact lens for the treatment of chronic dry eye syndrome, and the in vitro
results suggests that CyA delivery by silicone hydrogel contact lens can provide much
safer and more efficient drug delivery route compared to traditional commercial eye
drops. Finally, in Chapter 8 we demonstrated that timolol can successfully delivered to
glaucomatous dogs via drug-impregnated contact lenses without irritation of eye or any
other unwanted safety concern. By utilizing contact lens to deliver timolol to the eye,
the intraocular pressure in the treated eye decreased effectively to similar degree
compared to that by eye drop treatment, while it significantly reduced the risk of
systemic drug exposure.
In conclusion, silicone hydrogel contact lenses with Vitamin E are promising
candidates for extended ophthalmic drug delivery. The Vitamin loading inside the
silicone hydrogel matrix can significantly attenuate the drug delivery rate, reduce
wastage and provide safer treatment route. While the results presented have focused
on drug contact lenses, the novel approach of in situ creation of transport barriers in
silicone hydrogels could be used in other areas where extended release of solutes is
desired, such as pucta plugs, ophtha coils, retinal implants, transdermal patches, wound
healing patches, cornea replacement materials, etc.
227
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BIOGRAPHICAL SKETCH
Cheng-Chun Peng was born in Hsinchu, Taiwan in 1981. After graduating from
Taipei Municipal Chieh Kuo High School (Taipei, Taiwan) in 1999, he began his
undergraduate studies in National Taiwan University (NTU) in Taipei, Taiwan in
September of 1999. He received his Bachelor of Science degree in chemical
engineering in June of 2003. Shortly thereafter, he continued his graduate studies in
NTU in the Fall of 2003 and received his Master of Science degree in chemical
engineering in June of 2005. After serving as emergency medical technician in the Fire
Department of Taichung County in Taichung, Taiwan for 17 months, he joined the
Department of Chemical Engineering at the University of Florida in the Fall of 2007. In
October of 2007, he joined Dr. Aunj Chauhan‟s research group, where he has since
worked to complete his doctoral research on extended ophthalmic drug delivery by
silicone hydrogel contact lens.