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Development of novel biocompatible hydrogel coatings for implantable glucose sensors
h [electronic resource] /
by Chunyan Wang.
[Tampa, Fla] :
b University of South Florida,
Title from PDF of title page.
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Dissertation (Ph.D.)--University of South Florida, 2008.
Includes bibliographical references.
Text (Electronic dissertation) in PDF format.
ABSTRACT: Due to sensor -tissue interactions, currently none of the commercially available glucose sensors are capable of continuous, reliable monitoring of glucose levels during long-term implantation. In order to improve the lifetime of implanted glucose sensors, two series of biocompatible novel hydrogel coatings were designed, synthesized and the physical properties were measured. Different hydrogels with various 2,3-dihydroxypropyl methacrylate (DHPMA) compositions were coated onto glucose sensors. Results show that the higher freezable water content, swelling rate and uniform porosity that resulted from high DHPMA content increased the sensitivity and shortened the response time of glucose sensors. The linear range of a glucose sensor coated only with hydrogel is short, however, the range can be improved by coating the epoxy- polyurethane (PU) with a layer of hydrogel.Since the hydrogel minimizes the fibrosis and inflammation, it shows promise for use in implantable glucose sensors. However, the in vivo experiment shows only 25% of sensors still worked after 4 weeks. In order to overcome problems present in the first series of experiments, another series of novel hydrogels with various N-vinyl pyrolidone (VP) contents was developed. This study has provided a feasible approach to design and select the properties of the copolymer for coating implantable biosensors. The in vivo experiments demonstrate that a hydrogel coating significantly improved the performance of implanted glucose sensors. In order to suppress the acute inflammation caused by the surgery, dexamethasone-21 phosphate disodium salt (DX-21) was incorporated to a series of poly (HEMA-DHPMA-VP) hydrogels to investigate the drug delivery in vitro. All hydrogels showed a high initial release, followed by slow, long term release during the next hours to days.This release pattern is believed to be optimum for implanted glucose sensors suppressing the acute and chronic inflammation. Water structures in hydrogels swollen in different media water, PBS and DX-21 solution were also investigated. 1HEMA:1DHPMA copolymer and VP-HEMA-DHPMA copolymers imbibed higher freezable water fractions in DX-21 solution. The ratio of transporting water mass to DX-21 mass is 9.6 which is independent of the hydrogel composition.
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Advisor: Julie P. Harmon, Ph.D.
t USF Electronic Theses and Dissertations.
Development o f Novel Biocompatible Hydrogel Coatings f or Implantable Glucose Sensors by Chunyan Wang A dissertation submitted in partial fulfillment of the requirements for the degree of Doctor of Philosophy Department of Chemistry College of Arts and Sciences University of South Florida Major Professor: Julie P. Harmon Ph.D. Edward Turos, Ph.D. Xiao(Sheryl) Li, Ph.D. Roman Manetsch, Ph .D. Date of Approval: November 19, 2008 Keywords: poly(2 hydroxyethyl methacrylate), 2,3 dihydroxypropyl methacrylate N vinyl pyrolidone, freezing water, nonfreezing water, drug delivery Copyright 2008 Chunyan Wang
ACKNOWLEDGEMENTS Firstly, I would like to express my sincere gratitude to my dissertation advisor, Dr.Julie Harmon for her guidance, support, and encouragement in completing my Ph.D. research project. Thank you for giving me the freedom to pursue this p roject and for for your generosity, kindness, and wisdom over the years. I extend my appreciation to the members of my Ph.D. dissertation committee, Dr. Edward Turos, Dr. Rom ann Man etsch, Dr. Xiao (Sheryl) Li, Dr. Francis Moussy, Dr. Kirpal S. Bisht, who encouraged and leaded me to the right directions during my Ph.D. study. I would also like to thank Dr. David Rabson who served as the external committee chairperson for my di ssertation defense. I acknowledge Mr. Jay Bieber for his assistance with SEM analysis Benz Research & Development (Sarasota, FL) for the generous supply of ultra high purity monomers all of my fellows in P olymer Materials lab and my Chinese friends in th e USF for their support and encouragement. Last but certainly not least, I especially thank my parents and parents in law who have always supported and encouraged me with unconditional charity and their prayer. I am forever grateful to my husband Yangyang Zhang. He is always there with devotion, patience, love, and constant cheers. I am also grateful to my sweet little daughter, Yuxin Zhang, who is a blessing and a treasure of my heart.
i TABLE OF CONTENTS LIST OF TABLES v i LIST OF FIGURES vi ii LIST OF ABBREVIATIONS xv i ABSTRACT xvi ii C HAPTER 1: INTRODUCTION 1 1.1 Diabetes 1 1.2 Implantable Glucose S ensors 2 1.3. Biocompatibility of I mplanted G lucose S ensors 5 1.4. Strategies to I mprove S ensor B iocompatibility 10 1.5. Hydroge l 12 C HAPTER 2: SYNTHESIS AND PERFO RMANCE OF NOVEL HYDROGELS COATING FOR IMPLANTABLE G LUCOSE SENSORS 25
ii 2.1. Introductio n 25 2.2. Experimental Procedures 28 2.2.1. Materials 28 2.2.2. Synthesis of Poly (HEMA DHPMA VP EGDMA) Hydrogel Series 29 2.2.3. Scanning Electron Microscopy (SEM) 30 2.2.4. Sorption Experiments 30 2.2.5. Differential Scanning Calorimetry (DSC) 31 2.2.6. Sensor Preparation 32 2.2.7. In vitro Evaluation Method 32 2.2.8. In vivo Biocompatibility Studies 33 2.3. Resul ts and Discussion 35 2.3.1. Preparation of Hydrogels Sheets and Coatings 35 2.3.2. Inner Sample Morphology 35 2.3.3. Sorption Behavior 37 2.3.4. DSC Results 41 2.3.5. Sensor Performance of Hydrogel Coated Glucose Sensors 46 2.3.6 In vivo B iocompatibility 49 2.4. Conclusion 52 C HAPTER 3: USE OF HYDROGEL COATING TO IMPROVE THE PERFORMANCE OF IMPLANTED GLUCOSE SENSORS 53
iii 3.1. Introduction 53 3.2. Experimental 56 3.2.1. Chemicals and M aterials 56 3.2.2. Photopolymeri zation of P oly(HEMA DHPMA VP EG DMA) H ydrogels 56 3.2.3. Analysis of H ydrogels 57 3.2.4 Sensor P reparation 58 3.2.5. In vitro S ensor E valuation 61 3.2.6. Animals an d S urgical P rocedures 61 3.2.7. In vivo E valuation M ethod .. 62 3.3. Results and Discussion 65 3.3.1. Characterization of Hydrogels 65 3.3.2. In vitro P erformance of Pt/GOx/VP30 S e nsors 74 3.3.3 In vitro P erformance of Pt/GOx/EPU/VP30 S ensors 76 3.3.4. In vivo P erformance of Pt/GOx/EPU/VP30 S ensors 77 3.4. Conclusion 81 C HAPTER 4: EFFECT OF NOVLE HYDROGEL COMPOSITION ON TARGETED DEXAMETHASONE 21 PHOSPHATE DISO DIUM SALT DELIVERY 82 4.1. Introduction 82 4.2. Experimenta l 83 4.2.1. Chemicals 83
iv 4.2.2. Drug L oading by E quilibrium P artitioning 83 4.2.3. In vitro R elease S tudy 84 4.2.4. H PLC A nalysis of DX 21 and DX R elease 84 18.104.22.168 Preparation of Standard S olution 84 22.214.171.124 HPLC S ystem 85 126.96.36.199 Selection of Mobile Phase 85 188.8.131.52 Calibration C urve and Sample A ssay 85 4.3. Results and Discussion 86 4.3.1. The S election of the M obile P hase 86 4.3.2. Drug L oading by E quilibrium P artitioning 90 4.3.3. Drug Release 92 4.4. Conclusion 99 CHAPTER 5: WATER STRUCTURE IN HYDROGELS 100 5.1. Introduction 100 5.2. Experimental 102 5.2.1. Chemicals 102 5.2.2. Instrumentation 102 5.2.3. Method 103 5.3. Results and discussion 103 5.3.1. Equilibrium W ater F raction of H ydrogel B efore and A fter DX 21 R eleasing 103 5.3.2. Water S tructure of HEMA DHPMA C opolymer 106
v 5.3.3. Water S tructure of VP HEMA DHPMA C opolymer H ydrogels 118 5.4. Conclusion 126 C HAPTER 6: SUMMARY AND SUGGESTIONS FOR FUTURE STUDY 127 6.1. Summary 127 6.2. Suggestions for F uture W orks 132 LISTS OF REFERENCES 134 APPENDICES 146 A ppendix A: Chapter 2, DSC C urves 147 Appendix B: Chapter 5, DSC Curves 156 ABOUT THE AUTHOR End page
vi LIST OF TABLES Table 1.1 S ome biomedical applications of hydrogels 13 Table 1.2 M s 18 Table 2.1 Feed compositions of synthesized x erogels 30 Table 2.2 Dynamic s welling parameters 40 Table 2.3 T g s of different composition gels 41 Table 2.4 Water structures of different composition hydrogels 42 Table 2.5 Response s ensitiviti es (S) and t imes (T 90% ) of various hydrogel 46 Table 2.6 Response characteristics of various hydrogel EPU based glucose sensors 48 Table 3.1 Feed composition of synthesized hydrogel 57 Table 3.2 Influences of Coating Thickness on Sensor Response 75 Table 3.3 In vivo sensor sensitivity of Pt/GOx/EPU/VP3 0 sensors 78 Table 4.1 Mobile phase and Retention time 88 Table 4.2 Regression parameters for DX and DX 21 89 Ta ble 4.3 Drug loading percents & EWCs of hydrogels with DX 21 91 Table 4.4 DX 21 release exponents and coefficients 98 Table 5.1 Equilibrium water fractions of DX 21 loaded hydrogel s before and after the release of the drug in PBS solution 105 Table 5.2 EWF and NFWF of 1HEMA:1DHPMA copolymer 106
vii Table 5.3 Water melting points of hydrogels at different wate r fractions 110 Table 5.4 EWF of VP HEMA DHPMA copolymer 119 Table 5.5 Values of 21mass(g) 119 Table 5.6 NFWF of VP HEMA DHPMA copolymer 123 Table 5.7 FWF of VP HEMA DHPMA copolymer 123 Table 5.8 Molar ratios of non freezing bound water in different polymers 125
viii LIST OF FIGURES Figure 1.1 Schematic diagram of the sensor 4 Figure 1.2 The initial protein adsorption is followed by cellular adhesion onto the protein coated surface of the device 7 Figure1.3 SEM p hotographs of tips of glucose s ensors 8 Figure1.4 Light micrograph image of glucose sensor tip after 10 days of implantation in subcutis 9 Figure1.5 Various steps in a photoinitiated polym erization 21 Figure1.6 Types of photoinitiator 22 F Cleavage of alkylaryl ketones 23 Figure 2.1 S ensor photos show thin layers of hydrogel coating 34 Figure 2.2 SEM micrographs of inner g els 35 Figure 2.3 Swelling degree as a function of time for the different DHPMA fraction hydroge ls 39 Figure 2.4 Log versus logTime for the different DHPMA fraction hydroge ls 39 Figure 2.5 DSC heating curves of hydr ogels with different DHPMA content 45
ix Figure 2.6 Representative response curves of a Pt Ir/GOx/ HG D80 glucose sensor at day 7 and 28, respectively. 47 Figure 2.7 Calibration plots from a Pt Ir/GOx/HG D80 sensor and a Pt Ir/GOx/HG D80 sensor 48 Figure 2.8 Hi stology slides 51 Figure 3.1 Schematic diagram of hydrogel coated densing dlement of the glucose Electrode. 60 Figure 3.2 Photo of sensor coated with hydrogel 61 Figure 3.3 In vivo continuous glucose monitoring procedure 64 Fig ure 3.4 FTIR of four freeze dried samples 68 Figure 3.5 DSC of four freeze dried samples 69 Figure 3.6 Scanning electron mi croscopy images of four freeze dried hydrogel samples 71 Figure 3.7 Scanning electron microscopy images of VP15 and VP30 freeze dried hydrogel samp les with different magnification 73 Figure 3.8 Calibration plots for Pt/GOx and Pt/GOx/VP30 glucose sensors 75 Figure 3.9 Calibration plots for Pt/GOx/EPU/VP30 (0.4ul) sensors at day 3, 7, 16 a nd 28. 79 Figure 3.10 Hematoxylin and eosin stained sections of tissue surrounding glucose sensors implanted subcutaneously in rats for 28 days 80 Figure 4.1 Structure of dexamethasone 21 phosphate disodium salt(DX 21) 87
x Figure 4.2 Structure of dexamethasone 88 Figure 4.3 HPLC profiles with mobile phase methanol:water=6:4 (0.01MKH 2 PO 4 ) 89 Fig ure 4.4 DX 21cumulative release profiles from DHPMA hydrogels 95 Figure 4.5 DX 21Cumulative release profiles from VP hydrogels 96 Figure 4.6 Drug release kinetics log(Mt/Meq) versus log(Time ) 97 Figure 5.1 Photos of hydrogel loaded with DX 21(A) and hydrogel released DX 21(B) 105 Figure 5.2 DSC cooling and heating curves for 1HEMA:1DHPMA copolymer at various stages during desorption of deionized water, PBS solution and DX 21 115 Figure 5.3 Linear plot of total integrated endotherm area versus total water fraction for the 1 HEMA:1 DHPMA copolymer swollen in deionized water 117 Figure 5.4 Linear pl ot of total integrated endotherm area versus total water fraction for the 1 HEMA:1 DHPMA copolymer swollen in PBS solution 117 Figure 5.5 Linear plot of total integrated endotherm area versus total water fraction for the 1 HEMA:1 DHPMA copolymer swollen in DX 21 solution 117 Figur e 5.6 Linear plot of total integrated endotherm area versus total water fraction for the VP15 copolymer swollen in water, PBS and DX 21 120
xi Figure 5.7 Linear plot of total integrated endotherm area versus total water fraction for the VP15 copolymer swollen in water, PBS and DX 21 solution 121 Figure 5.8 Linear plot of total integrated endotherm area versus total water fraction for the VP15 copolymer swollen in water, PBS and DX 21 solution 122 Figure A 1 DSC curve for glass transition of D0 sample 147 Figure A 2 DSC curve for glass transition of D20 sample 147 Figure A 3 DSC curve for glass transit ion of D40 sample 148 Figure A 4 DSC curve for glass transition of D60 sample 148 Figure A 5 DSC curve for glass transition of D80 sample 149 Figure A 6 DSC curve for glass transition of D90 sample 149 Figure A 7 DSC curve for water structure of D0 with 48.8% EWC 150 Figure A 8 DSC curve for water str ucture of D20 with 52.0% EWC 150 Figure A 9 DSC curve for water structure of D40 with 60.1% EWC 151 Figure A 10 DSC curve for water structure of D60 with 65.5% EWC 151 Figure A 11 DSC curve for water structure of D80 with 69.4% EWC 152 Figure A 12 DSC curve for water structure of D90 with 69.9% EWC 152 Figure B 1 DSC curve for water structure of VP15 swollen in water with 1.8122 EWF 153 Figure B 2 DSC curve for water structure of VP15 swollen in water with 1.6144 EWF 153
xii Figure B 3 DSC curve for water structure of VP15 swollen in water with 1.3636 EWF 154 Figure B 4 DSC curve for water structure o f VP15 swollen in water with 1.0245 EWF 154 Figure B 5 DSC curve for water structure of VP15 swollen in water with 0.5908 EWF 155 Figure B 6 DSC curve for water structure of VP15 swollen in PBS with 1.6947 EWF 155 Figure B 7 DSC curve for water structure of VP15 swollen in PBS with 1.3817 EWF 156 Figure B 8 DSC curve for water structure of VP15 swollen in PBS with 1. 0981 EWF 156 Figure B 9 DSC curve for water structure of VP15 swollen in DX 21 with 2.8036 EWF 157 Figure B 10 DSC curve for water structure of VP15 swollen in DX 21 with 2.3937 EWF 157 Figure B 11 DSC curve for water structure of VP15 swollen in DX 21 with 1.3192 EWF 158 Figure B 12 DSC curve for water structure of VP15 swollen in DX 21 with 2.0100 EWF 158 Figure B 13 DSC curve for water structure of VP30 swollen in water with 1.6053 EWF 159
xiii Figure B 14 DSC curve for water structure of VP30 swollen in water with 1.0348 EWF 159 Figure B 15 DSC curve for water structure of VP30 swollen in water with 0.8266 EWF 160 Figure B 16 DSC curve for water structure of VP30 swollen in water with 0.5072 EWF 16 0 Figure B 17 DSC curve for water structure of VP30 swollen in PBS with 1.8709 EWF 161 Figure B 18 DSC curve for water structure of VP30 swollen in PBS with 1.5754 EWF 161 Figure B 19 DSC curve for water structure of VP30 swollen in PBS with 1.1570 EWF 162 Figure B 20 DSC curve for water structure of VP30 swollen in PBS with0.9071 EWF 162 Figure B 21 DSC curve for water structure o f VP30 swollen in PBS with 0.6337 EWF 163 Figure B 22 DSC curve for water structure of VP30 swollen in DX 21 with 2.9346 EWF 163 Figure B 23 DSC curve for water structure of VP30 swollen in DX 21 with 2.3633 EWF 164 Figure B 24 DSC cu rve for water structure of VP30 swollen in DX 21 with 1.5244 EWF 164
xiv Figure B 25 DSC curve for water structure of VP30 swollen in DX 21 with 1 .1222 EWF 165 Figure B 26 DSC curve for water structure of VP30 swollen in DX 21 with 0.7178 EWF 165 Figure B 27 DSC curve for water structure of VP45 swollen in water with 2.2428 EWF 166 Figure B 28 DSC curve for water structure of VP45 swollen in water with 1.5505 EWF 166 Figure B 29 DSC curve for water structure of VP45 swollen in water with 1.2132 EWF 167 Figure B 30 DSC curve for water structure of VP45 swollen in water with 0.8414 EWF 167 Figure B 31 DSC curve for water structure of VP45 swollen in PBS with 1.8890 EWF 168 Figure B 32 DSC curve for water structure of VP45 swollen in PBS with 1.6677 EWF 168 Figure B 33 DSC curve for water structure of VP45 swollen in PBS with 1.3592 EWF 16 9 Figure B 34 DSC curve for water structure of VP45 swollen in PBS with 1.2318 EWF 169 Figure B 35 DSC curve for water structure of VP45 swollen in PBS with 0.7746 EWF 170
xv Figure B 36 DSC curve for water structure of VP45 swollen in DX 21 with 3.1004 EWF 170 Figure B 37 DSC curve for water structure of VP45 swollen in DX 21 with 2.1888 EWF 171 Figure B 38 DSC curve for water structure of VP45 swollen in DX 21 with 2.1169 EWF 171 Figure B 39 DSC curve for water structure of VP45 swollen in DX 21 with 1.8013 EWF 172
xvi LIST OF ABBREVIATIONS NIH National Institute of Health PI Principal Investigator ADA American Diabetes Association Ag/AgCl Silver/silver chloride GOx Glucose o xidase SEM scanning electron microscopy PEG poly(ethylene glycol) HEMA hydroxyethylmethacrylate DLC diamond like carbon PU Polyurethane EPU Epoxy polyurethane PC phosphorylcholine MPC 2 methacryloyloxyethyl phosphorylcholine DX Dexamethasone VEG F Vascular endothelial growth factor VP N vinyl pyrolidone DHPMA 2,3 dihydroxypropyl methacrylate EGDMA ethylene glycol dimethacrylate IPN interpenetrating polymer network
xvii UV Ultra violet EWC Equilibrium water content HPLC High performance liquid ch romatography PVA poly(vinyl alcohol) PAAm poly(acrylamide) SD sorption degree S sensitivity T 90% response time W f % freezing water content W nf % Non freezing water content HG hydrogel T g Glass transition temperature FTIR Fourier transform infrared spectroscopy EWF equilibrium water fraction PBS phosphate buffer saline I.V. Intravascular IACUC Institutional Animal Care and Use Committee T c Crystallization temperature T m Melting temperature EWF Equilibrium water fraction FWF Freezing water fra ction NFWF Non freezing water fraction
xviii Development of Novel Biocompatible Hydrogel Coatings for Implantable Glucose Sensors Chunyan Wang ABSTRACT D ue to sensor tissue interactions c urrently none of the commercially available glucose sensors are capable of continuous, reliable monitoring of glucose levels during long term implantation. In order to improve the lifetime of implanted glucose sensors, two series of biocompatible novel hydrogel coatings were designed, synthesized and the physical pro perties were measured. Different hydrogels with various 2,3 dihydroxypropyl methacrylate (DHPMA) compositions were coated onto glucose sensors. Results show that the higher freezable water content, swelling rate and uniform porosity that resulted from hig h DHPMA content increased the sensitivity and shortened the response time of glucose sensors. The linear range of a glucose sensor coated only with hydrogel is short, however, the range can be improved by coating the epoxy polyurethane (PU) with a layer of hydrogel. Since the hydrogel minimizes the fibrosis and inflammation, it shows promise for use in implanta ble glucose sensors. However, the in vivo experiment shows only 25% of sensors still worked after 4 weeks. In order to overcome problems present in the first series of experiments, another series of novel hydrogels with various N vinyl pyrolidone (VP) content s was developed. This study has provided a feasible approach to design and select the properties
xix of the copolymer for coating implantable bi osensors. The in vivo experiments demonstrate that a hydrogel coating significant ly improve d the performance of implanted glucose sensors. In order to suppress the acute inflammation caused by the surgery, dexamethasone 21 phosphate disodium salt (DX 21 ) was incorporated to a series of poly (HEMA DHPMA VP) hydrogels to investigate the drug delivery in vitro A ll hydrogels showed a high initial release, followed by slow, long term release during the next hours to days. This release pattern is believed to be optimum for implanted glucose sensors suppressing the acute and chronic inflammation. Water structures in hydrogels swollen in different media water, PBS and DX 21 solution were also investigated. 1HEMA:1DHPMA copolymer and VP HEMA DHPMA copolymers imbibed highe r freezable water fraction s in DX 21 solution The ratio of transporting water mass to DX 21 mass is 9.6 which is independent of the hydrogel composition.
1 CHAPTER 1 INTRODUCTION The aim of this project is to design, synthesize and characterize novel biocompatible hydrogels for the coatings of implantable glucose sensors and drug in viv o. This project has been funded by the National Institute of Health (NIH) The research has been conducted under the supervision of co PI, Dr. Julie Harmon of the Department of Chemistry (USF). 1.1. Diabetes Diabetes is a chronic disease which involves r egulatory problems with the hormone insulin. Insulin is released from the pancreas to regulate the amount of insulin, or the body does not respond appropriately to insulin, t he symptoms of diabetes occur. According to American Diabetes Association (ADA), there are 23.6 million children and adults in the United States, or 7.8% of the population, who have diabetes. ( Diabetes, Oct.20, 2008 ) There are two types of diabetes. Type 1 diab etes occurs when the insulin producing cells of the pancreas (called beta cells) are destroyed by the
2 insulin injections to control their blood glucose level. It is estimated t hat 5 10% of Americans who are diagnosed with diabetes have type 1 diabetes. Unlike people with type 1 diabetes, people with type 2 diabetes produce insulin. However, the insulin their pancreas secretes is either not enough or the body is unable to recog nize insulin and use it properly. When there isn't enough insulin or the insulin is not used as it should be, glucose can't get into the body's cells. About 90 to 95% people with diabetes have type 2 diabetes. ( Diabetes, Oct.21, 2008 ) Long term high glucose leve ls can cause a number of long term, sometimes life threatening complications including heart disease and stroke, high blood pressure, blindness, kidney disease, nervous system disease, dental disease; therefore it is crucial to control and monitor the bloo d glucose level. 1.2 Implantable Glucose S ensors In order to regulate the blood glucose concentration, it is important for diabetics to have the ability to monitor blood glucose concentration. Diabetes patients are used to self monitoring blood glucose l evels by finger pricking to obtain blood samples. This kind of monitoring is discontinuous since it depends on how frequently the blood glucose detection is performed. It is very difficult to record frequent enough determinations every day because the pa tients are unwilling to withstand the pain associated with frequent finger pricking. Currently, continuous glucose monitoring systems, glucose sensors, are being developed as an alternative to
3 the present method. Glucose sensors can continuously detect c hanges of blood glucose levels, and therefore provide information on how to optimize insulin therapy Since the first enzyme electrode was constructed for the measurement of blood sugar in the 1960s (Clark et al. 196 2) considerable amount of research has been devoted to the development of glucose sensors. Many kinds of glucose sensors have been studied including non invasive glucose sensors such as optical glucose sensors or sensors designed to detect the glucose co ncentration in tears (Baca, 2006) or urine (Shieh, 1997) mini invasive glucose sensors which can be implanted in the subcutaneous tissue, vascular bed (Armour et al. 1990 ; Frost and Meyerhoff, 2002) or transferring interstitial fluid outside the body and d etecting by micro dialysis device (Steinkuhl et al. 1996) A non invasive glucose sensor is the most optimum method for patients and because it overcomes the biocompatibility problems. However the precision of this method needs to be improved for clinica l applications. Micro dialysis needs recalibration at least once daily and has the time lag between 5 and 45 min caused by the length and diameter of the tube and the rate of the pumping. Due to the risk of inserting the glucose sensor in the intravascul ar compartment for a long period, most studies have focused on the development of needle type glucose sensors for subcutaneous glucose monitoring (Chen et al. 1992 ; Matthews et al. 1988)
4 Figure 1.1 Schematic diagram of the sensor Figure adapted f rom (Moussy et al. 1993)
5 Figure 1.1 shows the typical needle type implantable glucose sensor which is a two electrode system including a central platinum (Pt) anode as working electrode and a silver/ silver chloride(Ag/AgCl) reference electrode surr ounding around Pt anode. Glucose oxdiase( GOx ) is immobilized between the coiled Pt electrode. The entire sensor is coated with a Nafion membrane to protect other components from biological degradation. Once glucose passes through Nafion membrane in the presence of oxygen, it is oxidized by GOx and produces hydrogen peroxide (H 2 O 2 ). Hydrogen peroxide is oxidized by a polarization voltage of about +700 mV at Pt electrode surface, thus producing an electric current that is monitored (Abel and von Woedtke, 2002) The chemical reactions are: In vitro and animal experiments initially showed encouraging results. However the long term stability of the signal is insufficient due to the low bioco mpatibility of the electrode surface. 1.3 Biocompatibility of I mplanted G lucose S ensors The implanted glucose sensors are not reliable if they progressively lose function after several days in vivo. It is believed this is caused by the events that aff ect the sensor itself (e.g. degradation, fouling), as well as by changes in the tissue surrounding the sensor, caused by implantation (Moussy, 2000)
6 When glucose sensors are impla nted within the body membrane, biofouling processes start immediately. Low molecular weight solutes, such as ions, small organic molecules, and high molecular weight composites, such as proteins and enzymes, will deposit on the surface of the sensors. Surface deposition of the low molecular weight solutes on the metallic electr odes, such as Pt, can cause degradation or corrosion of the metallic component. These problems can be avoided by coating electrodes with polymers such as polyurethane, cellulose and cellulose acetate, and Nafion. The initial step of protein adsorption co uld result in an unstable protein interactions occur with protein coated sensors show n in Figure 1.2 (Fraser, 1997) These interactions lead to adhesion and activation of cells which characterize the inflammatory response. In such a case the diffusion distance of the glucose and oxygen is greatly increased, and the true glucose concentrat ion surrounding the implantation site is different from that in non inflam m ed subcutaneous tissue (Rebrin et al. 1992) Pickup et al. ( Pickup et al. 1993) reported protein and cellular accumulation on the tips of the non functioned glucose sensors after on ly five hours of implantation. Figure 1.3 shows scanning electron microscopy (SEM) Photographs of Tips of Glucose Sensors. The functioning sensor shows minimal biofouling, but the non functioning sensor shows significant protein and cellular accumulation.
7 When substances from the body are able to penetrate the outer membranes and alter the metal electrode surface, electrode fouling, sometimes referred to as electrode passivation, occurs and causes a decrease sensor signal (Reddya and Vadgamab, 1997; Linke et al. 1999) In the case of fibrosis, the smooth solid sensors become surrounded by the fibrous capsule. The fibrous capsule surrounding the sensor presents a structurally distinct barrier of collagen fibrils, which might have an effect on the substrate mass transport, resulting in a delay in sensor response and an underestimation of analytes. (Frazer, 1997) Figure 1.4 shows dense fibrous capsule tissue surrounding a glucose sensor tip after 10 days of implantation in subcutis (Ertefai and Gough, 1989) Figure 1.2 The initial protein adsorption is followed by cellular adhesion onto the protein coated surface of the device. Figure adapted from (Voskerician and Aderson, 2006)
8 Figure 1.3 SEM p hotographs of tips of glucose s ensors. (A) Control sensor not implanted; (B) Functioning sensor showing minimal biofouling; (C) Non functioning sensor showing significant protein and Cellular accumulation. Figure adapted from (Pickup et al. 1993)
9 Figure 1.4 Light micrograph image of g lucose s ensor t ip aft er 10 Days of implantation in s ubcutis. Figure adapted from (Ertefai and Gough, 1989).
10 1.4 Strategies to I mprove S ensor B iocompatibility It can be seen that modification of the inter face between sensor and tissue to control inflammation and fibrosi s will enhance a glucose sensor s biocompatibility, function and lifespan. Thus strategies to improve sensor biocompatibility include modifications to reduce protein adsorption and local drug delivery of tissue response modifiers. Modifications to R educe P rotein A dsorption Protein adsorption is controlled by the characteristics of the material including topography, charge density, distribution and mobility, surface groups (chain length, hydrophobicity, and hydrophilicity), structural ordering (soft to har d segment ratio and distribution), and the extent of hydration (Wisniewski et al. 2000) therefore the simple strategy to improve sensor biocompatibility is to reduce protein adsorption by surface modification. One approach to modify the surface of glucose sensors is to incorporate poly (ethylene glycol) (PEG) into a poly(hydroxyethylmethacrylate). The PEG chains tend to line up parallel to each other and perpendicular to the surface to present a water rich phase that resists penentration by many proteins ( Claesson, 1993 ). Compared to the same membranes without PEG, PEG in the outer membrane induced less fibrous encapsulation after subcutaneous implantation in rats. But there is no report about improvement in sensor performance (Quinn et al. 1995; Quinn et al. 1997) The diamond like carbon (DLC), so coat on glucose sensors to enhance hemocompatibility as determined by both
11 sensitivity change and surface deposition of blood components examined by SEM (Higson and Va dgama, 1995) Another strategy to reduce protein adsorption is to develop biomimetic coatings containing phosphorylcholine (PC) groups. It has been demonstrated that PC coated glucose sensor could diminish protein adsorption (Yang et al. 2000) One exam ple of making biomimic membrane is to self assemble phospholipids on the polymer surface containing a phosphorylcholine group, 2 methacryloyloxyethyl phosphorylcholine (MPC) and hydrophobic alkyl group. The surface shows an excellent resistance for both p rotein adsorption and blood cell adhesion. The relative output current of the sensor covered with this membrane was maintained as the initial level even after 14 days of subcutaneous implantation in a rat (Ishihara, 2000) Other methods to reduce protein adsorption include the use of alginate/ polyly sine gel (Shichiri et al. 1988) PHEMA/polyurethane (PU) (Shaw et al. 1991) NafionTM (perfluorosulphonic acid) membrane (Wilkins et al. 1995; Moussy et al. 1994 ) crosslinked albumin (Armour et al. 1990) cel lulose (Kerner et al. 1993) Local D rug D elivery S trategies Site specific, controlled delivery of tissue response modifiers can be used alone or combine with above surface modifications to help control the tissue reaction. The a nti inflammatory drug d exam ethasone (DX ) contained in microspheres is incorporated in to a hydrogel and coated onto glucose sensors. The in vivo experiment results show that eluting dexamethasone successfully controled negative
12 tissue reactions at the sensor tissue interface by redu cing the level of inflammation mediation cells to those observed in normal tissue (Patil et al. 2004; Norton et al. 2005) Controlled nitric oxide (NO) or vascular endothelial growth factor (VEGF) delivery also can suppress the tissue reactions and improve biosensor performance (Norton et al. 2005; Shin and Schoenfisch, 2006) 1.5 Hydrogel Hydrogels are three dimensional cross linked polymer networks which are designed to swell but not dissolve in water. They can be chemically cross linked by a covalen t bond with multi functional cross linkers or physically cross linked by physical cross links from entanglements, association bonds such as hydrogen bond, Van der Waals forces, hydrophobic forces or crystalline forces (Peppas, 1986, V1) Due to high water absorption, low interfacial tension, high permeability to small molecules, and the soft and rubbery nature hydrogles are well suited for biomaterials. Therefore, the biomedical applications of hydrogels, such as suture s contact lens es and drug delivery vehicles are of most interest. Table 1.1 lists some (Peppas, 1986)
13 Table 1.1 S ome biomedical applications of hydrogels Biomedical applications Example Materials Blood compatible appli cations Poly( vinyl alcohol)(PVA), Polyacrylammide(PAAm), Poly(N vinylpyrolidone)(P VP ),Poly(hydroxyl methacrylate)(P HEMA), poly(ethylene, oxide) ( PEO ) Poly(ethylene glycol)(PEG), cellulose Contact lenses PHEMA, P( VP co HEMA), acrylamide and acrylonitrile b ased hydrogels, P(HEMA co MMA) Artificial tendons Fiber (polyethylene terephathalate) reinforced hydrogel Controlled release devices PHEMA, P(HEMA methyl acrylate) copolymers, P(HEMA MMA)copolymers, P(HEMA polymethacrylic acid) copolymer, PVA, P VP Mem branes for biomacromolecular separations for plasmapheresis, artificial kidneys, artificial liver P(glycerolmethacrylate co MMA), P(Hydroxypropyl methacrylate MMA) pure PHEMA, PVA, P VP sutures PHEMA Artificial skins PVA formaldehyde foam, PHEMA
14 Hydroge ls commonly are classified into neutral hydrogels, ionic hydrogels and swollen interpenetrating polymeric networks (IPN). Neutral hydrogels are based on neutral monomers, while ionic hydrogels are based on the ionic monomers such as acryl acid, methacryli c acid. An interpenetrating polymer network (IPN) is a combination of two polymers one of which must be synthesized, or crosslinked in the immediate presence of the other. For an interpenetrating polymeric network (IPN) hydrogel system, there is no chemi cal bonding between two polymer networks but there are physical entanglements (Liu et al. 2003 ) Covalent crosslinked gels do not dissolve in organic solvents even upon the addition of heat; whereas non covalent crosslinked gel will eventually dissolve in solvents or melt upon the addition of heat (LaPorte, 1997) Thus physically crosslinked gel s are not stable in heat and solvent environment The properties of synthetic hydrogels can be controlled by chanaging the ratio of hydrophilic to hydrophobic mo nomer s N eutral hydrogels are more biocompatible than the ionic hydrogels since more proteins will absorb on the surface of ionic materials (Peppas, 1 986) and crosslinked neutral hydrogels are more stable than the interpenetrating hydrogels. In this proj ect the neutral hydrogels are polymerized for the sensor coatings because the nonionic hydrogels exhibit minimal protein build up and excellent pH stability (Gates and Harmon, 2001) Some common synthe t ic hydrogels being used as biomaterials include PHEMA poly(2,3 dihydroxypropyl methacrylate) ( P DHPMA)), PAAm, P VP and PVA.
15 Poly(hydroxyalkyl methacrylates)(PHEMA) H ydrogels PHEMA is the most studied hydrophilic polymer in the biomedical industry. Since Wichterle and Lim polymerized HEMA hydrogels in 1960 ( Wichterle and Lim, 1960) for contact lenses, PHEMA hydrogels have received a lot of interest. The biocompatibility of PHEMA hydrogels is primarily due to high water content absorbed into the hydrogel network. Initiators, reaction by products, residual mo nomer, and impurities can be extracted with successive aqueous washings, so the likelihood of toxic or local immune reactions decreases in biomedical applications. When used as an implant, the rubbery nature of the hydrated hydrogel reduces mechanical irr itation of surrounding tissue. The low interfacial tension between the hydrogel surface and aqueous solution minimizes protein adsorption and possibly cell adhesion (Ratner and Hoffman, 1976) Cross linker impurities are present in all methacrylate monom ers, so it is usual to obtain crosslinked gel instead of linear PHEMA. PHEMA hydrogels can be prepared by bulk polymerization or solution polymerization using heat initiators or photoinitiators. The most popular method for preparing PHEMA hydrogels is by solution polymerization of the monomer, HEMA, in the presence of a cross linking agent. Different solvents can be used for solution polymerization such as dimethyl su l foxide(DMSO), propanol, glycerol, ethylene glycol, cyclohexanol, toluene, and dimethyl formamide. Water is commonly used as a solvent in the preparation of PHEMA hydrogels. Below a critical solvent concentration (40 to 60% water by
16 weight), the gel thus formed is homogeneous, nonporous, and optically transparent. When the solvent concentra tion is above this critical value, the resulting gel is heterogeneous and has a macro porous structure, and is optically opaque (Laporte, 1997) The water content of PHEMA hydrogels is usually relatively low ( <45 wt%) which results in the limited applicati ons of these hydrogels (Lee and Lin, 2003) Furthermore, PHEMA does not have the stringent mechanical properties needed for many biomedical applications. It was found that mechanical properties could be improved by crosslinking with ethylene glycol dimetha crylate(E GDMA). Water content can also be increased by copolymerizing with other monomers such as DHPM or vinyl monomer, VP (Laporte, 1997) (Table 1.2) Poly(N Vinyl 2 pyr olidone)(P VP ) H ydrogels The backbone and ring structure of P VP provide hydrophobicit y while the highly polar amide group provides hydrophilicity similar to that of a protein, but is physiologically inactive. The size of VP repeating structure along the main carbon chain produces free volume in the material, resulting in flexibility and w ater solubility. P VP was first used as a blood plasma substitute and extender during World War II (Robinson et al. 1990) VP can also be copolymerized with other materials to modify properties for various medical applications (Laporte, 1997) P VP is a major component of some cont act lenses and has been proposed as a material that can enhance blood compatibility (Francois et al. 1996)
17 P VP hydrogels have also been used as cell culture substrates. Interpenetrating polymer networks (IPNs) composed of P VP and gelatine have been studied as non fouling materi als for devices in contact with blood (Lopes and Felisberti, 2003) It has been shown that 70% P VP exhibits optimum swelling degree and compression strength. In biocompatibility tests, 70%P VP hydrogels showed no deleterious effect to the cells but were not good as gelatin for cell growth. As it turns out, 50% P VP retains the best properties of both polymers with higher water swelling degree compression strength and biocompatibility. Risbud et al. (Risbud et al. 2000) reported a growth promoting effect in a P VP co chitosan (a linear polysaccharide) hydrogel as compared to chitosan gels alone. Janson et al. studied porous VP BMA (n butyl methacrylate) copolymers as scaffolds for bone tissue engineering In their study they reported that 50:50( VP : BMA) is less cytotoxic and more biocompatible than 7 0:30 counterpart in vivo (Janson et al. 2005) From these research results, it can be concluded that VP can enhance the water content, mechanical properties and porous structures of hydrogels at the same time, behavior unusual for a hydrogel. High VP per cent hydrogels show less biocompatibility, but 50% of P VP hydrogels are biocompatible Fu r thermore, P VP can be coated on the polyurethane surface and firmly attach to the surface (Francois et al. 1996) This property is essential for this project.
18 s and MW s Chemicals Structure Molecular Weight 2 hydroxyethyl methacrylate(HEMA) 130.14 2,3 dihydroxypropyl methacrylate(DHPMA) 160.17 N Vinyl 2 Pyrrolidinone( V P ) 111.14 Ethylene dim ethacrylate(EGDMA) 198.22 1 phenyl 2 hydroxy 2 meth yl propan 1 one (Benacure 1173) 164 .00
19 Poly (2,3 dihydroxypropyl methacrylate) ( P DHPMA)) H y drogels In comparison with P HEMA, P DHPMA is more hydrophilic and even soluble in water since it contains two hydroxyl groups in each of its repeat units (Hou et al. 2002) DHPMA is often cop o lymerized with HEMA or VP to make hydrogels for soft contact len s es (Wang et al. US patent) Clinical experiments show that hydrogels used as contact lenses will dehydrate especially at high water content s (McConville and Pope, 2001) Gates found that DHPMA exhibits decrease d water loss rate (Gates et al. 2003) Pho topolymerization of H ydrogel by U ltraviolet light (UV) Polymerization of h (UV) radiation with initiator s or they can self initiat e UV radiation is frequently used to induce cross linking of po lymers, because it requires less ti me, less energy, and generates little heat during polymerization (Laporte, 1997) In addition, heat sensitive substrates such as plastics, printed circuit boards, paper, and wood can be coated and cured by UV (Nichols et al. 2001; Koleske, 2002) There are three main components in UV curing system: monomers or oligomers, photo initiators and additives. First, photo initiators absorb high intensity UV light and are raised to an excited state. From their radiation excited state, the photo initiators photolyze or degrade directly or indirectly into free radicals. These free radicals initiate the very rapid polymerization and crosslinking of monomers (Lee et al. 2006) Steps of UV polymerization are listed in Figure 1.5.
20 A ccording to the process for forming the initiating radicals, photo initiators can be divided into two classes. Norrish Type I photo initiators (cleavage) photolyze through a homolytic fragmentation mechanism or through a cleavage, and thereby directly for m free radicals capable of initiating polymerization. ( Figure 1.6) Norrish Type II photo initiators (H abstraction) are activated with radiation and form free radicals by hydrogen abstraction or electron extraction from a second compound (coinitiator) that becomes the actual initiating free radical. In addition, the Norrish Type II reactions are intra molecular, so Type II photo initiators are more easily affected by the environment such as oxygen and viscosity (Lee et al. 2006) The majority of Type I phot o initiators are aromatic carbonyl compounds cleavage. The most important fragmentation in photo cleavage of the carbon carbon bond between the carbonyl group and the alkyl residue in alkyl aryl ketones ( Figure 1.7). 1 phenyl 2 hydroxy 2 methyl propan 1 one (Benacure 1173) is a very popular photo initiator (structure shown in Table 1) because it is highly reactive, thermally stable, does not yellow and can be used in water based formulations (Crivello and Dietliker, 1998 )
21 Figure 1.5 Various steps in a photoinitiated p olymerization
22 Type I Photoinitiator: unimolecular fragmentation Type II photoinitiator: bimolecular fragmentation Figure 1.6 Types of photoinitiator ( Crivello and Dietliker, 1998 )
23 Figure Cleavage of alkylaryl k etones (Crivello and Dietliker, 1998 ) Prev ious investigations into the biocompatibility of PHEMA and PDHPMA as homo and copolymers by UV polymerization showed that hydrogels containing 80%DHPMA and 20%HEMA were found to induce minimal to no fibrosis when implanted subcutaneously in rats (Mohomed et al. 2006) Unfortunately, as it sorbed large amount s of water the mechanical stability of the high content DHPMA copolymers (80%DHPMA:20%HEMA) and the PDHPMA homopolymer decreased and the samples were easily fragmented. The biocompatible hydrogels w ere also coated onto the polyurethane (PU)/epoxy coated metal sensor by dip coating, or in situ polymerization It was noted that the PHEMA coating easily delaminated from the PU/epoxy coating once swollen in water ( Mohomed, 2006) To improve the lifespa n of glucose sensors, the objectives of hydrogel coatings are:
24 1) the coating must allow glucose, oxygen and hydrogen peroxide to diffuse freely, while causing minimal fibrosis and inflammation. 2) it should be easily coated onto the surface of implantable glucose sensors 3) it should attach firmly and permanently to the glucose sensors and strong enough not to break during the implantation period. 4) it must not deter the performance of the implantable glucose sensors. 5 ) it should deliver anti inflammator y drug and release drug to suppress both acute and chronic inflammatory. According to previous work by Mohomed the formulations of hydrogel have based on HEMA and DHPMA copolymer for its excellent biocompatibility, but well modified with another monomer to improve the coating properties.
25 CHAPTER 2 SYNTHESIS AND PEFORMANCE OF NOVEL HYDROGELS COATING FOR IMPLANTABLE GLUCOSE SENSORS 2.1. Introduction State of the art implantable glucose sensors do not work reliably and have a short life after i mplantation (Moussy 2002 ) This in vivo loss of function is caused by tissue reactions surrounding the sensor such as fibrosis and inflammation ( Mang et al. 2005 ) In order to improve the lifetime of implantable glucose sensors, the biocompatibility of the coatings must be monitored ( Gerritsen et al. 2000 ) Many biocompatible coatings have been prepared for sensors such as hydrogels ( Praveen et al. 2003 ; Galeska et al. 2001 ; JP 60171140 ; USP 5786439 ) silica based hybrid materials ( Kros et al. 2001 ) sol gel coatings ( Gerrit sen et al. 2000 ) diamond like carbon films and diamond like carbon coated anodized aluminum oxide nanoporous membranes ( Narayan et al. 2007 ) All of these coatings showed some potential for use in sensors, but formulations need to be optimized at this tim e. The outer interfacial coatings of implantable glucose sensors must be stable in biological fluids, ensure the transport of glucose to the sensors, resist the deposition of the proteins and other interferents and minimize fibrous encapsulation. Three di mensional cross linked polymer networks known as hydrogels have been shown to
2 6 meet these requirements and have the added benefit of swelling in water without dissolving ( Praveen et al. 2003 ; Kros et al. 2001 ; Gerritsen et al. 2000 ) Poly(hydroxyethyl metha crylate) (PHEMA) hydrogels were first polymerized and studied for biological use by Wichterle and Lim in the early 1960s ( Wichterle and Lim, 1960 ; Wichterle and Lim, 1961 ) Since they are biocompatible, PHEMA hydrogels have been used in biomedical and pharm aceutical applications such as contact lenses, implants and drug delivery systems. PHEMA was first used in glucose sensors by John Christopher, et al in 1990 ( Pickup and Claremont, 1990 ) Sensors coated with PHEMA were stable for 24hrs in vitro at 37 However, much longer life times are required. PHEMA hydrogels are clear, exhibit smooth surfaces and are biocompatible ( Rozakis et al. 2005 ) Drawbacks of using PHEMA gels are that their mechanical properties are poor and equlibrium water contents are relatively low, <45 wt% ( Lee and Lin, 2003 ) ; water uptakes of 200 wt% are needed for optimum operation ( Van Antwerp et al. 1998 ) Mechanical properties can be increased by preparing PHEMA hydrogels with ethylene glycol dimethacrylate cross linker (EGDMA). Equilibrium water contents can be increased by preparing PHEMA hydrogels with more hydrophilic monomers such as 2, 3 dihydroxypropyl methacrylate (DHPMA) ( Yasuda et al. 1966 ; Macret and Hild, 1982 ) or N vinyl pyrolidinone ( VP ) (Laporte, 1997) According to p revious investigations of hydrogel (80%DHPMA:20%HEMA) coating for implantable glucose sensors, it was found that the biocompatibility and equilibrium water content are improved (Mohomed. 2006)
27 Unfortunately, as it sorbed large amount of water the mechan ical stability of the high content DHPMA copolymers (80%DHPMA: 20%HEMA) decreased and the samples were easily fragmented. The biocompatible hydrogels were also coated onto the polyurethane (PU)/epoxy coated metal sensor by dip coating, or in situ polymeri zation. It was noted that hydrogel coatings easily delaminated from the PU/epoxy coating once swollen in water (Mohomed, 2006) Since P VP was first used as a blood plasma substitute and extender during World War II (Robinson et al. 1990) VP can be copolymerized with other materi als to modify properties for various medical applications (Laporte, 1997) VP can enhance the water content, mechanical properties and porous structures of hydrogels at the same time, behavior unusual for a hydrogel. High VP percent hydrogels show less bi ocompatibility, but 50% of P VP hydrogels are biocompatible (Lopes and Felisberti, 2003; Risbud et al. 2000) Fu r thermore, P VP can be coated on the polyurethane surface and firmly attach to the surface (Francois et al. 1996) This property is essential for this project. The goal of this study was to develo p a novel hydrogel copolymer with HEMA, DHPMA, VP and EGDMA which was facile to coat on the glucose sensors, did not deter the detection of glucose and exhibits improved biocompatibility. In order to optimize the compositions of hydrogels swelling, SEM and water structure of hydrogels were characterized. Different formulations hydrogels were coated on the implantable glucose sensors by UV curing. Glucose sensors used for this study were
28 based on the long term excess enzyme loading coil type implantable ampe rometric sensors ( Yu et al. 2005 ; Yu et al. 2006 ) In vitro and in vivo evaluations of the hydrogel coated glucose sensors with and without PU were also undertaken. 2.2 Experimental Procedures 2.2.1 Materials. 2 Hydroxyethyl methacrylate (HEMA) and 2, 3 di hydroxypropyl methacrylate (DHPMA) were donated by Benz R & D (Sarasota, FL USA). They were used as received without further purification. N Vinyl 2 Pyrrolidinone ( VP ) (99.9+%) was purchased from Sigma Aldrich Co.( St. Louis, MO, USA) and purified by vacuum distillation to obtain a colorless liquid according to reference (Jansen e t al. 2005) Ethylene glycol dimethacrylate (EGDMA), glucose oxidase ( GOX ) (EC 184.108.40.206, Type X S, Aspergillus Niger, 157,500U/g), benzoic acid, gentamicin, ATACS5104 epoxy adhesive, Fluka Selectophore 81367 polyurethane (PU) and tetrahydrofunan (THF) were obtained from Sigma Aldrich Co.(St.Louis, MO, USA). 1 phenyl 2 hydroxy 2 methyl propan 1 one (Benacure 1173) from Mayzon Corporation (Rochester, NY, USA) was used as received. Dextrose (D glucose), bovine serum albumin (BSA), and glutaraldehyde (50%) wer e obtained from Fisher Scientific (Pittsburgh, PA). 0.125 mm Teflon covered platinum iridium (9:1 in weight) and silver wires were obtained from World Precision Instruments, Inc. (Sarasota, FL USA). High temperature mini round glue sticks were supplied b y Adhesive Tech (Hampton, NH, USA). Hematoxylin and eosin were obtained from
29 Fisher Scientific (Pittsburgh, PA) and trichrome stains were purchased from Newcomer Supply (Middleton, WI). 2.2.2 Synthesis of Poly (HEMA DHPMA VP EGDMA) Hydrogel Series Various molar ratios of HEMA and DHPMA were weighed and mixed in 20 ml glass vials. 9 mol % of VP 1 mole% EGDMA and 1.5 mol% initiator Benacure 1173 were added to the solutions (Table 2.1). The solutions were diluted 1:1 with distilled water. After deoxygenatio n with argon gas, the solutions were injected into hand made glass cells measuring 55mm 25mm 1mm. The sides of the cells were secured with a glue gun. The solution filled cells were placed under a UV lamp (Spectro nics Corporation, Westbury, NY, USA) at a wavelength of 254nm for 100min. Argon gas was used to purge the system in order to minimize oxygen inhibition of the reaction. Transparent uniform sheets were obtained and these xerogels were swollen in distilled water to yield hydrogels. The water was refreshed every day for one week to remove any unreacted monomers. The hydrogels were freeze dried for various studies using a Virtis Freezemobile 12XL freeze dryer (The Virtis Company Inc, Gardiner, NY, USA) to ke ep the pores inner morphology.
30 Table 2.1 Feed c ompositions of s ynthesized x erogels Mole% HEMA DHPMA VP EGDMA DHPMA0 90 0 9 1 DHPMA20 70 20 9 1 DHPMA40 50 40 9 1 DHPMA60 30 60 9 1 DHPMA80 10 80 9 1 DHPMA90 0 90 9 1 2.2.3 Scanning Electron Micro scopy (SEM) The hydrogel samples obtained by the above method were cut and freeze dried to determine the inner morphology. The morphology of the freeze dried gels were studied via a Hitachi Scanning Electron Microscope S 800 ( Hitachi High Technologies, Pl easanton, CA) with 20nm gold coating by a HummerX sputter coater (Anatech, Ltd, Springfield, VA). The working distance between the sample and electron (WD) was 5mm with 25.0 kV. 2.2.4 Sorption Experiments Hydrogel samples were cut with a 12mm diameter cor k bore to obtain uniform shapes. All sheets were freeze dried for one day and stored in a desiccator at room temperature. The sorption behavior of hydrogels was monitored by detecting the increase in mass of the samples at different time intervals by Sar torius BP211D balance ( 0.01mg, Sartorius Corporation, Edgewood,NY, USA). In a typical sorption experiment, a pre weighed dry gel sheet was immersed into water at 241
31 in a Fisher Scientific Isotemp water bath (Pittsburgh, PA). At a prescribed time intervals the hydrogel was taken out of solution and weighed after wiping off the excess water from the surface with Kimwipe paper (Kimberly Clark Professional). The sorpti on degree, SD, of hydrogels was defined as follows: SD%= 100= (2 1) Where is the weight of the dry gel, is the weight of w et hydrogel at each time interval, and is the gain in the weight of the dry gel at time t. The equilibrium water content (EWC) of the hydrogel was determined using the following equation: EWC%= 100= (2 2) where and are the weight and weight gain of the swollen hydrogel at equilibrium separately. 2.2.5 Differential Scanning Calorimetry (DSC) All calorimet ric data were obtained via a TA Instruments 2920 differential scanning calorimeter (DSC, TA Instruments, New Castle, Delaware). Nitrogen gas was passed through the instrument at a flow rate of 70ml/min. Before measurement the DSC was calibrated from 100 to 250 at a heating rate of 5 /min with an indium standard. All the sample masses for DSC ranged from 4mg to 10mg. For the determination of glass transition temperature, samples were freeze dried for one day and kept in a desiccator. The samples were q uickly weighed, sealed in aluminum pans, and immediately scanned from 30 to 200
32 Samples were cooled and rescanned; the glass transition was determined from the second heating cycle in order to minimize any aging effects. The reported glass transition te mperature was determined as midpoint at half heat flow or heat capacity. To measure the water structure of equilibrium swollen hydrogels, first samples were cut and weighed accurately after wiping off the surface water with Kimwipe paper and immediately s ealed in Al pans. The DSC curves were measured heating from 100 to 20 with a heating rate 5 /min. Endotherm areas were determined. 2.2.6 Sensor Preparation Hydrogel coated glucose sensors used for this study were based on the coil type implantable sensor previously described ( Yu et al. 2005 ; Yu et al. 2006 ) The detail fabrication method of Pt/GOx and Pt/GOx/Epoxy polyurethane base sensors can be found in those papers. The coating process was achieved by applying a ~ 0.4 L of 50% hydrogel solution to either the enzyme layer or the epoxy polyurethane layer, using a 10 L micropipette, and then curing under an argon atmosphere and a 254 nm UV light for 100 min. Membrane surfaces were observed and photographed using an Olympus BX41 microscope (Quantitative Imaging Co., Canada). The cured hydrogel coating was trans parent and firmly attached to the sensor as shown in Figure 2.1. 2.2.7 In vitro Evaluation Method Electrochemical measurements were performed with Apollo 4000 4 Channel Potentiostats (World Precision Instruments, Inc., Sarasota, FL). Newly prepared
33 glucos e biosensors were conditioned in a 5mM glucose /PBS (NaCl 8.76g, KH 2 PO 4 3.53g, Na 2 HPO 4 3.40g, Benzoic acid 2g, water 10000ml) (ionic strength = ~ 0.16M) for at least 2 hours and then continuously polarized at +0.7V vs. Ag/AgCl until a stable background cur rent was reached. The response was determined as the time 90% min) the maximum current when the glucose concentration changed from 5 mM to 15 mM. The sensitivity (S) was determined using a two point meth od and can be calculated by S (nA/mM) = (3) where I 15 mM and I 5 mM represent the sensor response currents obtained in 15 mM and 5 mM glucose solutions, respectively. The sensors for long term observation were stor ed in PBS. Calibration plots were obtained by stepwise adding 100 mM glucose solutions to 8.0 mL PBS. All measurements were performed at 24 1C 2.2.8 In vivo Biocompatibility Studies The epoxy PU coated Silastic tubing (Dow Corning, Midland, MI) ( = ~ 4 mm, length = ~ 50 mm), a DHPMA 80 bar ( = ~5 mm, length = ~50 mm), an epoxy PU coated glucose sensor and a DHPMA 80 coated glucose sensor were implanted subcutaneously in rats for 28 days. The tissues surrounding the implants were removed at day 28 and p rocessed for histology with hematoxylin, eosin and trichrome stains (Galeska et al. 2001)
34 Figure 2.1 S ensor photos show thin layers of hydrogel coating: A. hydrogel on Pt/GOx sensor B. hydrogel on Pt/GOx sensor after swell ing C. hydrogel on Pt/GOx/Epoxy polyurethane base sensors D. hydrogel on Pt/GOx/Epoxy polyurethane sensors after swelling 1 mm A B C D
35 2.3 Results and Discussion 2.3.1 Preparation of Hydrogels Sheets and Coatings A type I photoinitiator, Benacure 1173, undergoe s unimolecular bond cleavage upon irradiation in argon atmosphere. It produced well cured samples, including the thin layers used on the sensors ( Lewis et al. 1975 ; Maillard et al. 1983 ) The water present in the polymerization mixture minimized the heat of polymerization and bubble free samples were obtained. The amount of solvent in the reaction mixture did not exceed the critical solvent conc entration which would result in heterogenous samples ( Peppas, 1986 ) In order to avoid producing bubbles and ob tain the homogenous transparent hydrogel, monomers were dissolved in water with the weight ratio around 1:1. Even if there was a little heat produced, it would diffuse readily, and would not yield big bubbles inside the hydrogel. In this work, homogenous hydrogel sheets and coatings of implantable glucose sensor were obtained by UV cured. 2.3.2 Inner Sample Morphology The scanning electron micrographs shown in Figure 2.2 depict the dramatic effect of DHPMA on morphology. The sample with no DHPMA exhibits irregular pores. At 20% DHPMA the number and regularity of the pores increase. This trend increases up to 80% DHPMA. At 90% DHPMA the structure is less regular, possibly due to the high swelling stresses which can induce tearing in the sample. Coatings for implantable glucose sensors must effectively transport water, glucose
36 Figure 2.2 SEM micrographs of inner gels ( 350 magnification): D 0 ,D 20 D 40 D 60 D 80 and D 90
37 and oxygen. From the morphology of different gels it can be seen that 80% DHPMA gives the most uniform, porous structure and therefore, might be an optimum candidate for permeable sensor coatings. 2.3.3 Sorption Behavior The sorption versus time profile is shown in figure 2.3. Both the equilibrium water conte nt and the sorption rate increase with DHPMA content. This is due to the extra hydroxy group in DHPMA which increases the hydrophilicity. It has also been shown that increasing DHPMA content can slow the dehydration rate and increase the rehydration rate of the sample ( Gates et al. 2003 ; Pescosolido et al. 199 0 ) This is ideal for the implantable glucose sensors, since dehydration reduces biocompatibility ( McConville and Pope, 2001 ) It has been stated that glucose sensor coatings should exhibit the equ ilibrium sorption degrees of over 120%, and, preferably over 200% ( Van Antwerp et al. 1998 ) Figure 2.3 demonstrates that the 80% DHPMA, again, is an optimum formulation for sensors. The water diffusion coefficient can be easily calculated when diffusion controls the transport process. It is well known that sorption behavior in polymers is quite complex ( Harmon et la. 1988 ; Harmon et al. 1987 ) At one extreme, Fickian diffusion controls kinetics. However, since swelling involves segmental relaxation, th e relaxation rate can influence transport. If the relaxation rate is slower than the diffusion rate, relaxation can dominate transport kinetics. In reality, sorption behavior can range from pure Fickian behavior to pure relaxation behavior. In pure
38 Fick ian behavior, sorption versus the square root of time curves are linear up until about 60% of the equilibrium water content is imbibed ( dominates transport, sorption is linear with respect to time. The extent of influence from Fickian diffusion and relaxation is characterized by the time exponent in the following equation ( Franson and Peppas, 1983 ; Alfrey et al. 1966 ) : = Kt n (2 4) Where, t is time, k is a constant and n is the time expone nt. The time exponent, n, varies from 0.5 for pure Fickian control to 1 for pure relaxation control. Table 2.2 and Figure 2.4 show that transport kinetics are controlled by diffusion in these hydrogel systems, since n is close to 0.5 (Wang and Wu, 2005)
39 Figure 2.3 The swelling degree as a function of time for the different DHPMA fraction hydroge ls Figure 2.4 log versus logTime for the different DHPMA fraction hydroge ls
40 Table 2.2 Dynamic Swelling Parameters Sample n R 2 D(cm 2 S 1 )*10 5 DHPMA0 0.57 0.072 0.569 0.993 0.5 DHPMA20 0.52 0.094 0.557 0.994 1.6 DHPMA40 0.47 0.111 0.554 0.996 2.0 DHPMA60 0.45 0.135 0.532 0.993 2.4 DHPMA80 0.41 0.263 0.589 0.991 3.9 DHPMA90 0.32 0.343 0.568 0.994 4.5 A ll sample shapes are similar and the ratio of thickness (L) to diameter (D i ) is about 1:10. It is assumed that water diffuses into a planar sheet at a constant rate. When and time w ill fit equation 5 ( Crank 1978 ) = ( ) 0.5 (2 5) The diffusion coefficient (D) values are estimated according to equation (5) and shown in Table 2.2. From Table 2.2 it can be seen that (D) increases with DHPMA content. All the swelling kinetics data show that when DHPMA content is increased the swelling ratio and water diffusion rate increased. Since water acts as the diffusion medium between body f luid and the glucose sensor the faster the water transport characteristics of the hydrogel coating the better the detection of the glucose concentration.
41 2.3.4 DSC Results Table 2.3 shows that the T g s of different xerogels composition do vary with DHPMA c ontent. This trend was noted this effect in an earlier study ( Gates et al. 2003 ) The dihydroxypropyl side chain in DHPMA imparts greater free volume, which can facilitate the mobility and decrease the T g However, this effect is offset by the extra hyd roxyl group which forms hydrogen bonds in the network structures and restricts mobility Table 2.3 T g s of different composition gels DHPMA% Tg( )(H) (1) 0 118 (2) 20 119 (3) 40 120 (4) 60 117 (5) 80 117 (6) 90 120 It is well known that the biocompatibility of hydrogels not only depends on the equilibrium water content but also on the state of water in the hydrogels (Kim et al.,1980;Mirejo vsky et al. 1991; GOx a et al. 2006;Jhon and Andrade, 2004) Som e techniques employed to study water structure in hydrogels are NMR ( Chowdhury et al. 2004 ; Capitani et al. 2001 ) dilatometry and electrical conductivity ( Lee et al. 1975 ; Choi et al. 1977 ) d ielectric relaxation spectroscopy ( Kyritsis et al. 1995 )
42 dynamic mechanical spectroscopy ( Lustig et al. 1991 ) and differential scanning calorimetry ( Murphy et al. 1988 ; Pedley and Tighe, 1979 ) The classification of water in the hydrogels depends on the techniques used. According to the most common method, DSC, water in the hydrogel is classified as freezing free water, freezing bound water and non freezing water. Free water does not take part in hydrogen bonding with polymer molecules. It has a simila r transition temperature, enthalpy and DSC curves as pure water ( Nakamura et al. 1983 ) Freezing bound water interacts weakly with polymer molecules and has a phase transition temperature lower than 273K. Non freezing bound water is complexed with the po lymer chain through hydrogen bounds and has no detectable phase transition over the temperature range from 200K to 273K ( Higuchi et al. 1984 ) Since the freezing water effects transport and biocompatibility of the hydrogels ( Mirejovsky et al. 1991 ; God a e t al. 2006 ) it is important to detect water structure inside hydrogel for optimization of material properties in biomedical applications. Table 2.4 Water stru ctures of different composition hydrogels Sample EWC% W f % W nf % W f /W nf W f /EWC% D0 48.8 21.6 27. 2 0.79 44.3 D20 52.0 30.0 22.0 1.36 57.7 D40 60.1 39.9 20.2 1.98 66.4 D60 65.5 50.7 14.8 3.43 77.4 D80 69.4 56.2 13.2 4.26 81.0 D90 69.9 41.2 28.7 1.44 41.1
43 In this work the water structure in equilibrium swollen hydrogel was studied by DSC. Since t he peak shape and the maximum peak temperature of heating and cooling process are dependent on the heating rate ( Higuchi and Iijima,1985 ; Chan et al. 1992 ) and different heating and cooling cycles can cause the evaporation of water from the hydrogel and co ndensation inside the sample pan ( Roorda et al. 1988 ) only the results of the first heating runs were used to calculate relative content of freezing water (W f %) and the relative content of non freezing water (W nf %). The W f % was calculated by the melting en thalpy of the hydrogel divided by the melting enthalpy of pure water. The observed melting enthalpy is calculated by the area under the endothermic curve. The melting enthalpy of bulk water is taken as 333.3J/g ( Ahmad et al. 1994 ) W f % and W nf % are cal culated by the following equation: W f %=W free %+W f bond %= (2 6) W nf %=EWC W f % J/g is the observed melting enthalpy. is the melting enthalpy of b ulk water. The relative content of non freezing water is obtained by the difference between the equilibrium water content, EWC, and the relative content of freezing water. The EWC, W f % and W nf % are listed in Table 2.4. Figure 2.5 shows the DSC curves obtain ed during heating run from 100 to 20 From Table 2.4 we can see that when the DHPMA content was increased from 0 to 80% the content of freezing water increased and non freezing water decreased.
44 The ratio of freezing water to non freezing water reached a peak at D80. The freezing water increase may be due to greater free volume because of the dihydroxypropyl side chain while the non freezing water decrease may be due to more hydrogen bonds formed between the side chains, which results in less hydrogen b ond formed between water and hydroxyl groups. According to reference ( Liu et al. 2000 ) hydrogel porosity increased with NaCl concentration in the swelling medium. This leads to an increase in freezing water content and a decrease in non freezing water co ntent and was thought to stem from the increase in porosity, as seen in SEM pictures, and DHPMA content. However, hydrogels without HEMA showed a decrease in freezing water and increase in non freezing water. From the results it can be seen that 80% DHPMA shows the highest freezing water content. Since the content of freezing water determines the transport properties of the hydrogel ( Mirejovsky et al. 1991 ) it can be expected that 80% DHPMA is the best for transport of water, glucose and oxygen to the bio sensor according to reference ( G oda et al. 2006 ) higher freezing water content also results in less protein deposition on the polymer surface. Future studies will characterize the effect of swelling and water structure on mechanical relaxations. Previous publications documented the fact that hydrophilic groups on the polymer backbone influence secondary relaxations and swelling responses ( Gates et al. 2003 )
45 Figure 2.5 DSC heating curves of hydrogels with different DHPMA content, performed at the heating rate of 5.0 min 1
46 2.3.5 Sensor Performance of Hydrogel Coated Glucose Sensors Hydrogels of various compositions were used for coating glucose sensors and the characteristics of the corresponding sensors were examined in vitro. Table 5 gives the measured results of sensitivity (S, nA/mM) and response time (T 90% min) at Day 7 and Day 28. For different compositions of hydrogels a higher percentage of DHPMA resulted in the higher sensitivity and shorter response time in accordance with the results of SEM m orphology, swelling behavior and water structure observations. Among five hydrogels, D0 showed the best long term stability but its response was too slow. D20 to D60 failed to gave a stable sensor response at day 28. Comparatively, the sensor with a D80 c oating showed the best performance. At day 28, the sensor response curve with stepwise glucose increases was still good (Figure 2.6) despite a 40% loss of the sensitivity at day 28. Table 2.5 Response s ensitivities (S) and t imes (T 90% ) of various hydroge l (HG) based glucose sensors (mean SD) Pt Ir/GOx/HG (n=4) D0 D20 D40 D60 D80 Day 7 S (nA/mM) 6.12.6 6.32.2 10.51.7 16.20.8 18.42.1 T 90% (min) 12.72.2 10.52.0 8.61.2 6.91.3 4.61.0 Day 28 S (nA/mM) 3.90.4 Low response 10.9 2.1 T 90% (min) 3.51.5 4.31.2
47 Figure 2.6 Representative response curves of a Pt Ir/GOx/HG D80 glucose sensor at day 7 and 28, respectively.
48 Table 2.6 Response characteristics of various hydrogel EPU based glucose sensors* (mean SD) Pt Ir/GOx/EPU/HG (n=4) Control D0 D20 D40 D60 D80 Day 7 S (nA/mM) 3.31.1 0.60.4 0.60.4 1.20.8 3.01.4 3.20.8 T 90% (min) 2.91.0 27.03.4 24.32.7 18.47.4 17.42.2 9.11.4 *S: response sensitivity of sensor; T 90% (min): th e time to reach 90% of the maximum response when glucose varies from 5 mM to 15 mM; Figure 2.7 Calibration plots from a Pt Ir/GOx/HG D80 sensor and a Pt Ir/GOx/HG D80 sensor. Sampling time was between 4 to 5 min after each stepwise addition of a concen trated glucose solution.
49 Figure 2.7 shows the calibration lines of a Pt Ir/GOx/HG D80 sensor and a Pt Ir/GOx/HG D80 sensor. The linear range of Pt Ir/GOx/HG sensors was in the range of 1 to 15 mM. It was thought that the poor linearity at high concentr ation could be significantly improved by adding an epoxy polyurethane membrane in between the enzyme layer and the hydrogel coating. Such a sensor, i.e. Pt Ir/GOx/EPU/HG, extended the linear range to 40 mM (see Figure 6). Table 2.6 summarized the sensitivi ty and response time of various Pt Ir/GOx/EPU/HG sensors. Since the permeability of the epoxy polyurethane membrane is much lower than that of the hydrogel coating layer, the epoxy polyurethane membrane played the major role in limiting diffusion. Therefo re, when the hydrogel was more permeable, the sensitivity was closer to that of the control sensors (i.e. Pt Ir/GOx/EPU). However, the additional hydrogel layer caused a significant response delay, especially for D0, D20, D40 and D60. D80 was determined t o meet the design requirement of implantable glucose sensors. 2.3.6 In vivo B iocompatibility PU coated Silastic tube ( = ~ 4 mm, length = ~ 50 mm), 80% DHPMA hydrogel bar ( = ~ 4 mm, length = ~ 50 mm), PU coated glucose sensor and 80% DHPMA hydrogel coa ted glucose sensor were implanted subcutaneously in rats for 28 days. Figure 2.8 shows the histology results of the PU coated Silastic tube (A), hydrogel bar (B), PU coated glucose sensor (C) and hydrogel coated glucose sensor (D). It can be seen that abou t 200 m and more than 500m in thickness fibrous
50 capsule formed around the silastic tubing and PU coated glucose sensor with obvious inflammation of the cells. While only 30~ 60 m in thickness, the fibrous capsule formed at the hydrogel bar and hydrogel coated sensor, respectively. 80% DHPMA hydrogel can reduce the fibrosis and inflammation and it should make the lifetime of glucose sensor longer. A sensor with a hydrogel coating would not only have all advantages of an epoxy polyurethane glucose senso r (e.g. long term stable, wide response range) but also minimize the tissue reactions occurring around the implanted sensor. This study was limited to 28 days. There is no reason to expect that the biocompatibility of this material would change after thi s period. As previously stated in the introduction, the main innovation of this hydrogel compared to similar materials is the increase in mechanical strength which makes it suitable to coat biosensors for long term implantation.
51 Figure 2.8 Histology slides: (A) epoxy polyurethane coated silastic tubing ( = ~ 4 mm, length = ~ 50 mm), (B) one hydrogel bar ( = ~ 4 mm, length = ~ 50 mm), (C) one PU coated sensor (D) one hydrogel coated sensor (rat 4, left) implanted subcutaneously in rats for 28 days. 100um m A B C D
52 2.4 Conclusion In this study different hydrogels with various DHPMA content were successfully synthesized and coated on to glucose sensors by UV polymerization. The higher freezable water content, swelling rate and uniform porosity that result from high DHPMA content increase the sensitivity and shorten the response time of the glucose sensors. The linear range of a glucose sensor only coated with hydrogel is short. However, it can be improved by coating the epoxy P U with a layer of hydrogel because PU can control the diffusion of glucose molecules. Since the hydrogel minimizes the fibrosis and inflammation, it shows promise for use in implantable glucose sensors.
53 CHAPTER 3 USE OF HYDROGEL COATING TO IM PROVE THE PERFORMANCE OF IMPLANTED GLUCOSE SENSORS Foreword: 8 D80 coated Pt/GOx/epoxy polyurethane glucose sensors were implanted into 4 rats (2 sensors/rat). The original attempt of in vivo implanted glucose sensors was not successful, because only 25% o f the sensors were working at 4 weeks. In order to further improve the efficiency of implanted glucose sensors, another series of novel hydrogels were developed based on copolymer of HEMA and DHPMA. 3.1. Introduction The material tissue interaction during sensor implantation, i.e. so called biofouling which may include protein/platelet deposition, or the attachment of inflammatory response related cells, is one of the major causes in unpredictable and unexplainable behaviors of implanted glucose biosensors The performance of implanted biosensors can greatly benefit from the use of more biocompatible outer most coatings such as hydrophilic alginate (Shichiri et al., 1989) phosphorylcholine modified polyurethane (Yang et al.,2000), 2 methacryloyloxyethyl phosphorylcholine co n butyl methacrylate (Nishida et al., 1995; Yasuzawa et al., 2000) and hydrogels (Suri et al., 2003). Hemocompatible synthetic copolymers containing phosphorylch oline significantly improved the performance of the sensors both ex vi vo in whole blood and in vivo. Unfortunately, these polymers are mechanically weak and can be easily dislodged in flowing blood or in the subcutaneous tissue and are thereby are not
54 useful for long term application. Hydrogels are a kind of network polyme r possessing a degree of flexibility very similar to natural tissue due to their significant water content. When in physiological solutions, hydrogels can reduce protein adsorption and the subsequent inflammatory response. In the medical field, hydrogels have been extensively used as scaffolds in tissue engineering and sustained release drug delivery systems (Norton et al., 2005) The properties of some hydrogels make them attractive candidates for us e in glucose sensors. A glucose sensitive phenylboroni c acid based hydrogel used pressure stimuli to detect glucose (Lei et al., 2006) Similarly, a polyacrylamide hydrogel coated cantilever sensor was constructed and responded to swelling which directly correlated with glucose concentration (Ji et al., 2005 ) Glucose oxidase was immobilized in polyacrylamide gels via crosslinking. Swelling and viscoelastic properties were used along with amperometric measurements to evaluate the materials for use in biosensors (Fernandez and Lopez, 2005) Phenylboronic ac id containing hydrogels were shown to exhibit red shifts in holograms which quantify glucose concentrations (Kabilan et al., 2005) Another group of researchers used photonic properties of boronic acid based hydrogels to sense glucose (Alexeev et al., 2004 ) Berner et al. (1998) determined that the iontophoretic properties of hydrogels can be used to monitor glucose. Certain hydrogels can be doped with glucose reactive fluorescent dyes and the light that is emitted changes with glucose concentration (Thoni yot et al., 2006) Hydrogels can potentially be used as the outermost coating of implanted
55 glucose sensors to reduce foreign body reaction surrounding the sensor, thereby improving the in vivo sensor performance. Polyethylene oxide based hydrogels were use d as diffusion limiting membranes in glucose sensors and short term studies were implemented by subcutaneously implanting hydrogel coated glucose sensors in Sprague Dawley rats (Quinn et al., 1997, 1995). Most hydrogels like pHEMA are too weak to maintain their membrane structure and function during implantation, therefore, long term studies are difficult. Hydrogels include neutral hydrogels, ionic hydrogels and swollen interpenetrating polymeric networks. However, neutral hydrogels can better meet the sens or coating requirements of biocompatibility and structural stability. A neutral hydrogel, P HEMA and DHPMA (Wichterle and Lim, 1960) was previously developed for contact lenses. This gel was shown to exhibit an increased resistance to dehydration (Gates et al., 2003). The copolymer exhibited high equilibrium water content and good mechanical properties. These copolymers have great potential for producing more durable coatings for glucose sensors. In this study, novel HEMA DHPMA based hydrogels were develope d and used to improve the biocompatibility of implantable glucose biosensors. The porosity and water content of the copolymer were improved by the addition of a structural strengthener, N vinyl 2 pyrrolidinone ( VP ; McConville and Pope, 2001). The compositi ons of monomer mixtures were optimized and the resulting hydrogels were characterized. A selected hydrogel was used for coating an implantable glucose sensor
56 (Yu et al., 2006). The in vitro and in vivo sensor performance of hydrogel coated sensors were exa mined and discussed. 3.2 Experimental 3.2.1. Chemicals and M aterials Monomers HEMA and DHPMA were donated by Benz R & D (Sarasota, FL). Other chemicals including VP EGDMA and Benacure 1173 were obtained from Sigma Aldrich (St. Louis, MO). VP was purifi ed by vacuum distillation prior to polymerization according to reference (Jansen et al. 2005) 0.020 in ch Silastic tubing (i.d. o.d. = 0.51mm 0.94 mm) was purchased from Dow Corning (Midland, MI). Size 3 0 Prolene (polypropylene) were obtained from Ethic on (Somerville, NJ). I.V. catheters (14 ga.) with a needle were obtained from Terumo Medical Corporation (Somerset, NJ). Sprague Dawley outbred rats (male, 375 399 g) were purchased from Harlan (Dublin, VA). VetBondTM glue came from 3M (St. Paul, MN). 3 .2.2. Photopolymerization of P oly(HEMA DHPMA VP EGDMA) H ydrogels Pure monomers (all liquid state) were first combined to prepare the following four formulations (termed VP 0, VP 15, VP 30 and VP 45) listed in Table 3.1. The mixtures were diluted 1:1 with deio nized water, 1% (w/w) Photoinitiator (Benacure 1173) was added, and the solutions mixed at room temperature. Polymerization was performed as follows. After purging each solution with argon for 2 min, each solution was transferred into a glass mold (L W H =55mm 25mm 1 mm). The
57 molds were placed under a 254 nm UV light for 100 110 min in an argon atmosphere at 25 C to polymerize the hydrogels. The hydrogels were removed from the molds and placed into 500 ml of deionized water at 25 C. The water was changed daily for one week to remove soluble residues. Table3.1 Feed composition of synthesized hydrogel Molar ratio VP 1HEMA:1DHPMA EGDMA VP 0 0 99 1 VP 15 15 84 1 VP 30 30 69 1 VP 45 45 54 1 For formation of the hydrogel coating on the glucose sensor tip, 0.4, 0.8 or 1.2 l of the argon purged VP 30 solution (including additional water and photo initiator) was pi petted onto the sensor tip. The solution was allowed to spread evenly over the surface of the tip (bare or EPU coated). After exposing to UV light under argon (as above), the sensors were stored in phosphate buffered saline until further use. 3.2.3. Ana lysis of H ydrogels The purified hydrogel samples were cut and freeze dried for 24 h, and then stored in dessicator at room temperature. The inner morphology of cutting section was examined using an S 800 Hitachi Scanning Electron Microscopy System (Hitach i High Technologies, Pleasanton, CA).
58 The glass transition temperature ( T g ) was measured using a model 2920 differential scanning calorimeter (DSC, TA Instruments, New Castle, DE) under a pressurized nitrogen atmosphere to avoid sample degradation. In thi s process, a 4 10mg freeze dried sample was sealed in aluminum pans. The samples were scanned from 30 to 200 C at 5 C per min. Each sample was tested at least twice. Fourier transform infrared spectroscopy (FT IR) was carried out on the rectangular samples that were dessicated at room temperature after freeze dried, using a Nicolet Avatar 320 FT IR Spectrophoto meter (Nicolet, Madison, WI) with a 64 scan per sample cycle at a resolution of 4 cm 1. For the determination of equilibrium water fraction ( EWF ), hydrogel samples were lyophilized to obtain dry mass ( M d ). Xerogels were then swollen in a pH 7.4 phosphat e buffer solution (PBS, ionic strength = ~ 0.16 M) at 37 C for 2 days. The swollen hydrogels were pat dried with filter paper and weighed every 2 h until the weight reached a constant value ( M eq). The water Fraction ( EWF ) at swelling equilibrium was calcula ted by: EWF (%) = ( M eq M d) 100/ M d (3 1) 3.2.4. Sensor P reparation Glucose sensors were based on coil type design (Yuet al., 2006)(Figure3.1). Briefly, the coil type glucose biosensors were prepared by winding the top 10mm of a 40 50mm long platinum wi re ( 0.125mm, including 10% iridium in weight) along a 30 gauge needle to form a coil like cylinder which was filled with cotton. The enzyme layer was formed over the cylinder by coating an aqueous solution of
59 10mg/ml GO x 30 40mg/ml BSA and 0.6% (v/v) gl utaraldehyde. In this study, the novel membrane configurations were used on Pt/ GO x or Pt/ GO x /EPU sensors. The outermost hydrogel coating was formed by applying a 0.4 solution to the Pt/ GO x or Pt/ GO x /EPU sensors as described above. All sensors used in this study were visually inspected for defects under 40 magnification using a dissection microscope (Figure3.2) and then stored in PBS.
60 Figure 3.1 Schematic d iagram of hydrogel coated s ensing e lement of the g lucose e lectrod e. Figure adapted from Ju (Ju, 2006) (1) Teflon covered Pt Ir wire; (2) Ag/AgCl reference wire; (3) Collagen scaffold; (4) Electrically insulating sealant; (5) Epoxy Pu outer membrane; (6) Enzyme layer; (7) Stripped and coiled Pt Ir wire; (8) Cotton fiber with GOX gel.
61 Figure 3.2 Photo of sensor coated with hydrogel (40 magnification using a dissection microscope) 3. 2.5. In vitro S ensor E valuation The in vitro performance of the sensors was examined in glucose/PBS to determine the response t ime, sensitivity and linear range of sensors at different days. The response time represented the required time to reach 90% of the maximum response when the glucose concentration increased from 5 to 10mM. The sensitivity of the Pt/GOx/hydrogel sensors wa s measured using 5 and 10mM glucose/PBS. The linear range was obtained from a stepwise glucose increase from 2 to 30mM. All test solutions were prepared using pH 7.4 PBS (ionic strength = ~ 0.16M) and amperometric measurements were performed at room temper ature at 0.7V vs. Ag/AgCl. 3. 2.6. Animals and S urgical P rocedures The sensors were disinfected during polymerization of the hydrogel coating under a 254 nm UV light. Sensors used for implantation were pre conditioned in Ag/AgCl Electrode 1 mm Sensing Element with Hydrogel Coating
62 sterile saline for 2 days. The sen sor immobilization method for implantation was according to paper (Long et al., 2005) and further improved. Briefly, the sensor wires were covered with a section of 0.02 in. s ilastic tubing (2 cm in length). An overhand knot then was made in the middle of this section to use as a suturing site (i.e. for anchoring) during implantation and to prevent the sensor moving out of the skin. For each sensor placement (two/rat), a ~ 1.5 cm length longitudinal incision was made 1.5 cm laterally to the dorsal midline, 3 4 cm caudally from the neck. A subcutaneous pocket was created by using blunt surgical scissors before sensor wire placement. A 14 ga. I.V. catheter was inserted through the subcutaneous tissue area to the incision from the low back. The needle was w ithdrawn leaving the cannula in the subcutaneous tissue. The sensor tip and wires were carefully fed into the cannula through the incision. The sensor was secured to the skin by passing a 3 0 Prolene suture through the knot of the silastic tubing covered sensor wires and the incision was closed using 3 0 Prolene sutures. After implantation, a ~ 4mm length of Silastic tubing containing a ~ 8mm length of sensor wires was allowed to protrude from the skin. All protocols were approved by the University of Sou th Florida Institutional Animal Care and Use Committee (IACUC). 3. 2.7. In vivo E valuation M ethod During the course of in vivo testing, a continuous flow anesthesia system was used to deliver 1.5% isofluorane to the rats with 1.0 l/min oxygen. A maximum of 8 sensors were continuously monitored for 2 3 h by using two Apollo 4000
63 potentiostats (4 channels each; World Precision Instruments, Sarasota, FL). Response currents (nA) were acquired at 1 s intervals every second and dynamic current (nA) vs. time (s ) curves for all sensors were simultaneously displayed. After a run in period of approximately 1 h, a stable signal was obtained from the sensors. Dextrose (0.5 0.8 ml of a 50% (w/v) solution) was administered intraperitoneally using a 27 ga. needle. Fig ure 3.3 shows the in vivo continuous glucose monitoring procedure. Following the injection, a drop of blood was sampled every 6 8 min from the transected distal rat tail and blood glucose level was determined using a portable glucometer (FreeStyleTM; Ther aSense, Alameda, CA). The sensor sensitivity was calculated by dividing the current change by the blood glucose difference between the initial (before dextrose injection) and the peak status (after dextrose injection).
64 Figure 3.3 In vivo Con tinuous Glucose Monitoring Procedure Figure adapted
65 3.3 Results and Discussion 3.3.1. Characterization of Hydrogels After polymerization, hydrogels VP 0, VP 15, VP 30 and VP 45 were transparent and colorless in appearance and had a uniform and smooth surface, indicating that addition of VP did not result in phase separation or deterioration in the uniformity of the cured hydrogels. FT IR spectrometry (shown in Figure 3.4) revealed specific carbonyl adsorption peaks at to polymerization of HEMA o rDHPMA) and VP ) (Yaung and Kwei, 1998). As VP concentration was raised, the VP carbonyl peak increased. FTIR results indicate that VP was successfully co p olymerized with HEMA and DHMPA. The measured glass transition temperature values ( T g) of VP 0, VP 15, VP 30 and VP 45 were 111.88, 121.72, 133.86 and 143.75 o C, respectively (shown in Figure 3.5). Due to formation of hydrogen bonds between the carbonyl group in the P VP network and OH groups on the P HEMA or P DHPMA chains, VP effectively enhances the interactions between polymer chains and restricts the mobility of chains (Jin et al., 2006). Consequently, T g values increased with increasing VP concentration. VP im proved porosity of the hydrogels. Without VP the copolymer of HEMA and DHPMA did not show pores in Figure 3.6 while the hydrogels with VP were highly porous. The pore size increased when VP increased in the range of 0 45%.
66 According to higher magnificati on SEM photos (Figure3.7), it can be seen that nets start to appear inside the gel on the surface of the pores from VP 15 and VP 30. There are more nets insede VP 30 than VP 15. When it is magnified to 10K times, it can be seen that some big pieces of structu res combined with nets present on the surface of the pores. The diameter of net string is about 50nm. However, there are no nets on the surface of pores of VP 45. Equilibrium water fraction EWF values for VP 0, VP 15, VP 30 and VP 45 were 146%, 149%, 166% and 217% (wt%), respectively. In consideration of biocompatibility, the hydrogel with a large EWF and large pores would be preferred. In order to detect the durability of these novel hydrogels, hydrogel samples ( n = 4/hydrogel) were stored in water at room temperature for 4 weeks. Fortunately, no significant weight loss was observed. However, VP 45 spontaneously broke into several smaller pieces. From SEM pictures ( Figure 3.6 and 3.7) nets are present inside the gel on the surface of the pores of VP 15 and V P 30. These nets may strengthen polymer matrix with high water content. Since there are no nets on the surface of pores of VP 45, VP 45 was observed fragility probably due to swelling stresses as a result of the high water content. Balancing the properties of water content, porosity and durability of various hydrogels, VP 30 was selected for coating sensors for more in depth in vitro / in vivo performance testing.
68 Figure 3.4 FTIR of four freeze dried samples (the thickness of samples are around 0.5mm, taken by reflected light)
69 Figure 3.5 DSC of four freeze dried samples
70 Figure 3.6 A SEM of sample VP0 Figure 3.6 B SEM of sample VP15
71 Figure 3.6 C SEM of sample VP30 Figure 3.6 D SEM of sample VP45 Figure 3.6 Scanning electron micro scopy images of four freeze dried hydrogel samples (magnification: 350
72 Figure 3.7 A SEM of sample VP15(magnification ) Figure 3.7 B SEM of sample VP30(magnification )
73 Figure 3.7 C SEM of sample VP30(magni fication ) Figure 3.7 D SEM of sample VP30(magnification ) Figure 3.7 Scanning electron microscopy images of VP 15 and VP 30 freeze dried hydrogel samples with different magnification
74 3.3.2. In vitro P erformance of Pt/ GO x / VP 30 S ensors In order to improve the long term biocompatibility of the sensors in vivo, the hydrogel coating must be thick enough to be durable. On the other hand, if the coating is too thick and/or relatively less porous, the sensor response will be very slow. The wet coating thickness was estimated by the volume of the coating solution ( V c) VP 30 solution on the enzyme layer of Pt/ GO x sensors ( n = 4). In this case, the hydrogel layer served as the diffusion limiting membran e. Compared to the control sensors (Pt/ GO x ), the hydrogel layer did not improve the sensor response linearity; neither did it significantly reduce sensitivity except for the thicker coating (Table 3.2). The calibration plots from Pt/ GO x / VP 30 sensors with coating layers formed from 0.4, 0.8 and 1.2 VP 30 solutions are shown in Figure 3.8. The thicker coating did not improve linearity but reduced sensitivity. After the sensors had been in water for 28 days, all coatings were still firmly attached to the sensors but the sensitivity decreased ~ 50 %.
75 Table 3.2 Influences of Coating Thickness on Sensor Response Sensor (n=4) VP 30 Coating Volume (L) S (nA/mM) T 90% (min.) Pt/ GO x 0.0 48.26.3 0.220.07 Pt/ GOx / VP 30 0.4 47.62.8 1.10.1 0.8 45.41.5 1.40.2 1.2 34.71.8 1.90.3 Pt/ GO x /EPU 0 .0 3.02.1 1.41.1 Pt/ GO x /EPU/ VP 30 0.4 3.01.6 3.40.9 0.8 2.92.2 6.43.0 1.2 2.70.6 8.12.2 Figure 3.8 Calibration plots for Pt/GOx and Pt/GOx/ VP 30 glucose sensors. Data are the means of measurements from 4 sensors at day 3. Error bars are sta ndard error.
76 3.3.3 In vitro P erformance of Pt/ GO x /EPU/ VP 30 S ensors Due to the poor linearity, Pt/ GO x / VP 30 sensors might not be able to meet the basic requirement of in vivo measurements for accuracy and precision. Though non linear calibration can be perf ormed in an effort to improve measurement reliability, a better solution is to add an additional diffusion limiting membrane in between the hydrogel layer and the enzyme layer, for example, a low permeability epoxy polyurethane membrane (Yu et al., 2006). Such sensors with hydrogel coating, i.e. Pt/ GOx /EPU/ VP 30, were prepared using similar methods as those used for preparing Pt/ GO x / VP 30 sensors and tested in glucose/PBS. It was found that hydrogels could firmly attach to the epoxy polyurethane membrane. The sensitivity was decreased to about 10% of the sensitivity of Pt/ GO x / VP 30 sensors, which did not change with an increase in hydrogel coating (Table 3.2). This was because the epoxy polyurethane was much less permeable than the hydrogel and played the d ecisive diffusion limiting role. significantly improved by the epoxy polyurethane membrane and was at least in the range of 2 30mM. The sensor response was dramatically delayed when the hydrogel coating was thic ker (the response time is proportional to the square root of membrane thickness (Baronas et al., 2003). coating, the response time was 3.4 0.9, 6.4 3.0 and 8.1 2.2 min, respectively. Thus, VP 30 coating. The Pt/ GO x /EPU/ VP decline in either sensitivity or
77 linearity over 28 days ( Figure 3.9). The coating was still visible and firmly attached to the sensor after 28 days. 3. 3.4. In vivo P erformance of Pt/ GOX /EPU/ VP 30 S ensors Eight Pt/ GOx /EPU/ VP eously implanted in four rats (two sensors/rat) and the in vivo sensitivity of each sensor was tested at days 7, 14, 21 and 28. The testing results are summarized in Table 3.3. All sensors kept functioning in the first 3 weeks and produced an in vivo res ponse sensitivity of ~ 4 nA/mM. At day 28 after implantation, 3 of 8 implanted sensors still performed. It has been reported that 68 Pt/ GOX /EPU sensors were tested in vivo and the sensor survival at day 21 was 13.5% (Long et al., 2005). This suggests that the addition of hydrogel coating apparently improves the long term performance of the implanted glucose sensors. The improvement might be attributable to the excellent biocompatibility of the hydrogel coating. Representative histological sections, prepa red from tissue surrounding a Pt/ GO x /EPU/ VP 30 sensor and a Pt/ GO x /EPU sensor at day 28 after implantation, are shown in Figure 3.10. Only a few inflammatory cells were observed in the tissue surrounding hydrogel coated sensors ( Figure 3.10B). Without hydr ogel coating, many inflammatory cells appeared ( Figure 3.10A). The thickness of the fibrous tissue around the Pt/ GO x /EPU/ VP 30 sensor was 50 the thickness of 100 Pt/ GO x /EPU sens ors (Long et al., 2005). The reason for the functional failure of the other five sensors at day 28 in this study is unknown. Histological results did not
78 show any apparent differences between the tissue surrounding the surviving sensors and the tissue sur rounding the failed sensors. Table 3.3 In vivo sensor sensitivity of Pt/ GO x /EPU/ VP 30 sensors (n=8) Rat No. 1 2 3 4 Average 1 Sensor left right left right left right left right Standard Error Day 7 12.0 5.5 2.6 2.1 4.6 1.2 2.6 1.4 4.0 1.3 Day 14 4.4 7.3 1.5 5.8 3.4 1.9 4.2 2.1 3.8 0.7 Day 21 6.6 4.2 1.5 5.6 5.3 2.1 2.0 5.4 4.1 0.7 Day 28 0 0 0 0.9 0 0 4.2 2.2 N/A
79 Figure 3.9 Calibration plots for Pt/ GOx /EPU/ VP 30 (0.4 u l) sensors at day 3, 7, 16 and 28. Data are the means of meas urements from 4 sensors. Error bars are standard error.
80 Figure 3.10 Hematoxylin and eosin stained sections of tissue surrounding glucose sensors implanted subcutaneously in rats for 28 days. The photo (A) shows the capsular tissue formed a round a Pt/ GO x /EPU sensor. The photo (B) shows thecapsular tissue formed around a Pt/ GO x /EPU/ VP A B Hydrogel Debris Inflammatory Cells
81 3.4. Conclusion This study has provided a feasible approach to design and select the properties of the copolymer for coating implantable bio sensors. For the first time, animal experiments were used to demonstrate that a hydrogel coating was effective in minimizing tissue reactions surrounding implanted minimally invasive needle type glucose biosensors. With a hydrogel outermost coating and an appropriate diffusion limiting layer, significant improvements of the in vivo performance of implanted glucose sensors can be anticipated.
82 CHAPTER 4 EFFECT OF NOVLE HYDROGEL COMPOSITION ON TARGETED DEXAMETHASONE 21 PHOSPHATE DISO DIUM SALT DELIVERY 4.1. Introduction When implants such as biosensors, pacemakers and bioartificial organs are implanted in the body, acute inflammation can occur within seconds or may be delayed for days. Without suppressive agents, chronic inflammatio n and fibrosis usually follows. The implants will lose normal functions with these tissue reactions (Hickey et al. 2002). According to recent reports ( No rton et al. 2006 ; Patil et al. 2004 ; Yoon et al. 2003 ; Klueh et al. 2007), the local release of DX can su ppress the acute inflammation around the biosensors and may improve biocompatibility and prolong sensor lifetime. The dexamethasone concentration around the implant depends on the drug release characterization from the matrix and must be high enough to pr event inflammation (Moussy et al. 2006). The drug release from eq u ilib r ium swollen hydrogel can be affected by both drug and materials used (Karlgard et al. 2003). Norton et al (Norton et al. 2005 ; Gallardo et al. 2001) used HEMA VP PEG hydrogel as a matr ix for DX release, but there is no report about the effect of hydrogel compositions on DX releasing. It has been reported that polymer structure and composition greatly influenced drug release, so it was a matter of importance to focus
83 on the effect of hy drogel composition on the targeted delivery of DX 21. In this phase of the research, a serious of porous poly(HEMA DHPMA VP ) matrixes are prepared in the membrane form according to previous two chapters. The water soluble DX 21 was loaded in to the hydroge l followed by the study to determine drug load and releas e The simple HPLC method was developed to detect both DX and DX 21 using the same mobile phase. 4.2 Experimental 4.2.1 Chemicals 2 Hydroxyethyl methacrylate (HEMA) and 2, 3 dihydroxypropyl methacr ylate (DHPMA) were donated by Benz R & D (Sarasota, FL USA)and used as received without further purification. N Vinyl 2 Pyrrolidinone ( VP ) (99.9+%) was purchased from Sigma Aldrich Co.( St. Louis, MO, USA) and purified by vacuum distillation to obtain a colorless liqui d (Janson et al. 2005) Ethylene glycol dimethacrylate (EGDMA), Dexamethasone 21 phosphate disodium salt (98%, powder) and monobasic potassium phosphate ACS grade from Sigma Aldrich Co. (St. Louis, MO, USA) were used as received. Methanol HPLC grade fr om Fisher Scientific (Fair Lawn, NJ, USA) and Benacure 1173 from Mayzon Corporation (Rochester, NY, USA). Phosphate buffered saline (PBS) 10X solution from Fisher Scientific (Fair Lawn, NJ, USA)was diluted to 1X with pH 7.40.1.
84 4.2.2 Drug loading by E qu ilibrium P artitioning The homemade swollen hydrogels were cut by the core bore (diameter, 5.73mm) and freeze dried for one day to obtain xerogel pellets which are kept in the descicator until used. Xerogel pellets were weighed, wrapped in the net and plac ed in individual vials with 3mL of 25mg/mL dexamethasone 21 phosphate disodium (DX 21) solutions. The pellets were immersed in the solutions which were stirred and incubated in 37 oil bath for 24 hours. 4.2.3 In vitro R elease S tudy Drug loaded hydrogels were removed from drug solution. The excess solution on the surface of the hydrogels was removed by using Kimwipes. Drug loaded hydrogel were wrapped again in the net and placed into 3mL fresh PBS solution which was stirred and incubated in 37 oil bath. PBS at pH7.40.1 at 37 resembles the environment of the drug inside the human body. At different time intervals, hydrogels were removed and placed in fresh PBS solutions to kee p the sink condition. The drug solutions were kept in the refrigerator prior to detect the DX 21 and DX concentration. Three in vitro release studies were performed under the same conditions for each hydrogel sample. The means were calculated and graphed for each different time interval.
85 4.2.4 HPLC A nalysis of DX 21 and DX R elease 220.127.116.11 Preparation of Standard S olution An accurately weighed quantity of DX 21 was dissolved in different mobile phase (Table 4.1)to obtain a known concentration of about 5 35.20ug/mL. Stock solution was diluted to obtain 5.35, 10.70, 21.41, 42.82, 53.52, 64.22, 85.63 and 107.04 ug/mL solution. A known amount of DX 21 was weighted and dissolved with mobile phase to obtain a known concentration of 98.80ug/mL. The stock solu tion was diluted to obtain 79.00, 59.30, 49.40, 39.50, 19.75, 9.88 and 4.94 ug/mL solution. 18.104.22.168 HPLC S ystem The concentration of the solution was monitored by HPLC with Shimadzu system controller, UV vis detector, and liquid chromatograph. The column is Shimadzu premier C18 5um 150*4.6mm. The wavelength was 241nm, and the flow rate was 1 mL/min. 22.214.171.124 Selection of the Mobile Phase Different mobile phases were prepared according to table 4.1 and degassed by ultrasonic. The mobile phase ran about ha lf hour until the base line was stable. First, 39.5 ug/ml DX and 42.82 ug/ ml DX 21 were separately injected. Second, the mixture of 10ml 79.00 ug/ml DX and 10ml 85.63 ug/ ml DX 21 was injected. De ionized water is used to clean the tube and column for two hours. After that a
86 new mobile phase ran about one hour before the drug solution was injected. The same procedure was repeated to test different mobile phase. 126.96.36.199 Calibration C urve and Sample A ssay The peak area under the curve was plotted to t he standard concentration to obtain the calibration line. The formula derived from the line was used to calculate the drug concentration of the samples. The samples were injected to the 20uL auto valve. Every sample was injected at lease twice. If the err or was beyond 5%, a third injection was done. The concentration of the solution was taken as the average value. The mass of the drug equalled the average of the triple experiments. 4.3 Results and D iscussion 4.3.1The S election of the M obile P hase Dexamet hasone 21 phosphate disodium (DX 21)( Figure 4.1), prodrug of dexamethasone, can convert to dexamethasone(DX) ( Figure 4.2)in biological fluids in vitro (Blackford et al. 2000) During the release experiment, some DX 21 molecules convert to DX too. Different mobile phases have developed to determine DX 21 and DX separately (Blackford et al. 2000) It was too costly and time consuming to use different mobile phase to detect DX and DX 21 separately; therefore, a single mobile phase that could be used to detect both was necessary. Table 4.1 shows different mobile phases are used to separate DX 21 and DX. According to Tabel 4.1, it is seen that DX 21 is eluted earlier than DX because
87 DX 21 is more polar than DX since reverse HPLC was used. When the polari ty of mobile phase decreases, retention times of DX 21 and DX are extended. When the ratio of methanol to water arrives at 6:4, DX 21 and DX can be separated. They can further separa ted by mobile phase of methanol : water (5:5). However, retentions time of DX 21 (2.850 and 3.175) are so short that they are close to the dead time of the column which is about 2min. When the 0.01M KH 2 PO 4 is added to the mixture, it makes the retention time of DX 21 longer, while the retention time of DX shorter. Both of the two mobile phases of methanol:water (6:4) and methanol :water(5:5) can separate DX 21 and DX too. Furthermore, the peak of DX 21 can be separated from dead volume of column. In order to save time, methanol: water (6:4) with 0.01M KH 2 PO 4 was preferred. Figure 4.3 shows HPLC profiles of DX 21 and DX with mobile phase of methanol:water(5:4, 0.01KH 2 PO 4 ). Figure 4.1 structure of Dexamethasone 21 phosphate disodium salt(DX 21)
88 Figure 4.2 Structure of dexamethasone Table 4.1Mobile phase and Retention t ime Mobile phase Retention time DX 21 (min) Retention time DX (min) 100% methanol 2.075 2.117 Methanol :water=7:3 2.658 3.392 Methanol:water=6:4 2. 850 6.725 Methanol: water=5:5 3.175 17.342 Methanol:water=5:5(0.01MKH 2 PO 4 ) 9.942 16.625 Methanol:water= 6:4(0.01MKH 2 PO 4 ) 4. 533 6. 558
89 Figure 4.3 A Figure 4.3 B Figure 4.3 C Figure 4.3 HPLC profile s with mob ile phase Methanol:water=6:4(0.01MKH 2 PO 4 ) : A.DX 21 B. DX C. DX 21 and DX Table 4.2 Regression parameters for DX and DX 21 formula R 2 n* DX 21 3 E 5x 1.0425 0.9986 21 DX 1E 5x+0.0707 0.988 0 21 n* is the number of experiments: seven samples with three times injection each
90 Table 4.2 shows the calibration formulas of the DX 21 and DX, y=3E 5x 1.0425 and y=1E 5x+0.0707 separately, x is the area und er the peak, and y is drug concentration. The concentrations of DX 21 and DX are calculated by these two formulas with the area under the peak. 4.3.2 Drug L oading by E quilibrium P artitioning DX has often been loaded to hydrogels by putting certain amount of DX in the monomer solution before polymerization because it is not water soluble (Norton et al. 2005). polymerization due to the loss of drug, and the drug may not stable during th e polymerization. For the current research effort, DX 21, pro drug of DX 21, was chosen for its high water solubility. It has been reported that highly hydrophilic hydrogels are useful for the release of water soluble drugs that can be immobilized in the hydrophilic matrix by physical entanglement (Barbu et al. 2005) Furthermore, hydrogels can be loaded with water soluble drugs by partitioning it into the polymer matrix when the purified xerogel swells in the drug water solution after polymerization (Ki m et al.1992)
91 Table 4.3 drug loading percent s & EWCs of hydrogel s with DX 21 sample Load(%) a EWC% D0 4.75 52.8 D20 8.95 64.4 D40 11.16 71.8 D80 15.73 77.2 VP 0 13.3 73.5 VP 15 13.4 73.7 VP 30 21.2 74.6 VP 45 20.1 b* 75.6 a*Load(%)=Mass of tota l DX 21/Mass of Xerogel b* VP 45 tears during releasing experiment The DX 21 loading percent is calculated by mass of cumulated release amount of DX 21 and DX divided by mass of xerogel. When xerogel is immersed in the drug solution, drug enters to the pol ymer due to three types of driving forces: a drug concentration gradient, a polymer stress gradient and the osmotic forces (Brazel and Peppas, 1999). Therefore, a higher degree of swelling is caused by the polymer stress gradient and osmotic forces which allow greater amounts of drug imbibed in the polymer matrix. From above table 4.3, it can be seen that DX 21 loading percent increased with the swelling degree of hydrogels in each series. p took about 6713ug DX 21 and the release could not be evaluated due to the low concentration of drug released. The lowest maximum uptakes observed for DX 21 can be
92 explained by the high aqueous solubility of DX 21 which would remain preferentially in so lution rather than within the lens material (Karlgard et al. 2003). According to results of table 4.3, the drug uptakes of different formulations of hydrogels are much The high aqueous solubility of DX 21 does not appea r to be the reason for the lowest uptakes of DX 21. This result is consistent with Pepp as It is reported that the amount of drug loaded into xerogel is a function of the partition coefficient between the drug in sol ution and the gel itself, hydrophilic drugs which are more soluble in the loading solutions are loaded to higher concentration into polymer samples (Brazel and Peppas, 1999). 4.3.3 Drug Release After surgical implantation of the glucose sensors, the acute inflammatory response normally takes place within seconds. During the initial stage, proteins and inflammatory cells adsorb to the sensor surface. Phagocytic cells ( e.g. neutrophils, monocytes, and macropages) then surround the device within minutes to hours in an effort to destroy it. Such membrane biofouling is detrimental to sensor function resulting in restriction of analyte diffusion to the sensor and/or degradation of the sensor membrane (Shin and Schoenfisch, 2006) If the acute inflammation can not be suppressed, the chronic inflammation will be sustained from weeks to lifetime. This may result in the forming of fibrosis around the implanted glucose sensors and enable the implanted glucose sensors lose the valid function (Shin and Schoenfisch, 2 006). In order to prevent the acute inflammation and prolong the lifetime of the implanted
93 glucose sensors, both penetration rate and concentration of anti inflammatory drug around the tissue should be high enough to suppress the inflammation ( Moussy et a l. 2006 ). According to the results from figure 4.4 and 4.5, all hydrogels showed a high initial release, followed by slow, long term release during the next hours to days. This function is believed to be good for the implanted glucose sensors to suppress the acute inflammation and chronic inflammation. Drug release kinetics from hydrogels can be expressed by the following equation (4 1): =kt n (4 1) Where M t is the amount of release at time t, is the initial d rug loading, is the fraction of drug released, t is the release time, n, drug release exponent, and k is a constant incorporating the structural and geometrical characteristics of the release device. It has been reported that n=1 is Case II release kinetics and n=0.5 corresponds to the Fickian release kinetics (Ritger and Peppas, 1987). The values of n were derived from the linear regression slopes of the release profiles (Mt/ ) <0.6 using above equation shown in Figure 4.5. It shows that n is about 0.5 which indicates the release mechanism is Fickian diffusion. Average DX 21 release rates from hydrogels can be expressed by the diffusion coefficient. All sample shapes are similar and the ratio of thickness (L) t o diameter (D i ) is about 1:6.5. It is assumed that water diffuses into a planar sheet at a constant rate. When and time will fit
94 equation 5 (Kim et al. 1980). = ( ) 0.5 (4 2) The diffusion coefficient (D) values were estimated according to equation (5) and shown in Figure 4.6 and Table 4.4. From Table 4.4 it can be seen that ave rage DX 21 release rates from DHPMA series hydrogels increases with DHPMA content due to higher equilibrium water content, initial DX 21 loaded percent and more pores. The higher initial drug loaded percent causes higher concentration gradient between hydr ogel and fresh PBS solution which acts as the driving force for the drug diffusion. The pores of polymers work as a tunnel for drug transport. Since water is the vehicle to transport the drug out of polymer, higher water content will facilitate drug diffu sion. Factors influencing drug releasing behavior is related to the drug uptaking behavior: concentration gradient, equilibrium water content and pores which are dictated by polymer composition. For VP series hydrogels, average release rates of VP 15 and VP 30 are lower than VP 0 and VP 45. According to SEM figures 3.4, there are nets that appear on the surface of pores. These nets may work to block DX 21 transport out of the hydrogel. There are more nets on the surface of VP 30 pores than in VP 15. The aver age release rate of VP 30 is also slower than VP 15. This confirms that the polymer inner morphology is another factor that greatly influences drug diffusion.
95 Figure 4.4 A initial burst release Figure 4.4 B low level long term release Figure 4.4 DX 21Cum ulative release profiles from DHPMA hydrogels
96 Figure 4.5 A initial burst release Figure 4.5 B low level long term release Figure 4.5 DX 21Cumulative release profiles from VP hydrogels
97 A. DHPMA series B. VP series Figure 4.6 Drug release kinetics log( Mt/Meq) versus log Time
98 Table 4.4 DX 21 release exponents and coefficients n R 2 8 cm 2 S 1 ) D0 0.4394 0.9958 2.600.08 D20 0.4964 0.997 0 13.800.13 D40 0.4177 0.9987 21.761.03 D80 0.4780 0.9969 62.364.04 VP 0 0.5649 0.998 0 38.371.12 VP 15 0.5022 0.9995 33.702.54 VP 30 0.4422 0.9999 27.670.23 VP 45 0.4193 0.9999 52.302.72
99 4.4. Conclusion The mobile phase methanol:water=6:4(0.01MKH 2 PO 4 ) could separate DX and DX 21 at the same run so it is available to detect both DX and DX 21 by a simple HPLC method. DX 21 was successfully loaded in to a series of hydrogels by equili brium partitioning. The drug loaded percents show that D80 could uptake the maximum DX 21 due to high equilibrium water content and more pores. The drug release kinetics shows that DX 21 releasing mechanism from novel hydrogels is Fickian diffusion. The factors influence drug releasing behavior is relative to drug uptaking behavior: concentration gradient, inner morphology and water content which are decided by polymer composition.
100 CHAPTER 5 WATER STRUCTURE IN HYDROGELS 5.1. Introduction The study of w ater structure in the hydrogel s can as sist in determin ing the biocompatibility of the hydrogel and affect the transportation of the molecules between the matrix and intermediate. An investigation of water structure in a polymer yields valuable information on the sorption, diffusion and permeation properties of molecular species in hydrophilic polymers.(Ping et al. 2001) Most research on water structure in hydrogels has focused on the water in hydrogel s using deionized water or PBS solution as m edia ( Goda et al.2006; Liu and Huglin, 2003; Ping et al. 2001 ; Ahmad and Huglin, 1994). Research conducted by Stevenson and Gates compared water st ructures in hydrogel s at equilibrium in different media such as water and PBS (Stevenson and Sefton,1988; Gat es and Harmon, 2001) However, there is no report about water structure in hydrogel s which swell to reach equilibrium when immersed in drug solution W ater st ructure in hydrogels can be detected by differential scanning calorimetry (DSC) ( Murphy et al. 1988 ; Pedley and Tighe, 1979 ) ,NMR ( Chowdhury et al. 2004 ; Capitani et al. 2001 ) dilatometry electrical conductivity ( Lee et al. 1975 ; Choi et al. 1977 ) dielectric relaxation spectroscopy ( Kyritsis et al. 1995 ) and
101 dynamic mechanical spectroscopy ( Lustig et a l. 1991 ) The most common method is DSC. Using DSC, water in a hydrogel is classified as free water, freezing bound water and non freezing water. Free water does not take part in hydrogen bonding with polymer molecules. It has a similar transition temperat ure, enthalpy and DSC curves as pure water ( Nakamura et al. 1983 ) Freezing bound water interacts weakly with polymer molecules and has a phase transition temperature lower than 273K. Non freezing bound water is complexed with the polymer chain through hy drogen bonds and has no detectable phase transition over the temperature range from 200K to 273K ( Higuchi et al. 1984 ) I t was previously discussed that equilibrium water content of the hydrogel is one important factor affect ing drug diffusion. It is kno wn that freezing water content is essential for predicting efficient diffusion of salts and macromolecule s (Pedley D.G. and Tighe 1979) H ydrogel coatings loaded with DX 21will be implanted and come in contact with different body fluid s which contain many small water soluble molecules. When these small molecules diffuse into hydrogels and DX 21 molecules diffuse out of hydrogels, the water content and structure in hydrogels may change too. Therefore, it is very important to investigate the water s tructure in hydrogel s at equilibrium in small molecu le media and DX 21 solution. In this chapter, 1HEMA:1DHPMA copolymer and 3 VP HEMA DHPMA copolymers were used to study the water st ructure PBS was chose n as the media for small molecules because it is similar to body fluid. The water st ructure s of these hydrogels were studied by DSC.
102 5.2. Experimental 5.2.1. Chemicals Four different polymers were prepared as described in chapter 3: one HEMA DHPMA copolymer and three VP HEMA DHPMA copolymers. Dexamethasone 21 phosp hate disodium salt (98%, powder) ACS grade was purchased from Sigma Aldrich Co. (St. Louis, MO, USA) was used as received. Phosphate buffered saline (PBS, 11.9mM Phosphate, 137mM Sodium Chloride, 2.7mM Potassium Chloride) 10X solution from Fisher Scientif ic (Fair Lawn, NJ, USA) was diluted to 1X with pH 7.40.1. 5.2.2. Instrumentation All water structure data were obtained by measuring heat flow with a TA Instruments 2920 differential scanning calorimeter (DSC). DSC measures heat flow and temperature betw een a sample and reference as a function of time and temperature. The reference is an empty hermetic aluminum pan of similar weight to that of the sample pan. The DSC was calibrated from 100 to 200 at 5 /min heating rate for baseline and temperature using an indium standard H ydrogels were analyzed by cooling the sample from 20 to 80 and a one minute isothermal period at 80 and heating the sample from 80 to 20 at 5 /min. The wate r fraction s for hydrated samples were reported as the water mass over xerogel mass. The total integrated area of the respective endotherm peak represents t he melting enthalpy. The melting enthalpy was plotted as a function of
103 total water fraction, and th e amount of nonfreezing water was determined by the water fraction intercept of the linear plot. 5.2.3. Method The drug loading and releasing method are the same as previously discussed The equilibrium water fraction (EWF) was determined by the mass of w ater divided by the mass of xerogel. Freeze dried were equilibrated swollen in de ionized water, 1X PBS solution, 25mg/ml DX 21 solution at atmospheric conditions, 23 and 55% relative humidity. The samples analyzed by DSC were cut from the swollen hydrog els with #1 cork bore and placed in media to reach equilibrium. For the partially hydrated samples, some water was allowed to evaporate and then the sample s w ere sealed in the pan s to equilibrate. The equilibrium samples were quickly weighed to 0.01mg and immediately analyzed by DSC. The sample mass ranged from 5 to 15mg. After DSC analysis, the samples were removed from the hermetic pans, oven dried to obtain the dry polymer mass. The water mass was determined by subtracting the xerogel mass from the mass of the hydrated DSC sample. 5.3. Results and discussion 5.3.1 Equilibrium W ater F raction of H ydrogel B efore and A fter DX 21 R elea e The e quilibrium water fraction of hydrogels were determined when sw ollen to equilibrium in DX 21 water solution and when DX 21 loaded hydrogels released
104 DX 21 completely in PBS solution at 37 This data are listed in Table 5.1. From Figure 5.1 it can be seen that when DX 21 i s released, the hydrogels shr i nk. Before the release of the drug, t he equilibrium water fraction of hydrogel is more than after the release From this it was apparent tha t when DX 21 is released from the hydrogel, some water also moves out of the hydrogel and causes shrinkage. It can be concluded that not only the composition of hydrogel but also solutes affect the water content of the hydrogel. Further research will focu s on study ing the water structure in hydrogel at the equilibrium state in different media.
105 Figure 5.1 P otos of hydrogel loaded with DX 21( A ) and hydrogel released DX 21( B ) Table 5.1 Equilibrium water fraction s of DX 21 loaded hydro gels at equilibrium before and after the release of the drug in PBS solution at 37 0 C (means SD, n=3) EWF(with DX 21) EWF(released DX 21) VP 15 2.910.17 1.490.11 VP 30 4.250.09 2.170.23
106 5.3.2 Water S tructure of the HEMA DHPMA C opolymer HEMA DHPM A copolymer was made by 1:1 molar ratio of HEMA and DHPMA. Table 5.2 shows EWFs of HEMA DHPMA copolymer swollen in different media. Table 5.2 EWF and NFWF of 1HEMA:1DHPMA copolymer sample EWF NFWF water 1.7465 0.5806 PBS 1.6913 0.4814 DX 21 2.7737 0. 4239 Accordi ng to table 5.2 results, HEMA DHPMA copolymer exhibit s a lower EWF in PBS solution than in deionized water. This is different from the results reported by other researchers (Stevenson, 1988 ; Gates, 2001 ). Stevenson found that EW F s of polyHEM A, and poly(HEMA co MMA) were much higher in PBS than in distilled water. Gates reported that HEMA DHPMA copolymer exhibited the same equilibrium water content in de ionized water as well as in saline solution. These results are not consistent which may imply that the equilibrium water content of hydrogel in PBS solution depends on the gel composition and crosslink ratio. HEMA DHPMA copolymer imbibed more equilibrium water in the DX 21 solution than in PBS solution and de ionized water. This may be due t o incorporation of large molecule DX 21 which weakens the seco ndary bonds within the hydrogel, enlarg ing the distance between the polymer chains and caus ing an uptake of water. Thus the equilibrium water content of hydrogel in DX 21 is much
107 more than in P BS solution and water. This result is similar to the above result when DX 21 released from hydrogel along with water. HEMA DHPMA copolymer at equilibrium and at various states of partial hydration was analyzed by DSC ( Figure 5.2 ) When hydrogels were sw ollen in de ionized water, t wo crystallization exotherm peaks, assigned as T c1 at higher temperature and T c2 at lower temperatu re, are shown in cooling curves. The integrated area of lower temperature exotherm is greater than the higher temperature at equ ilibrium in deionized water. At initial dehydration stage, a third exotherm peak is seen at T c3 which is lower than T c1 and T c2, and both T c1 and T c2 are the same as equilibrium. When the material continued desorption, all crystallization temperatures T c 1 T c2 and T c3 decreased. These results are consistent with previous research ( G. Gates and J. Harmon, 2001 ) When hydrogels were swollen in PBS solution, t here is one crystallization exotherm (T c1 ) with a lower temperature shoulder at equilibrium. The s houlder progressively separates from the T c1 and moved to lower temperature during desorption. There is only one crystallization exotherm ( T c1 ) at equilibrium during initial dehydration in DX 21 solution. T c1 moved to lower temperature and one new crysta llization peak (T c2 ) appeared when dehydration process continued. In summary, the crystallization of water inside the hydrogel appears to be dependent o f the water fraction of hydrogel. At high water fraction, only one
108 transition peak is present, and more transitions appear as water fraction decreas es At high er water fraction s there is only one sharp peak indicating the perfect crystallization of water. With lower water fraction, one or two more crystallizing peaks appeared. The reduced swelling and sh rinkage during the process of dehydration can restrict the mobility of the water due to stronger interactions with polymer or water. This makes it harder to form perfect crystals, so broader peaks appeared in lower temperatures instead of sharp peaks. The re are two transitions observed in the heating curves at different water fraction s when HEMA DHPMA copolymer is swollen in water and PBS solution. The temperature T m1 (transition near 273K) was assigned for free freezing water, and temperature T m2 for fre ezable bound water (transition near 263K) ( Ahmad and Huglin ,1994 ) From the transitions of different water fraction s swollen in DX 21, it can be seen that when EWF is over 2.36, the two peaks merge progressively to become a broad peak. Such a broad endot hermic peak was also observed in other hydrophilic polymers ( Ping et al. 2001 ; Wycisk and Trochimczuk,1992 ; Tasaka et al. 1988 ; Zhang et al. 1989 ; Fushimi et al. 1991 ; Hatada et al. 1982 ) Ahmad M. B. reported that the appearance of the second transition at T m2 seemed to be dependent o f the amount of water absorbed by the hydrogels while there is no evidence for T m1 to deviat ing from 273 K. Ping reported that the melting point of freezable bound water increased linearly with the water content in PVA hydrogels an d explained that the lowering of the freezable bound water melting point was due to
109 the increase in the overall hydrogen bonding ability of the polymer (Ping et al. 2001) Baker reported that the thermal transition temperature for free freezing water was f ound to decrease with decreasing water content (Fushimi et al. 1991) For the present results, both peak maximum temperatures T m1 and T m2 decreased with decreasing water fraction. These heating results are consistent with our cooling results because in the cooling process imperfect ice crystals formed during the dehydration process, were un stable compared to the perfect crystals and melt at lower temperatures. A decrease in separation (T m1 T m2 ) of peak maximum temperatures with increasing EWC has been r eported by ( Higuchi and Iijima,1985), but the separation did not decrease consistently with increasing EWC was also reported by Ahmad M. B. ( Ahmad,1994 ) These reports were not in agreement with each other, which may be due to different polymer compositio ns with different EWC. In present study, the separation was studied based on the same formulation with different water contents. The results showed that separation (T m1 T m2 ) (Table 5.3) increased with increasing water fraction contrary to what was report e d by Higuchi (Higuchi and Iijima,1985)
110 Table 5.3 Water melting points of hydrogels at different water fraction s WF T m1 ( ) T m2 ( ) T m1 T m2 VP 0(water) 1.74 5.62 5.13 10.75 VP 0 1(water) 1.58 3.45 5.67 9.12 VP 0 2(water) 1.20 2.03 5.84 7.87 VP 0 3(water) 0.88 0.22 6.13 5.91 VP 0(PBS) 1.69 3.06 VP 0 1(PBS) 1.51 3.14 7.63 10.77 VP 0 2(PBS) 1.41 1.84 8.15 9.99 VP 0 3(PB S) 1.27 0.61 8.67 9.28 VP 0 4(PBS) 0.89 0.22 8.93 8.71 VP 0(DX21) 2.77 6.49 VP 0 1(DX21) 2.36 4.96 VP 0 2(DX21) 1.62 2.96 6.20 9.16 VP 0 3(DX21) 1.14 7.64 VP 0 4(DX21) 0.81 0.24 8.37 8.13
115 Figure 5.2 DSC cooling an d heating curves for 1HEMA:1DHPMA copolymer at various stages during desorption of deionized water, PBS solution and DX 21
116 Since the nonfreezing water fraction can be defined as the limiting value of water fra c tion at zero enthalpy of fusion (Vaquez et al 1997), the enthalpy (integrated area) of melting endotherms from the DSC data was used to determine the nonfreezing water fraction by extrapolation. Linear plots of the total endotherm area as a function of water fraction are presented in figure 5.3, 5. 4, 5.5. The water fraction intercept was used as the fraction of nonfreezing water (NFWF) and shown in Table 5.2. fractions in water and PBS solution are almost the same. This may be due to different PBS solution and different crosslink ratio of 1HEMA:1DHPMA copolymer.
117 Figure 5.3 Linear plot of total integrated endotherm area versus total water fraction for the 1 HEMA:1 DHPMA copolymer swollen in deionized water Fi gure 5.4 Linear plot of total integrated endotherm area versus total water fraction for the 1 HEMA:1 DHPMA copolymer swollen in PBS solution Figure 5.5 Linear plot of total integrated endotherm area versus total water fraction for the 1 HEMA:1 DHPMA co polymer swollen in DX 21 solution
118 5.3.3 Water S tructure of VP HEMA DHPMA C opolymer H ydrogels Based on 1HEMA:1DHPMA copolymer, different molar ratios (15, 30, 45) of VP were copolymerized with 1HEMA:1DHPMA monomers to obtain VP HEMA DHPMA copolymer, VP 15, VP 30, VP 45. Equilibrium water fractions of VP 15, VP 30 and VP 45 were determined by mass of water divided by mass of xerogel and listed in Table 5.4. Water structure s of VP HEMA DHPMA copolymer swollen in de ionized water, PBS solution and DX 21 solution w ere studied by DSC. Linear plots of the total endotherm area as a function of water fraction are presented in figure 5.6, 5.7, 5.8. The water fraction intercept was used as the fraction of nonfreezing water (NFWF) and shown in Table 5.6. Freezing water fraction was determined by subtracting nonfreezing water fraction from equilibrium water fraction (Table5.7). All VP HEMA DHPMA copolymers exhibited highe st equilibrium water fractions in DX 21 and l owest in PBS solution. This result is the same as the 1HEMA:1DHPMA copolymer. The low equilibrium water content of hydrogel is usually improved by copolymerizing more hydrophilic monomers. It seems that large water soluble salts such as DX 21 also can increase the equilibrium water content of hydrogel. Table 5.5 shows values of the difference of water mass between the hydrogels at equilibrium in DX 21 and PBS solution divided by mass of DX 21 loaded in hydrogel. For DX 21 all values are about 9.6 which is independent of polymer compositions. This offers a new and simple method to estimate the drug loading amount in hydrogels.
119 Table 5.4 EWF of VP HEMA DHPMA copolymer sample PBS water DX 21 VP 15 1.6947 1.8122 2.8036 VP 30 1.8709 2.0100 2.9346 VP 45 1.8890 2.2428 3.1004 Table 5.5 V er mass(g) / DX 21mass(g) Hydrogel DX 21mass(g) VP 15 9.6 VP 30 9.6 VP 45 9.6
120 Figure 5.6 Linear plot of total integrated endotherm area versus total water fraction for the VP 15 copolymer swollen in water, PBS and DX 21 s olution
121 Figure 5. 7 Linear plot of total integrated endotherm area versus total water fraction for the VP 15 copolymer swollen in water, PBS and DX 21 solution
122 Figure 5. 8 Linear plot of total integrated endotherm area versus total water fract ion for the VP 15 copolymer swollen in water, PBS and DX 21 solution
123 Table 5.6 NFWF of VP HEMA DHPMA copolymer sample PBS water DX 21 VP 15 0.4581 0.4120 0.4756 VP 30 0.4291 0.3791 0.4798 VP 45 0.4301 0.3186 0.5112 Table 5.7 FWF of VP HEMA DHPMA copolym er sample PBS water DX 21 VP 15 1.2366 1.4002 2.3280 VP 30 1.4418 1.6309 2.4548 VP 45 1.4589 1.9242 2.5892 When the hydrogels are at equilibrium in PBS solution and water, non freezing water content decreased with the increase of the VP content even tho ugh the whole water content increase. A similar decrease has also been found for the poly( VP MMA) hydrogels ( Liu and Huglin, 1995 ) Liu tried to explain that the decrease of nonfreezing water content was due to the increase of water content according to di fferent swelling time. This was an obvious fact because the same formulation swell s at different time when nonfreezing water content is expressed relative to the mass of the hydrogel. However, this explanation is not reasonable when nonfreezing water conte nt is expressed relative to the mass of xerogel. This explanation also does not explain the relation of the nonfreezing water content and VP
124 content, t hus there is no definit ive explanation for this trend. However, following are two possible explanations. One explanation for the observed trend may be that when VP content increased more H bond formed between the polymer chains as evident by the increase of the glass transition temperature (chapter 3). Since non freezing water is the H bond water formed betwe en water and polymer chain, the inter chain H bond reduced the total available sites to form H bond with water molecules. Thus, the amount of non freezing bond water decreases with increasing VP content. Another possible explanation for present polymers i s that different average number of non freezing water molecules per site. From the following table, it can be seen that average number of non freezing water molecules per VP is 4.2, per HEMA is 4.1 and per DHPMA is 6.4. Since the molar ratio of HEMA and DH PMA is 1:1, so the average number of non freezing water molecules per (HEMA DHPMA) is 5.25 which is larger than VP Thus when VP content increased, the number of non freezing water molecules decreased which lead to the decrease of nonfreezing water content
125 Table 5.8 Molar ratios of non freezing bound water in different polymers polymers N nf /N p PVP* 4.2 PHEMA* 4.1 PDHPMA* 6.4 *PHEMA and PDHPMA results were calculated from the reference ( Gates and Harmon, 2001 ) *P VP result was obtained from the r eference ( Ping et al. 2001 ) From table 5.6 it also can be seen that all VP contained samples showed higher nonfreezing water content in PBS solution than in de ionized water which is different from 1HEMA:1DHPMA copolymer. Without VP the salt ions interfe re with the OH groups on the polymer chain thereby reducing the number of H bonding sites available to water. VP in the polymer chain blocks the salts from interfering with the H bonding sites thus allowing water to H bond with the polymer. Thus more H bon ding sites are created and higher nonfreezing water content was seen When VP contained gels were swollen to reach equilibrium i n DX 21, hydrogels exhibited more FWF and NFWF than in de ionized water and PBS solution The additional nonfreezing water cont ent could be from H bonding water with the DX 21 since DX 21 molecule s contain polar groups such as O, F, and OH When more DX 21 molecules entered in to the polymer matrix, they also bring in more free water to the matrix. The incorporation of large DX 2 1 molecules can strengthen the distance of the polymer chain and let more freezable water enter into the matrix.
126 5.4. Conclusion Equilibrium water fraction of drug loaded hydrogel is bigger than that of drug released hydrogel, which is confirmed by the shr inkage of hydrogel. This result implies that both the composition of the hydrogel and the solutes can affect the water structure 1HEMA:1DHPMA copolymer and VP HEMA DHPMA copolymers imbibed highest equilibrium water fraction in DX 21 solution, lowest equ ilibrium water fraction in PBS solution. During dehydration process, crystallization exotherms and melting endotherms of 1HEMA:1 DHPMA copolymer in all different media moved to lower temperatures due to the shrinkage of hydrogel causeing imperfect ice cry stals to form, and a decrease in separation (T m1 T m2 ). For VP HEMA DHPMA copolymer, the non freezing water fraction decreases with an increase of VP at equilibrium state in water and PBS solution even though the equilibrium water fraction increase d This trend is explained by the decrease of H bond sites with water and less molar ratio of nonfreezing water per molecule when compared to 1HMEA:1DHPMA. Both non freezing water fraction and freezing water fraction of VP contained hydrogels are higher in DX 21 solution than in PBS and water. The information offers a new way to increase the low water content hydrogel by swelling xerogels in media containing large water soluble molecules. Since DX 21 molecule s contain polar groups such as OH F and O it can bri ng some H bond water in to the matrix. When the distance between the polymer chains is
127 increased more free water will be incorporated. Since the ratio of transporting water mass to DX 21 mass is 9.6 which is independent of hydrogel compositions, the DX 21 load can also be estimated by the difference of equilibrium water content of hydrogel swollen in DX 21 solution and PBS solution.
128 CHAPTER 6 SUMMARY AND SUGGESTIONS FOR FUTURE STUDY 6.1. Summary State of the art implantable g lucose sensors do not work reliably and have a rather short life after implantation ( Moussy 2002 ) This in vivo loss of function is caused by tissue reactions surrounding the sensor such as fibrosis and inflammation (Mang et al, 2005). It is believed th at the performance of implanted glucose sensors can greatly benefit from the use of more biocompatible outer most coatings. Therefore, in order to improve the lifetime of implantable glucose sensors, two different novel series of hydrogel coatings were de signed, synthesized and characterized for use on implantable glucose sensors. During this effort novel hydrogel polymers with various DHPMA content were prepared, characterized, and coated on to implantable glucose sensors and then tested in vitro and in v ivo The effects of 2,3 dihydroxypropyl methacrylate (DHPMA) on the swelling, morphology, glass transition(T g ), an d water structure were studied. The results show that the degree of swelling increases with increasing DHPMA content. Scanning electron mic roscopy (SEM) studies identified uniform, porous structures in samples containing 60 90 mole % DHPMA. Glass transition temperatures did not change significantly with DHPMA content, but the ratio of
129 freezing to nonfreezing water tended to increase with DHP MA content. Sensors coated with different hydrogels were prepared and in vitro evaluations were performed. The 80% DHPMEA hydrogels exhibited optimum sensitivity, response and stability when coated directly onto the sensor or top of a polyurethane (PU) l ayer. The histology results show that 80% DHPMA samples exhibit reduced fibrosis and inflammation. Eight D80 coated Pt/ GOX /epoxy polyurethane glucose sensors were implanted into 4 rats (2 sensors/rat). The original attempt of in vivo implanted glucose se nsors was not successful, because only 25% of the sensors were working at 4 weeks. In order to further improve the efficiency of implanted glucose sensors, another series of novel hydrogels were developed based on copolymer of HEMA and DHPMA. The porosit y and mechanical properties of the hydrogels were improved using VP and EGDMA. The results of SEM, DSC, and FT IR analysed show that the hydrogel (VP30) produced from a monomeric mixture of 69 % (1 HEMA :1 DHPMA ) 30% VP and 1% EDGMA (mol%) had a more uniform pore structure and net structure high water content at swelling equilibrium ( EWF = 166% by mass) and acceptable mechanical properties. Two kinds of VP30 coated sensors, Pt/ GO x /VP30 and Pt/ GO x /epoxy polyurethane (EPU)/VP30 sensors were examined in glucose solutions during a period of 4 weeks. The Pt/ GO x /VP30 sensors produce large response currents but the response linearity was poor. Therefore, further studies were focused on the Pt/ GO x /EPU/VP30 sensors. With a diffusion limiting epoxy polyurethane membr ane, the linearity was improved (2 30 mM) and the response time was within 5
130 min. Eight Pt/ GO x /EPU/VP30 sensors were subcutaneously implanted in rats and tested once per week over 4 weeks. All of the implanted sensors kept functioning for at least 21 day s and 3 out of 8 sensors still functioned at day 28. Histology revealed that the fibrous capsules surrounding hydrogel coated sensors were thinner than those surrounding Pt/ GO x /EPU sensors after 28 days of implantation. It was discussed that the biocompat ibility of glucose sensors was improved by hydrogel. However, there is still a small amount of inflammation and fibrosis occuring within the tissue around the implanted glucose sensors after 4 weeks. According to recent reports (Norton et al, 2006; Patil et al, 2004; Yoon et al,2003; Klueh et al, 2007) the local release of DX can suppress the acute inflammation around the biosensors and may improve biocompatibility and prolong sensor lifetime. In order to further suppress the biofouling process, Dexametha sone 21 phosphate disodium salt (DX 21) was incorporated to the hydrogel to do the targeted drug delivery research. DX 21 was successfully loaded in to a series of hydrogels by equilibrium partitioning. The drug release kinetics shows that DX 21 releasing mechanism from novel hydrogels is Fickian diffusion. The factors influence drug releasing behavior is relative to drug uptaking behavior: concentration gradient, inner morphology and water content which are decided by polymer composition. A ll hydrogels showed a high initial release, followed by slow, long term release during the next hours to days. This function is believed to be good for the implanted glucose sensors to suppress the acute inflammation and chronic inflammation.
131 During the release experi ment, some DX 21 (Figure4.1) molecules convert to DX (Figure 4.2) Different mobile phases have been developed to determine DX 21 and DX separately (Blackford et al, 2000) It was too costly and time consuming to use different mobile phase to detect DX and DX 21 separately; therefore, a single mobile phase methanol:water=6:4(0.01MKH 2 PO 4 ) that could be used to detect both was created. The st ructure of water in the hydrogel can determine the biocompatibility of the hydrogel and affect the transportation of the molecules between the matrix and intermediate. The investigations on the st ructure of water in a polymer can give valuable information on the sorption, diffusion and permeation properties of molecular species in hydrophilic polymers (Ping et al. 2001) T he water structure in hydrogels swollen in different media including water, PBS and DX 21 solution were investigated. Results show that equilibrium water fraction of drug loaded hydrogel is larger than that of drug released hydrogel, which is confirmed by the shrinkage of hydrogel. This result implied both the composition of hydrogel and solutes can affect the water structure in the hydrogel. 1HEMA:1DHPMA copolymer and VP HEMA DHPMA copolymers imbibed highest equilibrium water fraction in DX 21 solution, lowest equilibrium water fraction in PBS solution. This information offers a new way to increase the low water content of hydrogel by swelling xerogels in large water soluble molecules media. Since the ratio of transporting water mass to DX 21 mas s is 9 .6 which is
132 independent of the hydrogel composition, the DX 21 load can be estimated by the difference of equilibrium water content of hydrogel swells in DX 21 solution and PBS solution. During dehydration process, crystallization exotherms and melting end otherms of 1HEMA:1 DHPMA copolymer in all the different media moved to lower temperatures due to the shrinkage of hydrogel which causes imperfect ice crystals, at the same time separation (T m1 T m2 ) decreased too. For VP HEMA DHPMA copolymer, the non freezi ng water fraction decreases with VP increases at equilibrium state in water and PBS solution even though the equilibrium water fraction increases. This trend is explained by the decrease of H bond sites with water and less molar ratio of nonfreezing water per molecule. Both non freezing water fraction and freezing water fraction of VP contained hydrogels are higher in DX 21 solution than in PBS and water. 6 .2. Suggestions for F uture W ork s The efficiency of implanted glucose sensors has been enhanced by novel hydrogel coatings. However, there is still some inflammation processes and fibrosis that occurs to the tissue around implant ed glucose sensor. The drug delivery in vitro study shows that DX 21 loaded hydrogel should effectively suppress acute inflam mation and chronic inflammation. In a future study, drug loaded hydrogel coated glucose sensors will be implanted into rats and determine the performance of
133 the implanted glucose sensors. It has been reported that controlled neovascularization around senso r by angiogenic growth factors such as VEGF can improve the performance of implanted glucose sensors (Klueh et al.2003,2005). The pores structures inside the hydrogel can be useful for locating nanoparticles or microspheres for controlled drug release sys tem. The future goal of more studies will be to incorporate VEGF loaded microspheres or nanoparticles to these novel pores matrix. At last a dual release system for both DX 21 in the hydrogel and VEGF in microspheres will be realized. It has been report ed that DX can lead to an anti angiogenesis effect along with an anti inflammatory response. The initial burse release of DX 21 from this new dual release system can suppress the acute inflammation. VEGF will be released slowly form microspheres and nano particle accompanied by low level release of DX 21 which c an not greatly affect the angiog enesis. Since these novel hydrogel coatings improved the biocompatibility of implanted glucose sensors, they have the potential use for other implantable biosensors. A future study can also be focused on the modification of these novel hydrogel coatings for other implantable biosensors.
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147 Appendix A: Chapter 2, DSC curves Figure A 1.DSC curve for glass transition of D0 sample Figure A 2. DSC curve for glass transition of D20 sample
148 Appendix A (continued) Figure A 3. DSC curve for glass transition of D40 sample Figure A 4. DSC curve for glass transition of D60 sample
149 Appendix A (continued) Figu re A 5. DSC curve for glass transition of D80 sample Figure A 6. DS C curve for glass transition of D90 sample
150 Appendix A (continued) Figure A 7. DSC curve for water structure of D0 with 48.8% EWC Figure A 8. DSC curve for water structure of D20 with 52.0% EWC
151 Appendix A (continued) Figure A 9. DSC curve for water structure of D40 with 60.1% EWC Figure A 10. DSC curve for water structure of D60 with 65.5%
152 Appendix A (continued) Figure A 11. DSC curve for water structure of D80 wit h 69.4% EWC Figure A 12. DSC curve for water structure of D90 with 69.9% EWC
153 Appendix B: Chapter 5 Figure B 1 DSC curve for water structure of VP15 swollen in water with 1.8122 EWF Figure B 2 DSC curve for water structure of VP15 swollen in water with 1.6144 EWF
154 Appendix B (Continued) Figure B 3 DSC curve for water structure of VP15 swollen in water with 1.3636 EWF Figure B 4 DSC curve for water structure of VP15 swollen in water with 1.0245 EWF
155 Appendix B (Continued) Figure B 5 DSC curve for water structure of VP15 swollen in water with 0.5908 EWF Figure B 6 DSC curve for water structure of VP15 swollen in PBS with 1.6947 EWF
156 Appendix B (Continued) Figure B 7 DSC curve for water structure of VP15 swollen in PBS with 1.3817 EWF Figure B 8 DSC curve for water structure of VP15 swollen in PBS with 1.0981 EWF
157 Appendix B (Continued) Figure B 9 DSC curve for water structure of VP15 swollen in DX 21 with 2.8036 EWF Figure B 10 DSC curve for water structure of VP15 swollen in DX 21 with 2.3937 EWF
158 Appendix B (Continued) Figure B 11 DSC curve for water structure of VP15 swollen in DX 21 with 1.3192 EWF Figure B 12 DSC curve for water structure of VP15 swollen in DX 21 with 2.0100 EWF
159 Appendix B (Continued) Figure B 13 DSC curve for water structure of VP30 swollen in water with 1.6053 EWF Figure B 14 DSC curve for water structure of VP30 swollen in water with 1.0348 EWF
160 Appendix B (Continued) Figure B 15 DSC curve for water structure of VP30 swollen in water with 0.8 266 EWF Figure B 16 DSC curve for water structure of VP30 swollen in water with 0.5072 EWF
161 Appendix B (Continued) Figure B 17 DSC curve for water structure of VP30 swollen in PBS with 1.8709 EWF Figure B 18 DSC curve for water structure of VP30 swol len in PBS with 1.5754 EWF
162 Appendix B (Continued) Figure B 19 DSC curve for water structure of VP30 swollen in PBS with 1.1570 EWF Figure B 20 DSC curve for water structure of VP30 swollen in PBS with0.9071 EWF
163 Appendix B (Continued) Figure B 21 DSC curve for water structure of VP30 swollen in PBS with 0.6337 EWF Figure B 22 DSC curve for water structure of VP30 swollen in DX 21 with 2.9346 EWF
164 Appendix B (Continued) Figure B 23 DSC curve for water structure of VP30 swollen in DX 21 with 2.3633 EWF Figure B 24 DSC curve for water structure of VP30 swollen in DX 21 with 1.5244 EWF
165 Appendix B (Continued) Figure B 25 DSC curve for water structure of VP30 swollen in DX 21 with 1.1222 EWF Figure B 26 DSC curve for water structure of VP30 swollen in DX 21 with 0.7178 EWF
166 Appendix B (Continued) Figure B 27 DSC curve for water structure of VP45 swollen in water with 2.2428 EWF Figure B 28 DSC curve for water structure of VP45 swollen in water with 1.5505 EWF
167 Appendix B (Continued) Fig ure B 29 DSC curve for water structure of VP45 swollen in water with 1.2132 EWF Figure B 30 DSC curve for water structure of VP45 swollen in water with 0.8414 EWF
168 Appendix B (Continued) Figure B 31 DSC curve for water structure of VP45 swollen in PBS with 1.8890 EWF Figure B 32 DSC curve for water structure of VP45 swollen in PBS with 1.6677 EWF
169 Appendix B (Continued) Figure B 33 DSC curve for water structure of VP45 swollen in PBS with 1.3592 EWF Figure B 34 DSC curve for water structure of V P45 swollen in PBS with 1.2318 EWF
170 Appendix B (Continued) Figure B 35 DSC curve for water structure of VP45 swollen in PBS with 0.7746 EWF Figure B 36 DSC curve for water structure of VP45 swollen in DX 21 with 3.1004 EWF
171 Appendix B (Continued) Figure B 37 DSC curve for water structure of VP45 swollen in DX 21 with 2.1888 EWF Figure B 38 DSC curve for water structure of VP45 swollen in DX 21 with 2.1169 EWF
172 Appendix B (Continued) Figure B 39 DSC curve for water structure of VP45 swollen in D X 21 with 1.8013 EWF
ABOUT THE AUTHOR Chunyan Wang received a Bachelor of Science in Chemistry Education from Shandong Normal University (China) in 2000 and a M.S. in Polymer Chemistry from Fujian Normal University (China) in 2003. After earning M.S., she worke d as a research scientist at the Natural Product Research Center in the National Oceanic Administration (China) until summer 2005. She entered the Ph.D. in Chemistry Program at the University of South Florida in fall 2005 and worked in the Polymer Material s Lab under the supervision of Dr. Julie Harmon.