lecithin-linker microemulsion-based gels for drug … · lecithin-linker microemulsion-based gels...

LECITHIN-LINKER MICROEMULSION-BASED GELS FOR

DRUG DELIVERY

By

Xiao Yue Xuan

A thesis submitted in conformity with the requirements for the degree of Master of Applied Science

Department of Chemical Engineering and Applied Chemistry University of Toronto

© Copyright by Xiao Yue Xuan 2011

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Lecithin-Linker Microemulsion-Based Gels for Drug Delivery

Xiao Yue Xuan

Master of Applied Science

Department of Chemical Engineering and Applied Chemistry University of Toronto

2011

ABSTRACT

Microemulsions have gained interest from the pharmaceutical industry due to their ability to co-

solubilize hydrophilic and lipophilic drugs, and to provide enhanced drug penetration. In this

work, thermosensitive gelatin- and poloxamer 407-stabilized microemulsion-based gels (MBGs)

were formulated using alcohol-free, low toxicity and low viscosity lecithin-based linker

microemulsions. The addition of gelatin to water-rich bicontinuous microemulsions induced the

formation of clear viscoelastic gels containing an oil-rich microemulsion as the gelatin seemed to

dehydrate the original microemulsion. The addition of poloxamer 407 to water-continuous

microemulsions produced MBGs with different gelation temperatures. High concentrations of

lipophilic components in the microemulsion, particularly the oil, reduced sol-gel transition

temperature, while hydrophilic components increased sol-gel transition temperature. Gelatin and

poloxamer MBGs provided desirable viscoelastic properties for ophthalmic and transdermal

applications with minimal impact on the transport properties of the original microemulsions.

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ACKNOWLEDGEMENTS

I would like to express my sincere gratitude to my supervisor Professor Edgar Acosta for his

advice, endless support and encouragement throughout the course of this project. His patience

and guidance has enabled me to complete this thesis work with an understanding and

appreciation of research. I am incredibly fortunate to have Professor Acosta as my advisor. I

would also like to thank Professor Yu-Ling Cheng for her advice, insightful comments, and

constructive criticisms at different stages of the research.

I want to acknowledge the NSERC 20/20 Ophthalmic Materials Network for the financial

support in pursuing this degree. I also want to express my appreciation to the Society of

Cosmetic Chemists (SCC) Ontario Chapter for providing me the opportunities to discuss my

work with professionals in the cosmetic industry, and for the economical support to present this

work in the Society of Cosmetic Chemists Annual Seminar in the United States.

My appreciation to my fellow researchers in the Laboratory of Colloid and Formulation

Engineering (LCFE): Americo Boza, Matthew Baxter, Jacquelene Chu, Oliver Chung, Sumit

Kiran, Suniya Quraishi and Ziheng Wang for their discussion and assistance in the laboratory. I

also want to acknowledge the assistance of the staff of Chemical Engineering and Applied

Chemistry Department with all administrative matters of this work.

Finally, I want to thank my father, Weigang Xuan, my mother, Hong Cui, and my fiancé, Zhanqi

Wu, for their love, support, encouragement and patience.

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TABLE OF CONTENTS

ACKNOWLEDGEMENTS ........................................................................................................ iii

TABLE OF CONTENTS ............................................................................................................ iv

LIST OF TABLES ....................................................................................................................... vi

LIST OF FIGURES .................................................................................................................... vii

CHAPTER 1 Introduction ............................................................................................................1 1.1 Overview ..............................................................................................................................1

1.2 Hypothesis............................................................................................................................3

1.3 Objectives and Scope ...........................................................................................................4

1.4 References ............................................................................................................................5

CHAPTER 2 Background .............................................................................................................8

2.1 Ophthalmic Drug Delivery ..................................................................................................8

2.2 Transdermal Drug Delivery ...............................................................................................11

2.3 Microemulsions..................................................................................................................14

2.4 Microemulsion-based Gels ................................................................................................19

2.5 References ..........................................................................................................................20

CHAPTER 3 Gelatin-Stabilized Microemulsion-Based Gels for Ophthalmic and Transdermal Lidocaine Delivery ...........................................................................................25

3.1 Abstract ..............................................................................................................................25

3.2 Introduction ........................................................................................................................25

3.3 Materials and Methods .......................................................................................................27

3.3.1 Materials ................................................................................................................27

3.3.2 Microemulsion preparation ....................................................................................28

3.3.3 MBG preparation ...................................................................................................29

3.3.4 Construction of ternary phase diagrams ................................................................29

3.3.5 Physiochemical characterization ............................................................................30

3.3.6 Rheological measurements ....................................................................................30

3.3.7 In vitro transport studies ........................................................................................30

3.3.8 Lidocaine quantification ........................................................................................32

3.4 Results and Discussion ......................................................................................................32

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3.4.1 Phase behaviour of lecithin-linker microemulsions ...............................................32

3.4.2 Phase behaviour of lecithin-linker MBGs ..............................................................36


3.5 Conclusions ........................................................................................................................46

3.6 References ..........................................................................................................................47

CHAPTER 4 Poloxamer 407-Stabilized Lecithin Microemulsion-Based Gels for Controlled Ophthalmic and Transdermal Drug Delivery ...................................................51

4.1 Abstract ..............................................................................................................................51

4.2 Introduction ........................................................................................................................51

4.3 Materials and Methods .......................................................................................................53

4.3.1 Materials ................................................................................................................53

4.3.2 Microemulsion phase behaviour and formulation selection ..................................53

4.3.3 MBG preparation ...................................................................................................55

4.3.4 Microemulsion characterization .............................................................................55

4.3.5 Rheological measurements ....................................................................................55


4.3.7 Dexamethasone quantification ...............................................................................56

4.3.8 Lidocaine quantification ........................................................................................56

4.4 Results and Discussion ......................................................................................................56

4.4.1 Phase behaviour of lecithin-linker microemulsions ...............................................56

4.4.2 Phase behavior of microemulsion-poloxamer 407 mixtures .................................59

4.4.3 Effect of microemulsion components on sol-gel transition temperature ...............60


4.5 Conclusions ........................................................................................................................70

4.6 References ..........................................................................................................................70

CHAPTER 5 Conclusions and Recommendations ...................................................................73 5.1 Conclusions ........................................................................................................................73

5.2 Recommendations for Future Work ...................................................................................74

APPENDIX Oil Solubilization Calculations ..............................................................................76

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LIST OF TABLES

Table 2.1 Comparison of emulsions and microemulsions. .......................................................... 14

Table 3.1 Microemulsion formulations. ....................................................................................... 29

Table 3.2 Simulated tear fluid composition [34].......................................................................... 32

Table 3.3 Conductivity measurements (μS/cm) of microemulsions with 2-10% SMO. .............. 33

Table 4.1 Summary of formulations. ........................................................................................... 54

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LIST OF FIGURES

Figure 2.1 Structure of the eye. ...................................................................................................... 8

Figure 2.2 Cross section of human cornea showing 5 layers: epithelium, Bowman’s membrane, stroma, Descemet’s membrane, and endothelium. Adapted from Hogan and Zimmerman [13]. 10

Figure 2.3 Schematic structure of the stratum corneum according to the brick and mortar model. The corneocytes are embedded in a lamellar structured lipid matrix. Adapted from Daniels [34]........................................................................................................................................................ 12

Figure 2.4 A schematic diagram of possible penetration pathways through the skin. Adapted from Hadgraft [31]. ....................................................................................................................... 13

Figure 2.5 Microemulsion phase scan, 1:1 volume ratio (oil:water). Changes in formulation morphology with increasing hydrophilic linker (formulation hydrophilicity): from w/o, bicontinuous, to o/w. Adapted from Acosta et al. [47]. ................................................................ 15

Figure 2.6 Schematic of the lipophilic linker effect. Adapted from Graciaa et al. [56]. ............. 17

Figure 2.7 Fluorescence microscopy images of cross sections of porcine ear skin showing penetration of Nile Red formulated in lecithin-linker microemulsion and in conventional emulsion (0.5h application + 5.5h release). Images provided by Dr. Acosta. .............................. 19

Figure 3.1 Proposed structure of gelatin-stabilized MBG. Adapted from Petit et al. [17]. ......... 26

Figure 3.2 Phase scan at room temperature. All formulations are consisted of 5% lecithin, 10% PEG-6 caprylic/capric glycerides and 4% decaglycerol monocaprylate/caprate. .................... 33

Figure 3.3 Viscosities of the microemulsions of Figure 3.2 containing 2-6% SMO at 25°C. ..... 34

Figure 3.4 Ternary phase diagrams of 6% SMO formulation at 25°C and 50°C. The surfactant vertex is a mixture of 1:2:0.8:1.2 weight ratio of lecithin: PEG-6-caprylic/capric glycerides : decaglycerol monocaprylate/caprate : SMO. The point “∆” indicates the composition of the base formulation containing 6% SMO. .................................................................................... 35

Figure 3.5 Microscopic image through cross polarizing lenses showing the presence of liquid crystals (bright spots) in concentrated aqueous surfactant solutions (20% water–80% surfactant). ................................................................................................................................ 35

Figure 3.6 SMO scan of lecithin-linker MBGs. All formulations contain 10% PEG-6-caprylic/capric glycerides, 4% decaglycerol monocaprylate/caprate, 5% lecithin, 36% of 0.9% NaCl solution and the balance of IPM. (a) Gel appearance. Elastic modulus (G’) vs. temperature profile of (b) 20% gelatin MBGs and (c) 10% gelatin MBGs containing (x) 2% SMO; (▲) 10% SMO; Control (10% or 20% gelatin in 0.9% NaCl solution) (●). ................. 37

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Figure 3.7 Fluorescence image of lecithin-linker MBG (20% gelatin) with 2% SMO labelled with (a) sodium fluorescein and (b) nile red. (c) Cross-polarized images of lecithin-linker MBG. ........................................................................................................................................ 38

Figure 3.8 Fluorescence image of lecithin-linker MBG (20% gelatin) with 10% SMO labelled with (a) sodium fluorescein and (b) nile red. Cross-polarized images of lecithin-linker MBG dyed with (c) sodium fluorescein and (d) nile red. .................................................................. 38

Figure 3.9 (a) Oil dilution path used to produce MBGs with 10% to 50% IPM. (b) Appearance of MBGs at room temperature. ................................................................................................ 40

Figure 3.10 Elastic modulus vs. temperature profile for the systems of Figure 3.9b: (●) 10% IPM; (□) 20% IPM; (▲) 30% IPM; (x) 40% IPM; (*) 50% IPM. Control (20% gelatin in 0.9% NaCl solution) is shown as (○). ...................................................................................... 40

Figure 3.11 (a) Water dilution path of MBGs containing 30% to 90% water. (b) MBG appearance at room temperature. ............................................................................................. 41

Figure 3.12 Cross-polarizer micrographs of MBG containing 80% water (system of Figure 3.11b). ...................................................................................................................................... 42

Figure 3.13 Elastic modulus vs. temperature profile for the systems of Figure 3.11b: (▲) 30% aqueous; (●) 50% aqueous; (□) 80% aqueous; (x) 90% aqueous. Control (20% gelatin in 0.9% NaCl solution) is shown as (○). ................................................................................................ 42

Figure 3.14 Effects of formulation components on elastic modulus of 20% gelatin gels: (●) control (20% gelatin in 0.9% NaCl solution); (x) lecithin + control; (▲) lecithin + SMO + control; (□) PEG-6-caprylic/capric glycerides + decaglycerol monocaprylate/caprate + control; (*) lecithin + SMO + PEG-6-caprylic/capric glycerides + decaglycerol monocaprylate/caprate + control. The surfactant/linker ratios were kept constant at 1:2:0.8:1.2 of lecithin:PEG-6-caprylic/capric glycerides:decaglycerol monocaprylate/caprate:SMO. ..... 43

Figure 3.15 Release profile of lidocaine from the skin after loading with: 20% gelatin MBG ), and lidocaine-loaded gelatin-free microemulsion (∆). ......................................................... 44

Figure 3.16 Permeation profile of lidocaine from in situ gelatin MBG (◆) and from lidocaine-loaded gelatin-free microemulsion (∆). .................................................................................... 46

Figure 4.1 Phase maps of lecithin-linker microemulsions formulated with equal volumes of aqueous and oil (IPM) phases at SMO/lecithin ratio of 1.3/1 at different temperatures. ........ 57

Figure 4.2 Oil solubilization contour diagram. The experimentally measured oil soluilization values are reported beside the “●”, and the contour lines represent those calculated using the HLD-NAC model. .................................................................................................................... 59

Figure 4.3 Behaviours of microemulsion-poloxamer mixtures. Only under certain conditions can thermosensitive poloxamer 407 MBGs be produced. The symbols indicate the formulations evaluated for this study (▲: non-thermosensitive gels; ■: non-thermosensitive liquids; *: thermosensitive MBGs). .......................................................................................................... 60

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Figure 4.4 Effects of microemulsion components on sol-gel transition temperature of systems containing 20% poloxamer 407: (◊) control (water); (x) 3.0% lecithin + water; (▲) 3.1% sorbitan monooleate + water; (□) 19.6% PEG-6-caprylic/capric glycerides + water. ............. 61

Figure 4.5 Elastic modulus vs. temperature profile for the systems containing 3.0% lecithin, 3.1% sorbitan monooleate, 2.5% (▲), 5.0% (■), 10.0% (◊), 15.0% (x), 20.0% (*) or 25.0% (○) PEG-6-caprylic/capric glycerides and balanced amount of water. .................................... 62

Figure 4.6 Elastic modulus vs. temperature profile for the systems containing 3.0% lecithin, 3.1% sorbitan monooleate, 19.6% PEG-6-caprylic/capric glycerides, 0% (*), 2.0% (□), 4.0% (▲), 6.0% (x) or 8.0% (◆) IPM and balanced amount of water. ............................................ 63

Figure 4.7 Elastic modulus vs. temperature profile for the systems containing 3.0% lecithin, 3.1% sorbitan monooleate, 25% PEG-6-caprylic/capric glycerides, 4.0% (□), 8.0% (▲) or 10.0% (x) IPM and balanced amount of water. ....................................................................... 64

Figure 4.8 Behaviours of microemulsion-poloxamer mixtures with constant PEG-6-caprylic/ capric glycerides (PEG-6-CCG) and oil contour lines. ............................................................ 65

Figure 4.9 Release profile of dexamethasone from the skin after loading with: 20% poloxamer 407 MBG (◆), and dexamethasone-loaded gelatin-free μE (∆). Dexamethasone loading in the donor compartment was 23.5±6.1 μg/cm2 for the MBG, and 28.5±8.9 μg/cm2 for the microemulsion. ......................................................................................................................... 67

Figure 4.10 Permeation profile of dexamethasone from in situ poloxamer 407 MBG (◆), dexaemthasone-loaded poloxamer-free microemulsion (∆), and poloxamer 407 hydrogel (●). Dexamethasone loading in the donor compartment was 281±8 µg/cm2 for the MBG, 352±8 µg/cm2 for the microemulsion, and 137±8 µg/cm2

for the hydrogel. ....................................... 68

Figure 4.11 Permeation profile of lidocaine from in situ poloxamer 407 MBG (◆) and lidocaine-loaded poloxamer-free microemulsion (∆). Lidocaine loading in the donor compartment was 1150±50 µg/cm2 for the poloxamer MBG, and 1380±40 µg/cm2 for the microemulsion. ......................................................................................................................... 69

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CHAPTER 1

Introduction

1.1 Overview

Ophthalmic drug delivery is a challenging task due to the unique anatomy, physiology, and

biochemistry of the eye that protects the organ from exogenous compounds [1]. Following

topical instillations, the applied drug barely penetrates the cornea due to the presence of the

lipophilic corneal epithelium which prevents the entry of large and hydrophilic solutes, while the

hydrophilic stroma below the epithelium limits the passage of lipophilic solutes [2]. Thus, for a

drug or delivery vehicle to cross the cornea effectively, it has to possess both hydrophilic and

lipophilic properties, and be of small size to pass through the tight junctions [1].

Similarly, in transdermal drug delivery, effective penetration of drugs through the stratum

corneum is also a major challenge [3]. Despite a thickness of only 15-20 µm, the presence of

lipid matrices render the stratum corneum a lipophilic barrier that restrict the permeation of large

and hydrophilic solutes [4,5]. Several strategies have been proposed to overcome the resistence

of the stratum corneum and to increase skin permeation. Microemulsions, which are clear,

thermodynamically stable mixtures of oil, water and surfactant [6], have been shown to be able

to deliver drugs through the skin better than aqueous solutions or creams [7,8,9,10].

Compared to conventional drug delivery vehicles, formulations based on microemulsions

provide the ability to co-solubilize oil-soluble and water-soluble drugs. Moreover, these systems

have high surface areas resulting from nano-scale droplets (1-100 nm) [6]. As a result,

microemulsions formulated from biodegradable and biocompatible surfactants have become

promising delivery systems for oral, topical, and transdermal administration. Eye drops

formulated using microemulsions have also been proposed as an alternative [11,12,13]. These

microemulsions are potential ophthalmological carriers because of their transparency, and their

improvement in drug solubility and stability, which could potentially provide increased

bioavailability especially for poorly water-soluble drugs.

Among the various microemulsion systems, lecithin microemulsions are especially desirable

since lecithin is a naturally occurring nontoxic biological surfactant with Generally Recognized

as Safe (GRAS) status [14]. However, lecithin cannot produce microemulsions when utilized as

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the sole surfactant because of its tendency to form liquid crystalline phases [10]. Earlier lecithin

microemulsions made use of medium-chain alcohols, such as butanol, to prevent the formation

of liquid crystals and promoting microemulsion formation [10]. Unfortunately, these medium-

chain alcohols tend to disrupt cell membranes [15], making the formulations cytotoxic. Alcohol-

free lecithin microemulsions have been formulated using linker molecules to replace alcohols

[14,16,17]. The addition of linkers helps increasing the surfactant-oil (lipophilic linkers) and

surfactant-water (hydrophilic linkers) interactions and also offers enhanced solubilization

capacity [18,19]. Compared to conventional alcohol-based lecithin microemulsions, linker-based

lecithin microemulsions have a low toxicity, since all ingredients involved in microemulsion

preparation are food or pharmaceutical grade [17,20].

Compared to other pharmaceutical-grade microemulsions, lecithin-linker microemulsions have

an exceptionally low viscosity (< 40 mPa·s) that together with their affinity towards hydrophilic

and lipophilic environments, and small drop size (<10 nm) makes them a suitable vehicle to

penetrate epithelial tissue and use these tissues as a depot for drug delivery [16,20]. While the

low viscosity of lecithin-linker microemulsions is suitable for spray or roll-on applications, these

linker microemulsions spread well beyond the intended area when applied as drops [20]. The low

viscosity is also an undesirable feature in periocular ophthalmic delivery where higher viscosities

are desirable to increase the residence time and improve the effectiveness of ophthalmic delivery

formulations [21,22]. The objective of this project is to introduce gelling agents, particularly

gelatin and poloxamer 407, into lecithin-linker microemulsions with the purpose of increasing

the viscosity of the delivery system while retaining the transport properties observed in low-

viscosity microemulsions.

Microemulsion-based gels (MBGs) are stimuli-responsive (e.g. pH, temperature, etc.), high

viscosity systems prepared from microemulsions with the addition of one or multiple gelling

agents [23]. These systems have been reported in literature as potential topical delivery vehicles

for a variety of drugs, including dexamethasone [24], clotrimazole [25], and ibuprofen [26].

Thermosensitive, in situ-gellling MBGs is a category of MBGs that can undergo sol-gel phase

transition upon variation in temperature. These systems are potential drug carriers which can

offer the advantages of both liquid and semi-solid (i.e. gel) delivery systems. As a result, there is

the possibility that these MBGs can be used to prepare in situ-gelling eye drop formulations and

in situ-forming skin patches. For example the MBGs can be applied in their liquid state for easy

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application and faster drug penetration, and then undergo phase transition to form high viscosity

gels in the cul-de-sac (ophthalmic delivery) or inside the epidermis layer of human skin

(transdermal delivery), thus providing longer residence time and the desired extended release

effect.

The preparation of MBGs was first reported in 1986 and there are various articles that describe

their physical/structural characteristics [23,27,28,29,30]. The gelling agents used to formulate

these MBGs include natural polymers such as gelatin [31], κ-carrageenan [31],

hydroxypropylmethyl cellulose (HPMC) [32], and non-ionic surfactant poly(ethylene oxide)-

poly(propylene oxide)-poly(ethylene oxide), commercially known as Poloxamers (ICI) or

Pluronics (BASF) [29]. Among these, gelatin-stabilized MBGs are the most reported systems.

However, the major limitation of gelatin MBGs is the toxicity associated with the surfactants

used to formulate these microemulsions. In most cases, the presence of anionic sodium dioctyl

sulfosuccinate (AOT), either as the surfactant alone or in combination with other non-ionic

surfactants, is required for the formation of gelatin MBGs [33,34]. Gelatin MBGs prepared from

lecithin microemulsions have also been reported however the formulations contain short or

medium chain alcohols [31,32,35]. The presence of AOT and/or alcohols in the formulations

poses cytotoxicity concerns when utilizing gelatin MBGs as potential drug delivery vehicles.

Poloxamer 407-stabilized systems are thermosensitive, in situ-forming gels for extended drug

delivery. These systems, however, have only been studied as hydrogels [36,37,38] or MBGs with

low (≤ 3.5%) oil content [29]. For pharmaceutical applications, a larger quantity of solubilized

oil in the microemulsion is ideal for the delivery of a higher dosage of oil-soluble drugs.

1.2 Hypothesis

Thermosensitive, gelatin- and poloxamer 407-stabilized MBGs with viscosities suitable for

ophthalmic and transdermal applications can be prepared using alcohol-free, low toxicity

lecithin-linker microemulsions. Moreover, it is expected that these gels will not interfere with the

ability of the microemulsions to penetrate epithelial tissue and use the tissue as drug depot.

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1.3 Objectives and Scope

The overall objectives of this study are to produce thermosensitive microemulsion-based gels

(MBGs) using low toxicity lecithin-linker microemulsions and investigate their potential as

ophthalmic and transdermal delivery systems. The specific aims of this work include:

1. To produce gelatin- and poloxamer 407-stabilized MBGs using alcohol-free lecithin-

linker microemulsions;

2. To evaluate the effect of specific formulation components on the overall properties of

gelatin and poloxamer 407 MBGs;

3. To explore the effectiveness of gelatin and poloxamer 407 MBGs as trans-epithelial

delivery vehicles.

This thesis is organized in five chapters. Chapter 1 presents a brief introduction of the research

topic, hypothesis, objectives and scope. Chapter 2 provides an overview of the literature.

Background information including ophthalmic and transdermal drug deliveries, microemulsions,

linker molecules, lecithin-linker microemulsions and microemulsion-baesd gels are covered.

The main contributions of this study are included in Chapters 3 and 4. Chapter 3 reports the

production of alcohol-free gelatin MBGs using lecithin-linker microemulsions and their potential

to act as opththalmic and transdermal delivery vehicles. The food/pharmaceutical grade

microemulsions consist of a combination of lecithin (surfactant), PEG-6-caprylic/capric

glycerides and decaglycerol monocaprylate/caprate (hydrophilic linkers), sorbitan monooleate

(lipophilic linker), and isopropyl myristate (IPM, “carrier” oil for lipophilic drug, lidocaine).

Study of rheological properties of gelatin MBGs show that the addition of gelatin to oil-in-water

microemulsions produces clear viscoelastic gels of bicontinuous morphology, and the addition of

gelatin to microemulsion systems containing excess water produces milky but strong gels

containing emulsified microemulsions. The in vitro transport studies demonstrate release of

lidocaine from gelatin MBGs for more than 12 hours with minimal effect on the release profile.

This contribution has been submitted to the online journal Pharmaceutics.

Chapter 4 describes the preparation of poloxamer 407-stabilized MBGs using lecithin-linker

microemulsions. Rheological characterizations of poloxamer 407 MBGs with different

surfactant-linker-oil-water compositions reveal that the interactions among poloxamer 407,

hydrophilic PEG-6-caprylic/capric glycerides, and lipophilic components (oil, lipophilic linker

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and lecithin) control the gelation temperature and overall gel properties. The in vitro transport

studies show that poloxamer 407 MBGs provide sustained and near zero-order release of

dexamethasone for 48 hours. This contribution has been prepared for submission to the

International Journal of Pharmaceutics.

In Chapter 5, the overall conclusions drawn from this study and some recommendations for

future directions are presented.

1.4 References

1. Mitra, A.K.; Anand, B.S.; Duvvuri, S. Drug Delivery to the Eye. Advances in Organ Biology Volume 10: The Biology of the Eye; Elsevier B.V.: Amsterdam, The Netherlands, 2006; pp 307-351.

2. Davies, N.M. Biopharmaceutical considerations in topical ocular drug delivery. Clin. Exp. Pharmacol. P. 2000, 27, 558-562.

3. Walters, K.A. Dermatological and Transdermal Formulations; Marcel Dekker: New York, NY, USA, 2002; pp 4.

4. Prausnitz, M.R.; Langer, R. Transdermal drug delivery. Nat. Biotechnol. 2008, 26, 1261-1268.

5. Hadgraft, J. Skin, the final frontier. Int. J. Pharm. 2001, 224, 1-18.

6. Kogan, A.; Garti, N. Microemulsion as transdermal drug delivery vehicles. Adv. Colloid Interface Sci. 2006, 123-126, 369-385.

7. Lawrence, M.J.; Rees, G.D. Microemulsion-based media as novel drug delivery systems. Adv. Drug Delivery Rev. 2000, 45, 89-121.

8. Bagwe, R.P.; Kanicky, J.R.; Palla, B.J.; Patanjali, P.K.; Shah, D.O. Improved drug delivery using microemulsions: rationale, recent progress, and new horizons. Crit. Rev. Ther. Drug Carrier Syst. 2001, 18, 77-140.

9. Kreilgaard, M., Pedersen, E.J., Jaroszewski, J.W. NMR characterisation and transdermal drug delivery potential of microemulsion systems. J. Controlled Release 2000, 69, 421-433.

10. Tenjarla, S. Microemulsions: an overview and pharmaceutical applications. Crit. Rev. Ther. Drug 1999, 16, 461-521.

11. Vandamme, Th.F. Microemulsions as ocular drug delivery systems: recent developments and future challenges. Prog. Retin. Eye Res. 2002, 21, 15-34.

12. Ligório Fialho, S.; da Silva-Cunha, A. New vehicle based on a microemulsion for topical ocular administration of dexamethasone. Clin. Exp. Ophthalmol. 2004, 32, 626-632.

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13. Haße, A.; Keipert, S. Development and characterization of microemulsions for ocular application. Eur. J. Pharm. Biopharm. 1997, 43, 179-183.

14. Acosta, E.J.; Nguyen, T.; Witthayapanyanon, A.; Harwell, J.H.; Sabatini, D.A. Linker-based bio-compatible microemulsions. Environ. Sci. Technol. 2005, 39, 1275-1282.

15. McKarns, S.C.; Hansch, C.; Caldwell, W.S.; Morgan, W.T.; Moore, S.K.; Doolittle, D.J. Correlations between hydrophobicity of short-chain aliphatic alcohols and their ability to alter plasma membrane integrity. Toxicol. Sci. 1997, 36, 62-70.

16. Acosta, E.; Chung, O.; Xuan, X.Y. Lecithin-linker microemulsions in transdermal delivery. J. Drug Deliv. Sci. Technol. 2011, 21, 77-87.

17. Yuan, J.S.; Ansari, M.; Samaan, M.; Acosta, E.J. Linker-based lecithin microemulsions for transdermal delivery of lidocaine. Int. J. Pharm. 2008, 349, 130-143.

18. Acosta, E.J.; Harwell, J.; Sabatini, D. Self-assembly in linker-modified microemulsions. J. Colloid Interface Sci. 2004, 274, 652-664.

19. Uchiyama, H.; Acosta, E.J.; Tran, S.; Sabatini, D.; Harwell, J. Supersolubilization in chlorinated hydrocarbon microemulsions: Solubilization enhancement by lipophilic and hydrophilic linkers. Ind. Eng. Chem. Res. 2000, 39, 2704-2708.

20. Yuan, J.S.; Acosta, E.J. Extended release of lidocaine from linker-based lecithin microemulsions. Int. J. Pharm. 2009, 368, 63-71.

21. Ali, Y.; Lehmussaari, K. Industrial perspective in ocular drug delivery. Adv. Drug Deliver. Rev. 2006, 58, 1258-1268.

22. Bapatla, K.M.; Hecht, G. Ophthalmic ointments and suspensions. Pharmaceutical Dosage Forms: Disperse Systems; Marcel Dekker: New York, NY, USA, 2006; pp 357-397.

23. Luisi, P.L.; Scartazzini, R.; Haering, G.; Schurtenberger, P. Organogels from water-in-oil microemulsions. Colloid Polym. Sci. 1990, 268, 356-374.

24. Chandra, A.; Sharma, P.K.; Irchhiaya, R. Microemulsion-based hydrogel formulation for transdermal delivery of dexamethasone. Asian J. Pharm. 2009, 3, 30-36.

25. Bachhav, Y.G.; Patravale, V.B. Microemulsion-based vaginal gel of clotrimazole: Formulation, in vitro evaluation, and stability studies. AAPS PharmSciTech 2009, 10, 476-481.

26. Chen, H.; Chang, X.; Du, D.; Li, J.; Xu, H.; Yang, X. Microemulsion-based hydrogel formulation of ibuprofen for topical delivery. Int. J. Pharm. 2006, 315, 52-58.

27. Haering, G.; Luisi, P.L. Hydrocarbon gels from water-in-oil microemulsions. J. Phys. Chem. 1986, 90, 5892-5895.

28. Quellet, C.; Eicke, H.F.; Sager, W. Formation of microemulsion-based gelatin gels. J. Phys. Chem. 1991, 95, 5642-5655.

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29. Zhao, X.Y.; Xu, J.; Zheng, L.Q.; Li, X.W. Preparation of temperature-sensitive microemulsion-based gels formed from a triblock copolymer. Colloids Surf., A 2007, 307, 100-107.

30. Kantaria, S.; Rees, G.D.; Lawrence, M.J. Gelatin-stabilised microemulsion-based organogels: rheology and application in iontophoretic transdermal drug delivery. J. Controlled Release 1999, 60, 355-365.

31. Stamatis, H.; Xenakis, A. Biocatalysis using microemulsion-based polymer gels containing lipase. J. Mol. Catal. B: Enzym. 1999, 6, 399-406.

32. Blattner, C.; Zoumpanioti, M.; Kroner, J.; Schmeer, G.; Xenakis, A.; Kunz, W. Biocatalysis using lipase encapsulated in microemulsion-based organogels in supercritical carbon dioxide. J. Supercrit. Fluids 2006, 36, 182-193.

33. Patravale, V.B.; Date, A.A. Microemulsions: applications in transdermal and dermal delivery. Crit. Rev. Ther. Drug Carrier Syst. 2007, 24, 547-596.

34. Zoumpanioti, M.; Stamatis, H.; Xenakis, A. Microemulsion-based organogels as matrices for lipase immobilization. Biotechnol. Adv. 2010, 28, 395-406.

35. Nagayama, K.; Yamasaki, N.; Imai, M. Fatty acid esterification catalyzed by Candida rugosa lipase in lecithin microemulsion-based organogels. Biochem. Eng. J. 2002, 12, 231-236.

36. Qi, H.; Li, L.; Huang, C.; Li, W.; Wu, C. Optimization and physicochemical characterization of thermosensitive poloxamer gel containing puerarin for ophthalmic use. Chem. Pharm. Bull. 2006, 54, 1500-1507.

37. Escobar-Chávez, J.J.; López-Cervantes, M.; Naïk, A.; Kalia, Y.N.; Quintanar-Guerrero, D.; Ganem-Quintanar, A. Applications of thermo-reversible Pluronic F127 gels in pharmaceutical formulations. J. Pharm. Pharmaceut. Sci. 2006, 9, 339-358.

38. Ricci, E.J.; Bentley, M.V.L.B.; Farah, M.; Bretas, R.E.S.; Marchetti, J.M. Rheological characterization of Poloxamer 407 lidocaine hydrochloride gels. Eur. J. Pharm. Sci. 2002, 17, 161-167.

8

CHAPTER 2

Background

This chapter presents a review of the challenges and some recent advances in the field of

ophthalmic and transdermal drug delivery. Background information on formulation constraints

associated with lecithin-based microemulsions and the approach of using linker molecules are

also reviewed.

2.1 Ophthalmic Drug Delivery

The eye is an important organ with unique anatomy and physiology. Its protective barriers

restrict the entry of drug molecules at the required site of action, and offer many challenges to

the development of effective ophthalmic delivery systems. As a result, ophthalmic drug delivery

has remained as one of the most challenging task for pharmaceutical scientists [1].

Based on the structure of the eye, ophthalmic drug delivery can be broadly classified into two

categories: anterior and posterior segments. The anterior eye segment consists of the front one-

third of the eye, including pupil, cornea, iris, ciliary body, aqueous humour and lens (Figure 2.1)

[2]. The posterior segment is consisted of the back two-thirds of the eye that includes vitreous

humour, retina, choroid, macula and optic nerve (Figure 2.1) [3]. This review will focus on

delivery to the anterior segment of the eye.

Figure 2.1 Structure of the eye.

9

Currently, ophthalmic diseases are most commonly treated with topical medications

administered as eye drops. These conventional dosage forms account for 90% of the available

ophthalmic formulations due to their ease of administration, safety and low cost [1,4,5]. In fact,

eye drops were already used at the time of Cleopatra for the treatment of ocular conditions [6].

Nowadays, the majority of topical ophthalmic formulations are in the form of aqueous solutions,

which can be manufactured in large scale for the delivery of hydrophilic drugs [6]. Viscosity

modifying agents, including polyvinyl alcohol, hydroxypropyl cellulose, and natural polymers

such as hyaluronic acid and guar gum, are usually incorporated to enhance formulation viscosity

and retention time in the precorneal space in order to improve drug bioavailability [5]. Systane®

(aqueous eye drop containing 0.4% w/v polyethylene glycol 400, 0.3% w/v propylene glycol;

Alcon Laboratories, Inc., USA) is a commercial product available in both liquid and gel (gelling

agent: hydroxypropyl guar) forms for temporary relief of burning and irritation due to dryness of

the eye. Moreover, with the discovery of many new lipophilic drugs, ophthalmic suspensions are

becoming an important dosage form for the delivery of poorly water-soluble drugs. The drugs

are typically homogeneously suspended in aqueous solutions due to the presence of many

inactive ingredients in the formulation, such as dispersing and wetting agents, suspending agents

and buffers [6]. The average particle size for the suspensions are typically less than 10 µm [6].

Maxidex® (Alcon Laboratories, Inc., USA) is an example of ophthalmic suspension containing

0.1% w/v dexamethasone. However, despite their popularity, aqueous eye drops (solutions and

suspensions) present very low drug absorption in the cornea, and the drugs are usually lost into

the systemic circulation which lead to undesirable systemic side effects [7,8].

Ophthalmic ointments are another topical dosage form mainly for night-time use and when

prolonged therapeutic actions are required [6]. GenTeal® night-time lubricant eye ointment

(Norvartis International AG, Switzerland), Bausch & Lomb moisture eyes PM ointment (Bausch

& Lomb Inc., USA), and Systane® night-time lubricant eye ointment (Alcon Laboratories, Inc.,

USA) are examples of commercially available products. While ophthalmic ointments improve

drug bioavailability by providing extended contact time of the drug with the corneal tissues,

patients suffer from inaccurate dosing, matted eyelids and blurred vision due to the refractive

index differences between the tears and the non-aqueous nature of the ointment base [9].

The anterior segment of the eye has various protective mechanisms for maintaining visual

functions that contribute to low drug bioavailability. For example, after instillation of an

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ophthalmic drug, the majority of it is rapidly eliminated from the precorneal area due to drainage

via the nasolacrimal system into the nasopharynx and the gastrointestinal tract [10]. This

drainage typically takes place when the volume of fluid in the cul-de-sac exceeds the normal

lacrimal volume of 7-10 µl [6,10]. As a result, approximately 90% of the 50-100 µl dose

(corresponding to one to two drops) administered as eye drops is cleared within 2 min, providing

a maximum of 2 min contact time of the drug with the absorbing surfaces (cornea and sclera)

[6,10,11]. Additional factors such as blinking, dilution by lacrimation and tear turnover

(approximately 1 µl/min), protein binding and metabolic degradation in the tear film also

contribute to drug loss [10], where less than 5% of the administered dose permeates the eye [12].

In addition, the cornea itself is a highly selective barrier with five different layers that exclude

compounds from the eye (Figure 2.2) [6,10]. The most significant barriers in ocular drug

absorption are the lipophilic corneal epithelium and the hydrophilic stroma. The epithelium is

consisted of 5-6 layers of squamous stratified cells with a total thickness of around 50-100 µm

and a turnover rate of about one cell layer per day [10]. The cells of all the layers of the

epithelium are separated by a 10-20 nm intercellular space [10]. The superficial epithelial cells

are connected by adherent junctions and tight junctions, establishing a lipophilic diffusion barrier

in the surface of the epithelium [10]. The stroma, while accounts for approximately 90% of the

corneal thickness (450 µm), is composed of about 80% of water and has a relatively open

structure that will normally allow the diffusion of hydrophilic solutes [10]. The epithelium and

stroma combined lead to poor corneal absorption of both hydrophilic and lipophilic drugs.

Indeed, for a drug to cross the cornea effectively, it has to posses both hydrophilic and lipophilic

properties, and be of low molecular size to pass through the tight junctions [10].

Figure 2.2 Cross section of human cornea showing 5 layers: epithelium, Bowman’s membrane, stroma, Descemet’s membrane, and endothelium. Adapted from Hogan and Zimmerman [13].

11

Many strategies have been implemented to overcome the disadvantages associated with eye drop

formulations and to increase ocular absorption of ophthalmic drugs. Most of them aim at

increasing drug residence time in the cornea and conjunctiva sac, and at reducing precorneal

drug loss for the purpose of maximizing ocular drug absorption and permeability [4]. Examples

of these ophthalmic drug delivery systems include micro- and nanoparticles [14], liposomes [15],

and collagen shields [16]. On the other hand, microemulsions provide a promising alternative to

topical ophthalmic drug delivery [17,18,19]. Microemulsions are transparent, thermodynamically

stable systems that can be prepared and sterilized simply and inexpensively. Due to the co-

existence of aqueous and oil domains, microemulsions offer the ability to solubilize both

hydrophilic and lipophilic drugs in the formulation, and provide enhanced drug solubilization.

Siebenbrodt and Keipert [20] demonstrated improved solubilization of indomethacin,

chloramphenicol and sodium diclofenac in microemulsions when compared to aqueous eye

drops. Microemulsions in ophthalmic delivery can also provide prolonged dose effects. Naveh et

al. [21] and Hasse and Keipert [22] showed longer intra-ocular pressure reduction effect and

enhanced bioavailability of pilocarpine from oil-in-water microemulsions. The possibility of

prolonged drug release from microemulsions makes them attractive delivery vehicles for ocular

administration due to the potential of decreased daily eye drop instillation frequency [23].

The ideal ophthalmic dosage form would be one that provides sustained drug release and remains

in contact with the cornea for extended periods of time [24]. If the precorneal residence time of a

drug can be improved from a few minutes to a few hours, improved local bioavailability, reduced

dose concentrations and dosing frequency, and improved patient acceptability can potentially be

achieved. While liquid dosage forms offer advantages such as ease of administration and better

patient compliance, semi-solid systems such as gels provide extended corneal contact time and

hence improved bioavailability. In situ-gelling, microemulsion-based drug delivery systems have

the potential to offer the advantages of both liquid and semi-solid systems. A detailed review of

these systems will be provided in Section 2.4.

2.2 Transdermal Drug Delivery

Transdermal drug delivery (TDD) refers to the diffusion of drugs through skin into the systemic

circulation for distribution and therapeutic effects [25]. Compared to oral and parenteral delivery

routes, transdermal delivery of drugs offers many advantages, including avoidance of first-pass

metabolism, minimizing undesirable side effects, non-invasiveness, better patient compliance,

12

and potential for continuous or controlled delivery [26]. In addition, TDD provides a convenient

route of administration for a variety of clinical indications and there is reversibility of drug

application [26]. Current TDD dosage forms (chemical approach) include skin patches, creams,

gels, ointments or sprays [27]. This review does not include physical approaches such as micro-

needles, ultrasound, radio waves, etc. [28].

As the largest organ of the human body, the skin provides a painless and compliant interface for

systemic circulation [29]. One of the most important functions of the skin is to prevent

penetration of external pathogens and toxins into the body [30]. Associated with this protective

function is the barrier property of the skin. The human skin is consisted of two main layers: the

epidermis and the dermis. The epidermis is the outermost layer and is typically 100 µm thick.

However, the impermeability of the skin is mainly provided by the uppermost layer of epidermis,

the stratum corneum, which is approximately 15-20 µm thick [29,31].

The stratum corneum is composed of dead, flattened and interlocked keratin-rich cells, called

corneocytes, embedded in a lipid matrix. The structure of stratum corneum can be roughly

described by a “brick and mortar” model [32], where the corneocytes of hydrated keratin are the

“bricks” and the epidermal lipids fill the intercellular spaces like mortar (Figure 2.3). The lipids

found in the stratum corneum are mainly ceramides, fatty acids, cholesterol and cholesterol

esters, and they typically comprise 10 to 30% of the total volume of the stratum corneum [33]. In

addition, the lipids are organized as multiple lipid bilayers which form regions of semi-

crystalline gel and liquid crystals [34]. Due to the presence of these lipid matrices, the stratum

corneum is lipophilic in nature, giving the skin the selective permeability (i.e. poor diffusion of

large and hydrophilic molecules). The barrier function is further facilitated by the continuous

desquamation, with a total turnover of the stratum corneum occurring once every 2-3 weeks [30].

Therefore, the stratum corneum is recognized as the rate controlling hydrophobic barrier of the

skin [25,29].

Figure 2.3 Schematic structure of the stratum corneum according to the brick and mortar model. The corneocytes are embedded in a lamellar structured lipid matrix. Adapted from Daniels [34].

13

There are three potential pathways through which drug molecules in contact with the skin can

penetrate (Figure 2.4): transceullar, intercellular and appendageal/shunt (through sweat ducts,

hair follicles and sebaceous glands). Scheuplein and colleagues [35,36,37] proposed that the

appendageal route was responsible for the pre-steady state permeation of polar molecules and the

flux of large polar molecules or ions that have difficulty diffusing across the intact stratum

corneum. Nevertheless, the appendages occupy a relatively low fraction of the skin area (~0.1%),

and their contribution to the steady state flux of most drugs is minimal [38]. It is thought

conventionally that lipophilic compounds (octanol/water partition coefficient, log Pow > 2)

diffuse through the lipid matrices between the keratin filaments (intercellular route) while

relatively hydrophilic compounds diffuse within the aqueous regions near the outer surface of

intracellular filaments (transcellular route) [26]. However, recent research work has shown that

the intercellular route is now considered to be the major pathway for permeation of most drugs

across the stratum corneum [39]. As a result, most of the current techniques in optimizing drug

permeation across the skin aim at manipulating solubility in the lipid domain or altering the

ordered structure of the stratum corenum.

Figure 2.4 A schematic diagram of possible penetration pathways through the skin. Adapted from Hadgraft [31].

Delivering sufficient quantities of a drug transdermally to reach the systemic circulation in the

right time frame is not a simple task [40]. Researchers are challenged to come up with

formulations that increase the permeability of the drug without irreversibly altering the stratum

corneum [41]. Potential mechanisms to enhance drug permeation through the skin include

directly affecting the skin and modifying the formulation so the partition, diffusion, or solubility

of the drug can be altered [27,31]. Chemical penetration enhancers such as surfactants (e.g.

Azone® molecules, Tween, Span and phospholipids), fatty acid/esters (e.g. oleic acid), terpenes

(e.g. limonene) and solvents (e.g. dimethylsulfoxide and ethanol) are used in the formulations to

increase drug diffusion across the skin [27,42]. In recent years, colloidal drug carriers such as

14

emulsions, micelles, liposomes and other deformable vesicular structures have been studied

extensively in topical delivery of cosmetic and dermatological agents [43]. Another modern

colloidal carrier, microemulsion, is a promising alternative in providing enhanced transdermal

drug delivery due to their spontaneous formation and ability to co-solubilize high concentrations

of both water-soluble and oil-soluble pharmaceutical ingredients [42].

2.3 Microemulsions

The concept of microemulsion was introduced in the early 1940s by Hoar and Schulman in

which transparent water-in-oil systems were observed [44]. However, the term “microemulsion”

was first used later in 1959 by Schulman et al. to describe a multiphase system consisting of

water, oil, surfactant and alcohol which forms a transparent solution [45]. By definition,

microemulsions are clear, macroscopically isotropic fluid mixtures of water, oil and surfactants

[46]. In addition, the thermodynamic stability and nano-structure (droplet sizes of 1-100 nm) of

microemulsions are important characteristics that differentiate them from normal coarse

emulsions, which are not thermodynamically stable. Appearance-wise, emulsions are typically

cloudy; they are kinetically stabilized dispersions that tend to phase separate over time.

Microemulsions, on the other hand, are typically transparent and are stable over a long range of

time. Table 2.1 lists the main differences between emulsions and microemulsions.

Table 2.1 Comparison of emulsions and microemulsions.

Property Emulsion Microemulsion

Droplet Size > 500 nm

(Nanoemulsions: 1-100 nm) 1-100 nm

Appearance Milky Transparent/translucent

Stability Thermodynamically unstable (kinetically stable) Thermodynamically stable

Formation External energy input required Spontaneous; minimal energy input required

Interfacial Tension ~ 50 mN/m Towards 0 mN/m

Investigating the phase behavior of a microemulsion system is important in understanding the

changes in microemulsion morphology by changing certain conditions. This phase behaviour can

be explored by using phase scans with different combinations of formulation variables, such as

15

salinity, surfactants, and temperature. This is accomplished by keeping all formulation variables

(e.g. type of oil, surfactant, temperature, electrolyte concentration, etc.) constant, except for one

(the scanning variable). For example, the water and oil volume ratios in a phase scan is typically

fixed at 1:1 while the hydrophilic linker is increased along the scan as shown in Figure 2.5.

Depending on the surfactants and the formulation, four types of microemulsion systems can be

classified: Winsor Type I oil-in-water (o/w) system coexisting with an excess oil phase, Type II

water-in-oil (w/o) microemulsion coexisting with an excess aqueous phase, and Type III/IV

bicontinuous microemulsion systems. For a Type III system, the microemulsion coexists with

excess oil and water phases whereas in a Type IV system both the water and oil are completely

solubilized in a single isotropic phase (Figure 2.5). In addition, for non 1:1 (water:oil) volume

ratios, Types I and II may or may not have any excess phases. For example an 8:2 (water:oil)

volume ratio may produce a single phase microemulsion that is potentially a Type I system.

Figure 2.5 Microemulsion phase scan, 1:1 volume ratio (oil:water). Changes in formulation morphology with increasing hydrophilic linker (formulation hydrophilicity): from w/o, bicontinuous, to o/w. Adapted from Acosta et al. [47].

Microemulsions have several advantages as drug delivery vehicles in pharmaceutical

applications [42,48,49]:

1. Enhanced solubilization capacity: due to the co-existence of oil, water and amphiphilic

surfactants, microemulsions enable the entrapment of lipophilic, hydrophilic and

amphiphilic drug entities;

16

2. Thermodynamic stability: this feature helps to provide an extended shelf-life for the

pharmaceutical agent incorporated in the microemulsions;

3. Improved bioavailability: microemulsions have been shown to improve bioavailability of

several drugs, especially in oral drug delivery;

4. Ease of manufacture: the spontaneous formation of microemulsions facilitates their

manufacture.

The major disadvantage of microemulsions for drug delivery arises from the high concentrations

of surfactants and cosurfactants necessary for stabilizing the nanodroplets. High levels of

surfactants and cosurfactants in the formulation can be toxic and may trigger unwanted immune

responses such as irritation and allergies. However, these side effects can be reduced by

replacing ionic surfactants with nonionic or zwitterionic surfactants [49].

Lecithin (phospholipids) is an example of a zwitterionic surfactant. It is naturally derived from

eggs, soybean, etc. and has Generally Recognized as Safe (GRAS) status [50]. Lecithin is a

byproduct of the degumming process of vegetable oils [47]. Soybean is typically the preferred

source for lecithin. Soybean lecithin is a mixture of phosphatidylcholine (main component),

phosphatidylethanolamine, phosphatidylinositol, some free oil and other phospholipids [51].

Since phospholipids are present in cell membranes and can be found in all plants and animals,

they do not introduce the toxicity and sensitivity problems associated with other surfactants, even

when administered in high concentrations [47]. As a result, lecithin microemulsions are

especially desirable and advantageous in topical drug delivery.

However, it is difficult to produce microemulsions with lecithin alone. Chemically, lecithin is

phosphatidylcholine. When many phosphatidylcholine molecules are placed in water, they tend

to self-assemble in rigid bilayers [52]. Moreover, due to their tendency to form lamellar or other

liquid crystal phases, it is necessary to use medium-chain alcohols, such as butanol and pentanol,

as cosurfactants to induce the formation of lecithin microemulsions [53]. The disadvantage to

using medium-chain alcohols is that they disrupt cell membranes, making the formulations more

cytotoxic [54].

An alternative to using medium-chain alcohols is the use of hydrophilic and lipophilic linker

molecules, which improve surfactant-water and surfactant-oil interactions, respectively [55]. The

concept of linker molecules was introduced by Graciaa et al. with the introduction of the

17

lipophilic linker [56,57,58]. Lipophilic linkers are molecules that present in the oil phase, which

orientate along the surfactant tails to help promote orientation of the oil molecules. In other

words, lipophilic linkers serve as a link between oil molecules and the surfactant tails [56].

Figure 2.6 illustrates the orientation effect caused by lipophilic linkers [56]. The main difference

between lipophilic linkers and cosufactants, which are typically medium-chain alcohols (4-9

carbons) and amines [59,60], is that while both of them interact strongly with the oil molecules,

lipophilic linkers do not absorb at the interface whereas cosurfactants do [58]. However, there

were two limitations associated with the lipophilic linker approach: 1) “saturation” or plateau

phenomenon observed at moderate to high lipophilic linker concentrations [61]; 2) high

formulation lipophilicity where water-continuous and biocontinuous microemulsion with polar

oils became too difficult to produce [47]. To deal with these limitations, the use of hydrophilic

linker molecules to coadsorb with the surfactant at the oil/water interface was proposed by

Sabatini et al. [62].

Figure 2.6 Schematic of the lipophilic linker effect. Adapted from Graciaa et al. [56].

The concept of hydrophilic linker was introduced by Uchiyama et al. [63] to balance the effects

of the lipophilic linker. This hydrophilic linker was designed to have a strong hydrophilic head

group, with a short hydrophobic tail, to adsorb and/or segregate near the water-oil interface near

the polar head group of the surfactant. Compared to lipophilic linkers that fit in between

surfactant molecules, hydrophilic linkers open up spaces on the oil side of the interface in which

the lipophlic linker segregates.

18

When hydrophilic and lipophilic linkers are combined, a special synergism can be observed. This

combination of hydrophilic linker and lipophilic linker was found to enhance solubilization

capacity of microemulsions, as well as an increased interfacial activity of each. Furthermore, the

combined linkers also have some capacity to solubilize oil, which therefore suggests a self-

assembly in the linkers themselves to form “surfactants” at the interface [60,63].

The main motivations behind the development of linker molecules were two-fold: first, they are

less toxic and more biocompatible than medium-chain alcohols (cosurfactant) [50]; second, the

combination of hydrophilic and lipophilic linkers limits the formation of undesirable liquid

crystal phases that form at high surfactant concentrations [55].

Recently, alcohol-free, linker-based lecithin microemulsions have been formulated as potential

vehicles for transdermal delivery of the lipophilic drug lidocaine [55]. Mixtures of sodium

caprylate and caprylic acid were used as hydrophilic linkers, sorbitan monooleate was used as

the lipophilic linker, and isopropyl myristate (IPM) was the oil [55]. Compared to conventional

alcohol-based lecithin microemulsions, linker-based lecithin microemulsions offer lower

toxicity, since all ingredients involved in microemulsion preparation are food or pharmaceutical

grade. This is in comparison to using alcohols, which are cytotoxic to skin cells. Moreover,

linker microemulsions can provide twice the absorption and penetration of lidocaine through the

skin when compared to conventional emulsions [47]. Figure 2.7 presents the fluorescence

microscopy images of cross sections of porcine ear skins illustrating the penetration of the

lipophilic solute (Nile Red) formulated in lecithin-linker microemulsion and in conventional

emulsion (cream). These images support the idea that microemulsions can penetrate the porous

network of the skin carrying the lipophilic solute (e.g. Nile Red, lidocaine, etc.) with it [64]. In

addition, lecithin-linker microemulsions demonstrate sustained transdermal delivery of lidocaine

for 12h [55,64]. Furthermore, a recent trial on the delivery of Deepaline PVB (palmitoyl

hydrolyzed wheat protein), an anti-wrinkle active [65], using a formula similar to that of Yuan et

al. [55,64,66] has also produced comparable effect in wrinkle reduction (~15% reduction in

wrinkle area) to an optimized surfactant formulation containing the same level of active [65].

19

Figure 2.7 Fluorescence microscopy images of cross sections of porcine ear skin showing penetration of Nile Red formulated in lecithin-linker microemulsion and in conventional emulsion (0.5h application + 5.5h release). Images provided by Dr. Acosta.

2.4 Microemulsion-based Gels

Under certain conditions, oil-continuous, water-continuous and bicontinuous microemulsions can

be transformed into highly viscous gels with the addition of certain gelling agents [67].

Microemulsions that exhibit this viscoelastic gel characteristic are called microemulsion-based

gels (MBGs). MBGs can be considered sub families of organogels and hydrogels, where the

main components of the gels are organic-based and water-based respectively [67,68]. Compared

to conventional gel systems, MBGs incorporate microemulsion systems to provide enhanced

solubilization of both hydrophilic and lipophilic drugs. An important property of MBGs is their

sensitivity to environmental stimuli such as UV-radiation and changes in pH or temperature to

form in situ gels [69]. In pharmaceutical applications of MBGs, in situ-forming gel systems are

of particular interest for their ability to form gels within the body [69]. In drug delivery for

example, MBGs could provide better mass transfer control over liquid microemulsions due to an

increase in viscosity, resulting in increased drug residence time after administration.

Furthermore, among the different gelation stimuli, thermosensitivity is advantageous because it

does not require the use of organic solvents, co-polymerization agents, or externally applied

triggers for gelation. Several thermosensitive gelling agents have been quoted in literature,

including polysaccharides (e.g. cellulose derivatives, xyloglucan, chitosan), N-

isopropylacrylamide, gelatin, and block copolymer surfactants (poloxamers) [67,70,71]. A more

detailed review of gelatin and poloxamer 407 MBGs will be provided in Chapters 3 and 4,

respectively.

20

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52. Bagwe, R.P.; Kanicky, J.R.; Palla, B.J.; Patanjali, P.K.; Shah, D.O. Improved drug delivery using microemulsions: rationale, recent progress, and new horizons. Crit. Rev. Ther. Drug Carr. Syst. 2001, 18, 77-140.

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25

CHAPTER 3

Gelatin-Stabilized Microemulsion-Based Gels for Ophthalmic and

Transdermal Lidocaine Delivery

3.1 Abstract

Low viscosity biocompatible microemulsions containing lecithin and hydrophilic (PEG-6-

caprylic/capric glycerides, decaglycerol monocaprylate/caprate) and lipophilic (sorbitan

monooleate) surfactant additives (linkers) have been shown to quickly penetrate epithelial tissue

and slowly release from the tissue. To improve the viscosity of these formulas for transdermal

and ophthalmic applications, gelatin was added as gelling agent. Upon addition of gelatin to oil-

in-water microemulsions coexisting with excess oil phase, clear viscoelastic gels of bicontinuous

microemulsions were obtained. The addition of gelatin to microemulsion systems containing

excess water produced milky but strong gels containing emulsified microemulsions. Despite the

fact that the viscosity of the clear lecithin microemulsion-based gels (MBGs) is at least one order

of magnitude larger than the original microemulsions, it only produced a minor retardation in the

loading and release from tissues and from the gel, which is explained by transport models in gel

networks with large polymer fiber separation.

3.2 Introduction

Recently, linker-based lecithin microemulsions have been introduced by Yuan et al. [1,2] as an

alternative to alcohol-based lecithin microemulsions. These lecithin-linker microemulsions have

a characteristic low viscosity (< 40 mPa·s) while suitable for spray or roll-on methods of

applications on the skin, they spread well beyond the intended area when applied as drops [3]. In

addition, in periocular ophthalmic delivery, higher viscosities are desirable to increase the

residence time and improve the effectiveness of ophthalmic delivery formulations [4,5]. The

objective of this study is to introduce gelling agents, particularly gelatin, into lecithin-linker

microemulsions with the purpose of increasing the viscosity of the delivery system while

retaining the transport properties observed in low-viscosity microemulsions.

Under certain conditions, oil-continuous microemulsions can be transformed into highly viscous

gels with the addition of certain gelling agents [6]. Since the main component is an organic

26

solvent, these gels can be referred to as organogels [7]. A subgroup within the organogels is the

microemulsion-based gels (MBGs), introduced in the 1980s [8]. Generally, MBGs undergo

liquid to gel transition when subjected to environmental stimuli such as changes in pH,

temperature and electrolyte concentration [9]. Compared to their hydrogel counterparts, MBGs

incorporate microemulsion systems that provide a suitable environment for the solubilization of

hydrophilic, lipophilic and amphiphilic drugs. Furthermore, due to their rheological properties,

such as thermo-reversibility, MBGs have been proposed as potential vehicles for sustained drug

and vaccine delivery [10], and as enzyme entrapment media (e.g. lipase) for esterification and

catalysis [11]. The gelling agents used to formulate these MBGs include natural polymers such

as gelatin [12], κ-carrageenan [12], hydroxypropylmethyl cellulose (HPMC) [13], and block

copolymer surfactants such as poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide),

commercially known as Poloxamers or Pluronics [14]. Out of the many different MBGs

reported, the majority were prepared using gelatin with a Type II w/o micreomulsion system.

Many of the reported microemulsions are systems of sodium dioctyl sulfosuccinate

(AOT)/isooctane/water. Isopropyl myristate (IPM) have also been used as the oil phase instead

of isooctane. With the addition of 10 to 20% w/w solid gelatin, AOT microemulsions can be

transformed into high viscosity, transparent thermosensitive gels [6,8,9].

The gelation mechanism of gelatin-stabilized MBGs have been studied using various techniques,

including SAXS, SANS, 1H and 13C NMR, SEM, TEM, DSC and electrical conductometry [15].

Atkinson et al. has proposed a gelation model based on small angle neutron scattering [16]. The

proposed gelatin MBG structure consists of hydrated gelatin forming cylinders in connected

water channels, which are surrounded/stabilized by a monolayer of AOT surfactant molecules,

accompanied by “conventional” w/o microemulsion droplets (Figure 3.1) [12,17].

Figure 3.1 Proposed structure of gelatin-stabilized MBG. Adapted from Petit et al. [17].

27

A major limitation of gelatin MBGs is the toxicity associated with the anionic surfactant, AOT,

used to formulate these microemulsions [18,19]. Gelatin MBGs prepared from lecithin

microemulsions have been reported [12,13,20]. However, these previous lecithin MBGs contain

short or medium chain alcohols. While gelatin MBGs have been used as reaction media in

enzymatic reactions [20,21,22], their use as drug delivery vehicles is limited since the presence

of AOT and/or alcohols in these formulations poses cytotoxicity concerns.

In the present study, we hypothesize that gelatin-stabilized MBGs can be prepared using low

toxicity linker-based lecithin microemulsions and that the improvement in the viscoelastic

properties of the formulation will not affect significantly the ability of the microemulsion to

permeate through epithelial tissue or membranes. To produce these lecithin-linker

microemulsions, the formulation of Yuan et al. [1] was slightly modified by replacing the

hydrophilic linkers sodium octanoate and octanoic acid with a milder combination of PEG-6-

caprylic/capric glycerides and decaglycerol monocaprylate/caprate that has been confirmed to be

non-irritant to human skin, non-mutagenic and highly biocompatible [3]. To understand the

effect of formulation components (lecithin, linkers, gelatin, oil and water) on the reological

properties of the resulting MBGs, phase scans and ternary phase diagrams were obtained and

used to guide a set of rheological studies. To evaluate the effect of gelatin on drug transport

through epithelial tissue, the release profile of lidocaine topically adsorbed in the skin from

lecithin-linker gelatin MBGs, and from the corresponding microemulsion was obtained. To

evaluate the effect of gelatin on the transport properties through the gel, a membrane permeation

study was conducted to compare the fraction of lidocaine permeated when using gelatin MBGs

and microemulsions.

3.3 Materials and Methods

3.3.1 Materials

3.3.1.1 Chemicals

The following chemicals were purchased from Sigma-Aldrich (Canada) and were used as

received: gelatin (from porcine skin, type A, 300g Bloom), isopropyl myristate (IPM, 98%),

sorbitan monooleate (SMO, Span® 80, 99%), sodium chloride (99%+, Fluka brand), Dulbecco’s

phosphate buffered saline (PBS) and lidocaine powder (base form, 98%+). Decaglycerol

28

monocaprylate/caprate (Drewpol 10-1-CC) was a gift from Stepan Company and PEG-6-

caprylic/capric glycerides (Softigen 767) was donated by Sasol. Laboratory grade soybean lecithin

was purchased from Fisher Scientific (Fairlawn, NJ, USA). Soybean lecithin is a mixture of

phospholipids (mainly phosphatidylcholines) obtained by acetone purification of soybean gum

residues. The average composition of soybean lecithin has been reported elsewhere [23].

Acetonitrile (HPLC grade) was purchased from Caledon Laboratories Ltd. (Ontario, Canada),

sodium phosphate monobasic, monohydrate (ACS grade) was purchased from EMD Chemicals

Inc. (Darmstadt, Germany), and were used as received. Anhydrous ethyl alcohol was purchased

from Commercial Alcohols Inc. (Ontario, Canada). Unless otherwise stated, the composition is

expressed on a mass basis (i.e. wt%.) throughout this paper.

3.3.1.2 Skin

Ear skin from domestic pigs (approximately 6 months old) was used as a surrogate for human

epidermis [24]. Porcine ears were purchased from local market and frozen overnight. They were

partially thawed by rinsing with running water for 10 seconds at room temperature to soften the

skin for smoother cuts. The skin of the external side of the ear was dermatomed to a thickness

ranging from 700 to 900 μm [25], and was cut into circles with 11.4 mm diameter ready for use.

The skins were equilibrated to room temperature prior to use.

3.3.2 Microemulsion preparation

Microemulsions were prepared in flat-bottom tubes by mixing lecithin (surfactant), decaglycerol

monocaprylate/caprate and PEG-6-caprylic/capric glycerides (hydrophilic linkers), sorbitan

monooleate (SMO, lipophilic linker), isopropyl myristate (IPM) and 0.9% sodium chloride

solution at various compositions (Table 3.1). The concentration of lecithin, PEG-6-

caprylic/capric glycerides and decaglycerol monocaprylate/caprate were kept at 5 wt%, 10 wt%

and 4 wt%, respectively (giving a 1:2:0.8 weight ratio of lecithin : PEG-6-caprylic/capric

glycerides : decaglycerol monocaprylate /caprate). This may not be the optimal composition but

it corresponds to the maximum amount of lecithin and minimum amount of hydrophilic linkers

that can be used to produce microemulsion systems without the formation of meta-stable phases.

In addition, a minimum of 2% SMO was also required to prevent these meta-stable phases. The

phase scan of the microemulsion consisted of increasing SMO concentration from 2% to 10% (g

29

SMO/g formulation), thus increasing the hydrophobicity of the formulation. After introducing all

ingredients, the tubes were thoroughly vortexed and left to equilibrate for 2 weeks.

Table 3.1 Microemulsion formulations.

Component Composition (wt%)

0.9% sodium chloride in de-ionized water 36

PEG-6-caprylic/capric glycerides 10

Decaglycerol monocaprylate/caprate 4

Lecithin 5

Sorbitan monooleate (SMO) 2-10

Isopropyl myristate (IPM) To 100%

3.3.3 MBG preparation

MBGs were prepared by addition of gelatin powder to o/w microemulsions at various

surfactant/linker/oil/water ratios, according to the method of Kantaria et al. [9]. Briefly, the

mixture was first stirred at room temperature for 45 min to swell the solid gelatin, and then

heated to 50oC and stirred for 20 min until the gelatin was completely dissolved in the

microemulsion. Agitation was then stopped and the sample was allowed to cool in ice bath for 30

minutes. MBGs with various hardness and opacity were obtained.

3.3.4 Construction of ternary phase diagrams

Phase behaviour studies were performed by constructing ternary phase diagrams at room

temperature (25oC) and gelatin activation temperature (50oC) using the water titration method

[26,27]. The “surfactant” vertex of the ternary phase diagrams was a mixture of 1:2:0.8:1.2

weight ratio of lecithin: PEG-6-caprylic/capric glycerides : decaglycerol monocaprylate/caprate :

SMO. Mixtures of lecithin+linkers and IPM were prepared; then an aqueous solution of 0.9%

sodium chloride was added until the desired composition along the dilution line was achieved.

After each titration the flat-bottom tubes were vortexed for 3 minutes to ensure thorough mixing.

The phase behaviours of each of the tubes at room temperature were observed after 2-week

equilibration time and recorded. The ternary phase diagram at 50oC was generated by heating the

tubes in constant-temperature water bath at the desired temperature for 2 weeks allowing the

systems to reach the new equilibrium. The changes in phase behaviour were then recorded.

30

3.3.5 Physiochemical characterization

The conductivity of the microemulsions was measured at room temperature using a VWR

bench/portable conductivity meter equipped with a custom-build OEM conductivity

microelectrode (Microelectrodes Inc., Bedford, NH, USA). Viscosity measurements of the

microemulsion samples were obtained (in triplicate) using a CV-2200 falling ball viscometer

(Gilmont Instruments, Barrington, IL, USA) at room temperature. The hydrodynamic diameter

measurements of the microemulsion aggregates was determined via dynamic light scattering at a

90° angle, using a Brookhaven 90Plus particle size analyzer equipped with a 35mW diode laser

(wavelength ~674 nm).

3.3.6 Rheological measurements

Rheological measurements of the MBGs were obtained using a stress and strain controlled CSL2

500 rheometer (TA Instruments Ltd.). The measuring system used was the 4 mm diameter

stainless steel cone and plate geometry (cone angle 2o). The sample volume was approximately 1

ml. Oscillation experiments were performed at 1.00 Hz frequency and 0.177 Pa oscillation stress

over the temperature range of 20 to 50oC. At the maximum temperature the formulations were

liquid and gelled upon cooling. The elastic moduli (G’) of the samples were recorded.

The G’ values reflect the elastic behaviour of viscoelatic systems, and are particularly useful

when characterizing gels (e.g. identifying sol-gel transition temperatures) [28]. Many researchers

have also used G’ to characterize the rheological behavior of topical formulations [9,14,29,30].

Hence, the elastic moduli, instead of viscosity (which reflect the viscous behaviour of the

system), were measured and compared in this study for the rheological characterization of

lecithin-linker MBGs.

3.3.7 In vitro transport studies

To evaluate the performance of gelatin-stabilized MBGs as delivery vehicles, a lipophilic drug,

lidocaine base, was chosen as the model drug in this work. Lidocaine is an anesthetic that has

been used in topical formulations as a pain reliever in the treatment of minor burns, sunburn,

insect bites, after various laser skin surgeries and during cataract surgery [31,32,33]. It was

incorporated in the lecithin-linker microemulsions and gelatin MBGs by pre-dissolving 10% w/w

lidocaine in IPM.

31

3.3.7.1 Transport in the skin

The transdermal in vitro extended release experiments for a selected microemulsion and its

gelatin MBG were conducted using MatTek Permeation Device (MPD) supplied by MatTek

Corporation (Ashland, MA, USA) fitted with porcine ear skin as membrane. The experiment was

carried in laboratory conditions that simulated the topical application of these formulations.

Briefly, porcine ear skin tissues were placed into the MPD with the epidermis layer facing up.

The exposed tissue area in the MPD is 0.256 cm2. After assembling the device, 400 μl of test

microemulsion (25oC) or liquefied gelatin MBG (initially at 50oC) were applied in the donor

compartment. During the transfer process, the liquefied gelatin MBG cooled down to 40oC or

less, making it unlikely for the MBG to affect the structure and permeability of the stratum

corneum. The receptor compartment was filled with 5 ml of PBS. The donor microemulsion or

the gelatin MBG was withdrawn 30 minutes after application [1,2]. The skin surface in the MPD

was blotted dry with Kimwipes and was then used for extended release. At predetermined times

(1, 3, 6, 12, 24 and 48h), the receiver solution was withdrawn completely from the receptor

compartment and was immediately replaced with fresh PBS solution to maintain sink condition.

The experiment was terminated at 48h. At the end of the experiment, the porcine ear skin from

each of the MPD were collected and used to determine the final concentration of lidocaine

absorbed in the skin (and hence the total amount of lidocained absorbed in the skin). Prior to

analysis, the skin pieces were rinsed with a few droplets of PBS solution and placed into 2 ml of

methanol for 48h for the extraction of lidocaine [1,2]. All experiments were conducted in

quadruplets at room temperature.

3.3.7.2 Transport in the gel

In vitro permeation experiments were conducted as described in Section 3.3.7.1 with some

modifications that would simulate the release from the MBG and its corresponding

microemulsion onto a solution that simulate tear fluid as a simple model for periocular delivery

from MBGs. Briefly, cellulose acetate membranes (Harvard Apparatus, MWCO 100 kDa, ~10

nm pore size) were placed into the MatTek Permeation Device with exposed area of 0.256 cm2.

50 µl of test microemulsion (25oC) or liquefied MBG (50oC) were applied to the donor

compartment after assembling the device and was kept in the donor compartment throughout the

48h permeation study. The receptor compartment was filled with 5 ml of simulated tear fluid

(Table 3.2). At predetermined times (1, 2, 3, 6, 12, 24 and 48h), the receiver solution was

32

withdrawn completely and replaced with fresh medium immediately. All permeation

experiments were conducted in quadruplets at 34oC (ocular surface temperature).

Table 3.2 Simulated tear fluid composition [34].

Components Weight

CaCl2 · 2H2O 0.008 g

NaCl 0.658 g

H2O (pH 7.2) To 100 g

3.3.8 Lidocaine quantification

The concentration of lidocaine in the skin (determined after extraction with methanol) and in

receiver solutions was quantified using a Dionex ICS-3000 liquid chromatography system

equipped with an AS40 automated sampler, AD25 absorbance detector and a reverse phase

column (Genesis, C18, 4μm, 150 mm × 4.6 mm). The UV detector was set to 230 nm. A mixture

of acetonitrile and 0.05M NaH2PO4·H2O (pH 2.0) (30:70, v/v) was used as the mobile phase at

flow rate of 1.0 ml/min. The retention time of lidocaine under the described conditions was

approximately 2.7 min, and the calibration curve for area under the peak vs. concentration was

linear (R2 = 0.9997).

3.4 Results and Discussion

3.4.1 Phase behaviour of lecithin-linker microemulsions

3.4.1.1 Sorbitan monooleate (SMO) phase scan

The compositions of linker-based lecithin microemulsions considered in the SMO scan are

shown in Table 3.1. A picture of the vials employed in this phase scan is presented in Figure 3.2,

showing the transition from Type I o/w microemulsion (bottom phase) with excess oil phase (top

phase) obtained at 2% and 4% SMO to single phase Type IV microemulsion at 6% SMO, and

finally to coexisting w/o + o/w microemulsions at 8% and 10% SMO. This split microemulsion

phases obtained at 8 and 10% SMO are not common, but can occur when hydrophilic and

lipophilic linkers partition in different phases. The types of microemulsions produced were

confirmed using conductivity measurements (Table 3.3).

33

Yuan et al. evaluated Type I, IV and II lecithin linker microemulsions as lidocaine delivery

vehicles and found that while all these types improved lidocaine loading in the skin and its

transdermal flux, Type IV systems produced the largest lidocaine loading in the skin [1].

Furthermore, bicontinuous (Type IV) microemulsions produce the largest co-solubilization of oil

and water. Considering the advantages of Type IV formulations, the 6% SMO system was

selected as the base composition to construct ternary phase diagrams and to evaluate in vitro

transport of lidocaine in skin and in the gel.

Figure 3.2 Phase scan at room temperature. All formulations are consisted of 5% lecithin, 10% PEG-6 caprylic/capric glycerides and 4% decaglycerol monocaprylate/caprate.

Table 3.3 Conductivity measurements (μS/cm) of microemulsions with 2-10% SMO.

% SMO 2% 4% 6% 8% 10%

Top 0 0 1500

40 30

Bottom 1660 1600 1320 1220

The viscosity of microemulsion formulations containing 2 to 6% SMO are presented in Figure

3.3. The increase in viscosity from 20 to 120 mPa·s with increasing SMO concentration can be

explained on the basis that when approaching the Type I-IV transition (increasing SMO), oil-

swollen micelles grow larger and turn cylindrical, which increases the viscosity of the

formulation [1,35]. The data in Figure 3.3 confirms the relatively low viscosity of lecithin-linker

microemulsions when compared to medium-chain alcohol-based lecithin microemulsions that

reach viscosities as high as 1000 mPa·s [7], and to commercial topical creams such as Lanocort

34

10 that have viscosities of approximately 1300 mPa·s [36]. The low viscosity of lecithin-linker

micormulsions makes them suitable for spray and roll-on applications, but it is a disadvantage

for gel-type topical and ophthalmic applications.

The mean hydrodynamic diameters for micellar systems containing 2% and 4% of SMO ranged

from 1 to 3 nm, comparable to values obtained in previous linker-lecithin systems [1]. The small

droplet size of lecithin-linker microemulsions has been associated with the use of hydrophilic

linkers that increase the interfacial area and reduces the size of oil-swollen micelles and water-

swollen reverse micelles [37,38].

Figure 3.3 Viscosities of the microemulsions of Figure 3.2 containing 2-6% SMO at 25°C.

3.4.1.2 Ternary phase diagrams

Ternary phase diagrams at 25°C and 50°C (gel activation temperature) were constructed using

the surfactant composition corresponding to the 6% SMO formulation of Figure 3.2. These phase

diagrams are presented in Figure 3.4. At both temperatures, the surfactant mixture is not

completely soluble in either the aqueous (0.9% NaCl) solution or in isopropyl myristate (IPM).

Liquid crystalline (LC) or surfactant + oil + precipitate phases (S+O+P) were found in systems

containing less than 10% IPM or 5% water. Figure 3.5 presents a micrograph of a surfactant-

water system under cross-polarizer, where the white or light gray parts correspond to anisotropic

liquid crystal phases. Other phases observed in the ternary phase diagrams include isotropic,

single phase microemulsions (µE), microemulsion in equilibrium with dispersed liquid

crystalline phases (µE+LC), microemulsions in equilibrium with excess oil (µE+oil),

microemulsion in equilibrium with excess water (µE+water), and microemulsion coexisting with

35

excess oil and excess water phases (µE+oil+water). Please note that the excess oil or aqueous

phases are not necessarily pure oil or water; instead, they are most likely mixtures with linkers

that partition into these phases. The main difference in the ternary phase diagram at 50oC

compared to that at 25oC is the larger µE+water region at 50oC. This can be explained on the

basis that the hydrophilic linker PEG-6-caprylic/capric glycerides is temperature sensitive (due

to the presence of ethylene oxide groups) [39]. As the temperature increases from 25oC to 50oC,

hydrogen bonding between the ethylene oxide groups of the hydrophilic linkers and the water

molecules weaken (dehydration) and the formulations become more hydrophobic, hence the

ability for the microemulsion to solubilize water decreases, resulting in the formation of larger

µE+water region [39,40,41].

Figure 3.4 Ternary phase diagrams of 6% SMO formulation at 25°C and 50°C. The surfactant vertex is a mixture of 1:2:0.8:1.2 weight ratio of lecithin: PEG-6-caprylic/capric glycerides : decaglycerol monocaprylate/caprate : SMO. The point “∆” indicates the composition of the base formulation containing 6% SMO.

Figure 3.5 Microscopic image through cross polarizing lenses showing the presence of liquid crystals (bright spots) in concentrated aqueous surfactant solutions (20% water–80% surfactant).

36

3.4.2 Phase behaviour of lecithin-linker MBGs

3.4.2.1 Sorbitan monooleate (SMO) phase scan

The systems of Figure 3.2 were used to produce MBGs formulated with 10% and 20% gelatin.

Figure 3.6a presents pictures of these MBGs. MBGs containing less SMO and 20% gelatin were

translucent. The elastic moduli (G’) of MBGs produced with 2% and 10% SMO are presented in

Figure 3.6b and 3.6c as a function of temperature for formulations containing 10% and 20%

gelatin, respectively. Higher G’ values are obtained in systems formulated with 20% gelatin,

indicating the formation of stronger gels, likely due to a denser gel network structure.

Furthermore, gels with lower SMO produced weaker, albeit clear, gels.

The electrical conductivity of the clear gels produced with 2-6% SMO ranged from 50 to 200

μS/cm, suggesting that although the original microemulsions were Type I systems, the structure

of the final gel was closer to a bicontinuous microemulsion rich in oil. This observation is

consistent with the work of Atkinson et al. and Petit et al. on AOT microemulsions and the

structure proposed in Figure 3.1 [16,17]. One explanation for the fact that the presence of gelatin

induced the transition from Type I microemulsions to bicontinuous oil-rich systems, is that

gelatin “dehydrated” the microemulsion and used that water to form the gel network. According

to the ternary phase diagrams of Figure 3.4, removing water from the microemulsion shifts the

formulation into a region of single phase microemulsions rich in oil. For the MBGs containing 8

and 10% SMO, their marginally translucent appearance might be a sign that the final

microemulsion phase in the MBG is closer to a phase transition.

Fluorescent dyes, sodium fluorescein (hydrophilic) and nile red (hydrophobic), were separately

added to the 2% and 10% SMO MBGs produced with 20% gelatin to assess the phase continuity

and structure of the gels using florescence and polarized light microscopy. According to Figure

3.7, the MBG formulated with 2% SMO show apparent continuity in both aqueous (3.7a) and oil

(3.7b) phases as shown by the continuous green color from fluorescein (water soluble) and nile

red (oil soluble), respectively. Areas showing different colour intensity are typically indicative of

bubbles or gelatin fibers in the sample. The structure of these fibers is evidenced by the bright

strands observed under cross polarizers in Figure 3.7c. Similarly, the MBG containing 10% SMO

also seem to be continuous in both phases (Figure 3.8), and also has a three-dimensional network

of gelatin fibers. According to Figure 3.8c and 3.8d, the fibers seem to be thicker than those

37

shown in Figure 3.7c for 2% SMO MBG. This difference in fiber thickness would explain the

higher gel strength (higher G’) and turbidity of MBGs prepared with 10% SMO.

Figure 3.6 SMO scan of lecithin-linker MBGs. All formulations contain 10% PEG-6-caprylic/capric glycerides, 4% decaglycerol monocaprylate/caprate, 5% lecithin, 36% of 0.9% NaCl solution and the balance of IPM. (a) Gel appearance. Elastic modulus (G’) vs. temperature profile of (b) 20% gelatin MBGs and (c) 10% gelatin MBGs containing (x) 2% SMO; (▲) 10% SMO; Control (10% or 20% gelatin in 0.9% NaCl solution) (●).

38

Figure 3.7 Fluorescence image of lecithin-linker MBG (20% gelatin) with 2% SMO labelled with (a) sodium fluorescein and (b) nile red. (c) Cross-polarized images of lecithin-linker MBG.

Figure 3.8 Fluorescence image of lecithin-linker MBG (20% gelatin) with 10% SMO labelled with (a) sodium fluorescein and (b) nile red. Cross-polarized images of lecithin-linker MBG dyed with (c) sodium fluorescein and (d) nile red.

39

3.4.2.2 Ternary phase diagrams

Gel formation in gelatin MBGs occurs between room temperature (25oC) and 50oC. At 50oC, the

gelatin activation temperature, the native helical structure of gelatin is denatured, and gelatin

exists as flexible random coils in solution [42]. Upon cooling, gelatin recovers its native helical

structure, producing a gelatin hydrogel network [42]. Using the ternary phase diagrams of Figure

3.4, the appearance and rheological properties of MBGs along oil and water dilution lines that

passed through the optimal 6% SMO formulation (containing 36% water and 39% oil) were

evaluated.

Figure 3.9a shows the oil dilution path at 25oC and 50oC. Figure 3.9b presents a picture of the

gels prepared along the oil dilution line. For systems containing between 10% and 30% IPM, the

gels have a milky appearance, suggesting the presence of multiple phases in the gel. Considering

the ternary phase diagrams of Figure 3.9a and the earlier discussion that gelatin “dehydrates” the

original microemulsion, it is likely that the gelation process for these 10-30% IPM systems led to

the formation of a µE+LC system embedded in the gelatin hydrogel. Although the system

containing 50% IPM looks clear, it is really a 2-phase system of microemulsion and a MBG. The

system of 40% IPM is nominally the same as the 6% SMO gel of Figure 3.6a. Figure 3.10

presents the temperature dependence of the elastic modulus (G’) of the gels of Figure 3.9b.

According to Figure 3.10, increasing the oil (IPM) content in the MBG from 10 to 40% reduces

the strength of the gel, albeit increasing its clarity. To obtain a measurement for the 50% IPM

MBG, only the gel portion of the 2-phase system was evaluated, which explains the high strength

of this gel.

For the 40% IPM lecithin-linker MBG, the zero shear viscosity was close to 3 Pa·s at 25°C and 1

Pa·s at 37°C, which represents one order of magnitude increase in viscosity with respect to the

original microemulsion. Viscosities in the 1-10 Pa·s range are comparable to some topical

creams [36]. The gel strength of the system with 40% IPM is lower than other commercial

lidocaine creams and gels (measured G’ at 25°C for EMLA® cream and Topicaine gel were

approximately 500 Pa and 200 Pa, respectively). As shown in Figure 3.6, MBGs with high G’

values can be obtained with higher gelatin and SMO content.

40

Figure 3.9 (a) Oil dilution path used to produce MBGs with 10% to 50% IPM. (b) Appearance of MBGs at room temperature.

Figure 3.10 Elastic modulus vs. temperature profile for the systems of Figure 3.9b: (●) 10% IPM; (□) 20% IPM; (▲) 30% IPM; (x) 40% IPM; (*) 50% IPM. Control (20% gelatin in 0.9% NaCl solution) is shown as (○).

41

Figure 3.11a presents the water dilution path used to produce MBGs containing 30 to 90% water.

Figure 3.11b shows that with increasing water content the MBG becomes more turbid. These

milky gels reflect the presence of an emulsified phase within the gel. These emulsion gels are

likely produced when, at 50°C, a microemulsion phase coexist with an aqueous solution, as

shown in Figure 3.11a. The gelatin gel is likely formed within the excess aqueous phase and,

upon cooling, the microemulsion is emulsified within the gel. Figure 3.12 show cross-polarizer

micrographs of the MBG prepared with 80% water, showing the gelatin network (Figure 3.12a)

and the drops of the emulsion entrapped in the gel (Figure 3.12b).

Figure 3.13 presents the elastic modulus (G’) for the MBGs of Figure 3.11b, as a function of

temperature. In general, increasing the water content in the MBGs increases the strength of the

gels. However, the system containing 30% water, similar to the 50% oil MBG, consists of a

liquid oil-rich microemulsion that coexists with a strong gel.

Figure 3.11 (a) Water dilution path of MBGs containing 30% to 90% water. (b) MBG appearance at room temperature.

42

Figure 3.12 Cross-polarizer micrographs of MBG containing 80% water (system of Figure 3.11b).

Figure 3.13 Elastic modulus vs. temperature profile for the systems of Figure 3.11b: (▲) 30% aqueous; (●) 50% aqueous; (□) 80% aqueous; (x) 90% aqueous. Control (20% gelatin in 0.9% NaCl solution) is shown as (○).

3.4.2.3 Lecithin-linker gels

Lecithin-linker gels (not MBGs) were prepared using selected mixtures of lecithin and linkers

using the ratios corresponding to the 6% SMO formulation of Figure 3.2. Figure 3.14 presents

the elastic modulus of these formulations as a function of temperature. Mixtures of liphophilic

amphiphiles (lecitihin, SMO) and gelatin tend to increase the strength of the gels, even at 50°C.

This observation is consistent with the fact that lecithin and SMO can also produce organogels

on their own [10]. On the other hand, introducing the hydrophilic amphiphiles (linkers) PEG-6-

caprylic/capric glycerides and decaglycerol monocaprylate/caprate tend to decrease the gel

strength and its transition temperature from 50°C to values close to 45°C. This observation

suggests that the ethylene glycol and glycerol groups of the hydrophilic linkers interfere with the

43

self-assembly of the collagen strands during gelation, thus reducing the strength of the resulting

gel [42].

Figure 3.14 Effects of formulation components on elastic modulus of 20% gelatin gels: (●) control (20% gelatin in 0.9% NaCl solution); (x) lecithin + control; (▲) lecithin + SMO + control; (□) PEG-6-caprylic/capric glycerides + decaglycerol monocaprylate/caprate + control; (*) lecithin + SMO + PEG-6-caprylic/capric glycerides + decaglycerol monocaprylate/caprate + control. The surfactant/linker ratios were kept constant at 1:2:0.8:1.2 of lecithin:PEG-6-caprylic/capric glycerides:decaglycerol monocaprylate/caprate:SMO.


The 20% gelatin MBG prepared with 6% SMO (Figure 3.6a) was evaluated as a delivery vehicle

to load lidocaine in the skin, and use the skin as a lidocaine reservoir (in situ patch) for extended

release. Using a permeable synthetic membrane, instead of tissue, the gel itself was evaluated as

a reservoir for the extended release of lidocaine. The corresponding 6% SMO microemulsion

was used as a control to compare the transport of lidocaine in both scenarios.


The total mass of lidocaine loaded in the skin from the microemulsion (control) and the MBG

were determined by adding the mass of lidocaine recovered from methanol extraction of the

porcine ear skin at the end of the experiment and the mass of lidocaine recovered from the

receiver solutions at different times. Slightly less lidocaine (0.8±0.3 mg/cm2) was loaded in the

skin from the MBG than from the microemulsion (1.3±0.5 mg/cm2), suggesting that lidocaine

permeation into skin is slower in the MBG. Figure 3.15 shows the lidocaine release profile for

44

the MBG- and the microemulsion-loaded skins. Both release profiles are similar, both akin to a

first order release, although there was little release after 24 hours. Half of the lidocaine loaded

was released between 5 and 7 hours in these formulations. These observations are consistent with

those of Yuan and Acosta, only that the time to release 50% of lidocaine in their work was close

to 7.5 hours for Type II microemulsions and 11 hours for Type I formulations [2]. These

formulations offer the potential for longer lasting pain relieving effect when compared to

commercial lidocaine creams such as EMLA (emulsion containing 2.5 wt% lidocaine and 2.5

wt% prilocaine), whose action only lasts between 2 and 4 hours [43,44].

The release profiles of Figure 3.15 suggest that other than a minor reduction in lidocaine loading,

the addition of gelatin significantly improved the viscoelastic behaviour of lecithin-linker

microemulsions with miminal impact on the release of lidocaine from the skin. These results

support the initial hypothesis that the gelatin gel network should not affect the loading or release

of drugs on the skin, but produce MBG systems with viscoelastic properties suitable for

topical/transdermal applications.

Figure 3.15 Release profile of lidocaine from the skin after loading with: 20% gelatin MBG (◆), and lidocaine-loaded gelatin-free microemulsion (∆).


Although the evaluation of transdermal loading and release is relevant to the method of

topical/transdermal delivery, it does not provide clear information about the transport of

lidocaine in the gel. Gel permeation is important if the gel itself, and not the epithelial tissue is

45

used as drug depot, particularly in periocular ophthalmic delivery. As a simplified model for this

gel transport, the skin was replaced by a cellulose acetate membrane with a molecular weight cut

off (MWCO) of 100 kDa, which corresponds to an approximate pore size of 10 nm. The use of

synthetic ultrafiltration membranes, particularly for scleral transport has been reported in the

literature [45,46,47]. Various ranges of equivalent pore size of the sclera have been reported in

the literature, but for paracellular transport the pore size has been estimated to be in the order of

2-4 nm, although pore sizes approaching 10 nm have been proposed as well [48,49].

According to the gel permeation profiles presented in Figure 3.16, the microemulsion system

showed slightly higher, although not significantly different, lidocaine flux through the membrane

than the MBG. This is consistent with the earlier observation that the loading of lidocaine in the

skin from the MBG and the microemulsion were similar. The permeation of lidocaine through

the gel in both processes can be interpreted using obstruction-diffusion models applied to

heterogeneous hydrogels [50]. According to a simplified form of the equation of Amsden, the

diffusivity in the gel ≈ diffusivity in the solvent*exp(-π/4*(thickness of the fiber/distance

between fibers)2) [50]. Considering the distance between the fibers, and the thickness of the

fibers in Figures 3.7 and 3.8, and the expression of Amsden, one simply concludes that the fibers

are too loosely packed to interfere with the diffusion of drug molecules through the gel.

However, the packing is strong enough to increase the macroviscosity of the microemulsion.

The macroviscosity reflects the motion (flow) of a system at the macroscopic scale and is usually

evaluated via flow (e.g. viscosity) or rheological (e.g. elastic modulus G’, loss modulus G’’ and

complex viscosity η*) measurements [51,52,53,54]. Drug diffusion rates in aqueous dispersions

of polymers are typically governed by the restricitive effects of the polymer on drug mobility,

whether due to a reduction in free volume or an increase in the macroviscosity of the medium

[55,56,57]. However, for some systems, the effects of the polymer molecules on the macroscopic

flow properties of the systems do not necessarily correlate with effects on diffusion (i.e.

movement at the microscopic scale). For such systems, the microviscosity (i.e. a measure of

viscosity at the microscopic scale), in addition to the macroviscosity, should also be considered

when predicting drug diffusion rates [54,58].

In this study, the addition of gelatin to lecithin-linker microemulsions produced porous gelatin

networks which only modified the macroviscosities of the microemulsions (i.e. significant

46

improvement in viscoelastic properties) while the microviscosities remained mostly unaffected

(i.e. miminal impact on transport properties).

Figure 3.16 Permeation profile of lidocaine from in situ gelatin MBG (◆) and from lidocaine-loaded gelatin-free microemulsion (∆).

3.5 Conclusions

Gelatin microemulsion-based gels (MBGs) have been produced using a biocompatible

formulation containing lecithin, sorbitan monooleate, PEG-6-caprylic/capric glycerides and

decaglycerol monocaprylate/caprate. When gelatin was added to oil-in-water and bicontinuous

microemulsions, it used some of the water in the microemulsion to produce a dispersed network

of fibers embedded in an oil-rich bicontinuous microemulsion. Higher gelatin content was

desirable in order to obtain stronger and clear gels. When gelatin was added to oil-rich

microemulsions, the gel network was not able to absorb all the microemulsion in the gel network

and an excess liquid microemulsion fraction was obtained. When gelatin was added to

microemulsions containing an excess aqueous phase, milky emulsion gels were obtained.

When a bicontinuous, Type IV, system was used as the “parent” microemulsion for the MBG, a

clear gel with viscoelastic properties suitable for topical applications was obtained. This MBG

has comparable drug transport properties to the original microemulsion as the separation between

the gelatin fibers is too large to interfere with the free diffusion of the drug (lidocaine).

47

3.6 References







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10. Murdan, S. Organogels in drug delivery. Expert Opin. Drug Del. 2005, 2, 489-505.

11. Rees, G.D.; Jenta, T.R.; Nascimento, M.G. Use of water-in-oil microemulsions and gelatin-containing microemulsion-based gels for lipase-catalysed ester synthesis in organic solvents. Indian J. Chem. 1993, 32B, 30-34.

12. Stamatis, H.; Xenakis, A. Biocatalysis using microemulsion-based polymer gels containing lipase. J. Mol. Catal. B: Enzym. 1999, 6, 399-406.

13. Blattner, C.; Zoumpanioti, M.; Kroner, J.; Schmeer, G.; Xenakis, A.; Kunz, W. Biocatalysis using lipase encapsulated in microemulsion-based organogels in supercritical carbon dioxide. J. Supercrit. Fluids 2006, 36, 182-193.


15. Caldararu, H.; Timmins, G.S.; Gilbert, B.C. The structure of gelatin-water/oil microemulsion sols and gels. An EPR spin-probe and spin-labelling study. Phys. Chem. Chem. Phys. 1999, 1, 5686-5695.

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16. Atkinson, P.J.; Robinson, B.H.; Howe, A.M.; Heenan, R.K. Structure and stability of microemulsion-based organogels. J. Chem. Soc. Faraday T. 1991, 87, 3389-3397.

17. Petit, C.; Zemb, Th.; Pileni, M.P. Structural study of microemulsion-based gels at the saturation point. Langmuir 1991, 7, 223-231.

18. Patravale, V.B.; Date, A.A. Microemulsions: applications in transdermal and dermal delivery. Crit. Rev. Ther. Drug Carrier Syst. 2007, 24, 547-596.

19. Zoumpanioti, M.; Stamatis, H.; Xenakis, A. Microemulsion-based organogels as matrices for lipase immobilization. Biotechnol. Adv. 2010, 28, 395-406.

20. Nagayama, K.; Yamasaki, N.; Imai, M. Fatty acid esterification catalyzed by Candida rugosa lipase in lecithin microemulsion-based organogels. Biochem. Eng. J. 2002, 12, 231-236.

21. Jenta, T.R.-J.; Batts, G.; Rees, G.D.; Robinson, B.H. Biocatalysis using gelatin microemulsion-based organogels containing immobilized Chromobacterium viscosum lipase. Biotechnol. Bioeng. 1997, 53, 121-131.

22. Uemasu, I.; Hinze, W.L. Enantioselective esterification of 2-methylbutyric acid catalyzed via lipase immobilized in microemulsion-based organogels. Chirality 1994, 6, 649-653.

23. Scholfield, C.R. Composition of soybean lecithin. J. Am. Oil Chem. Soc. 1981, 58, 889-892.

24. Jacobi, U.; Kaiser, M.; Toll, R.; Mangelsdorf, S.; Audring, H.; Otberg, N.; Sterry, W.; Lademann, J. Porcine ear skin: An in vitro model for human skin. Skin Res. Technol. 2007, 13, 19-24.

25. Bronaugh, R.L.; Maibach, H.I. In Vitro Percutaneous Absorption: Principles, Fundamentals and Applications; CRC Press: Boca Raton, FL, USA, 1991; pp 139.

26. Aboofazeli, R.; Lawrence, M.J. Investigations into the formation and characterization of phospholipid microemulsions. I. Pseudo-ternary phase diagrams of systems containing water-lecithin-alcohol-isopropyl myristate. Int. J. Pharm. 1993, 93, 161-175.

27. Moreno, M.A.; Ballesteros, M.P.; Frutos, P. Lecithin-based oil-in-water microemulsions for parenteral use: Pseudoternary phase diagrams, characterization and toxicity studies. J. Pharm. Sci. 2003, 92, 1428-1437.

28. Mezger, T.G. The Rheology Handbook: For users of rotational and oscillatory rheometers, second ed.; Vincentz Network: Hannover, Germany, 2006; pp 118.


30. Edsman, K.; Carlfors, J.; Peterson, R. Rheological evaluation of poloxamer as a in situ gel for ophthalmic use. Eur. J. Pharm. Sci. 1998, 6, 105-112.

31. Tetzlaff, J.E. Clinical Pharmacology of Local Anesthetics; Butterworth-Heinemann: Boston, MA, USA, 2000; pp 85.

49

32. Behndig, A; Linden C. Aqueous humor lidocaine concentrations in topical and intracameral anesthesia. J. Cataract. Refr. Surg. 1998, 24, 1598-1601.

33. Assia, E.I.; Pras, E.; Yehezkel, M.; Rotenstreich, Y.; Jager-Roshu, S. Topical anesthesia using lidocaine gel for cataract surgery. J. Cataract. Refr. Surg. 1999, 25, 635-639.

34. van Haeringen, N.J. Clinical biochemistry of tears. Surv. Ophthalmol. 1981, 26, 84-95.

35. Kiran, S.K.; Acosta, E.J. Predicting the morphology and viscosity of microemulsions using the HLD-NAC model. Ind. Eng. Chem. Res. 2010, 49, 3424-3432.

36. Adeyeye, M.C.; Jain, A.C.; Ghorab, M.K.; Reilly, W.J., Jr. Viscoelastic evaluation of topical creams containing microcrystalline cellulose/sodium carboxymethyl cellulose as stabilizer. AAPS Pharm SciTech 2002, 3, E8.


38. Acosta, E.; Uchiyama, H.; Sabatini, D.A.; Harwell, J.H. The role of hydrophilic linkers. J. Surfactants Deterg. 2002, 5, 151-157.

39. Tyrode, E.; Johnson, C.M.; Rutland, M.W.; Claesson, P.M. Structure and hydration of poly(ethylene oxide) surfactants at the air/liquid interface. A vibrational sum frequency spectroscopy study. J. Phys. Chem. C 2007, 111, 11642-11652.

40. Tyrode, E.; Johnson, C.M.; Kumpulainen, A.; Rutland, M.W.; Claesson, P.M. Hydration state of nonionic surfactant monolayers at the liquid/vapor interface: Structure determination by vibrational sum frequency spectroscopy. J. Am. Chem. Soc. 2005, 127, 16848-16859.

41. Acosta, E.J.; Bhakta, A.S. The HLD-NAC model for mixtures of ionic and nonionic surfactants. J. Surfactants Deterg. 2009, 12, 7-19.

42. Atkinson, P.J.; Robinson, B.H.; Howe, A.M.; Pitt, A.R. Characterisation of water-in-oil microemulsions and organo-gels based on sulphonate surfactants. Colloids Surf., A 1995, 94, 231-242.

43. Pasero, C.L.; Davies, P.; Faucett, J.; Wong, D.L. Pain control. Am. J. Nurs. 1995, 95, 22-24.

44. Skaaret, I. Extended Release of Lidocaine Hydrochloride from a New Formulation. Department of Pharmacy, Uppsala University, 2006, <http://www.farmfak.uu.se/farm/galfarm-web/DiplomaWork/Ingrid Skaaret.pdf>.

45. Qian, Y.; Wang, F.; Li, R.; Zhang, Q.; Xu, Q. Preparation and evaluation of in situ gelling ophthalmic drug delivery system for methazolamide. Drug Dev. Ind. Pharm. 2010, 36, 1340-1347.

46. Chan, J.; Maghraby, G.M.M.E.; Craig, J.P.; Alany, R.G. Phase transition water-in-oil microemulsions as ocular drug delivery systems: In vitro and in vivo evaluation. Int. J. Pharm. 2007, 328, 65-71.

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47. Bonferoni, M.C.; Chetoni, P.; Giunchedi, P.; Rossi, S.; Ferrari, F.; Burgalassi, S.; Caramella, C. Carrageenan-gelatin mucoadhesive systems for ion-exchange based ophthalmic delivery: In vitro and preliminary in vivo studies. Eur. J. Pharm. Biopharm. 2004, 57, 465-472.

48. Hämäläinen, K.M.; Kananen, K.; Auriola, S.; Kontturi, K.; Urtti, A. Characterization of paracellular and aqueous penetration routes in cornea, conjunctiva, and sclera . Invest. Ophthalmol. Vis. Sci. 1997, 38, 627-634.

49. Trier, K. The Sclera. Advances in Organ Biology Volume 10: The Biology of the Eye; Elsevier B.V.: Amsterdam, The Netherlands, 2006; pp 353-373.

50. Amsden, B. Solute diffusion within hydrogels. Mechanisms and models. Macromolecules 1998, 31, 8382-8395.

51. Ullrich, S.; Metz, H.; Mäder, K. Sucrose ester nanodispersions: Microviscosity and viscoelastic properties. Eur. J. Pharm. Biopharm. 2008, 70, 550-555.

52. Barreiro-Iglesias, R.; Alvarez-Lorenzo, C.; Concheiro, A. Incorporation of small quantities of surfactants as a way to improve the rheological and diffusional behavior of carbopol gels. J. Control. Release 2001, 77, 59-75.

53. Alvarez-Lorenzo, C.; Gómez-Amoza, J.L.; Martínez-Pacheco, R.; Souto, C.; Concheiro, A. Microviscosity of hydroxypropylcellulose gels as a basis for prediction of drug diffusion rates. Int. J. Pharm. Sci. 1999, 180, 91-103.

54. De Smidt, J.H.; Crommelin, D.J.A. Viscosity measurement in aqueous polymer solutions by dynamic light scattering. Int. J. Pharm. Sci. 1991, 77, 261-264.

55. Sarisuta, N.; Parrott, E.L. Diffusivity and dissolution rate of polymeric solutions. Drug Dev. Ind. Pharm. 1983, 9, 861-875.

56. Kumar, S.; Himmelstein, K.J.; Modification of in situ gelling behavior of carbopol solutions by hydroxypropyl methylcellulose. J. Pharm. Sci. 1995, 84, 344-348.

57. Suh, H.; Jun, H.W. Physicochemical and release studies of naproxen in poloxamer gels. Int. J. Pharm. Sci. 1996, 129, 13-20.

58. Al-Khamis, K.I.; Davis, S.S.; Hadgraft, J. Microviscosity and drug release from topical gel formulations. Pharm. Res. 1986, 3, 214-217.

51

CHAPTER 4

Poloxamer 407-Stabilized Lecithin Microemulsion-Based Gels for

Controlled Ophthalmic and Transdermal Drug Delivery

4.1 Abstract

In this work, a temperature sensitive block copolymer surfactant, poloxamer 407, was introduced

in biocompatible lecithin-linker microemulsions with the objective of producing thermosensitive

lecithin-linker MBGs for ophthalmic and transdermal delivery applications. After exploring

various regions of the formulation phase map, it was determined that only selected formulations

were capable of producing the desired thermosensitive MBGs. It was also determined that

lipophilic components (lecithin, sorbitan monooleate, isopropyl myristate) facilitate the

formation of worm-like poloxamer micelles, decreasing the gel transition temperature and

inducing the formation of temperature insensitive gels. Furthermore, the presence of hydrophilic

linker tends to increase the gelation temperature. In the in vitro transport studies, thermosensitive

poloxamer 407 MBGs provided sustained and near zero-order release of dexamethasone for 48h,

and first-order release of lidocaine for more than 24h. These low toxicity poloxamer 407-

stabilized lecithin-linker MBGs demonstrate the potential for controlled delivery of a wide range

of hydrophilic and lipophilic actives.

4.2 Introduction

Microemulsion-based gels with thermosensitivity can be prepared using gelling agents such

gelatin, N-isopropylacrylamide, and certain polysaccharides, and are promising topical delivery

vehicles for a variety of drugs [1,2,3]. Poloxamer triblock copolymers were introduced in the late

1950s and have been utilized in many pharmaceutical applications (topical, ophthalmic, rectal,

subcutaneous, etc.) as biocompatible and thermosensitive gelling agents [4,5,6]. This group of

copolymers consist of ethylene oxide (EO) and propylene oxide (PO) blocks arranged in a

triblock structure [(EO)a-(PO)b-(EO)a], where a and b vary depending on the type of triblock

copolymer (a ~ 106, b ~ 70 for poloxamer 407). Of special interest is poloxamer 407 as it has

been largely studied as a gelling and thickening agent in medical, pharmaceutical and cosmetic

industries due to its low toxicity [4,5].

52

It has been proposed that at low temperature, the poloxamer 407 unimers coexist with

conventional o/w microemulsion droplets, but as temperature increases, hydrogen bonds between

the central poly(propylene oxide) (PPO) blocks of poloxamer 407 and water are minimized, and

the dehydration of PPO drives the unimers to form micelles [7]. These micelles are comprised of

hydrophobic PPO blocks as the core and hydrophilic poly(ethylene oxide) (PEO) blocks as the

hydrated corona.

These poloxamer 407-based systems, however, have only been studied as hydrogels [5,8,9] or

MBGs with low (≤ 3.5%) oil concentrations [7]. For the purpose of this study, and for

pharmaceutical applications, a larger quantity of solubilized oil in the microemulsion is ideal for

the delivery of higher doses of lipophilic drugs. Currently, the role of microemulsion

components, such as surfactants, oil and water, in producing these poloxamer 407 MBGs is

unclear, and guidelines for optimal formulation conditions are also missing in the literature.

To address these gaps, this work aims at producing thermosensitive poloxamer 407-stabilized

MBGs using biocompatible lecithin-linker microemulsions that have high oil solubilization

capacity, and at evaluating the effect of formulation components (surfactant, linkers and oil) on

gel properties. The microemulsions presented here were based on the lecithin-linker

microemulsions previously reported by Yuan et al. [10]. The hydrophilic linkers sodium

caprylate and caprylic acid used by Yuan et al. were replaced with mild nonionic PEG-6-

caprylic/capric glycerides. Phase maps were constructed for these new lecithin-linker

microemulsions and were used as guidance for the selection of microemulsion systems for

subsequent introduction of poloxamer 407. The effects that microemulsion components have on

the gelation temperature of poloxamer 407 MBGs were also explored by conducting rheological

studies on microemulsion-poloxamer 407 mixtures as a function of oil and hydrophilic linker

content. Finally, the potential of poloxamer 407-stabilized MBG as controlled and sustained

trans-epithelial delivery vehicle was evaluated via in vitro transport studies of dexamethasone

and lidocaine through porcine ear skin and through the MBG itself.

53

4.3 Materials and Methods

4.3.1 Materials

4.3.1.1 Chemicals

Laboratory grade soybean lecithin was purchased from Fisher Scientific (Fairlawn, NJ, USA).

PEG-6-caprylic/capric glycerides (Softigen 767) was received as a gift from Sasol. Poloxamer

407 (Lutrol F127 NF prill surfactant) was kindly donated by BASF. The following chemicals

were purchased from Sigma-Aldrich and were used as received: isopropyl myristate (IPM, 98%),

sorbitan monooleate (SMO, Span® 80, 99%), sodium chloride (99%+, Fluka brand), Dulbecco’s

phosphate buffered saline (PBS) and lidocaine powder (base form, 98%+). Dexamethasone

powder (base form, 99%+) was purchased from BioShop® Canada and acetonitrile (HPLC grade)

from Caledon Laboratories Ltd. Unless otherwise stated, the composition is expressed on a mass

basis (i.e. wt%.) throughout this paper.

4.3.1.2 Skin

Porcine ears were purchased from local supermarket and the preparation procedures are similar

to those reported in Section 3.3.1.2.

4.3.2 Microemulsion phase behaviour and formulation selection

Phase behaviour studies of lecithin-linker microemulsions were conducted using equal volumes

of oil and aqueous solution (5 ml each) in flat-bottom tubes. Systems containing 2 to 5% lecithin

(evaluated at intervals of 1.0%) were investigated at constant temperature (24±2 oC, room

temperature), electrolyte concentration (0.9% NaCl in aqueous solution) and ambient pressure (1

atm). The weight ratio of sorbitan monooleate, the lipophilic linker, to lecithin was kept constant

at 1.3:1. For each lecithin concentration, the ratio of hydrophilic linker PEG-6-caprylic/capric

glycerides to lecithin was gradually increased from 0 to 3. Increasing the hydrophilic linker to

lecithin ratio produces Winsor Type II-III-I and II-IV-I transitions [10]. Table 4.1(A) presents a

summary of the microemulsion formulations considered in this work. After preparing the

formulation, the tubes were thoroughly vortexed and left to equilibrate for 2 weeks. All vials

were kept in an enclosed cabinet away from sunlight at room temperature.

54

The microemulsion system with 3.0% lecithin, 3.1% sorbitan monooleate and 19.6% PEG-6-

caprylic/capric glycerides were selected for drug loading and in vitro transport studies. Table

4.1(B) lists the composition of the microemulsion system. This microemulsion is characterized

as a water-continuous Windsor Type I (o/w) system. It was selected among a set of Type I

microemulsion systems capable of producing thermosensitive poloxamer 407 MBGs since this

system solubilized the highest amount of oil (7.5%) in the microemulsion phase.

Table 4.1 Summary of formulations.

(A) Microemulsions used in phase behaviour studies (%, w/w).

% LE % SMO % PEG-6-CCG % NaCl* % Water % IPM

2.0 2.6 1.5 to 5.0 0.45 49.1 – %PC 45.4

3.0 3.9 3.5 to 7.0 0.45 49.1 – %PC 43.1

4.0 5.2 3.0 to 9.0 0.45 49.1 – %PC 40.8

5.0 6.5 5.0 to 12.0 0.45 49.1 – %PC 38.5 * 0.9% over the aqueous phase (1:1 of aqueous phase and oil phase).

(B) Microemulsion used in drug loading and in vitro transport studies (%, w/w).

% LE % SMO % PEG-6-CCG* % NaCl % Water % IPM*

3.0 3.1 19.6 0.9 49.5 23.9

LE, lecithin; SMO, sorbitan monooleate; PEG-6-CCG, PEG-6-caprylic/capric glycerides; IPM, isopropyl myristate * PEG-6-caprylic/capric glycerides is used for dexamethasone loading, and IPM used for lidocaine loading.

Amphiphilic dexamethasone base (poorly oil soluble) was selected as a model drug and was

introduced in the formulation by pre-dissolving the drug in PEG-6-caprylic/capric glycerides to a

concentration of 0.8%, producing microemulsion with dexamethasone content (0.2%) twice that

of commercial products (0.1%). It is worth mentioning that the solubility of dexamethasone in

de-ionized water (0.12 mg/ml) and IPM (0.23 mg/ml) was significantly lower than that in PEG-

6-caprylic/capric glycerides (8.26 mg/ml).

Another amphiphilic drug, lidocaine base (oil-soluble), was chosen as the second model drug for

this study. It was incorporated in the formulations by pre-dissolving 10% of the drug in IPM,

giving a final concentration of 0.75% (g lidocaine/g microemulsion).

55

4.3.3 MBG preparation

MBGs were prepared by addition of appropriate amount of poloxamer 407 (20%) to Winsor

Type I (o/w) microemulsions. For water-continuous microemulsions with excess oil, the excess

oil was removed prior to mixing with poloxamer 407. The poloxamer 407-microemulsion

mixture was stored in a refrigerator at 4 oC with constant mixing until the poloxamer 407

dissolved completely (approximately 24 hours). Two types of MBGs were produced: 1)

thermosensitive MBGs which transformed from liquid to gel upon heating beyond the gelation

temperature; 2) non-thermosensitive MBGs (i.e. remain gel-like from 4-45 °C). The order of

mixing poloxamer 407 into the solution is also critical. Pre-dissolving poloxamer 407 into the

aqueous phase followed by mixing with the oil phase would produce pluronics lecithin organogel

(PLO)-like systems rather than MBGs (unpublished observations). For this reason, it is necessary

to obtain the microemulsion prior to mixing with poloxamer 407.

4.3.4 Microemulsion characterization

Conductivity, viscosity, and hydrodynamic diameter of the microemulsions were measured

according to the description in Section 3.3.5.

4.3.5 Rheological measurements

Rheological measurements of the MBGs were obtained according to the description in Section

3.3.6, with some modifications. Briefly, oscillation experiments were performed using a 6 mm

diameter cone and plate geometry (cone angle 2o) at 1.00 Hz frequency, 1 s-1 shear rate and

1.77% oscillation strain over the temperature range of 4 to 45 oC (heating rate ~3 oC/min). At the

minimum temperature the formulations were liquid and gelled upon heating. The elastic moduli

(G’) of the samples were recorded and were used to characterize the rheological properties of the

poloxamer 407 MBGs.


To evaluate the performance of poloxamer 407 MBGs as delivery vehicles, amphiphilic model

drugs dexamethasone base (poorly oil soluble) and lidocaine base (oil soluble) were chosen in

the transport studies.

56


The in vitro transdermal extended release experiments were conducted as described in Section

3.3.7.1, except that liquefied poloxamer 407 MBG was applied at approximately 4 oC instead of

~50 oC for the gelatin MBG.


In vitro permeation experiments were conducted as described in Section 3.3.7.2, except that

liquefied poloxamer MBG was applied at approximately 4 oC.

4.3.7 Dexamethasone quantification

The concentration of dexamethasone in the skin (determined via the skin extract) and receiver

solutions was quantified using a Dionex ICS-3000 liquid chromatography system equipped with

a reversed-phase C18 column (Genesis, 4μm, 150mm×4.6 mm) and AD25 absorbance detector.

The mobile phase was a 45:55 (v:v) mixture of acetonitrile and double distilled water, flow rate

was set to 1.0 ml/min. At wavelength 240 nm, the dexamethasone retention time was 3.9-4.0

min. The calibration curve for area under the peak vs. concentration was linear (R2 = 0.9998).

4.3.8 Lidocaine quantification

The concentration of lidocaine was quantified using Dionex ICS-3000 liquid chromatography

system under conditions described in Section 3.3.8.

4.4 Results and Discussion

4.4.1 Phase behaviour of lecithin-linker microemulsions

Figure 4.1 presents the “phase maps” of lecithin-linker microemulsions where the phase

boundaries for different microemulsion transitions (Type II-III, II-IV, III-I or IV-I) are plotted in

terms of lecithin concentration (y-axis) and PEG-6-caprylic/capric glycerides-to-lecithin ratio (x-

axis), while holding a constant sorbitan monooleate (SMO) to lecithin weight ratio of 1.3:1. The

use of hydrophilic linker/lecithin ratios to define microemulsion phase boundaries for lecithin-

linker systems is justified [11]. It is worth mentioning that the straight “X” shape of the phase

boundaries in Figure 4.1 deviated from the slanted “X” boundaries reported by Acosta et al. [11]

and Yuan et al. [12]. In both articles, the slanted shape was associated with the selective

57

partition/segregation of hydrophilic linkers (C6 or C8 tail) into the aqueous phase. It is possible

that these partition/segregation effects were not significant in this work because the PEG-6-

caprylic/capric glycerides has a significant fraction of less hydrophilic C10 glycerides.

Conductivity measurements confirmed the microemulsion types in Figure 4.1. The conductivity

in Type II microemulsions was approximately 0 μS/cm. The conductivity of Type III and IV

systems was approximately 200 μS/cm, and between 900 and 3100 μS/cm for Type I

microemulsions.

The mean hydrodynamic diameter of the microemulsion aggregates (Type I and II) is between 2

to 10 nm. These sizes are comparable to the value of 6 nm obtained by Yuan et al. [10]. The

small droplet size of lecithin-linker microemulsions has been associated with the use of

hydrophilic linkers that increase the interfacial area and reduces the size of oil-swollen micelles

and water-swollen reverse micelles [13,14].

The viscosity of the microemulsions at room temperature was measured to be approximately 10

to 60 mPa·s. These values are substantially lower compared to the viscosities of conventional

medium-chain alcohol lecithin microemulsions (~1000 mPa·s) [15] and commercial topical

creams such as Lanocort 10 (~1300 mPa·s) [16]. The low viscosity of lecithin-linker

microemulsions makes them suitable for spray or roll-on applications, however this viscosity

needs to be modified for gel-type applications.

Figure 4.1 Phase maps of lecithin-linker microemulsions formulated with equal volumes of aqueous and oil (IPM) phases at SMO/lecithin ratio of 1.3/1 at different temperatures.

58

According to Figure 4.1, at 24±2 oC (room temperature), the intersection of the “X” occurs at

~4.4% lecithin. This implies that a single phase, biocontinuous (Type IV) microemulsion with

equal volumes of oil and aqueous phase can be obtained with approximately 4.4% lecithin, 5.7%

sorbitan monooleate and 6.8% PEG-6-caprylic/capric glycerides, giving a total of ~17%

surfactant concentration. This quantity is substantially lower than the 40-60% surfactants

required to produce single phase systems with polymeric surfactants [17,18], and show the larger

solubilization capacity of lecithin-linker microemulsions. The phase map at 4 oC is also plotted

in Figure 4.1. With a decrease in temperature (i.e. 24 to 4 oC), the position of the “X” shifted

towards the left hand side of the graph, indicating that lower PEG-6-caprylic/capric glycerides

concentration is required to produce a bicontinuous microemulsion. As temperature decreases,

the hydrogen bonding between the ethylene oxide groups of the hydrophilic linker and the water

molecules becomes stronger (hydration) and the formulations become more hydrophilic, thus

requiring a lower concentration of the hydrophilic linker to produce a biocontinuous

microemulsion [19,20,21].

In addition, the amount of oil solubilized in the microemulsions (i.e. oil content) was also

determined. The oil solubilization in some of the formulations was determined by measuring the

corresponding phase volumes. The hydrophlic-lipophilic difference, net-average curvature

(HLD-NAC) model was used to construct the oil solubilization contour lines in Figure 4.2. This

figure compares the experimental solubility values to the calculated solubility ranges (using the

HLD-NAC model) that will be used later to interpret the formation and properties of MBGs. The

HLD-NAC model has been applied to lecithin-linker microemulsions as well as microemulsions

formulated with ionic and nonionic surfactants [21,22,23,24]. For linker-based lecithin

microemulsions of IPM, the net curvature (HN) for Type I microemulsions, away from the Type

I-III and I-IV transitions, can be simplified as [22,24]:

/ / (1)

where RO is the equivalent solubilization radius of oil, L is a length scaling parameter that has

been found to be proportional to the extended length of the surfactant tail (lipophilic group),

HL/Le is the ratio of the hydrophilic linker (PEG-6-caprylic/capric glycerides) to lecithin for any

given formulation, and HL/Le* is the ratio that produces the optimum formulation (equal

volumes of oil and water solubilized in the middle phase, intersection of the X in Figure 4.1 at 24

59

oC). The “b” term in the equation is a constant that was adjusted in order to produce

solubilization contour lines consistent with experimental solubilization data. A more detailed

calculation is presented in the Appendix.

Figure 4.2 Oil solubilization contour diagram. The experimentally measured oil soluilization values are reported beside the “●”, and the contour lines represent those calculated using the HLD-NAC model.

From Figure 4.2, the oil solubilization contour lines calculated using the HLD-NAC model

approximately match the oil contents measured experimentally, with a few exceptions at low

PEG-6-caprylic/capric glycerides-to-lecithin (P/L) ratios (e.g. P/L < 3). At low P/L ratios, the

microemulsions are closer to the Type I-III and I-IV transitions (P/L ~ 1.5), hence some of the

assumptions used to simplify the HLD-NAC equation may not be valid. Nevertheless, this model

still allowed us to predict the oil solubilization lines relatively accurately.

4.4.2 Phase behavior of microemulsion-poloxamer 407 mixtures

Type I microemulsion systems with various lecithin, PEG-6-caprylic/capric glycerides, and

solubilized isopropyl myristate (IPM) content were studied for their ability to form poloxamer

407 MBGs. Microemulsions with intermediate to high lecithin concentration (thus higher oil

solubilization potential) and a wide range of P/L ratios (systems become more hydrophilic at

higher ratio) were selected for this investigation.

60

Microemulsion-poloxamer 407 mixtures with thermosensitivity (thermosensitive MBGs) and

non-thermosensitivity (non-thermosensitive gels or liquids) were obtained (Figure 4.3). The

thermosensitive MBGs are systems that undergo sol-gel transition within the temperature range

of 4-45 oC. On the other hand, within the specified temperature range, no sol-gel or gel-sol

transitions were observed for the non-thermosensitive liquids or gels, respectively.

Figure 4.3 Behaviours of microemulsion-poloxamer mixtures. Only under certain conditions can thermosensitive poloxamer 407 MBGs be produced. The symbols indicate the formulations evaluated for this study (▲: non-thermosensitive gels; ■: non-thermosensitive liquids; *: thermosensitive MBGs).

The results from this preliminary study indicate that producing thermosensitive poloxamer 407

MBGs is a complex task and it is difficult to predict the best formulation without a more

systematic study of the influence of each microemulsion component on gel properties. For this

reason, additional rheological studies were performed to investigate the effect of microemulsion

components on sol-gel transition temperature of microemulsion-poloxamer 407 mixtures. The

results are reported in Section 4.4.3.

4.4.3 Effect of microemulsion components on sol-gel transition temperature

In order to understand the effect of microemulsion components and their intermediate mixtures

on the sol-gel transition temperature of microemulsion-poloxamer 407 mixtures, rheological

61

studies were performed on selected combinations of microemulsion components as indicated in

Figure 4.4, 4.5 and 4.6. The sol-gel transition temperature was defined as the point where the

elastic modulus (G’) was half way between G’ for the liquid solution and G’ for the gel [25].

From the rheological results shown in Figure 4.4, the addition of lipophilic components such as

sorbitan monooleate and lecithin to water (control) reduced the sol-gel transition temperature. On

the other hand, the addition of PEG-6-caprylic/capric glycerides to water increased the sol-gel

transition temperature to greater than 40 oC. Because of the significant effect of PEG-6-

caprylic/capric glycerides on gel properties, as shown in Figure 4.4, a more detailed study was

undertaken.

Figure 4.4 Effects of microemulsion components on sol-gel transition temperature of systems containing 20% poloxamer 407: (◊) control (water); (x) 3.0% lecithin + water; (▲) 3.1% sorbitan monooleate + water; (□) 19.6% PEG-6-caprylic/capric glycerides + water.

4.4.3.1 PEG-6-caprylic/capric glycerides content

The elastic modulus of aqueous formulations containing 20% poloxamer 407, 3.0% lecithin,

3.1% sorbitan monooleate, and 2.5 to 25.0% PEG-6-caprylic/capric glycerides was evaluated as

a function of temperature.

The results showed that the addition of PEG-6-caprylic/capric glycerides caused an increase in

sol-gel transition temperature and the effect was concentration dependent: the higher the

concentration of hydrophilic linker, the greater the increase in transition temperature as shown in

62

Figure 4.5. As PEG-6-caprylic/capric glycerides content increased from 2.5 to 15.0%, the sol-gel

transition temperature increased gradually from 6 to 14 oC. However, significant increases in

temperature occurred at 20 and 25% PEG-6-caprylic/capric glyceride, where no gelation was

observed even at 45 oC.

Figure 4.5 Elastic modulus vs. temperature profile for the systems containing 3.0% lecithin, 3.1% sorbitan monooleate, 2.5% (▲), 5.0% (■), 10.0% (◊), 15.0% (x), 20.0% (*) or 25.0% (○) PEG-6-caprylic/capric glycerides and balanced amount of water.

Similar concentration-dependent modification effects were observed by Gilbert et al., where

water-soluble PEG polymers were reported to increase the transition temperature of poloxamer

407 MBGs [26]. The PEG molecules may form mixed micelles with poloxamer molecules,

hindering the poloxamer-poloxamer association [26].

4.4.3.2 Isopropyl myristate (IPM) content

To investigate the effect of IPM content on sol-gel transition temperature, the elastic modulus of

formulations containing 20% poloxamer 407, 3.0% lecithin, 3.1% sorbitan monooleate, 19.6%

PEG-6-caprylic/capric glycerides and 0 to 8% IPM (just over the solubilization capacity of 7.5%

IPM) was evaluated (Figure 4.6).

63

Figure 4.6 Elastic modulus vs. temperature profile for the systems containing 3.0% lecithin, 3.1% sorbitan monooleate, 19.6% PEG-6-caprylic/capric glycerides, 0% (*), 2.0% (□), 4.0% (▲), 6.0% (x) or 8.0% (◆) IPM and balanced amount of water.

At 2 to 6% IPM, water-continuous (Type I) single phase microemulsions were obtained. The

addition of poloxamer 407 to the microemulsions produced transparent liquid systems at 4 oC. At

8% IPM, a Type I microemulsion with excess oil phase was obtained (solulbilization capacity

7.5%). The addition of 20% poloxamer to the microemulsion+excess oil formulations produced

translucent-opaque gels at 4 oC, where the poloxamer 407 micelles and their micellar

associations helped to solubilize the excess oil that co-existed with the microemulsion.

The rheological studies demonstrated that the addition of IPM caused a reduction in the sol-gel

transition temperature of poloxamer 407-based systems, and that this effect was also dependent

on the oil content. Figure 4.6 shows that the gelation temperature for the IPM-free formulation

was above 45 oC and decreased to 28 oC in 2% IPM formulations, and to 9 oC in 6% IPM system.

At 8% IPM, the gelation temperature was further reduced to ~5 oC. Furthermore, the elastic

modulus (G’) tends to increase with increasing oil content.

According to Figure 4.5, both 20% and 25% PEG-6-caprylic/capric glycerides significantly

increased gelation temperature (to beyond 45 oC) in the absence of IPM. Since the addition of

IPM reduced the gelation temperature of formulations containing 20% PEG-6-caprylic/capric

glycerides (Figure 4.6), this effect was investigated with systems containing 25% PEG-6-

caprylic/capric glycerides (Figure 4.7).

64

Figure 4.7 Elastic modulus vs. temperature profile for the systems containing 3.0% lecithin, 3.1% sorbitan monooleate, 25% PEG-6-caprylic/capric glycerides, 4.0% (□), 8.0% (▲) or 10.0% (x) IPM and balanced amount of water.

From Figure 4.7, gelation temperatures of the formulations decreased from ~15 oC for 4% IPM,

to 7 oC for 8% IPM and below 4 oC for 10% IPM. These are consistent with the results from

formulations containing 20% PEG-6-caprylic/capric glycerides. In addition, the effect of IPM on

gelation temperature seemed to be more significant than the effect of PEG-6-caprylic/capric

glycerides, since at 4% IPM, the gelation temperature was almost the same for systems

formulated with 20% and 25% PEG-6-caprylic/capric glycerides.

In order to understand the combined effect of microemulsion components, it is important to keep

in mind that the gelation mechanism of poloxamer 407 involves the change in micelle

morphology from small spherical (high curvature) micelles to a network of worm-like (low

curvature) micelles with increasing temperature [4,7,27]. Microemulsion components that

increase the lipophilicity of the formulation (lecithin, sorbitan monooleate and IPM) tend to

reduce the curvature (increase radius) of the micelles, thus facilitating the formation of worm-

like micelles and reducing the transition temperature. The only hydrophilic component in the

formulation, PEG-6-caprylic/capric glycerides, tends to reduce the solubilization of oil and

increases the curvature of micelles, which results in an increase in the gelation temperature.

While lecithin-linker microemulsions are highly efficient in solubilizing a wide range of oils, the

restrictions described above limit the ability to take full advantage of this high solubilization

65

capacity. Still, the level of oil solubilization is approximately twice than that reported by Zhao et

al. [7].

Superimposing the oil solubilization contour lines (2 to 8% IPM, consistent with the oil range in

Figure 4.6) onto Figure 4.3, this new figure (Figure 4.8) helps visualizing why thermosensitive

MBGs and non-thermosensitive gels/liquids are produced. From Figure 4.8, it can be seen that

thermosensitive MBGs tend to form with microemulsions containing less than 8% solubilized

IPM, and with 20-25% PEG-6-caprylic/capric glycerides content. Formulations with high oil

content and relatively low PEG-6-caprylic/capric glycerides result in the formation of non-

thermosensitive gels. High PEG-6-caprylic/capric glycerides content and relatively low oil

content tend to produce temperature-insensitive liquids. These observations support the argument

that the concentrations of the lipophilic and hydrophilic components influence the overall

properties of poloxamer 407 MBGs.

Figure 4.8 Behaviours of microemulsion-poloxamer mixtures with constant PEG-6-caprylic/capric glycerides (PEG-6-CCG) and oil contour lines.

An additional factor which may have affected the overall properties of poloxamer 407 MBGs

was the MBG preparation method. As previously described, Type I microemulsions coexisting

with excess oil phases were formulated and the microemulsions were then extracted (by

removing excess oil) for subsequent mixing with poloxamer 407 (refer to Section 4.3.3). This

method assumed that the excess oil phase was composed of pure IPM only, which may not be

66

valid since some of the surfactant and lipophilic linker may have partitioned into this excess

phase. As a result, the actual compositions of the extracted microemulsions were probably

different than the compositions reported, leading to complications in the characterization of the

MBGs. For this reason, the MBG preparation method was slightly modified for the rheological

studies (i.e. investigations of PEG-6-caprylic/capric glycerides and IPM effects), where

microemulsions with oil contents less than or equal to the maximum oil solubilization capacity

(i.e. without adding excess amounts of oil during formulation) were formulated. This method

prevented the formation of excess oily phases (along with the partitioned surfactant and

lipophilic linker) and the possible change in microemulsion compositions, and is the

recommended MBG preparation method.


The 20% poloxamer 407 MBG prepared with 3.0% lecithin, 3.1% sorbitan monooleate, 19.6%

PEG-6-caprylic/capric glycerides, 0.9% NaCl solution, 49.5% water and 7.5% solubilized IPM

was evaluated as a delivery vehicle to load dexamethasone in the skin, and use the skin as a

dexamethasone reservoir (in situ patch) for extended release. Using a permeable synthetic

membrane, instead of tissue, the gel itself was evaluated as a reservoir for the extended release of

dexamethasone and lidocaine. The corresponding poloxamer 407-free microemulsion was used

as a control to compare the transport of dexamethasone and lidocaine in both scenarios.


The total mass of dexamethasone loaded in the skin from the microemulsion (control) and the

poloxamer MBG were determined considering the residual mass of dexamethasone in skin and

the dexamethasone permeated to the receiver solution. Slightly less dexamethasone penetrated

the skin from the poloxamer 407 MBG (23.5±6.1 μg/cm2) than from the microemulsion

(28.5±8.9 μg/cm2) during the initial 30 min loading period, suggesting that dexamethasone

permeation into skin is slower from the poloxamer 407 MBG. Release of dexamethasone from

the skin was observed after removal of the microemulsion and the poloxamer MBG from the

donor compartment, confirming that both formulations acted as in-skin drug depot for the release

of dexamethasone.

Figure 4.9 illustrates the dexamethasone release profile for the microemulsion- and MBG-loaded

skins. For the microemulsion, approximately 75% of the dexamethasone in the skin was released

67

during the first 24h, and the system showed first-order release, confirmed with log (-dm/dt) vs.

log (m) graph, slope (order) = 1.18±0.10, r2 > 0.97. On the other hand, the poloxamer 407 MBG

demonstrated near zero-order release of dexamethasone (order = 0.25±0.026, r2 > 0.97). A lag

time was observed during the first 3h. Beyond this time, the release increased and maintained at

an almost constant rate. At the completion of the study (48h), approximately 87% of the

dexamethasone initially loaded in the skin was released. Both the microemulsion and the MBG

were able to provide extended delivery of dexamethasone for 24h.

The release profiles of Figure 4.9 confirmed that dexamethasone in the lecithin-linker

microemulsion and its poloxamer 407 MBG was absorbed in the skin to act as in situ drug depot

for extended drug delivery.

Figure 4.9 Release profile of dexamethasone from the skin after loading with: 20% poloxamer 407 MBG (◆), and dexamethasone-loaded gelatin-free μE (∆). Dexamethasone loading in the donor compartment was 23.5±6.1 μg/cm2 for the MBG, and 28.5±8.9 μg/cm2 for the microemulsion.


As explained in Section 3.4.3.2, cellulose acetate membrane with a molecular weight cut off

(MWCO) of 100 kDa, which corresponds to an approximate pore size of 10 nm, was used as a

simplified model for gel transport study. Shortly after application (< 1 min), the poloxamer 407

MBG (at 4 oC) gelled inside the donor compartment of the MatTek Permeation Device. This

property is desirable in periocular ophthalmic delivery since the relatively rapid gel formation

could help to prevent formulation removal from the cul-de-sac by tear fluid and blinking, thus

providing extended residence time.

68

According to the gel permeation profiles shown in Figure 4.10, dexamethasone permeation from

the poloxamer 407-free microemulsion (control) was slightly faster than the poloxamer 407

MBG. In addition, permeation from the microemulsion followed a high order kinetics while it

was zero-order from the poloxamer MBG (order = 0.16±0.02, r2 > 0.99), consistent with the

release profiles from the skin transport study. The zero-order kinetics may be a result of the

interaction between the amphiphilic poloxamer 407 and the amphiphilic dexamethasone. To

explore the potential role of this interaction, an additional permeation study of dexamethasone

from a 20% poloxamer 407 hydrogel containing 10% PEG-6-caprylic/capric glycerides in

deionized water was conducted and compared to the permeation profiles from the poloxamer

MBG. Dexamethasone permeation from the poloxamer 407 hydrogel followed a slower but

higher order release kinetics (Figure 4.10). In addition, the different permeation profiles from the

microemulsion, poloxamer MBG, and hydrogel suggested different dexamethasone

permeation/release mechanisms from the three vehicles. Further study is required to investigate

the interactions between poloxamer 407, microemulsion and dexamethasone, and to propose

possible mechanisms for the zero-order kinetics from the poloxamer 407 MBG; this is beyond

the scope of the current study.

Figure 4.10 Permeation profile of dexamethasone from in situ poloxamer 407 MBG (◆), dexaemthasone-loaded poloxamer-free microemulsion (∆), and poloxamer 407 hydrogel (●). Dexamethasone loading in the donor compartment was 281±8 µg/cm2 for the MBG, 352±8 µg/cm2 for the microemulsion, and 137±8 µg/cm2

for the hydrogel.

In addition to dexamethasone (poorly oil-soluble), the permeation of lidocaine, an oil-soluble

amphiphilic drug, was also studied. Figure 4.11 presents the permeation profiles of lidocaine

from the poloxamer 407-free microemulsion (control) and the poloxamer 407 MBG. In contrast

69

to the linear profiles of dexamethasone from poloxamer 407 MBG, the permeation profiles for

lidocaine were of higher order from both the poloxamer 407-free microemulsion and the

poloxamer 407 MBG.

Figure 4.11 Permeation profile of lidocaine from in situ poloxamer 407 MBG (◆) and lidocaine-loaded poloxamer-free microemulsion (∆). Lidocaine loading in the donor compartment was 1150±50 µg/cm2 for the poloxamer MBG, and 1380±40 µg/cm2 for the microemulsion.

The different results from dexamethasone and lidocaine suggest that the release kinetics from

poloxamer 407 MBGs is probably dependent on the interaction between the drug and the

poloxamer micelles. The amphiphilicity of dexamethasone and its low solubility in IPM likely

segregated the drug in the amphiphilic surfactant phase, which interacted with the poloxamer 407

micelles.

Zero-order kinetics with poloxamer 407 systems was also observed by Bhardwaj and Blanchard

[28], where the authors reported a linear relationship between percentage Melanotan-I released

and poloxamer 407 dissolution in aqueous solutions. Similarly, the zero-order kinetics obtained

in this study could also result from the degradation of poloxamer 407 MBGs. In fact, during the

48h period of the permeation study, swelling of the poloxamer MBGs was observed. Following

instillation of the liquefied poloxamer MBG (4 oC), the formulation gelled in the donor

compartment after equilibration to the experimental temperature (34 oC). The MBG then began

to swell with water from the receiver solution. At the end of the 48h study, the poloxamer 407

MBG was completely degraded, and the donor compartment turned into a liquid.

70

In the case of lidocaine, the fact that it is highly soluble in IPM (solubility ~20%) indicates that a

great fraction of this drug is still associated with the oil in the microemulsion, and even in the

presence of poloxamer, its penetration is likely mediated by microemulsion transport.

4.5 Conclusions

In this study, thermosensitive poloxamer 407-stabilized microemulsion-based gels (MBGs) are

produced with water continuous, low toxicity lecithin-linker microemulsions containing lecithin,

sorbitan monooleate, PEG-6-caprylic/capric glycerides and isopropyl myristate (IPM). These

MBGs can be made optically transparent and have rigid gel structures (G’ ~ 10 kPa). The

rheological studies led to the conclusion that lipophilic lecithin, sorbitan monooleate and IPM

lower the gelation temperature of poloxamer 407 systems, likely by facilitating the formation of

worm-like poloxamer micelles. On the other hand, high concentrations of hydrophilic linker

PEG-6-caprylic/capric glycerides tend to increase the gelation temperature by reducing the

overall lipophilicity of the formulation. Overall, poloxamer 407 MBGs with upto 8% IPM could

be produced, more than twice the oil solubilization capacity reported for a previous poloxamer

MBG. Further optimization is still possible with systems with higher PEG-6-caprylic/capric

glycerides content.

When investigated as drug delivery vehicles, dexamethasone-loaded poloxamer 407 MBGs

showed zero-order release kinetics while the release from the “parent” microemulsions was first-

order. Transport studies with lidocaine demonstrated that drug properties (amphilicity,

lipophilicity, etc.) could also interfere with the release mechanisms from poloxamer 407 MBGs.

Further studies into the release mechanisms and gel degradation are necessary to optimize the

use of these formulations in ophthalmic and transdermal applications.

4.6 References



71

3. Ruel-Gariepy, E.; Leroux, J-C. In situ-forming hydrogels - review of temperature-sensitive systems. Eur. J. Pharm. Biopharm. 2004, 58, 409-426.

4. Dumortier, G.; Grossiord, J.L.; Agnely, F.; Chaumeil, J.C. A review of poloxamer 407 pharmaceutical and pharmacological characteristics. Pharm. Res. 2006, 23, 2709-2728.

5. Escobar-Chávez, J.J.; López-Cervantes, M.; Naïk, A.; Kalia, Y.N.; Quintanar-Guerrero, D.; Ganem-Quintanar, A. Applications of thermo-reversible Pluronic F127 gels in pharmaceutical formulations. J. Pharm. Pharmaceut. Sci. 2006, 9, 339-358.

6. Qian, Y.; Wang, F.; Li, R.; Zhang, Q.; Xu, Q. Preparation and evaluation of in situ gelling ophthalmic drug delivery system for methazolamide. Drug Dev. Ind. Pharm. 2010, 36, 1340-1347.


8. Qi, H.; Li, L.; Huang, C.; Li, W.; Wu, C. Optimization and physicochemical characterization of thermosensitive poloxamer gel containing puerarin for ophthalmic use. Chem. Pharm. Bull. 2006, 54, 1500-1507.






14. Acosta, E.; Uchiyama, H.; Sabatini, D.A.; Harwell, J.H. The role of hydrophilic linkers. J. Surfactants Deterg. 2002, 5, 151-157.

15. Sabatini, D.A.; Acosta, E.J.; Harwell, J.H. Linker molecules in surfactant mixtures. Curr. Opin. Colloid Interface Sci.2003, 8, 316-326.

16. Adeyeye, M.C.; Jain, A.C.; Ghorab, M.K.; Reilly, W.J., Jr. Viscoelastic evaluation of topical creams containing microcrystalline cellulose/sodium carboxymethyl cellulose as stabilizer. AAPS Pharm SciTech 2002, 3, E8.


72

18. Kreilgaard, M. Influence of microemulsions on cutaneous drug delivery. Adv. Drug Deliver. Rev. 2002, 54, S77-S98.

19. Tyrode, E.; Johnson, C.M.; Rutland, M.W.; Claesson, P.M. Structure and hydration of poly(ethylene oxide) surfactants at the air/liquid interface. A vibrational sum frequency spectroscopy study. J. Phys. Chem. C 2007, 111, 11642-11652.

20. Tyrode, E.; Johnson, C.M.; Kumpulainen, A.; Rutland, M.W.; Claesson, P.M. Hydration state of nonionic surfactant monolayers at the liquid/vapor interface: Structure determination by vibrational sum frequency spectroscopy. J. Am. Chem. Soc. 2005, 127, 16848-16859.

21. Acosta, E.J.; Bhakta, A.S. The HLD-NAC model for mixtures of ionic and nonionic surfactants. J. Surfactants Deterg. 2009, 12, 7-19.

22. Acosta, E.J. The HLD-NAC equation of state for microemulsions formulated with nonionic alcohol ethoxylate and alkylphenol ethoxylate surfactants. Colloid Surface A 2008, 320, 193-204.

23. Kiran, S.K.; Acosta, E.J. Predicting the morphology and viscosity of microemulsions using the HLD-NAC model. Ind. Eng. Chem. Res. 2010, 49, 3424-3432.

24. Acosta, E. Modeling and formulation of microemulsions: the net-average curvature and the combined linker effect. Doctoral Dissertation. University of Oklahoma, Norman, Oklahoma. 2004.

25. Edsman, K.; Carlfors, J.; Peterson, R. Rheological evaluation of poloxamer as a in situ gel for ophthalmic use. Eur. J. Pharm. Sci. 1998, 6, 105-112.

26. Gilbert, J.C.; Richardson, J.L.; Davies, M.C.; Palin, K.J.; Hadgraft, J. The effect of solutes and polymers on the gelation properties of pluronic F-127 solutions for controlled drug delivery. J. Controlled Release 1987, 5, 113-118.

27. Mortensen, K. Structural studies of aqueous solutions of PEO-PPO-PEO triblock copolymers, their micellar aggregates and mesophases; a small-angle neutron scattering study. J. Phys-Condens. Mat. 1996, 8, A103-A124.

28. Bhardwaj, R.; Blanchard, J. Controlled-release delivery system for the α-MSH analog Melanotan-I using poloxamer 407. J. Pharm. Sci. 1996, 85, 915-919.

73

CHAPTER 5

Conclusions and Recommendations

This chapter summarizes the knowledge gained in the individual sections of this thesis and

presents some recommendations for future work.

5.1 Conclusions

The overall purpose of this work was to produce thermosensitive gelatin- and poloxamer 407-

stabilized MBGs using biocompatible lecithin-linker microemulsions and to investigate the

influences of microemulsion morphology and formulation components on the overall properties

of lecithin-linker MBGs.

In Chapter 3 thermosensitive gelatin MBGs were produced using AOT- and alcohol-free

lecithin-linker microemulsions, for the first time. The addition of gelatin to oil-in-water,

bicontinuous, oil-rich and water-rich microemulsions induced microemulsion morphology

change and lead to the formation of MBGs with different rheological and physical properties. In

short, water-rich bicontinuous microemulsions with high gelatin content were desirable in order

to obtain clear and strong MBGs. These gelatin MBGs, with viscosities suitable for topical

applications, demonstrated drug transport properties comparable to their corresponding gelatin-

free microemulsions.

Chapter 4 explored the production of thermosensitive lecithin-linker MBGs using poloxamer 407

as the gelling agent. While lecithin-linker microemulsions are highly efficient at solubilizing a

wide range/amount of oils, high concentrations of lipophilic components in the formulation is

likely to induce the formation of worm-like poloxamer micelles, which reduce gelation

temperature and contribute to the formation of temperature-insensitive MBGs. It was also found

that the hydrophilic linker counterbalances this effect, likely by promoting the formation of

spherical poloxamer micelles, resulting in an increase in the MBG gelation temperature. Overall,

MBGs with up to 8% oil content were obtained (twice that reported in previous studies). The in

vitro transport studies showed that poloxamer 407 MBGs may have the potential as controlled

and extended delivery vehicles for trans-epithelial applications; however the release kinetics are

most likely dependent on the interactions between the drug and the poloxamer micelles. In

74

addition, the degradation of the MBG may represent an issue, particularly for ophthalmic

applications.

Furthermore, it was understood from these two studies (i.e. production of gelatin and poloxamer

407 MBGs) that the “systematic approach”, which examined the effect of individual formulation

components on the overall properties of the MBGs via rheological measurements, was a useful

method in characterizing the gelation properties of complex systems.

5.2 Recommendations for Future Work

In this thesis, we reported the production of thermosensitive MBGs using biocompatible lecithin-

linker microemulsions, the effect of formulation components on overall MBG properties, and the

in vitro transport of lidocaine and dexamethasone from lecithin-linker MBGs. Although these

results demonstrated that lecithin-linker MBGs are promising drug delivery vehicles, there are

some recommended studies that should be addressed for future research:

1. Investigating the effect of drugs on the release kinetics of lecithin-linker MBGs

The in vitro transport of amphiphilic drugs, lidocaine (oil-soluble) and dexamethasone

(poorly oil-soluble), have been demonstrated using lecithin-linker MBGs. However, as

discussed in Chapter 4, the release kinetics from poloxamer 407 MBGs are likely drug-

dependent. It is necessary to investigate a wider range of drugs (with different properties)

and their effects on the release kinetics of both gelatin and poloxamer MBGs.

2. Evaluating the in vivo efficacy of lecithin-linker MBGs

In this work, we have demonstrated the potential of lecithin-linker MBGs as effective in

vitro ophthalmic and trandermal delivery vehicles. However, these in vitro studies are

simplified versions of in vivo conditions, especially in the case of ophthalmic transport,

where blinking, tear clearance and formulation spreading effects are neglected. Future in

vivo studies are necessary to validate the in vitro results and to evaluate the efficacy of

lecithin-linker MBGs as ophthalmic and transdermal drug carriers.

3. Studying simpler formulations

A major limitation of lecithin-linker MBGs is the complexity associated with the

formulations. The microemulsions used in this work are multi-component systems

composed of oil, water, surfactant and linkers, which are likely to establish complex

75

interactions with the gelling agents. For more practical reasons (e.g. lower toxicity, easier

preparation, etc.), simpler systems are preferred over multi-component formulations. For

example, oil organogels, surfactant organogels or surfactant+oil organogels could be

considered.

76

APPENDIX

Oil Solubilization Calculations

In Chapter 4 Section 4.4.1, we provided the simplified HLD-NAC equation for Type I

microemeulsions away from Type I-III and IV transitions [1,2]:

/ / (1)

Rearranging Eq. 1:

/ / (2)

(3)

where HL/Le is the ratio of hydrophilic linker (PEG-6-caprylic/capric glycerides) to lecithin,

HL/Le* is the ratio which produces optimum formulation, RO is the equivalent solubilization

radius of the oil, L is the extended length of the surfactant tail (Llecithin ≈ 24 Å) [2], and the term

“b” is a constant which can be varied to adjust the location of the oil contour lines.

Moreover,

∑ 6.023 10 (4)

where As is the total surfactant interfacial area, nSi is the initial moles of the surfactant and

hydrophilic linker added to the formulation (MWlecithin ≈ 770 g/mol; MWPEG-6-CCG ≈ 550 g/mol),

and ai is the surface area per molecule of the surfactant (alecithin ≈ 90 Å2/molecule, aPEG-6-CCG ≈ 60

Å2/molecule) [1,2]. In Eq. 4, it has been assumed that the concentration of the surfactant in

monomer form (the critical micelle concentration) is negligible.

And the volume of oil solubilizd in the microemulsion (VO) can be calculated:

(5)

77

Total mass of formulation = 10 g ≈ 10 ml. At b = 0.29,

Le (g) HL (g) HL/Le HL/Le* HLD RO (Å) AS (Å2) VO (ml) % IPM1 2 2 1.5 -0.145 165.52 2.02E+22 1.11 11.13 1 3 3 1.5 -0.435 55.17 2.68E+22 0.49 4.92 1 4 4 1.5 -0.725 33.10 3.33E+22 0.37 3.68 1 5 5 1.5 -1.015 23.65 3.99E+22 0.31 3.14 1 6 6 1.5 -1.305 18.39 4.65E+22 0.28 2.85 1 7 7 1.5 -1.595 15.05 5.30E+22 0.27 2.66 1 8 8 1.5 -1.885 12.73 5.96E+22 0.25 2.53 1 9 9 1.5 -2.175 11.03 6.62E+22 0.24 2.43 1 10 10 1.5 -2.465 9.74 7.27E+22 0.24 2.36 1 11 11 1.5 -2.755 8.71 7.93E+22 0.23 2.30 2 4 2 1.5 -0.145 165.52 4.04E+22 2.23 22.27 2 6 3 1.5 -0.435 55.17 5.35E+22 0.98 9.84 2 8 4 1.5 -0.725 33.10 6.66E+22 0.74 7.35 2 10 5 1.5 -1.015 23.65 7.98E+22 0.63 6.29 2 12 6 1.5 -1.305 18.39 9.29E+22 0.57 5.70 2 14 7 1.5 -1.595 15.05 1.06E+23 0.53 5.32 2 16 8 1.5 -1.885 12.73 1.19E+23 0.51 5.06 2 18 9 1.5 -2.175 11.03 1.32E+23 0.49 4.87 2 20 10 1.5 -2.465 9.74 1.45E+23 0.47 4.72 2 22 11 1.5 -2.755 8.71 1.59E+23 0.46 4.61 3 6 2 1.5 -0.145 165.52 6.05E+22 3.34 33.40 3 9 3 1.5 -0.435 55.17 8.03E+22 1.48 14.76 3 12 4 1.5 -0.725 33.10 1.00E+23 1.10 11.03 3 15 5 1.5 -1.015 23.65 1.20E+23 0.94 9.43 3 18 6 1.5 -1.305 18.39 1.39E+23 0.85 8.54 3 21 7 1.5 -1.595 15.05 1.59E+23 0.80 7.98 3 24 8 1.5 -1.885 12.73 1.79E+23 0.76 7.59 3 27 9 1.5 -2.175 11.03 1.99E+23 0.73 7.30 3 30 10 1.5 -2.465 9.74 2.18E+23 0.71 7.08 3 33 11 1.5 -2.755 8.71 2.38E+23 0.69 6.91 4 8 2 1.5 -0.145 165.52 8.07E+22 4.45 44.54 4 12 3 1.5 -0.435 55.17 1.07E+23 1.97 19.68 4 16 4 1.5 -0.725 33.10 1.33E+23 1.47 14.71 4 20 5 1.5 -1.015 23.65 1.60E+23 1.26 12.58 4 24 6 1.5 -1.305 18.39 1.86E+23 1.14 11.39 4 28 7 1.5 -1.595 15.05 2.12E+23 1.06 10.64 4 32 8 1.5 -1.885 12.73 2.38E+23 1.01 10.12 4 36 9 1.5 -2.175 11.03 2.65E+23 0.97 9.74 4 40 10 1.5 -2.465 9.74 2.91E+23 0.94 9.44 4 44 11 1.5 -2.755 8.71 3.17E+23 0.92 9.21 5 10 2 1.5 -0.145 165.52 1.01E+23 5.57 55.67 5 15 3 1.5 -0.435 55.17 1.34E+23 2.46 24.60

78

5 20 4 1.5 -0.725 33.10 1.67E+23 1.84 18.38 5 25 5 1.5 -1.015 23.65 1.99E+23 1.57 15.72 5 30 6 1.5 -1.305 18.39 2.32E+23 1.42 14.24 5 35 7 1.5 -1.595 15.05 2.65E+23 1.33 13.30 5 40 8 1.5 -1.885 12.73 2.98E+23 1.26 12.65 5 45 9 1.5 -2.175 11.03 3.31E+23 1.22 12.17 5 50 10 1.5 -2.465 9.74 3.64E+23 1.18 11.80 5 55 11 1.5 -2.755 8.71 3.97E+23 1.15 11.52 6 12 2 1.5 -0.145 165.52 1.21E+23 6.68 66.81 6 18 3 1.5 -0.435 55.17 1.61E+23 2.95 29.52 6 24 4 1.5 -0.725 33.10 2.00E+23 2.21 22.06 6 30 5 1.5 -1.015 23.65 2.39E+23 1.89 18.87 6 36 6 1.5 -1.305 18.39 2.79E+23 1.71 17.09 6 42 7 1.5 -1.595 15.05 3.18E+23 1.60 15.96 6 48 8 1.5 -1.885 12.73 3.58E+23 1.52 15.18 6 54 9 1.5 -2.175 11.03 3.97E+23 1.46 14.60 6 60 10 1.5 -2.465 9.74 4.36E+23 1.42 14.17 6 66 11 1.5 -2.755 8.71 4.76E+23 1.38 13.82

References

1. Acosta, E.J. The HLD-NAC equation of state for microemulsions formulated with nonionic alcohol ethoxylate and alkylphenol ethoxylate surfactants. Colloid Surface A 2008, 320, 193-204.

2. Acosta, E. Modeling and formulation of microemulsions: the net-average curvature and the combined linker effect. Doctoral Dissertation. University of Oklahoma, Norman, Oklahoma. 2004.

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