Development of a drug delivery platform using multifunctional … · Development of a drug delivery platform using multifunctional polymeric scaffold for scar therapy . Vipul Agarwal,

Development of a drug delivery platform using

multifunctional polymeric scaffold for scar

therapy

Vipul Agarwal, MApplSc

This thesis is presented for the degree of Doctor of Philosophy at The University of

Western Australia

School of Chemistry and Biochemistry

2015

i

Contents Abbreviations .............................................................................................................. iv

Abstract ...................................................................................................................... vii

Acknowledgement ...................................................................................................... ix

Statement of candidate contribution ........................................................................... xi

1 Introduction and literature review ........................................................................ 1

1.1 Tissue engineering .............................................................................................. 2

1.2 Cellular scaffolds ................................................................................................ 3

1.3 Physical features of the cell microenvironment ................................................. 4

1.4 Design concepts and strategies ........................................................................... 5

1.4.1 Physical properties ...................................................................................... 6

1.4.2 Mechanical properties ................................................................................. 7

1.4.3 Surface properties ........................................................................................ 8

1.5 Skin, injury and wound healing .......................................................................... 9

1.5.1 Haemostasis and inflammation ................................................................. 11

1.5.2 Reepithelialisation ..................................................................................... 13

1.5.3 Formation of granulation tissue ................................................................. 16

1.6 Role of TGFβ in fibrosis and scarring .............................................................. 17

1.7 Mechanisms of activation of extracellular TGFβ ............................................. 19

1.8 Exogenous mannose-6-phosphate act as an antagonist of TGFβ1 activation ... 20

1.9 Structural requirements for mannose-6-phosphate recognition to M6P/IGFII

receptor ...................................................................................................................... 22

1.10 Mannose-6-phosphate analogues ..................................................................... 23

1.11 Scarring ............................................................................................................ 27

1.12 Current treatments for skin wound healing ...................................................... 28

1.13 Cell based therapies .......................................................................................... 29

1.14 Matrix based therapies ..................................................................................... 30

ii

1.14.1 Amniotic membrane ................................................................................ 30

1.14.2 Human cadaver allografts and xenografts ............................................... 30

1.14.3 Tissue engineered skin ............................................................................ 32

1.15 Nanofibrous scaffolds for skin tissue engineering ........................................... 34

1.16 Skin regeneration using electrospun scaffolds ................................................. 37

1.17 Electrospun hybrid materials ............................................................................ 41

1.18 Summary .......................................................................................................... 44

1.19 Hypotheses and Aims ....................................................................................... 46

2 Introduction to the series of papers .................................................................... 48

2.1 Development of a universal multifunctional scaffold ......................................... 48

2.2 Evaluation of mannose-6-phosphate analogues as potential anti-scarring agents ...

.......................................................................................................................... 51

2.3 Delivery of the lipophilic mannose-6-phosphate analogue PXS64 using an

electrospun PGMA scaffold ...................................................................................... 53

3 Series of papers ..................................................................................................... 56

A Functional Reactive Polymer Nanofiber Matrix .................................................... 57

Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor

Inhibitors by Intracellular Delivery ........................................................................... 62

Inhibiting the activation of transforming growth factor-β using a polymeric

nanofiber scaffold ...................................................................................................... 67

4 Conclusions and future work ............................................................................... 71

4.1 Tissue engineering of a nanoscaffold ............................................................... 71

4.2 Scar therapy and mannose-6-phosphate analogues .......................................... 73

4.3 Scaffold based delivery of analogue 2 ............................................................. 74

4.4 Future recommendations .................................................................................. 75

4.5 Final remarks .................................................................................................... 77

A Supporting information for papers .................................................................... 79

Supporting Information for ‘A Functional Reactive Polymer Nanofiber Matrix’ .... 80

iii

Supporting information for ‘Enhancing the Efficacy of Cation-Independent Mannose

6-Phosphate Receptor Inhibitors by Intracellular Delivery’ ...................................... 87

Supporting Information for ‘Inhibiting the activation of transforming growth factor-

β using a polymeric nanofiber scaffold’ .................................................................... 98

B Published work not directly included in the thesis ......................................... 103

References ................................................................................................................ 147

iv

Abbreviations

ANOVA Analysis of variance

bFGF Basic fibroblasts growth factor

CDMPR Cation dependent mannose-6-phosphate receptor

CIMPR Cation independent mannose-6-phosphate receptor

DBM Demineralized bone matrix

DCFH-DA 2’, 7’-dichlorodihydrofluorescein diacetate

ECM Extracellular matrix

EGF Epidermal growth factor

ES Electrospun

FTIR Fourier transform infra-red

GAG Glycosaminoglycan

HCAS Human cadaver allograft skin

HDF Human dermal skin fibroblasts

HPLC High pressure liquid chromatography

HSF Human scar fibroblasts

LAP Latency associated peptide

LCST Lower critical solution temperature

LTGFβ Latent transforming growth factor β

M6P Mannose-6-phosphate

M6P/IGF II Mannose-6-phosphate/Insulin-like growth factor II

MMP Matrix metalloproteinase

v

MRI Magnetic resonance imaging

NMR Nuclear magnetic resonance

NP Nanoparticle

Pa Pascal

PBS Phosphate buffer saline

PCL Poly(ε-caprolactone)

Pd Palladium

PEG Poly(ethylene glycol)

PGA Poly(glycolic acid)

PGMA Poly(glycidyl methacrylate)

PLA Poly(lactic acid)

PLGA Poly(lactic-co-glycolic acid)

PNIPAM Poly (N-isopropyl acrylamide)

PU Polyurethane

ROS Reactive oxygen species

RT-qPCR Real time quantitative-polymerase chain reaction

SEM Scanning electron microscopy

SQUID Superconducting quantum interference device

TE Tissue engineering

TEM Transmission electron microscopy

TGFβ Transforming growth factor β

TSP Thrombospondin

UCNP Upconverting nanoparticles

vi

VEGF Vascular endothelial growth factor

vii

Abstract

Tissue engineering is a multidisciplinary approach and has been used to promote

tissue regeneration, wound repair and to enhance drug delivery. In the case of burn

injury, despite the advances in treatment leading to reduced mortality, the problem

of permanent, disfiguring scar formation following healing is far from being

resolved.

Scarring is the result of an imbalance of fibroblast activity resulting in an excess of

architecturally disorganised extracellular matrix protein deposition and is

predominantly mediated by the TGFβ pathway.

Whilst there are a number of promising therapeutic targets identified through our

increasing understanding of scar formation, there has been limited success in the

clinical translation. This can largely be attributed to difficulties with delivery,

stability and efficacy of treatments tested to date.

In this thesis the problems of drug delivery and stability have been addressed using a

combinatorial approach. First, a scaffold was developed to provide a platform for

drug delivery and stability. A novel multifunctional polyglycidyl methacrylate

(PGMA) scaffold was developed using electrospinning. The multifunctionality of

this polymeric scaffold was demonstrated for various applications. These include

surface functionalization of electrospun PGMA with poly (N-isopropyl acrylamide)

for stimuli response surfaces and development of multifunctional nanocomposites

with (NaGdF4:Yb, Er); Pd and Fe3O4 nanoparticles for upconversion fluorescence

imaging, sensing, and magneto-responsive properties. This was followed by

exploration of modified analogues of a potential therapeutic target as anti-scarring

agents. Mannose-6-phosphate (M6P) has been shown to ameliorate scarring by

inhibiting the activation of TGFβ1, a necessary step required for its receptor

recognition and function. However, therapeutic delivery of M6P to a wound is

currently ineffective due to the low stability of M6P and limited ability to maintain

M6P concentration at the site of injury. Therefore whilst targeting TGFβ activity

through M6P has significant potential in burn therapy, stability and delivery issues

must be addressed before this can become a therapeutic reality. Two M6P analogues,

viii

PXS25 (analogue 1) and PXS64 (analogue 2), have been developed in collaboration

with Pharmaxis Ltd., to overcome the metabolic vulnerability of M6P whilst

retaining the receptor recognition function. Initial work was carried out to

investigate the biocompatibility and cytotoxicity capabilities of both analogues

compared to M6P in human dermal skin fibroblasts. Subsequently they were

investigated for their proficiency in regulating the expression of the critical fibrotic

marker, Collagen I. Analogue 2 was shown to significantly inhibit TGFβ1 mediated

up-regulation of collagen I gene expression. However, the lipophilic analogue 2 has

limited bioavailability. The final chapter addresses this specific problem by

incorporating analogue 2 into the multifunctional scaffold and testing the efficacy of

the combined drug/scaffold therapy on human dermal skin fibroblasts.

The tissue engineering approach presented herein demonstrated the potential of

combinatorial scaffold mediated drug delivery method to progress some of the

existing therapeutic targets into clinical therapies. The future work using porcine

wound healing model is a necessary extension to establish the efficacy and potential

of this combinatorial approach in vivo before its clinical translation.

ix

Acknowledgement

I would like to begin by thanking my supervisors for their consistent support

throughout my candidature. I am really grateful to Prof Swaminathan Iyer for

offering me PhD scholarship and introduce me to a completely new stream of

research, to keep me on track and motivate me especially during the difficult times

and moulding me into a better researcher; Prof Fiona Wood for support and

guidance and brining clinical prospective into my work; Dr Mark Fear for bringing

biological insight, enabling me to resolve complex experimental problems and

improving my overall understanding. In addition, I would extend my gratitude

towards Prof Iyer and Dr Fear for going above and beyond to help me with my grant

application, you faith and support only project me to improve and become an

accomplished scientist. I would also like to thank my colleagues Dr Cameron Evans,

Dominic Ho, Diwei Ho, Dr Faizah Yasin, Alaa Munshi, Ruhani Singh, Ivan Lozic,

and Michael Bradshaw, Dr Rahi Varsani, Dr Tristan Clemons and Michael

Challenor for making me welcome and accepting me within the group with all the

humour and pranks and also to support me where possible in the lab. I would like to

acknowledge Dr Tristan Clemons, Dr Nicole Smith, Sumi Shrestha, Callum

Ormonde, Diwei Ho and Dr Marck Norret to proof reading this thesis. I would like

to thank the staff at CMCA especially Lyn Kirilak and Prof Paul Rigby for their

extended support with microscopy instrumentation and experiments. I would like to

also thank Pharmaxis Ltd and the team for inviting me into their lab and welcoming

me within their group, and also to consistently support me during my candidature.

Because of the multidisciplinary aspect to my project, I had the privilege to work

with some really good people including Prof Charlie Bond, Dr Bernard Callus,

Megan Finch, Benjamin Gully, Dr Foteini Hassiotou, Dr Ben Corry, Dr Natalie

Smith, Dulharie Wijeratne and Dr Lindsay Byrne, without whom this thesis would

not have been possible. For the motivation, persistent support and having faith in me

to pursue my research ambitions, I would like to thank Dr Karen Stack, Dr Stephen

Newbery, Dr Manab Sharma, Sinu Sharma and Murray Frith. For funding, I would

like to thank Australian Research Council (ARC), National Health & Medical

x

Research Council (NHMRC), Australian Nanotechnology Network (ANN),

Australian Academy of Science and The University of Western Australia.

Finally, I would like to thank my family, my girlfriend and my friends for their faith

and selfless support enabling me to pursue my goals.

xi

Statement of candidate contribution

This thesis contained published work and work prepared for publication, some of

which has been co-authored. The bibliographical details of the work and author

contributions are outlined below.

Refereed journal articles included in the series of papers

1. Agarwal, V., Ho, D., Ho, D., Galabura, Y., Yasin, F. M.D., Gong, P., Ye, W.,

Singh, R., Munshi, A., Saunders, M., Woodward, R. C., St. Pierre, T., Wood, F.M.,

Fear, M., Lorenser, D., Sampson, D. D., Zdyrko, B., Smith, N.M., Luzinov, I., Iyer,

K.S., A Functional Reactive Polymer Nanofiber Matrix, RSC Advances (Submitted)

VA and DH developed the initial concept of PGMA electrospinning. VA

collaborated with WY and DH to synthesize upconverting nanoparticles, AM to

synthesize palladium nanoparticles and acquired TEM images on nanoparticles, RH

to synthesize magnetite nanoparticles and acquired TEM images on fibers, PG to

perform NIR Room Temperature Emission Spectroscopy, FMDY to perform

hydrogen gas sensing analysis and RCW to perform squid analysis. YG, BZ, IL

provided the PGMA; remaining authors supervised the work. Contribution by VA:

75%

2. Agarwal, V., Toshniwal, P., Smith, N. E., Smith, N. M., Li, B., Clemons, T.

D., Byrne, L. T., Hassiotou, F., Wood, F. M., Fear, M., Corry, B., and Iyer, K. S.,

Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor

Inhibitors by Intracellular Delivery, Angewandte Chemie International Edition

(Submitted)

xii

PT and TDC carried out protein expression studies. NES and BC performed docking

experiments; rest all the authors supervised the work. Contribution by VA: 90%

3. Agarwal, V., Wood, F. M., Fear, M., and Iyer, K. S., Inhibiting the activation

of transforming growth factor-β using a polymeric nanofiber scaffold, Nanoscale

(submitted)

FMW, MF and KSI supervised the work. Contribution by VA: 90%

Refereed journal articles included in the appendix

1. Agarwal, V., Tjandra, E.S., Iyer, K. S., Humfrey, B., Fear, M., Wood, F. W.,

Dunlop, S. and Raston, C. L., Evaluating the effects of nacre on human skin and scar

cells in culture, Toxicology Research, 3, 223-227 (2014)

EST performed the viability and reactive oxygen species assay on HaCaTs; BH

provided the pearl shells; remaining authors supervised the work. Contribution by

VA: 90%

2. Eroglu, E., Chen, X., Bradshaw, M., Agarwal, V. , Zou, J., Stewart, S.G.,

Duan, X., Lamb, R.N., Smith, S.M., Raston, C. and Iyer, K. S., Biogenic production

of palladium nanocrystals using microalgae and their immobilization on chitosan

nanofibers for catalytic applications, RSC Advances, 3, 1009-1012 (2013)

VA performed and optimised protocol for electrospinning, performed

characterisation and contributed to the manuscript. Contribution by VA: 25%

3. Eroglu, E., Agarwal, V., Bradshaw, M., Chen, X., Smith, S. M., Raston, C.

and Iyer, K. S., Nitrate removal from liquid effluents using microalgae immobilized

on chitosan nanofiber mats, Green chemistry, 14 (10), 2682-2685 (2012)

VA performed and optimised protocol for electrospinning, performed

characterisation and contributed to the manuscript. Contribution by VA: 40%

xiii

Conference presentations

1. Agarwal, V., Schilter, H., Ho, D., Guo, L., Hassiotou, F., Jarolimek, W.,

Wood, F. M., Fear, M., Iyer, K.S., An Innovative Tissue Engineering Approach

Towards Scar Therapy, Fifth International NanoMedicine Conference, Sydney,

Australia, 30 June- 2 July 2014 (Oral presentation)

2. Agarwal, V., Ho, D., Schilter, H., Guo, L., Hassiotou, F., Jarolimek, W.,

Wood, F. M., Fear, M., Iyer, K.S., Novel Scaffold Approach Towards Scar Therapy,

International Conference on Nanotechnology and Nanoscience (ICONN2014),

Adelaide, Australia 2014 (Poster presentation)

3. Agarwal, V., Schilter, H., Guo, L., Jarolimek, W., Wood, F. M., Fear, M.,

Iyer, K.S., Scarless wound healing: A new approach, Sixth International Conference

on Advanced Materials and Nanotechnology (AMN-6), Auckland, New Zealand, 11-

15 February 2013 (Poster presentation)

4. Agarwal, V., Wood, F. M., Fear, M., Iyer, K.S., 6. Development of scarless

wound healing platforms, School of Chemistry and Biochemistry Research Forum,

University of Western Australia, Perth, Australia, 2 November 2012 (Poster

presentation award)

5. Agarwal, V., Wood, F. M., Fear, M., Iyer, K.S., Development of Hybrid

Hydrogel for Scarless Wound Healing, First International Conference on BioNano

Innovation (ICBNI 2012), Brisbane, Australia, 18-20 July 2012 (Oral presentation)

6. Agarwal, V., Wood, F. M., Fear, M., Iyer, K.S., Regulating the Migratory

Behaviour of Fibroblasts on a Hydrogel Scaffolds, International Conference on

xiv

Nanotechnology and Nanoscience (ICONN2012), Perth, Australia, 5-9 February

2012 (Oral presentation)

7. Agarwal, V., Eroglu, E., W., Wood, F. M., Fear, M., Iyer, K.S., In vitro

evaluation of electrospun pluronic F-127 dimethacrylate copolymer towards wound

healing in burn injuries, Australian Nanotechnology Network Early Career

Symposium, Sydney, Australia, 21–22 November 2011 (Oral presentation)

Vipul Agarwal Prof Swaminathan Iyer

Candidate Co-ordinating Supervisor

1

Chapter 1

Introduction and Literature Review

Each year 100 million people develop scars in the developed world alone,1 with

causes ranging from elective surgery to severe trauma. Scars result not only in

aesthetic deficits but also long-term functional and psychological problems.2

According to WHO estimates, 6 million people suffer from burn injury each year of

which over 300,000 succumb to their injuries.3 Further, chronic skin ulcers

contribute an additional 6 million patients annually.4 In attempts to address the

functional and aesthetic deficits, the field of tissue engineering has seen a surge of

interest and applications in recent years. However, current clinical treatment of

scarring still centres on a surgical approach. Small molecules such as mannose-6-

phosphate (and many other biological factors) have been shown to have significant

potential to promote healing and reduce scarring both in elective surgery and after

injury. However, the potential of these factors to deliver improvements in the clinic

has been significantly hampered by issues of stability and delivery. This thesis will

describe the preparation of a multifunctional universal nanofibrous scaffold with the

potential to be used as a delivery vehicle for wound repair. Subsequent work will

assess the potential of more stable mannose-6-phosphate analogues which may

overcome some of the limitations of the endogenous ligand that has previously been

trialled. Finally, a combinatorial approach will be investigated using the

scaffold/analogue combined.

In this chapter the literature detailing the application of tissue engineering and the

key considerations for tissue engineering approaches will be discussed. The

pathophysiology of skin injury and wound healing will also be reviewed. Finally, the

cell therapy and tissue engineering approaches to skin repair will be discussed and

the concept of multifunctional scaffolds, electrospinning and their potential for drug

delivery and enhanced healing will be introduced.

2

1.1 Tissue engineering

Tissue engineering (TE) has evolved as an interdisciplinary science combining

principles from nature and integrating them with material and engineering science

with the goal of developing functional substitutes for damaged tissues and organs.5

The materials serve as temporary scaffolds and promote the organisation of cells to

form a functional tissue with the cells providing a source for repopulation.6 The

general strategy behind TE is to seed cells within a scaffold that mimics the

architecture of the replacement tissue while providing environmental cues that

promote tissue regeneration. Tissue engineered skin equivalents have been in

clinical use since 1997.7

In tissue engineering, a scaffold is usually required to provide a platform or niche

that promotes the desired behaviour of cells (for example proliferation of epidermal

cells to promote wound coverage).8 The scaffolds are intended to replicate or

enhance the natural ECM environment in order to retain cell viability and control

cellular behaviour. For example, demineralized bone matrix (DBM, bone from

which mineral and cells have been removed, leaving only proteinaceous material)

have been implanted in the muscle to initiate bone formation in the neighbouring

muscle tissue. This led to the commercial production of recombinant DBM from

cadavers for implantation in bone defects.7

There are three important criteria to be considered for the development of tissue

engineering scaffolds: i) the 3-dimensional micro-structure of the scaffold such as

inter-connectivity of the network pores and subsequent porosity allowing the

exchange of gases, nutrients and waste and facilitating cell adhesion, spreading and

tissue formation, ii) mechanical parameters catering to specific tissue type such as

scaffold morphology like linearity, plasticity, flexibility or anisotropy, and iii)

successful delivery of cells, drugs, growth factors and/or cytokines.8

3

1.2 Cellular scaffolds

Cellular scaffolds have been developed using natural and synthetic materials and

their composites, harnessing the specific advantages of each component and

amalgamating them to achieve the desired application. Usually, natural materials are

based on purified ECM components (eg. collagen, gelatin, laminin, hyaluronic

acid)9-11, often containing multiple components to create a composite substrate

(Matrigel®, Integra®).12 Other natural materials are also commonly used, mainly

derived from plant or animal sources including silk, agarose and chitosan.13 The

main advantages associated with the use of natural materials are their biological

activity and biocompatibility. In contrast, synthetic materials are employed to

overcome the shortfalls experienced in the use of natural materials, for example

manufacturing and process variability and an inability to control their physico-

chemical properties.8

Single component scaffolds lack the complexity and often functionality of

biomaterial scaffolds, such as mechanical properties, electrical activity or cues for

cell-matrix interactions14.8 Composite materials are therefore generally preferred.

For example, bone is made up of collagen and ceramic like hydroxyapatite, based on

which polymer-ceramic composites have been widely used in bone tissue

engineering15-19.8 Hydrogels have also been explored as they can imitate the cross-

linked architecture of ECM through their cross-linked network of monomers,

oligomers or polymers.20 However, the major limitation in their use is their lack of

mechanical strength necessary for tissue engineering applications.8 Gelatin

methacrylate hydrogels for example are known to promote cell adhesion and

spreading but lack the required mechanical strength.10 Carbon nanotubes have been

used to reinforce the mechanical properties of gelatin methacrylate by significantly

enhancing the compressive modulus, while retaining the porosity and cell

adhesiveness21.8

In designing a scaffold, one of the most desirable features is biodegradability; they

must gradually degrade over time and simultaneously get substituted by the naturally

deposited ECM and newly formed tissue.8 Consequently, linear aliphatic polyesters

such as poly(lactic acid), poly(glycolic acid) and poly(ε-caprolactone) have been

4

extensively investigated owing to their biodegradability, achieved through the

hydrolysis of their ester bonds and the ability to control their degradation rate.8

Scaffold structure is another important feature that needs to be considered during

fabrication of cellular scaffolds. Initially the emphasis was on macroporous

structures to facilitate mass transfer of vital molecules, developed using

microspheres, salt leaches or gas foams.8, 22 Such micron-scale scaffolds however do

not recapitulate the nano-scale dimensions of the ECM.23 In order to generate such

nanometer-scale dimensions techniques such as electrospinning,24 molecular self-

assembly,25 soft lithography,26 and phase separation27 have been employed.8

1.3 Physical features of the cell microenvironment

The important physical properties to consider towards the cell’s microenvironment

are substrate mechanics and surface topography.8 Mechanical properties of the tissue

are dependent on their anatomical location. For example, the elastic modulus of the

soft tissue of the brain (0.5 kPa) is low compared to intermediate muscles and skin

(~ 10 kPa), and hard precalcified bone (>30 kPa).8, 28 One of the major limitations in

recreating the native cellular environment is to reproduce the intricate physical

features found in tissues which are viscoelastic with non-linear, anisotropic and

heterogeneous mechanical properties. It is an important consideration while

designing scaffolds because the optimal scaffold should completely replicate the

replacement tissue.8 For example, the biological nanostructure of the heart

influences its biochemical, electrical and mechanical functions and its ECM

comprised of dense elastic fibrillar collagen, elastin bundles and proteoglycans.5, 29 It

promotes mechanical coupling between cardiomyocytes resulting in aligned and

anisotropic cell bundles promoting not only intercellular interactions but also with

neighbouring capillaries and nerves.5 This bundled elongated cellular structure is

critical for the function of the cardiac syncytium, enabling the muscles with

exclusive rhythmic contraction and mechanical and electrical properties to facilitate

the blood pumping function of the heart.5 Cardiomyocytes have been shown to adopt

random morphology on flat surfaces as they lose their elongated morphology post

5

isolation. This is most likely because they lose the cues required to instruct their

formation and interactions with their environment.5 Strategies adopted to address

this problem of cell alignment in engineered cardiac tissue include applying

mechanical stretching,30 interstitial fluid flow, electrical stimulation31 and

microcontact printing,32 with limited success.5 Alternatively, a model cadiac tissue

was developed by controlling the nanotopography of the scaffold that mimics the

ECM matrix of the native myocardial tissue.5, 33 Polyethylene glycol hydrogel was

fabricated to yield grooved arrays with ridge widths ranging from 15 to 800 nm

which when cultured with cardiac cells resulted in their self-assembly aligned to the

direction of the scaffold.5 It resulted in elongation of cadiac cells to form

anisotrophic cell array which could result in rhythmic contraction necessary for

normal heart function.33 Despite the progress made, the challenge of patterning such

nano-arrays within 3D scaffolds towards anisotropically aligned tissues still

remains.5

1.4 Design concepts and strategies

Three-dimensional (3D) scaffold constructs have more demanding requirements for

efficient cell motility compared to the 2-D monolayer cultures to achieve tissue

uniformity and to avoid heterogeneous tissue growth.34 The supply of oxygen and

other nutrients along with waste removal and cell motility reply on mass transfer

properties of the scaffold.35 Considering the primary mechanism behind mass

transport is diffusion, scaffolds need to be designed after carefully considering its

diffusion characteristics.36 Cells in vivo reside within 100 µm of a capillary for

efficient nutrient supply.35 Neotissue growth on engineering scaffolds has been

reported to be preferentially limited to the peripheral regions (100-200 µm) of the

scaffold because of the limited oxygen supply.35, 37, 38 This is still a major limitation

in the engineering of thick 3D tissues because it limits the cellular distribution and

bioavailability of nutrients especially in the interiors of the scaffold as the cells

preferentially localise to the periphery. 35Therefore scaffolds for tissues with low

metabolic activity have been easier to synthesize. For example, cartilage, an

6

avascular tissue known to have low metabolic activity, has been successfully

fabricated into a thick 3D construct.39 Similarly, other structural tissues known for

their mechanical function such as blood vessels, heart valves, ligaments, tendons and

skin may have lower demand for high oxygen and nutrient transport and therefore be

easier to replicate.35, 40

1.4.1 Physical properties

One of the main features required facilitating nutrient and oxygen transport is the

pore size of the engineered construct. Pore size is critical as it influences tissue

ingrowth and cellular adhesion especially in the internal surface area of the matrix.

For example, small pores are impeded by the growing tissue preventing tissue

ingrowth and ECM production. Alternatively, big pores lack cellular recognition of

the surface topography preventing neovascularisation.35 Although the pore size is

susceptible to change in an in vivo environment,35, 39 it has been demonstrated that

optimal size varies with differing cell types and their respective architecture. For

example, where 5 µm is optimal for neovascularisation, fibroblasts ingrowth

demands 5-15 µm whereas hepatocytes require 20 µm, keratinocytes need 20-125

µm, and 100-700 µm size pores are required for bone regeneration.35, 39, 41, 42 Ma et

al. studied the effects of pore size in a 3-D polyethylene terephthalate nonwoven

fibrous matrix on long-term tissue development of human trophoblast ED27 cells in

terms of cell morphology and spatial organisation.43 They showed that human

trophoblast ED27 cells with average diameter of 14 µm were not able to bridge the

gap between fibres with a pore diameter of 20 µm. Instead they formed large

aggregates and started to differentiate in larger pores. On the contrary, cells rapidly

bridged the gap on the matrix with low porosity, a pore diameter of around 15 µm

and promoted cell spreading and proliferation. Further, the cell differentiation was

inhibited, highlighting the importance of pore size on cellular behaviour and

response, in terms of cell morphology, differentiation and proliferation, towards

scaffold design features. In yet another similar study, canine microvascular epithelial

cells formed a thin endothelial lining only when grown on scaffolds with average

pore size of 90 µm, whereas vascular smooth muscle cells preferred the pore size of

7

107 µm in order to form uniform tissue.35 Dermal fibroblasts, on the other hand, did

not respond to any pore size.35, 44 Similarly, higher pore size greater than 250 µm has

been shown to promote angiogenesis with faster cellular ingrowth when compared to

smaller pore sizes.35 Despite the elaborate evidence on the impact of pore size and

their distribution on angiogenesis, inflammation and tissue infiltration in vivo, the

optimal pore size and range is still fairly unpredictable, especially in the context of

regenerating complex tissues with multiple different cell types45, 46.35

In addition to pore size, cell transport mechanisms like diffusion, attachment and

migration is dependent on total porosity, pore interconnectivity and scaffold surface

area.35, 41 In order to achieve uniform cellular penetration and ingrowth, scaffolds are

required to be highly porous with open interconnected geometry and a large surface

area: volume ratio for example, scaffolds with over 90% porosity have been shown

to have optimal surface area for cellular-matrix interactions warranted for efficient

diffusion.35, 41, 47, 48 Highly porous scaffolds, however, lack the mechanical integrity

required for tissue regeneration.36, 49, 50 Therefore, a fine balance between porosity

and mechanical strength need to be established for the development of an optimal

replacement scaffold.35

1.4.2 Mechanical properties

Studies determining the mechanical properties of various tissue types revealed that

most tissues are heterogeneous, viscoelastic, nonlinear and anisotropic materials51.35

Scaffolds are required to be mechanically strong to carry out its function while

withstanding the hydrostatic or pulsatile pressures encountered in vivo.35, 41, 42 The

bulk properties of the constituents have been used to predict the mechanical

properties of the designed scaffolds. In addition, there has also been a direct

correlation between scaffold architecture and structural features such as pore

distribution, fibre diameter and orientation and mechanical strength of the

scaffold52.35 This is of particular interest in the case of scaffolds with high total

porosity and low material content, for example in hydrogels and electrospun

scaffolds. It has also been demonstrated that the cellular response, such as cell

contractility, motility, adhesion, spreading and differentiation, can be influenced by

8

the physical properties of the scaffold like softness and hardness53-55.35 Engler et al.

reported that matrix properties such as softness and elasticity has a profound

influence on stem cell lineage and their commitment to a particular phenotype.55 It

was shown that human mesenchymal stem cells display a more neurogenic

phenotype on soft substrates (0.1-1 kPa), a myogenic phenotype on moderately hard

matrices (8-17 kPa) and an osteogenic phenotype on hard substrates (25-40 kPa).35

In another study, soft substrates with Young’s moduli less than 1-1.5 kPa have been

reported to enhance the differentiation of neural progenitor cells.56 This property of

the substrates can be exploited to achieve tailored cell proliferation and

differentiation56.35

1.4.3 Surface properties

Cellular adhesion and growth has been shown to be correlated to the surface

properties of scaffolds including topology and chemical characteristics.57, 58

Chemical properties correspond to the cellular interactions including adhesion and

protein interactions at the material surface. Tamada and Ikada investigated 13

different polymeric surfaces for fibroblast adhesion, growth and collagen synthesis,

with a range of different surface energies using goniometer (contact angle

measurements).59 It was concluded that proliferation was independent of surface

chemistry, whereas cell adhesion which relies on cell protein-substrate interactions

was correlated to surface wettability. A positive correlation was observed between

the surface charge and density of adsorbed proteins which corresponded to better

cellular adhesion.35, 60 Furthermore, cellular behaviour was altered corresponding to

the surface topography from angular edges, abrupt grooves or other surface

indentations to smooth surfaces.61-64 A number of studies have reported the

directional motility of cells along the fibres and ridges fabricated on the substrate

surface which is explained by a phenomenon called “contact guidance”.35, 65

As highlighted above, an ideal scaffold is an amalgamation of chemical, mechanical

and structural properties. It should be biodegradable but structurally strong, should

be porous but also flexible. Therefore, an ideal scaffold should be multifunctional,

incorporating cues for cellular processes like adhesion, extracellular matrix

9

production and migration, but also have potential applications such as drug delivery

and multimodal imaging capabilities.

1.5 Skin, injury and wound healing

Skin is the largest organ on the human body. Adult skin consists of two main tissue

layers: a thin top layer called epidermis which is mainly composed of keratinocytes

while the lower epidermal layer is made up of melanocytes, cells responsible for

skin pigmentation and an underlying supporting layer called dermis which is mainly

composed of collagen synthesising cells, fibroblasts (Figure 1).66, 67 The main role of

epidermis is protection against microbes and to help regulate body temperature.68

The layer sandwiching these two layers is a 20 nm thick membrane called basement

membrane. It is made up of hemidesmosomal structures and its main role is to

mechanically stabilize the interaction between the epidermis and dermis.67

10

Figure 1: This diagram of human skin shows the two main layers of skin — the upper epidermal barrier layer

and the lower, much thicker, dermis. Figure and caption taken from MacNeil 2007.69

The main role of dermis which is 2-5 mm thick is to provide support to the

overlaying epidermis.67 It provides considerable tensile strength and elasticity to

skin mediated by the intertwined arrangement of collagen fibers, has specialized

components and structures.67 Collagenous architecture is composed of varying

amounts of interwoven elastin fibers, proteoglycans, fibronectin and other

components.67, 68 To avoid infection, injury to the skin demands rapid intervention

where the healing time is dependent on extent and depth of the injury. Epidermal

lesions mostly heal within a week.66 Wound healing is a dynamic and interactive

process involving synergistic interactions between blood cells, ECM, and

parenchymal cells.70 Wound healing is accomplished by three successive but

overlapping stages: inflammation, reepithelialisation and remodelling (Table 1).

Failure or delays in healing can lead to significant scarring and potentially fatal

sepsis and therefore interventions to promote healing are critical to reduce morbidity

and mortality associated with extensive skin injuries.

11

Table 1: Phases of wound healing, major types of cells involved in each phase, and selected specific events.

Taken from Falanga 2005.71

1.5.1 Haemostasis and inflammation

An insult to the skin causes the disruption of blood vessels and extravasation of

blood constituents. As the blood clots it re-establishes haemostasis and provides a

provisional ECM (Figure 2).70 As the platelets aggregate they not only facilitate the

formation of a hemostatic fibrin mesh but also secrete several mediators of wound

healing, such as platelet derived growth factor, a chemoattractant that recruits and

activate macrophages and fibroblasts.70 Coagulation of platelets generates numerous

vasoactive mediators and chemotactic factors and activates complementary

pathways by injured and activated parenchymal cells. This then results in the

recruitment of inflammatory leukocytes to the injury site.70

12

Figure 2: A Cutaneous Wound Three Days after Injury. Growth factors thought to be necessary for cell

movement into the wound are shown. TGF-β1, TGF-β2, and TGF-β3 denote transforming growth factor β1, β2,

and β3, respectively; TGF-α transforming growth factor α; FGF fibroblast growth factor; VEGF vascular

endothelial growth factor; PDGF, PDGF AB, and PDGF BB platelet-derived growth factor, platelet-derived

growth factor AB, and platelet-derived growth factor BB, respectively; IGF insulin-like growth factor; and KGF

keratinocyte growth factor. Reproduced with permission from A. J. Singer and R. A. F. Clark, N. Engl. J. Med.,

1999, 341, 738-746. Copyright Massachusetts Medical Society.

Next, neutrophils infiltrate the wounded area to initiate cleaning the injury site of

foreign particles and bacteria, which are then extruded with the eschar or

phagocytosed by macrophages.70 Chemoattractants such as fragments of ECM

proteins, transforming growth factor β (TGFβ) and monocyte chemoattractant

protein 1 (MCP1), attract monocytes to the injury site. These monocytes become

activated macrophages that release growth factors and additional cytokines to initiate

the formation of granulation tissue.70 Interaction of macrophages with the ECM

through their integrin receptors, stimulate them to phagocytose microorganisms and

fragments of ECM and simultaneously, stimulate monocytes to differentiate into

pro-inflammatory or reparative macrophages 72.70 This interaction also stimulates

both monocytes and macrophages to express colony-stimulating factor 1, a pro-

survival cytokine; tumor necrosis factor α, an inflammatory cytokine; and platelet-

13

derived growth factor, an important chemoattractant and mitogen for fibroblasts.70

Other key cytokines secreted by monocytes and macrophages include transforming

growth factor α, interleukin-1, transforming growth factor β, and insulin like growth

factor 173.70 Macrophages are an integral part of the wound healing process critical

to wound repair, with macrophage deficient animals shown to have markedly

defective wound healing.74, 75

Figure 3: Schematic diagram of wound reepithelialization models. (a) Basal KCs at the leading edge of the

wound that are firmly attached to surrounding basal and suprabasal KCs actively migrate to close the wound in

the “sliding” model of reepithelialization. Arrow indicates direction of movement of basal KC. (b) With the

basal KCs firmly attached to the BMZ, suprabasal KCs roll onto the wound matrix in the “rolling” model of

reepithelialization. Arrow indicates movement of suprabasal KCs as they tumble over basal KCs. Arrowheads

indicate initial cut edge. Figure and caption taken from Usui 2005.76

1.5.2 Reepithelialisation

Two mechanisms behind reepithelialisation have been proposed in the literature: the

“rolling” model and the “sliding” model (Figure 3). The “rolling” model postulates

that basal keratinocytes remain strongly attached to the basement membrane, while

suprabasal keratinocytes at the wound margin are activated to roll over into the

wound site77, 78.76 The “sliding” mechanism on the other hand, postulates that basal

keratinocytes are the principal cells mediating the migration and wound closure.

Both basal and suprabasal keratinocytes remain strongly attached to the leading edge

and basal keratinocytes are passively dragged along as a sheet79, 80.76 A variant is

also proposed where a large population of suprabasal cells migrate out of the wound

(Figure 4).76

14

Figure 4: Schematic diagram of a new model of reepithelialization. Both basal and suprabasal KCs are

activated to respond to the complex microenvironment created by a wound. Suprabasal KCs undergo

dramatic changes in response to injury and are possibly the primary source of cells available for wound closure.

Figure and caption taken from Usui 2005.76

Reepithelialisation is initiated spontaneously after injury.70 Epidermal cells from the

skin appendages initiate the cleaning process by removing the clotted blood and

damaged stroma.70 Simultaneously, cells alter their phenotype which includes

retraction of intracellular tonofilaments;78 dissolution of most intercellular

desmosomes, important for physical connections between the cells; and formation of

peripheral cytoplasmic actin filaments, crucial for cell motility81, 82.70 Furthermore,

the loss of basement membrane and hemidesmosomal links disrupts the interactions

between juxtaposed epidermal and dermal cells, allowing the lateral movement of

epidermal cells.70 Epidermal cells interact with a variety of ECM proteins via their

integrin receptors. These include proteins such as fibronectin and vitronectin which

are interspersed with stromal collagen I at the wound margin and interwoven with

the fibrin clot in the wound space83-85.70 The epidermal cells start to migrate and

dissect the wound to separate desiccated eschar from viable cells.70 Collagenase as

well as the activation of plasmin by plasminogen activator is produced by migrating

epidermal cells in order to degrade the ECM required for their migration between

the collagenous dermis and the fibrin eschar86.70 Plasminogen activator is also

15

known to activate collagenase (matrix metalloproteinase I) which propels the

degradation of collagen and ECM proteins.

Figure 5: A Cutaneous Wound Five Days after Injury. Blood vessels are seen sprouting into the fibrin clot as

epidermal cells resurface the wound. Proteinases thought to be necessary for cell movement are shown. The

abbreviation u-PA denotes urokinase-type plasminogen activator; MMP-1, 2, 3, and 13 matrix

metalloproteinases 1, 2, 3, and 13 (collagenase 1, gelatinase A, stromelysin 1, and collagenase 3, respectively);

and t-PA tissue plasminogen activator. Reproduced with permission from A. J. Singer and R. A. F. Clark, N.

Engl. J. Med., 1999, 341, 738-746. Copyright Massachusetts Medical Society.

Epidermal cell proliferation begin few days after injury.70 Local release of growth

factors, such as epidermal growth factor, transforming growth factor α and

keratinocyte growth factor;87 and absence of the neighbouring cells on the wound

margin are thought to be the principal reasons behind this epidermal response.70 The

reepithelialisation results in the synthesis of basement membrane proteins laid down

in an orderly sequence from the margins of the wound inwards, mimicking the

original architecture in a zipper like fashion88.70 Epidermal cells return to their

normal phenotype, establishing juxtaposition to the re-established basement

membrane and underlying dermis once the epidermal barrier has been restored.70

This also acts as a trigger for next step of inflammation and matrix remodelling.

16

1.5.3 Formation of granulation tissue

Four days post injury, granulation tissue starts to invade the wound space (Figure 5).

Macrophages, fibroblasts and blood vessels simultaneously start to infiltrate into the

wound space.70 Angiogenesis and fibroplasia is constantly supplemented by the

growth factor secreted by macrophages; while fibroblasts lay down ECM necessary

to support cell ingrowth and blood vessels carry oxygen and nutrients crucial to

sustain cell metabolism.70 Fibroblasts, nurtured by growth factors, especially

platelet-derived growth factor and transforming growth factor β1 (TGFβ1), in

addition to the ECM proteins, begin to proliferate and start to express integrin

receptors for their migration into the wound.70 Platelet derived growth factor, along

with many other similar factors, have been shown to accelerate healing of chronic

pressure sores and diabetic ulcers when applied exogenously to these wounds.89, 90

Newly synthesised provisional ECM promotes granulation tissue formation by

providing the conduit for cell migration88.70 It has been postulated that the

fibronectin architecture and fibronectin specific integrin receptors on fibroblasts

control the rate of granulation tissue formation.70 These fibroblasts are responsible

for the synthesis, deposition, and remodelling of the ECM. Cell migration into the

blood clot of crosslinked fibrin, or tightly woven ECM requires fibroblast-derived

enzyme proteases like plasminogen activator, collagenases, gelatinase A, and

stromelysin that can cleave the matrix to facilitate their migration91.70 Post migration

into the wound, fibroblasts initiates ECM synthesis.92, 93 It is believed that

transforming growth factor β1 stimulates the replacement of the provisional ECM

with a collagenous scar-type matrix92, 93.70 Once an abundant collagen matrix has

been deposited, fibroblasts stop their excessive collagen production triggering the

replacement of granulation tissue by a relatively acellular scar.70

17

1.6 Role of TGFβ in fibrosis and scarring

The TGFβ superfamily contains over 30 proteins found in both vertebrates and

invertebrates and encompassing a wide range of functions throughout the lifetime of

the animal.94 The three mammalian TGFβ gene isoforms are TGFβ1, 2 and 3 which

share 60-80% sequence homology and are encoded by three different genes.95, 96

The three mammalian isoforms of TGFβ are expressed in a cell specific and

developmentally regulated manner.97, 98 For example, TGFβ1 and TGFβ3 are

expressed during the morphogenesis stage of early development followed by TGFβ2

which is expressed later in mature and differentiating epithelium.98 All three

isoforms are highly conserved in mammals, but differ in their binding affinity

towards TGFβ receptors.98, 99 Altered and different phenotypes were resulted as a

result of the deletion of individual isoforms in mice.98-100 TGFβ1-null mice were

mostly embryonically lethal due to abnormal development of the yolk sac, and those

that did survive developed multi-organ autoimmunity and multi-focal inflammation

disease and died within 3 weeks of birth101.102 TGFβ2-null mice showed perinatal

mortality due to cyanotic heart disease and various other developmental

defects103.102 TGFβ3-null mice die due to cranial bone defects especially cleft

palate104.102

Figure 6: Schematic representation of latent TGF-β. The putative transglutaminase-mediated covalent bonding

between the N-terminus of LTBP and the ECM is indicated by a question mark (?). Disulfide bonds between

LAP monomers, between TGF-β monomers and between LAP and LTBP are indicted by thin lines. Carbohydrate

residues on LTBP are not shown. Abbreviations are in the text. Figure reproduced, and caption taken from

Munger 1997.105

18

Newly synthesised TGFβ is secreted in an inactive latent form. The non-covalent

association of TGFβ with a latency associated peptide (LAP) is the reason behind

the latency (Figure 6). LAP is a homodimer formed from the propeptide region of

TGFβ.105 This association of LAP and TGFβ is termed as latent TGFβ (LTGFβ). To

exert its function, TGFβ needs to be cleaved from its latent complex with LAP.105

Two of the N-linked carbohydrate residues contain mannose-6-phosphate (M6P)

groups. There are 3 cysteines in LAP-1 which form interchain disulfide bonds to

dimerise LAP monomers.105 This cysteine interaction is critical to maintain latency

because their replacement with serines at positions 223 and 225 of LAP-1results in

the secretion of active TGFβ1.106 LAP is also known to form a disulfide linkage to

another protein called latent TGFβ binding protein (LTBP).105 LTBP link latent

TGFβ to the ECM.

TGFβ has three high-affinity membrane receptors where type I and II are

transmembrane serine/threonine kinases that coordinate to facilitate each other’s

signalling,98, 107-109 while type III receptor, a transmembrane proteoglycan, has no

specific signalling function because of its highly conserved cytoplasmic domain.98,

110, 111 TGFβ interact with these different receptors to mediate specific functions. For

example, interaction with receptor type I promotes secretion and deposition of ECM,

while interaction with type II receptor mediates cell growth and proliferation. In

order to activate the type II receptor, TGFβ binds either to a type III receptor which

then facilitates its binding to type II receptor or it can bind directly to a type II

receptor.112 Upon activation by TGFβ, type II receptors recruit, bind and

transphosphorylate type I receptors, thereby promoting their protein kinase activity

and initiating its downstream signalling cascade.112 TGFβ signalling within the cells

is mediated by the Smad family of transcriptional activators and has been

extensively reviewed elsewhere.113-118

19

1.7 Mechanisms of activation of extracellular TGFβ

Downstream signalling of TGFβ is regulated by local activation of the LTGFβ

complex in vivo.102 There are various mechanisms of LTGFβ activation which have

been widely studied.2, 102, 119 M6P/IGFII receptor mediated activation of LTGFβ has

been illustrated as one of the most critical mechanism of activation.

One mechanism of activation involves integrins. Integrins are transmembrane

receptors that conjugate the cell cytoskeleton to the ECM and are critical for cell

adhesion, proliferation, migration and differentiation120-122.102 Integrins are made up

of α and β subunits.102 αv integrins recognise and bind to both TGFβ1 and LAP via

their RGD sequence. This RGD mediated interaction between αvβ6 integrin and

LAP-1 cause the activation of LTGFβ1 by altering its conformation.102 However, β6

integrin is required to interact with the actin cytoskeleton of the cell to ascertain

LTGFβ1 activation.102, 123 This alteration in the LTGFβ1 structural conformation

produces mature TGFβ which can interact with the TGFβ II receptor and therefore

allow it to mediate its function124.102

Thrombospondin-1 (TSP-1), a 300 kDa protein found in the α-granules in platelets

and ECM has also been shown to activate both small LTGFβ (LAP-TGF-β complex)

and large LTGFβ (LTBP-LAP-TGF-β) complexes in a non-proteolytic mechanism

both in vitro and in vivo125-127.102 It has been postulated that the TSP-1 interaction

with LAP induces a conformational change in relation to mature TGFβ thereby

unveiling it to the TGFβ receptor recognition site and facilitating TGFβ binding to

its receptor128.102

The proteases such as plasmin, matrix metalloproteinase (MMP) 2 and 9 have also

been reported to induce LTGFβ activation in vitro129, 130.102 Plasmin conditioned

media has also been shown to generate active TGFβ in fibroblasts and Chinese

hamster ovary cultures131.102 Alternatively, the activation of LTGFβ can be inhibited

in vitro by plasmin inhibitors or prolonged by neutralising antibody to plasminogen

activator inhibitor-1 as demonstrated in co-culture studies of endothelial cells and

pericytes or smooth muscle cells132.102

20

Another mode of in vivo activation is via the binding of the M6P residue present on

the latency associated peptide (LAP) to the M6P/IGFII receptor. This invokes a

conformational change in TGFβ bound to M6P/IGFII receptor, thus allowing the

proteolytic cleavage of active TGFβ from its latent complex. In vitro activation of

TGFβ has been investigated in detail in activated macrophage cultures and in co-

cultures of endothelial and smooth muscle cells where latent TGFβ is activated by a

complex process involving recognition and binding of the latent form to the

M6P/IGFII receptor and the concerted action of both transglutaminase and serine

protease plasminogen/plasmin.133, 134

Once activated TGFβ suppresses the inflammatory response and promotes the

formation of granulation tissue. TGFβ promotes keratinocytes migration by

upregulating fibronectin expression and various integrins necessary for keratinocyte

adhesion, whilst simultaneously inhibiting keratinocyte proliferation thereby

regulating reepithelialisation stage of wound healing.135 It has been reported that

overexpression of TGFβ1 in the epidermis results in delayed re-epithelialisation

response in transgenic mice model.136-138 Despite having an inhibitory effect on the

proliferation of endothelial cells, TGFβ has been reported to promote angiogenesis

in vivo.139, 140

In the dermis, TGFβ activates fibroblasts which then produce higher levels of matrix

molecules including collagen, fibronectin and glycosaminoglycans (GAG), matrix-

degrading proteases and protease inhibitors thereby resulting in an increase in the

production of matrix proteins and decreasing their proteolysis.141 Whilst these

changes are critical for rapid wound repair, excessive and/or extended matrix

deposition and excess TGFβ stimulation can lead to poor healing and scar outcomes.

1.8 Exogenous mannose-6-phosphate act as an antagonist of

TGFβ1 activation

Purchio et al., used radiolabeled [32P] glycopeptide and demonstrated the presence

of endogenous M6P as a carbohydrate unit on LAP (TGFβ precursor).142 Binding

21

studies demonstrated that TGFβ precursor has high binding affinity towards purified

M6P/IGFII receptor which can be hindered by exogenous M6P.142Miyazono and

Heldin studied the importance of carbohydrate in TGFβ1 activation and

demonstrated that interaction between endogenous M6P present on LTGFβ1

precursor and M6P/IGFII receptor can be result in TGFβ1 activation. In a model to

study potential anti-fibrotic agents Gosiewska et al. used macrophage-based system

for TGFβ1 activation and demonstrated that exogenous M6P inhibits the activation

of LTGFβ.143 Bates et al., used a rabbit flexor tendon in vitro and in vivo models to

investigate the role of exogenous decorin and M6P in TGFβ activation. It was

concluded that both decorin and M6P inhibits the stimulatory effects of TGFβ on

collagen production while even a single low intraoperative dose of M6P

significantly improved the range of motion of the operated tendon.144 Xia et al.,

studied inhibitory potential of M6P in three different cell types; sheath fibroblasts,

epitenon tenocytes and endotenon tenocytes from rabbit flexor tendon post TGFβ

stimulation. They reported that TGFβ induced collagen production was significantly

downregulated by the addition of exogenous M6P in a dose dependent manner in all

three cell types.145 It is clear that exogenous M6P is a potent inhibitor of TGFβ

signalling in multiple systems.

22

Figure 7: Chemical structure of a) mannose-6-phosphate showing the carbon numbers, b) an example of

isosteric phosphonate analogue and c) an example of non-isosteric analogue. Image taken from Vidal 2000.146

1.9 Structural requirements for mannose-6-phosphate recognition

to M6P/IGFII receptor

Tong et al., first elucidated the structural requirements of M6P recognition by both

the cation dependent mannose-6-phosphate receptor (CD/MPR) and cation

independent mannose-6-phosphate receptor (CI/MPR, also known as M6P/IGFII

receptor).147, 148 They investigated various ligands like M6P, pentamannose

phosphate, β-galactosides and a high mannose oligosaccharide with two

phosphomonoesters, and inferred that certain structural features are determinant in

binding to M6P/IGFII receptor such as:

a) The hydroxyl group at the C2-position (C2-OH) of the pyranose ring must be

axial (Figure 7a) to allow hydrogen bonding interaction between hydroxyl group and

Gln-348 and Arg-391 amino acids located on β-strands 3 and 7 of the receptor.149

23

Similar strong interaction was observed in case of fructose-1-phosphate (F1P) while

glucose-6-phosphate, with C2-OH in equatorial position is weakly recognised (2-

epimer of M6P).150

b) In addition to the orientation, C2-OH is also inherently required as its absence

results in diminished interaction with the receptor like in the case of 2-deoxy

glucose-6-phosphate (2dG6P).147

c) The substitution at anomeric centre does not influence the interaction of the

analogue to the receptor as F1P is recognised in similar propensity as M6P.

Furthermore, replacement of the hydroxyl group at anomeric position with a para-

nitrophenoxy group, slightly improves its recognition towards the M6P/IGFII

receptor. It is due to the lipophilic interactions between the aromatic moiety of the

ligand and the binding pocket of the receptor.150

d) Substitution at 5-position of the pyranoside ring does not influence the

interaction of analogue to the receptor.147

e) The distance between the negative charge (phosphate group) and C5 of the

pyranose ring should be four atoms.146

f) Only a single negative charge is necessary for the binding to M6P/IGFII

receptor since isosteric M6-phosphonate analogues (example in Figure 7b) are very

well recognised compared to non-isosteric M6-phosphonate analogues (example in

Figure 7c) which are very weakly recognised.151

1.10 Mannose-6-phosphate analogues

Due to the importance of M6P in inhibiting TGFβ and its potential to therefore

ameliorate scarring M6P has been studied in a double blind, placebo controlled,

randomised, phase 2 efficacy clinical trial (Renovo UK, Juvidex®).152, 153 The study

was carried out in almost 200 male and female subjects administering different doses

of the drug to split thickness skin graft donor sites. Intradermal delivery of the drug

was shown to significantly accelerate the rate of wound healing. However, no

significant improvement in the extent of scar formation was observed.154 This is in

part due to the subjective measurement and difficulties in measuring scar

improvement. However, the metabolic vulnerability of M6P against phosphatases

24

has also been postulated to be a major limitation in its clinical translation. Therefore

attempts have been made to develop M6P analogues with greater stability against

enzymatic degradation.

Vidil et al., reported the synthesis of M6P analogues with the aim of developing

analogues with enhanced affinity towards the M6P/IGFII receptor.155 It was

postulated that replacing the P-O bond at C6 position with P-C bond, as is the case

in phosphonate analogues, would enhance the stability of the analogues towards

hydrolases. They showed that the isosteric analogue of M6-phosphonate binds to the

M6P/IGFII receptor with higher affinity compared to the non-isosteric derivative

(with shorter chain length (1 atom) between phosphate group and pyranose ring) and

with similar affinity to M6P.155

Berkowitz et al., reported the synthesis of three mono and bivalent ligands bearing

M6P surrogates (malonyl ether, malonate, and phosphonate) and studied their

binding affinity towards the M6P/IGFII receptor. These surrogates were hydrolase

resistant phosphates (with a methylene bridge at C6 position bridging the pyranose

ring and respective functional groups) and mimic the M6P locked in the α-

configuration.156 It was concluded that phosphonate analogues have greater binding

affinity than malonyl ether and malonate analogues.156 In another example, Jeanjean

et al., developed two sulfonate (M6S) and one unsaturated phosphonate analogue of

M6P in an attempt to study their stability in human serum and their binding affinity

towards the M6P/IGFII receptor.157 The two isosteric sulfonated analogues were

synthesised with the view that higher chemical stability of conjugated (unsaturated)

analogues compared to M6S and unconjugated (saturated) analogue would have

elevated binding efficiency to the receptor. Conjugation studies were carried out to

decipher the binding of these analogues to the M6P/IGFII receptor. Binding studies

confirmed the vulnerability of M6P in human serum which dropped 5 fold while the

binding affinity of the three new analogues remained intact.157

25

Figure 8: Chemical structure of carboxylate analogues of M6P. Image taken from Jeanjean 2006.158

In a similar study, four new carboxylate analogues of M6P were investigated for

their receptor binding affinity and serum stability (Figure 8).158 Importance of a

negative charge on the M6P analogue was emphasised by the receptor binding assay

based on the ligand dependent elution of the receptor from PMP affinity columns.159

It was demonstrated that unsaturated isosteric carboxylate analogue had similar

binding affinity as M6P while non-isosteric carboxylate analogues had slightly

weaker affinity to the receptor.158 This difference in binding efficiency was

explained by structural proximity of the double bond, which in the case of

unsaturated isosteric analogue, could stabilise the analogue and its geometry would

promote its interaction with M6P binding sites on the receptor, in comparison to the

non-isosteric (saturated) analogue. The three analogues with greater binding affinity

were also shown to have prolonged stability against hydrolases. Analogues were

shown to retain their recognition potential even after 2 days incubation in human

serum, while 6 hour incubation was shown to reduce M6P binding affinity by 32%.

Christensen et al., synthesised a series of glycopeptide derivatives of M6P,

containing two 6′-O-phosphorylated mannose disaccharides linked either α(l → 2) or

α(l → 6) and 3-5 amino acids, in order to study the influence of structural variation

in mannose disaccharides on their binding affinity towards M6P/IGFII receptor.160

The analogue comprised of two 6’-O-phosphorylated α(1→2)-linked mannose

disaccharide showed higher binding affinity compared to α(1→6)- linked

analogue,160, 161 It was concluded that analogue receptor affinity was dependent on

terminal phosphorylated mannose disaccharides and was found independent of the

26

variation in invariant saccharide core from tri-, tetra- to penta-peptides. However,

removal of aminobenzoyl group bound to the lysine tail in one of the analogue was

shown to significantly diminish their receptor binding affinity.159 This demonstrates

that structural rigidity is required to maintain strong affinity of the peptide towards

the receptor160.159

It can be concluded that there are a number of key structural requirements to be

considered in M6P analogue design. These structural constraints are important for

binding affinity and specificity and also for stability in vivo. However, despite the

design of a number of analogues to date with promising data suggesting increased

stability and binding affinity, no clinical treatment targeted at the M6P inhibition of

TGFβ signalling currently exists. Therefore it is likely that other delivery

considerations may also be important.

27

Figure 9: An illustrated representation of raised dermal scar types, as commonly observed after a mid-

sternal incision post‐cardiac surgery. (a) Fine line scar, (b) Hypertrophic scar, (c) Intermediate raised dermal

scar, (d) Keloid scar. Image and caption taken from Sidgwick 2012.162

1.11 Scarring

There are a broad range of scar outcomes found in humans. Clinically scars are

quantified using scar assessment scales.163-165 Normal scars are usually fine line

scars which can be extended with mechanical stretch applied on the scar (Figure

9a).2, 162 Hypertrophic scars are raised dermal scars which are the result of

aggressive proliferation leading to excessive healing and matrix deposition162

(Figure 9b). Contracted scars as the name suggests result in contracture of the

granulation scar tissue around the joints following burn injury or surgery162 (Figure

9c). Finally, keloid scars are similar to hypertrophic scars but crucially extend

28

beyond the boundaries of the initial wound and are commonly considered to be more

similar to a benign tumour than a normal scar (Figure 9d).166-169

1.12 Current treatments for skin wound healing

The etiology of skin damage is diverse, from genetic disorders such as epidermolysis

bullosa through to acute trauma, tumours, chronic wounds (ulcers) and even surgical

intervention.170 One of the main factors behind a major loss of skin is burn injury.171

It is estimated that each year over 300,000 people die from burn-related injuries

worldwide.172 Many more people suffer from burn-related disabilities and

disfigurements.172

With the advent and development of new aggressive surgical interventions the rate

of patient mortality due to burn related injuries has receded considerably in recent

years. This has had a profound effect on patients, with the increase in survival being

mirrored by an increase in poor extensive scar outcomes.

Treatment of skin wounds, including burn injury, is dependent on the extent of the

injury. While superficial or incisional wounds heal with little or no intervention,

deeper wounds often require clinical intervention. The current gold standard for

replacement skin is via a split-skin graft - an autograft where skin is harvested from

an uninjured donor site on the patient, meshed to cover a larger area and patched

onto the injured site of the same patient. This approach is utilised preferentially in

full thickness and deep partial thickness wounds where both epidermal and some of

the dermal layer of the skin is harvested from an uninjured site and grafted on the

damaged site.173 The major limitation to this approach is that as the surface area of

the injury increases the availability of donor tissue decreases. Additionally, tissue

from different body sites is not equivalent and this can lead to poor aesthetic

outcomes (for example pigmentation). Finally, the use of donor sites creates an

additional wound which needs to be healed and can lead to significant donor site

morbidity.173, 174 In order to address these problems it is essential to develop new

29

interventions. Synthetic skin substitutes and engineered replacements are employed

to restrict the burden of disease by reducing the need for graft donor sites.175

1.13 Cell based therapies

Suspended keratinocytes cell based therapies are emerging as a leading therapeutic

strategy, aimed at utilising cells as replacement therapy to repair severely damaged

tissues176.173 In 1975, the first subcultures of human keratinocytes was developed

and manipulated to obtain epithelial sheets for subsequent grafting177.173 Their first

major clinical application was reported in 1981 where cultured autologous epithelial

sheets were used for the treatment of extensive third degree burns178.173 The time to

culture these sheets, approximately 3-5 weeks, was a major limitation leaving

patients susceptible for prolonged periods after the injury179-181.173 Some of the other

limitations included high production costs,182, fragility, variable engraftment rates,

difficult handling, storage and preservation of viable sheets.173, 183-186 It was the

development of an automated membrane bioreactor which brought down the delay

time to 2 weeks187.173 An alternative innovative approach was later developed, to

further reduce fabrication time, where cells extracted from a skin biopsy were not

cultured but rather directly sprayed onto the lesions (ReCell®; Avita Medical,

Australia and Spray®XP; Grace, USA)183, 188, 189.3 The biggest advantage in using

cultured keratinocytes, or non-cultured sprayed cells, was the coverage of larger

surface areas from relatively small biopsies in the range of 2-5 cm2 from an

uninjured site of the patient. It is unsuitable to be used as a sole treatment for

patients with deep dermal or full thickness injuries, as it predominantly caters to the

epidermal cell population while deeper injuries inevitably also demand dermal

support for efficient healing response.183 Nevertheless, autologous cells are in

widespread use today for the treatment of partial thickness burns. More recent

research has involved further developments of cell-based treatments, with the

potential to enhance these treatments through novel delivery mechanisms or the

addition of growth factors or other biologicals to promote better healing.3, 190-195

30

However, these advanced cell therapies remain in the research stage and are not

currently in widespread use for burn treatment.

1.14 Matrix based therapies

1.14.1 Amniotic membrane Since 1910, allogenic amnion has been used as a biological wound dressing.173 It

draws parallel to normal human skin allografts and was considered the most

effective dressing ever used, especially in preserving the healthy excised wound bed

and providing protection from pathogenic contamination196-199.173 Amniotic

membrane is derived from the innermost layer of the fetal membrane, it is a semi-

transparent tissue consisting of avascular stroma, thick continuous basement

membrane with a full complement of collagen IV and V, laminin and also contains

several protease inhibitors200.173 Advantages of human amniotic membranes include

promotion of epithelial regeneration by reduction in loss of proteins, electrolytes and

fluids, drug delivery and reducing the need and frequency of dressing changes and

pain associated with the process.173 The key disadvantages are its fragility,

technically it is difficult to handle, the risk of contamination and transmission of

disease and it is inadequate in deep dermal injuries where it disintegrates before

healing occurs201, 202.173

1.14.2 Human cadaver derived allografts and xenografts

Human cadaver allograft skin (HCAS) can be used as a temporary dressing in cases

where the availability of patient donor sites are limited203-205.173 Similar to amniotic

membranes, HCAS may also be used as wound dressing to cover widely meshed

autografts in massive burns.173, 206 There are serious problems associated with HCAS

use including limited supply, variable quality, possible contamination and immune

rejection.204, 206, 207 On the other hand, xenografts are derived from various animal

species including rabbit, dog and pig and have been used as a temporary replacement

31

skin for a long time.173 The major limitation in their use include immunologic

disparities and predetermined rejection over time, despite the obvious benefits of

biologically active dermal matrix they provide.203 These alternatives continue to be

considered strictly as only temporary wound cover with therapeutic strategies to

promote healing focused on living autograft cells and tissue.

Figure 10: (a), Cells are isolated from the patient and may be cultivated (b) in vitro on two-dimensional surfaces

for efficient expansion. (c), Next, the cells are seeded in porous scaffolds together with growth factors, small

molecules, and micro- and/or nanoparticles. The scaffolds serve as a mechanical support and a shape-

determining material, and their porous nature provides high mass transfer and waste removal. (d), The cell

constructs are further cultivated in bioreactors to provide optimal conditions for organization into a functioning

tissue. (e), Once a functioning tissue has been successfully engineered, the construct is transplanted on the defect

to restore function. Image and caption taken from Dvir 2011.5

32

1.14.3 Tissue engineered skin

In this approach skin cells are expanded in the laboratory before their application in

wound healing (Figure 10).69 In addition, for deeper injuries a number of dermal

scaffolds have also been developed that can be used in conjunction with the cell

based approaches to restore both dermal and epidermal layers.

All tissue engineered materials needs to have some essential characteristics

including low disease risk, promotion of healing, possess physical properties similar

to normal skin, it must attach well to the wound bed, be compatible to the

developing new vasculature, non-immunogenic and be convenient to use.69, 171 To

replace the dermal layer, predominantly collagen based matrices have been used,

with Integra the first demonstration of a bilayered and biocompatible dermal

scaffold used in the treatment of extensive burn injury to successfully induce the

synthesis of neodermis.141, 208

Epidermal tissue grafts consist of in vitro differentiated keratinocytes forming a

stratified epidermal layer.209 Keratinocytes and fibroblasts are cultured on various

biocompatible substrates as implantable scaffolds including both natural and

synthetic materials.209 Epicel® (Genzyme Corporation Offices, Cambridge, MA) is

the first commercialized epidermal autograft which is composed of cultured

keratinocytes extracted from 2 x 6 cm biopsy taken from the patient, which are

expanded in the laboratory and then cocultured in the presence of proliferation

arrested murine fibroblasts to form grafting sheets.209, 210 It is a front line treatment

for full thickness burns in US and Europe.209 The clinical outcome of epidermal

transplants fall short in cases with inadequate dermal tissue, which is required to

provide support for grafted epidermal sheets211. Alternatively, biocompatible and

biodegradable materials, such as benzyl-esterified analogues of hyaluronic acid

(Hyalograft 3D, Fidia Advanced Biopolymers, Padua, Italy) and polyglycolic acid

(Dermagraft, Shire Regenerative Medicine, San Diego, CA) have been used.209

Other examples include Transcyte (formerly Dermagraft-TC, Shire Regenerative

Medicine) which is made up of porcine dermal collagen coated nylon mesh seeded

with neonatal human fibroblast and externally supported by a silicone

membrane212,209 Dermagraft (Shire Regenerative Medicine), utilise a biodegradable

33

scaffold coated with an allogenic human-derived dermal fibroblasts culture, and

capable of producing several growth factors to stimulate angiogenesis, tissue growth

and reepithelialisation from the wound edge even after cryopreservation and

subsequent thawing213,209 and Apligraf (Organogenesis, Canton, MA) is an example

of a clinically approved allogenic dermoepidermal product (combinatorial therapy),

composed of cultured keratinocytes and a fibroblast dermal layer on collagen I

matrix, for the treatment of venous and diabetic ulcers214.209 Clinical trials revealed a

rate of healing of 63% in cases of ulcers as compared to 49% for compression

therapy and a vast improvement in healing times from 181 to 61 days under

Apligraf.215 Guo et al. developed a bilayer dermal equivalent by loading plasmid

DNA (pDNA) encoding vascular endothelial growth factor-165 (VEGF-

165)/N,N,N-trimethyl chitosan chloride into collagen-chitosan/silicon membrane

scaffold and tested for regenerative properties in a porcine full-thickness wound

model216.3 Full-thickness burn wounds in the pig model when treated with N,N,N-

trimethyl chitosan chloride/pDNA-VEGF dermal substitute showed the best

response towards dermal regeneration with highest density of newly-formed and

mature vessels compared to a blank scaffold and scaffolds loaded with naked

pDNA-VEGF and TMC/pDNA-eGFP.216 In another study, wound healing potential

of five acellular dermal skin substitutes (Integra®, Integra Life Sciences, NJ, USA;

ProDerm®, UDL Laboratories Inc., IL, USA; Renoskin®, Symatese, Ivry-le-Temple,

France; Matriderm® 2 mm, Ideal Medical Solutions, Wallington, UK; and

Hyalomatrix® PA, Addmedica, Paris, France) were investigated in a porcine full-

thickness wound model in a two-step process217.3 In the first step, skin substitutes

were implanted followed by the reconstruction of the epidermis using an autologous

split-thickness skin graft or cultured epithelial autograft (after 21 days, current

clinical treatment regime).3 Significant differences between the skin substitutes in

terms of dermis incorporation and early wound contraction was observed. However,

no significant long-term differences were in scar qualities between different dermal

substitutes and the control group.217 Dermoepidermal substitutes mimic both

epidermal and dermal layers of the skin and are considered the most advanced

substitutes available in the clinic.3 They made up of both keratinocytes and

fibroblasts incorporated into the scaffold which can provide temporary cover,

promote simultaneous dermal and epidermal regeneration and also resemble the

normal skin structure.3 The main limitations, however, are the high production costs

34

and their limited efficiency in terms of permanent wound closure due to allogeneic

cell rejection, as well as limited evidence that the tissue substitutes ultimately

improve the final outcome170.3

1.15 Nanofibrous scaffolds for skin tissue engineering

Tissue engineering (TE) and regenerative medicine is the amalgamation of cellular

components of a living tissue and functional biomaterials to develop functional

living tissue of sufficient size for clinical translation.218 Recently, cell patterning,

migration, proliferation and differentiation have been the point of focus for both TE

and regenerative medicine.218 Development of scaffold as ECM substitutes are

aimed to promote cells proliferation and differentiation for skin tissue regeneration

in vivo.218 Cells respond to ECM signalling in a multistep process which is initiated

by cell receptor-ECM ligand interactions, followed by the sequestration of growth

factors by ECM, spatial cues and the mechanical force transduction219.218

Nanofabrication techniques have been employed to develop complex multifunctional

porous, nanometer-sized fibrous scaffolds with surface morphology that can

determine and influence cell fate, regulate the expression of specific proteins and

encourage cell-specific scaffold remodelling.220.218 Techniques including nanoscale

surface pattern fabrication such as lithography, electrospinning and self-assembly

has been explored to fabricate scaffold topography down to nanoscale221.218 They

can incorporate a myriad biological cues such as growth factors, angiogenic factors,

cell surface receptors and spatial cues.218 These approaches can be used to influence

cell proliferation, migration, differentiation and 3D organisation including

ingrowth.3, 218 Electrospinning has gained considerable amount of interest because of

its operational simplicity, versatility, ability to process variety of materials and

propensity in producing nanofibers mimicking the ECM, for the fabrication of

nanoscaffolds for skin regeneration.3, 218

Electrospinning utilise electric field to produce a highly impermeable, non-woven

matrix of sub-micron to nanoscale fibers by pushing a highly charged polymer liquid

jet through a small diameter nozzle of a syringe24, 222.223 Conventionally, a Taylor

35

cone is formed when electrostatic charging of the fluid pulls the polymer solution at

the tip of the nozzle, ejecting a single fluid jet from the apex of the cone.223 This

polymer jet accelerates and thins in the electric field to form a nanofibrous scaffold.

Electrospun scaffolds with different morphologies from aligned to random have

been fabricated with fiber diameters down to the 100 nm range with various pore

sizes224.209 Electrospun scaffold with high porosity and large surface area to volume

ratio facilitate efficient mass transfer into the 3D structures of the scaffold.209

Nanofibrous scaffolds serve as a combinatorial approach towards regeneration of

various tissues such as skin, bone, cartilage, vascular and neural tissues and

subsequent drug delivery platform.209

Figure 11: Various fibrous structures fabricated via an electrospinning technique. (a) Core-shell structure,

(b) random/align directional fiber and (c) micro-/nano-sized fiber. Image and caption taken from Rim 2013.225

Nanofibers can be electrospun in various patterns including random, aligned and

core-shell (Figure 11).209 Each type has their own advantages; for example random

36

fibers are used in tissue engineering approaches, while aligned ones have been

explored in solar cells and shown to promote their efficiency225.209 In an another

approach, emulsion electrospinning is used to fabricate polymer fibers with an

integral core-sheath structure, promoting loading efficiency of the cargo, avoiding

the initial burst release and protect and help maintain bioactivity of the loaded

protein by avoiding the exposure of the cargo to the degrading physiological

environment226.209 These fibrous scaffolds can be further functionalised using

surface modification techniques such as plasma treatment, the wet chemical method,

surface graft polymerization, co-electrospinning, and the immobilization of bioactive

ligands227.209 In order to achieve controlled drug release the payload like drugs,

enzymes and cytokines are physically or chemically immobilised within the

fibers228.209

As a dermal substitute, highly porous 3D collagen scaffolds were fabricated using a

electrospinning and were evaluated in a keratinocyte/fibroblast co-culture in vitro

model for skin tissue engineering229.209 It resulted in well dispersed fibroblast

ingrowth, while keratinocytes migrated through the pore structure and differentiated

on the scaffold surface to form a stratum corneum and a 3D dispensed scaffold

similar to normal skin.229 A well interconnected porous network is required in 3D

scaffolds to achieve guided cell adhesion, cell growth favouring tissue development,

cell migration and subsequent transport of solutes.209 This highlights the importance

of mechanical properties of the engineered electrospun scaffolds towards their

functions.209 Parenteau-Bareil et al crosslinked collagen, derived from various

sources, with chitosan to produce ECM mimicking tissue engineering scaffolds230.209

Crosslinking of collagen increased mechanical strength and facilitated a similar

cellular response to normal ECM whilst maintaining the biocompatibility of the

scaffold. Nanofibrous scaffolds provide a high surface area allowing oxygen

permeability and avoid fluid accumulation at the wound site, while preventing

microbial infection by restricting their penetration through its small inter-fiber pores,

thereby making them ideal candidates for wound dressings.209

The aim of using an electrospun scaffold based TE approach is to facilitate and

promote the natural healing response of the body without inducing an immunogenic

response, allowing the body to utilise the scaffold to enhance regeneration of

“neonative” functional tissues.209 Electrospinning provides the flexibility to

37

coelectrospin polymers with drugs or proteins to form nanofibrous scaffolds for

applications in drug delivery.209 Various polymers have been electrospun including

synthetic ones such as poly(ε-caprolactone) (PCL), polylactic acid (PLA),

polyglycolic acid (PGA), PLGA, polystyrene, polyurethane (PU), polyethylene

terephthalate, poly(L-lactic acid)-co-poly(ε-caprolactone) (PLACL) and natural

polymers including collagen, gelatin and chitosan to obtain fibers with diameters

ranging from few nanometers to several microns24, 224, 228, 231.209

1.16 Skin regeneration using electrospun scaffolds

The high surface area to volume ratio, porosity and structural similarity to the ECM

architecture of the dermis makes nanopatterned electrospun fiber meshes ideal

scaffolds for skin regeneration232-234.3 The limitations associated with electrospun

scaffolds include poor mechanical properties, non-uniform thickness distribution and

poor integrity.3 For skin regeneration a myriad of both natural polymers,235-237

synthetic polymers,232, 238-240 and combinations of the two have been successfully

electrospun and tested.3 Poly(ε-caprolactone)/gelatin nanofibrous scaffold was

electrospun on top of a commercial polyurethane wound dressing (TegadermTM; 3M

Healthcare, USA) in an attempt to regenerate dermal wounds241.3 It was shown to

promote cell proliferation, adhesion and growth in human dermal fibroblasts. Yang

et al. employed emulsion electrospinning to fabricate ultrafine core sheath

poly(ethylene glycol)-poly(D,L-lactide) fibers loaded with basic fibroblast growth

factor (bFGF) and demonstrated its gradual release over 4 weeks232.3 An in vitro

study using mouse embryonic fibroblasts showed enhanced cell adhesion,

proliferation and secretion of ECM when cultured on a nanofibrous scaffold.232 In

vivo studies using a dorsal wound model in diabetic rats treated with bFGF/

poly(ethylene glycol)-poly(D,L-lactide) mats showed elevated rates of healing with

complete re-epithelialization and regeneration of skin appendages, while bFGF from

the scaffold promoted collagen deposition and its remodelling to normal

architecture232.3 In another study, potential of bone-marrow derived mesenchymal

stem cells differentiation into the epidermal lineage (keratinocytes) was investigated

38

on poly(L-lactic acid)-co-(ε-caprolactone) nanofibrous scaffold, with and without

collagen242.3 It was concluded that better results were obtained for collagen

containing poly(L-lactic acid)-co-(ε-caprolactone) nanofibrous scaffolds.242

Electrospun nanofiber scaffolds have also been explored as delivery systems for

variety of bioactive molecules including drugs, proteins and even RNA243-247.3

Ignatova et al. reviewed the use of nanofibrous scaffolds loaded with various

antibiotics and antibacterial agents including nanoparticles for wound dressing

applications248.249 Kenawy et al. electrospun poly(ethylene-co-vinyl acetate),

poly(lactic acid) and their blend with tetracycline hydrochloride as a model

drug250.249 It was reported that drug release behaviour was dependent on the polymer

carrier and drug loading. Scaffolds with a 50/50 blend and relatively low drug

loading (5 wt%) showed sustained release for over 5 days. Higher drug loading (25

wt%) resulted in an initial rapid burst release of the surface adsorbed drug that

quickly dissolved in tris buffer.249 Sodium alginate was electrospun with poly(vinyl

alcohol) containing different concentrations of ZnO nanoparticles as an antibacterial

agent in a mouse fibroblasts model251.3 They observed inverse correlation between

cell adhesion and spreading, and ZnO concentration in the scaffolds. The

antibacterial activity was evaluated in both gram positive and gram negative

bacterium cultures through diffusion disc experiments using Staphylococcus aureus

and Escherichia coli. The results showed direct correlation between antibacterial

activity of the scaffold and increasing concentration of ZnO nanoparticles251.3

Various strategies have been employed to control the drug release mechanism from

fibers with the aim to prevent or control the initial burst release of the drug. Dave et

al. developed electrospun enzyme-embedded antibiotic-releasing polymer scaffolds

using antibiotic gentamicin sulfate (GS) and a polymer degrading enzyme (lipase) in

PCL polymer to achieve endogenously triggered controlled drug release252.249The

GS release from the scaffold was shown to be dependent on lipase concentration. In

an another study, PDLLA was electrospun with Mefoxin to develop a nanofibrous

scaffold with the loading efficiency of 90%.249 However, most of the drug was

released in the first 3 h which reached its completion within 48 h, indicating that the

drug was adsorbed on the nanofiber surface.249 In order to block such initial burst

release of Mefoxin from PLGA nanofibers, Kim et al. used an amphiphilic block

copolymer (PEG-b-PLA) and reported prolonged drug release for up to 1 week253.249

39

In an another attempt, Yohe et al. used superhydrophobic polymer made up from

PCL doped with 0-50 wt% poly(glycerol monostreate-co-ε-caprolactone) (PGC-

C18) as a hydrophobic polymer and trapped air as a barrier to control the rate of

drug release from the scaffold by blocking the fiber pores254, 255.249 By using the

model bioactive agent SN-38 (7-ethyl-10-hydroxy campthothecin), PCL electrospun

fibers doped with 10 wt% PGC-C18 was shown to allow to have linear sustained

release over 60 days compared to initial burst release (10 days) observed in case of

PCL fibers alone which plateaued out at 20 days255.249Alternatively, covalent

conjugation of the drug to the fibers can be used as a method to control and regulate

the drug release mechanism from the fibrous scaffolds.249 Jiang et al. explored this

by covalently functionalising ibuprofen onto poly(ethylene glycol)-g-chitosan

(PEG-g-CHN) polymer and coelectrospinning it with poly(lactide-co-glycolide)

(PLGA) to yield nanofibrous scaffold with prolonged release of the drug for more

than 2 weeks256.249 Zou et al. copolymerised ε-caprolactone with PDLL, to induce

functional ketone groups for later functionalization into a PDLL backbone, and used

them to develop a fibrous scaffold257.249 Subsequently, the scaffold was surface

functionalised with heparin molecules and loaded with bFGF. It was reported that

bFGF release was dependent on the amount and molecular weight of heparin used.

One of the biggest limitations with this approach is the necessity to select carrier

polymers with desired functional groups257.249

In an alternate approach, polymer crosslinking has been explored as a parameter to

control the initial burst release of the drug.249 Some of the methods used include

UV-irradiation,258-260 dehydrothermal treatment,261 and chemical treatment including

glutaraldehyde,262, 263 formaldehyde264 and carbodiimide265, 266.249 Meng et al.

prepared Fenbufen-loaded poly(D,L-lactide-co-glycolide) (PLGA) and

PLGA/gelatin nanofibrous scaffolds267.249 Fenbufen release was found to be

dependent on the concentration of gelatin used, alignment of the fibers in the

scaffold and amount of crosslinking. It was shown that subsequent crosslinking

treatment of the fibers effectively curtailed the burst release of the FBF at the initial

release stage from a PLGA/gelatin (9/1) nanofibrous scaffold267.249 In an attempt to

control the formation of hypertrophic scars (HS), poly(l-lactide) was coelectrospun

with 20(R)-ginsenoside Rg3 (GS-Rg3).268 GS-Rg3 was used for its potential in

inhibiting the formation of HS in vivo as measured using H&E staining and

40

apoptosis of fibroblasts to control the later stage HS hyperplasia by inhibiting

inflammation and down-regulating VEGF expression.268 The in vivo wound rabbit

ear HS model suggested sustained release of the drug for 3 months which was

reported to be dependent on the drug concentration in the electrospun fibers. GS-

Rg3/ poly(l-lactide) scaffold was shown to significantly inhibit HS formation with

decreased expression of collagen fibers and microvessels.268

Bioactive molecules such as proteins, DNA, RNA and growth factors have also been

encapsulated in polymer fibers mostly using modified techniques such as blend

electrospinning and coaxial electrospinning.249 As the name suggests, blend

electrospinning involve premixing of bioactive molecules with polymers prior to

electrospinning increasing the surface localization efficiency of bioactive

molecules.249 In an attempt to study if the activity of an encapsulated protein can be

maintained after electrospinning, human β-nerve growth factor (hNGF) was

electrospun along with BSA as carrier protein, into a partially aligned nanofibrous

scaffold using poly ε-caprolactone (PCL) and poly(ethyl ethylene phosphate)

(PEEP) polymers269.249 While the protein aggregates were heterogeneously

distributed within the fibers, its bioactivity was demonstrated in a neurite outgrowth

assay in PC12 cells. It was concluded that the prolonged release of an active protein

can be achieved over the period of 3 months269.249 The same group also

demonstrated the efficacy of this polymer system in the delivery of small interfering

RNA (siRNA) and transfection reagent (TKO) complexes270.249 The siRNA/TKO

complexes showed better gene knockdown efficiency than siRNA alone both of

which were shown to be able to release from the scaffold for almost a month.

Coaxial electrospinning was developed to fabricate core-shell fibrous scaffolds

where both polymer and biomolecules are coaxially and simultaneously

electrospun.249 The strategy involves use of the shell polymer to protect the cargo

encapsulated in the core from the physiological degradation whilst allowing slow

sustained release of the therapeutic.249 A nanofibrous scaffold encapsulating plasmid

DNA (pDNA) within the PEG core and the non-viral gene carrier poly(ethyleimine)-

hyaluronic acid (PEI-HA) within the PCL sheath was prepared by the coaxial

electrospinning technique271.249 They also studied the effects of various processing

parameters using fractional factorial design. Extended release of the non-viral gene

delivery vector from the sheath was observed over a period of 60 days quantified

41

using EGFP transfection expression and was shown to be dependent on pDNA

loading into the fibers. Mickova et al. studied if liposome activity can be maintained

post electrospinning process. They studied liposomes blended within nanofibers and

polyvinyl alcohol-core/poly-ε-caprolactone-shell nanofibers with embedded

liposomes prepared using coelectrospinning technique272.249 They concluded that

liposomes encapsulated in the core/shell fibers retained enzymatic activity of

encapsulated horseradish peroxidase while blended liposome became inactive.

1.17 Electrospun hybrid materials

Despite the evidence of the capabilities of electrospun scaffolds in tissue engineering

and drug delivery this technology has not reached its full potential in terms of

translation. One of the biggest limitations associated with electrospun scaffolds

developed from a single polymer has been the lack of multi-functionality and

universality. They require specific surface modification approaches which are

limited to specific polymers and can be difficult to translate into other polymer

models. In an alternate approach, nanoparticles have shown tremendous potential

and versatility not only in drug delivery applications but also in the fabrication of

various functional materials which has already been translated into commercial

products. Incorporation of these two technologies could open new avenues by

utilising the potential of both technologies. In the following sections the potential of

this combinatorial approach is demonstrated for applications in material science.

This approach can be extrapolated towards tissue engineering using multifunctional

scaffolds.

Among the various fibers prepared using electrospinning, the nanoparticles (NPs)

containing electrospun fibers exhibit a huge variety of potential applications.273 The

composite fibrous mats show flexibility, are free standing,273 and incorporate the

advantages of both starting materials i.e. polymer and the NPs. There are three main

ways specified in the literature to fabricate NPs-electrospun fiber composite

materials.

42

1. Electrospun fibers with surface functionalised nanoparticles: This is an

efficient method where pre-treated electrospun scaffold is immersed into the

colloidal NP solution to facilitate surface adsorption of NPs on fibrous scaffold.273

Various kinds of NPs such as metal, metal oxide, carbon, polymer, fluorescent NPs

and even cells have been successfully combined with electrospun fibers to form

composite materials15, 274, 275.273 For example, a water stable poly(vinyl alcohol)

(PVA)-AuNPs composite fibrous mat was fabricated by pretreating the surface of

the fibers with 3-mercaptopropyltrimethoxysilane, allowing the adsorption of

AuNPs on the surface of the fibers via Au-S bonds276.273 In cases where the

electrospun fibers are stable but do not readily adsorb NPs, an in situ reduction

method is used.273 For example, Au adsorbed TiO2 composite nanofibers were

prepared by organic capping agent mediated photocatalytic reduction of

HAuCl4277.273 It was reported that shape of the nanomaterial is dependent on type

and concentration of the capping reagent.277 It provides the means to fabricate

specific nano-structures on the nanofibers for respective applications in chemical

and biological sensing277.273 Various NPs have been synthesised on electrospun

fibers using this technique including Au, Ag, Pd, Pt, TiO2, WO3 and SnO2278-282.273

2. NPs encapsulating electrospun fibers: In this method NPs are synthesised

within the polymer fibers.273 Typically, the polymer solution is electrospun with a

metallic or ceramic acetate precursor which is then annealed to reduce the precursor

and form NPs using methods such as gas-solid reaction, calcination or laser

ablation.273 Yang et al. used AgNO3 as a precursor by electrospinning it with

poly(acetonitrile) (PAN) which upon exposure to HCl yielded AgCl NPs both on the

surface and interior of the fibers283.273 Similarly, copper nitrate was used as a

precursor and electrospun with poly(vinyl butyral) (PVB) to give fibrous scaffold

encapsulating copper nitrate. Polymer layer was removed by annealing the scaffold

at 450 °C in an air atmosphere for 2 h. Subsequent heating at 300 °C for 1 h under

hydrogen atmosphere was then applied to obtain copper fibers284.273This technique

can also be used to fabricate more specialised structures.273 For example, Sn NPs

were used as a precursor and were electrospun with porous multichannel carbon

microtubes. The resulting scaffold underwent calcination to yield Sn-Carbon

nanofibers285.273

43

3. NPs-electrospun fibers: This is a one-pot fabrication technique where NPs

are homogenised in a polymer solution and electrospun together to obtain composite

fibers.273 This technique relies on the homogeneity of the polymer-NP solution to

yield scaffold with uniformly distributed NPs.273 Inorganic NPs can easily be

encapsulated within the electrospun fibers because of their high electron density.273

Heterogeneous solution of NPs results in NP cluster formation within the

electrospun fibers, which can be address by surface treatment of the fibrous

scaffold.273 Both hydrophobic and hydrophilic polymers have been explored and

electrospun. Some of the examples include (poly(ethylene oxide) (PEO), poly(vinyl

alcohol) (PVA), poly(vinylpyrrolidone) (PVP), polyacetonitrile (PAN), PLGA and

poly(L-lactide) (PLLA).273

Magnetite (iron oxide (Fe3O4) nanoparticles), another interesting class of

nanoparticles have been explored for their magnetic resonance properties. These

properties have been shown to be retained in the electrospinning process opening

avenues for various applications.273 It can be used as intelligent fabrics in defence

clothing, ultrahigh-density data storage, sensors and in health care.273 Problem in

their use lies in their aggregation which tends to curb their magnetic response. It is

understood that magnetite nanoparticles form clusters to reduce their energy because

of their high surface area to volume ratio.273 Stabilisers have been employed to

overcome this problem. The other alternatives include the use of electrostatic

surfactant and steric polymers286-288.273 Coaxial electrospinning was used to address

this problem of aggregation of nanoparticles. Core-shell Fe3O4-poly(ethylene

terephthalate) magnetic composite nanofibers were electrospun where Fe3O4NPs

form extended aligned structures within the core of the nanofibers and were further

shown to have retained their super-paramagnetic behaviour289.273 The core-shell

nanofibers demonstrated higher mechanical properties and responded to externally

applied magnetic field due to the dipole-dipole interaction between magnetic NPs in

magneto-rheological fluid.273 Other composites using similar approach include

polyarylene ether nitriles-Fe-phthalocyanine-Fe3O4, Fe3O4-poly(ethylene oxide) and

Fe3O4-poly(ethylene terephthalate)273, 289, 290, 291.

Many metal and inorganic fluorescing NPs have also been electrospun for their

luminescence properties including boron carbonitride oxide, sodium yttrium

44

fluoride, quantum dots and rare earth metals292, 293.273 Some of the examples includes

Sm+3/TiO2, NaYF4:Yb+3, Er+3/SiO2 and YVO4: Eu+3/PEO for rare earth metals,294-296

CdTe and CdS quantum dots. #498;, #499}.273

As demonstrated, electrospun polymer/nanoparticle composite materials showcase

hyphenated properties of the two constituents. This approach has great potential not

only in material science but also in tissue engineering applications. Electrospun

scaffolds can be functionalised with different types of nanoparticles to induce multi-

functionality, for example magnetite and upconverting nanoparticles can be

incorporated within the polymer scaffold to achieve sustained drug delivery and

simultaneous imaging using MRI and fluorescence microscopy.

1.18 Summary

With surgery continuing to be the mainstay for treating burn injury and

reconstructing scars and limited evidence for the efficacy of cell therapies alone,

there is a need for the development of alternative treatment modalities. M6P has

been explored as an anti-scarring agent and its exogenous delivery has been shown

to be antagonistic towards the activation of LTGFβ. TGFβ is a key cytokine that

influences scar outcome. Upon activation, TGFβ upregulates the expression and

synthesis of ECM proteins and molecules including collagen I, fibronectin,

hyaluronic acid and α smooth muscle actin. The extent to which these molecules are

expressed closely correlates to the extent of scarring. It has been proposed that

clinical trials for commercial M6P (Juvidex®) could not meet the final end point

because of the delivery mechanism adopted and the metabolic vulnerability of M6P.

The mode of M6P delivery is critical as it gets rapidly metabolised and appropriate

vehicles have to be employed to achieve and maintain requisite tissue levels over the

post-wounding or healing period. This shortcoming of M6P needs to be addressed

for its development as a potent anti-scarring drug and can potentially be solved

through the development of more stable analogues.

45

Phosphonate analogues are metabolically more stable than phosphate analogues and

they also project higher affinity towards the receptor (M6P/IGFII receptor). Two

novel phosphonate analogues of M6P, analogue 1 and analogue 2 were investigated

for their toxicity and efficacy in inhibiting the expression of collagen I under TGFβ1

stimulation in a primary skin dermal fibroblast in vitro model. With analogue 2,

whilst it is able to penetrate cells due to its lipophilicity, this is also a limitation in its

sustained topical delivery and therefore new modalities need to be explored to

achieve local, controlled and consistent administration of the drug.

Conventional skin substitutes are still prevalent in clinic use. However, they suffer

from various disadvantages including low adhesion, creation of a new injury site (in

autografts) and immune rejection (in allografts). To address these problems

alternatives have been developed. Firstly, cell-based epidermal substitutes, efficient

in accelerating the reepithelialisation phase of wound healing but only applicable in

superficial deep wounds without additional therapies have been widely used.

Dermal substitutes, both natural and synthetic have also been developed and used in

deeper injuries. However, these dermal substitutes have problems with inadequate

vascularisation, low mechanical strength, uncontrolled degradation profiles and high

costs. Finally, the most advanced dermo-epidermal substitutes aimed at

simultaneous regeneration of both epidermal and dermal layers by using constructs

incorporating respective cells from both layers of the skin have more recently been

trialled. The major limitations have been high production costs, poor vascularisation

and failure to achieve permanent wound closure due to allogeneic cell rejection.3

Nanotechnology based techniques such as nanoparticles or nanofibrous scaffold

based delivery systems have been shown to facilitate the delivery of hydrophobic

drugs. Nanofibrous scaffolds are preferred because they mimic ECM, providing

congenial environment for cellular ingrowth, and simultaneously deliver drugs to

promote regeneration. In addition, the fractal-like geometries of fibrous scaffolds

have been of interest recently to design and develop personalised diagnostic devices

using fractal models of complex biological processes. For example, fractals have

been used to describe the architecture of tumor microenvironments. Most of the

polymers used thus far have been designed for specific applications and require

specific approaches for surface modifications which are complex and laborious in

order to induce any multifunctionality. PGMA, which has an epoxy group in each

46

monomer unit, would be an ideal polymer to address the concerns with

multifunctionality. A nanofibrous fractal-like scaffold using PGMA was developed

with an ability to incorporate multimodality and multifunctionality.

It has been demonstrated that existing clinical treatments for wound healing have

various shortcomings and there is a need for improved therapeutic approaches. TE

inspired approaches including electrospun scaffolds, although having shown promise

in addressing the shortcomings with conventional approaches, are marred with their

own limitations restricting their clinical translation. These include poor mechanical

strength, inadequate vascularisation and cellular ingrowth, difficult to control drug

diffusion and limited ability to be functionalised. The inadequacy of these scaffolds

for improving scar formation leaves a current unmet medical need. In this thesis, a

multifunctional scaffold has been developed and characterised, with data describing

its use in a range of applications presented. In addition, data showing the potential

efficacy of a stable analogue of mannose-6-phosphate for wound healing has been

demonstrated. This has led to initial work on a combinatorial approach using both

scaffold and analogue to overcome many of the limitations previously discussed and

potentially enhance wound healing and reduce scar formation in patients into the

future.

* PXS25 would be referred to as analogue 1 and PXS64 would be referred as

analogue 2 except in the paper #3

1.19 Hypotheses and Aims

Hypotheses

1. The epoxy functionalised polymers can be used to develop multifunctional

fractal-like substrates

2. Stable isosteric analogues of mannose-6-phosphate (M6P) can be used to

reduce the overexpression of collagen I gene

47

Aims

1. Optimise the parameters for electrospinning PGMA and study its efficacy as a

multifunctional polymeric substrate

2. Investigate the biocompatibility and potential of two M6P analogues as anti-

scarring agents in human dermal skin fibroblasts

3. Investigate the potential of a combinatorial scaffold:drug approach to limit

fibrosis

48

Chapter 2

Introduction to the series of papers

2.1 Development of a universal multifunctional scaffold

The first paper presented as part of this thesis presents the preparation of universal

multifunctional fractal-like scaffold. Fractals are structures that maintain their

structural characteristics with successive variation of scale.297 In addition to their

applications in material science, fractals also find applications in cancer biology.

They have been used to understand tumor architecture and morphology with

implications in tumor growth and angiogenesis.298

Polyglycidyl methacrylate (PGMA), which contains a reactive epoxy unit per GMA

monomer, can be functionalized with dyes, other polymers and drugs using SN2

nucleophilic substitution reactions. Amines, alcohols, carboxylic acids, alkyl halides,

thiols and acid anhydrides are examples of highly reactive species that interact with

PGMA.299

Different solvent systems such as acetone, chloroform and methylethyl ketone

(MEK) were tested in the electrospinning of PGMA. Even though PGMA has high

solubility in these solvents, acetone and chloroform were limited in their use because

of their high vapor pressure and low conductivity. Therefore PGMA was electrospun

in methyl ethyl ketone (MEK) to obtain a nanofibrous scaffold matrix (ES-PGMA).

ES-PGMA was heated at 80 °C overnight to allow inter-crosslinking of the polymer.

Fiber morphology was determined using scanning electron microscopy (SEM). The

49

PGMA-MEK system yielded uniform nanofibers with an average diameter of 0.69 ±

0.04 µm over a large area.

PGMA is a hydrophobic polymer which needs to be functionalized to induce

hydrophilicity to the backbone. This is usually carried out by reacting PGMA

polymer backbone with hydrophilic reacting groups such as carboxylic acids and

amines. ES-PGMA film was end-grafted with carboxylic acid-terminated poly (N-

isopropyl acrylamide) (PNIPAM-COOH) to introduce ‘switchability’ in terms of

thermally induced hydrophilic-hydrophobic behavior. PNIPAM is a thermo-

responsive polymer which undergoes reversible phase transition in water at 32 °C

(lower critical solution temperature (LCST)). This is caused by the transition from a

swollen hydrated state to a shrunken dehydrated state above the LCST. Surface

properties of the scaffolds were examined by contact angle measurements. End

grafting of ES-PGMA with PNIPAM-COOH (ES-PGMA-g-PNIPAM-COOH

nanofibers) resulted in dramatic temperature dependent changes in the contact angle,

confirming the surfacing grafting of the hydrophilic PNIPAM moiety.

Polymer composites comprising both polymers and nanomaterials have been

regarded as a new class of hybrid materials with many functions. Lack of

universality however is the major limitation in the development of a ubiquitous

polymeric functional platform.

PGMA was explored as a potential ubiquitous polymeric platform. PGMA was

coelectrospun with three different types of nanoparticles: (NaGdF4:Yb, Er);

palladium (Pd) and magnetite (Fe3O4) to develop polymer nanoparticle composite

materials. These fibrous composite materials were characterized using SEM,

transmission electron microscopy (TEM) and X-ray microanalysis.

Coelectrospun polymer-nanoparticle composites were tested for their respective

applications in upconversion fluorescent imaging in the case of (NaGdF4:Yb, Er),

hydrogen gas sensing in the Pd composite and magneto-responsive behavior in the

case of Fe3O4 composite matrix. Upconversion properties of the ES-

PGMA/(NaGdF4:Yb, Er) nanocomposite was evaluated using NIR room temperature

emission spectroscopy. This measures the efficiency of the nanocomposites in

converting near infra-red excitation wavelength into visible emission. ES-

PGMA/(NaGdF4:Yb, Er) nanocomposites showed prominent upconversion

50

properties with three major emission peaks observed in the visible region at 521,

541, and 655 nm.

Pd has been shown to demonstrate high selectivity and significant adsorption of

hydrogen gas. Because of this feature Pd has been used in hydrogen gas sensors. Pd

readily absorbs hydrogen gas which diffuses as atomic hydrogen into the lattice to

form palladium hydride, PdHx, resulting in to phase transition and a

corresponding change in the lattice spacing. Such changes in phase and lattice

spacing cause a measurable resistance change in Pd material. Mostly, Pd has been

adsorbed onto functionalized electrospun scaffolds. In this case, however, Pd was

coelectrospun with PGMA to yield the fibrous composite material. ES-PGMA/Pd

composites were evaluated for their hydrogen gas sensing capacity using alternating

concentrations of nitrogen and hydrogen gas. Change in current was measured using

current-voltage (I-V) sweeps. Response time 90 of ~14 seconds was observed for a

hydrogen gas concentration range of 1-10% from 0.16 ng of Pd as ES-PGMA/Pd

fibrous composite mounted on 650 x 900 µm IED platform.

More recently magnetic materials and matrices were developed for potential

applications in monetary currency and defense fabric materials. ES-PGMA/Fe3O4

nanocomposite fibers were evaluated for magnetization properties using SQUID

magnetometry (Superconducting Quantum Interference Device). SQUID analysis

confirmed superparamagnetic behavior of the composite material at room

temperature. The mass specific saturation magnetization (Ms) of the fibers was 4.0

emu g−1.

Results are presented in Agarwal, V., Ho, D., Ho, D., Galabura, Y., Yasin, F. M.D.,

Gong, P., Ye, W., Singh, R., Munshi, A., Saunders, M., Woodward, R. C., St. Pierre,

T., Lorenser, D., Wood, F. W., Fear, M., Sampson, D. D., Zdyrko, B., Smith, N.M.,

Luzinov, I., Iyer, K.S., A Functional Reactive Polymer Nanofiber Matrix, RSC

Advances (Submitted)

51

2.2 Evaluation of mannose-6-phosphate analogues as potential anti-

scarring agents

M6P has been shown to control and inhibit the secretion of extracellular matrix

proteins including collagen I, one of the key proteins in granulation tissue and

scars.144, 145, 300 M6P inhibits TGFβ1 activation, a primary cytokine regulating wound

healing and subsequent scar formation, by competitively binding to M6P/IGFII

receptor.301 Recombinant M6P (Juvidex®) has been examined in phase II

randomised human clinical trials for its efficacy in wound healing and scar reduction

at split thickness skin graft donor sites.153 Intradermal delivery of Juvidex showed an

accelerated wound healing response. However, no significant reduction was

observed in scar formation post healing. It has been postulated that the metabolic

vulnerability and delivery mechanism for M6P could have been the limiting factors

behind its translation into the clinic.

Two new bioisosteric phosphonate analogues of M6P were investigated (Aim 2,

analogues were kindly provided by Pharmaxis Pvt Ltd and used without any further

modification). Phosphonate drugs were selected because of their inherently higher

stability towards phosphatase enzymes.159 Analogue 1 was designed by replacing the

P-O bond at C6 with the methylene bridge at C6 in M6P to enhance its stability

against phosphatases present in serum. In addition, the anomeric hydroxyl was

substituted with m-xylene group as they have been reported to improve their

recognition by the receptor.150 Analogue 2 was designed as a prodrug by further

derivatization of analogue 1 with bis(pivaloyloxy)methyl (POM) linkers to yield a

neutral product. Neutral analogues are interesting because they get rapidly

internalized into the cell, where the POM linkers are gradually hydrolysed by

microsomal esterases to release an active analogue 1. Hence analogue 2 was

developed to achieve intracellular targeting of the M6P/IGFII receptor.

In vitro behaviour of the analogues was assessed in primary human dermal skin

fibroblasts (HDF) and compared against M6P. Although binding to the M6P/IGFII

receptor is one of the key mechanisms behind LTGFβ activation, there are other

alternate modes of activation of LTGFβ in vivo. Therefore, it is also important to

investigate the efficacy of the new analogues in the presence of recombinant active

52

human TGFβ1. To assess the dose response of the analogues, HDF cells were

incubated with varying concentrations of the analogues both in the presence and

absence of TGFβ1.

The MTS assay was used to assess the cytotoxicity of the analogues. Cells were

incubated with the analogues, both in presence and absence of TGFβ1, over a period

of 72h. At the specified time, MTS reagent was added to the culture and further

incubated to allow metabolically active cells to bioreduce the MTS reagent and form

the coloured formazan product which is measured by spectrometry at 490 nm.

Colour intensity is correlated to the number of metabolically active cells in culture

i.e. proliferating cells. TGFβ1 has been reported to curtail the proliferation of HDF as

was observed here. However, no add on effect of the analogues was observed at any

of the concentrations studied. Both analogue 1 and 2 behaved similarly to M6P in

terms of inherent toxicity. Analogue 2 however, showed toxicity at high

concentrations (> 50 µM) due to the formation of formaldehyde, which is one of the

degradation products produced as a result of esterase hydrolysis of POM linkers in

analogue 2.

A complimentary live/dead assay was carried out to study the effects of the

analogues on cell viability over a period of 72 h. Cells were incubated with varying

concentrations of analogues. At the specified time-point calcein AM/ ethidium

bromide I dyes were added. Calcein stains the intracellular esterase in viable cells

and fluoresces green whereas ethidium bromide stains non-viable cells and

fluoresces red due to the binding of the ethidium homodimer with nucleic acids

penetrated by ethidium through the compromised cell membranes. No loss was

observed in cell viability.

Cell morphology was also assessed. TGFβ1 stimulation significantly inflated the cell

body area compared to the non-treated control cells which was similar to what has

been reported previously. Interestingly, this inflation in cell area came down to

normal levels post analogue treatment. No apparent alteration in cell body area was

observed in cells treated with varying concentrations of analogue alone without

TGFβ1 stimulation.

The potential of the analogues as anti-scarring agents was evaluated in terms of their

ability to inhibit the overexpression of key fibrotic indicators such as collagen I gene

53

expression. Collagen I is a key fibrotic marker which has major implications in

wound healing and formation of granulation tissue. In order to curb the secretion of

collagen matrix it is crucial to control its expression at the transcriptional level.

TGFβ1 is known to up-regulate the expression of collagen I in HDF.

Real time-quantitative polymerase chain reaction (RT-qPCR) analysis was carried

out to quantitate the expression of collagen I post-treatment. TGFβ1 stimulation has

been reported to significantly up-regulate the expression of collagen I gene

expression as was observed here compared to untreated control cells. Efficiency of

the analogues was tested in terms of their potential in inhibiting this increase in

collagen I mRNA. 10 µM analogue 2 significantly reduced the TGFβ1 mediated

increase in collagen I gene expression. This response was similar to that observed

with M6P treatment. The hydrophilic analogue 1 did not show any response.

It can be deduced that analogues behave similar to M6P. Further, intracellular

targeting of the M6P/IGFII receptor, using analogue 2 as a prodrug of analogue 1, is

more efficacious compared to intercellular targeting using analogue 1.

Results are presented in Agarwal, V., Toshniwal, P., Smith, N. E., Smith, N. M., Li,

B., Clemons, T. D., Byrne, L. T., Hassiotou, F., Wood, F. M., Fear, M., Corry, B.,

and Iyer, K. S., Enhancing the Efficacy of Cation-Independent Mannose 6-

Phosphate Receptor Inhibitors by Intracellular Delivery, Angewandte Chemie

International Edition (Submitted)

2.3 Delivery of the lipophilic mannose-6-phosphate analogue PXS64

using an electrospun PGMA scaffold

One of the biggest limitations in the clinical translation of lipophilic drugs has been

their delivery and associated serum stability. Rapid degradation, particularly in a

wound environment, is a common issue. In order to achieve prolonged stability,

drugs are encapsulated in polymer shells to protect them. One of the niche ways to

achieve controlled delivery of sensitive lipophilic drugs is by encapsulating them in

nanofibrous scaffolds developed using electrospinning.302 Drugs get released

54

through a diffusion mechanism in a controlled rate over time sufficient to be taken

up by the cells before degradation.303

Analogue 2 was coelectrospun with PGMA to fabricate a nanofibrous scaffold (Aim

3). Despite proven toxicity at high concentrations, analogue 2 was selected for its

potential in reducing the expression of fibrotic marker, collagen I. In addition, it was

rationaled that the release of the drug would be slow and within the tolerated limit.

The electrospun scaffold was characterised using scanning electron microscopy

(SEM) and drug loading was estimated using HPLC. Release studies were also

attempted using HPLC. It is important to conduct release studies in aqueous

conditions to mimic the in vivo environment. No drug release was detected from the

fibers in the HPLC analysis. It was implicated that either there is no drug in the

fibers or it may have been degraded during the electrospinning process.

Alternatively, it was not being released.

HPLC confirmed the loading of intact drug at ~76 %. Cell-matrix interactions were

assessed by fluorescent microscopy and SEM. In vitro studies were carried out on

human dermal fibroblasts (HDF) using electrospun PGMA scaffold as a control

matrix. It was hypothesised that cells may be more sensitive to the released drug and

that despite the lack of release data obtained there would be a detectable

physiological effect of the drug even at very low concentrations.

Cytotoxicity studies using the MTS assay showed no apparent toxicity of either

control or test scaffolds. Interestingly, no reduction was observed in cell

proliferation incubated on the scaffolds post TGFβ1 stimulation. Cell viability study

using calcein AM/ethidium bromide I dyes confirmed cell viability. Both florescent

and electron microscopy revealed positive interactions between cells and the fibrous

scaffolds. Cells were seen to take their extended spindle shape morphology as

confirmed by SEM images. RT-qPCR studies were carried out to study the effect of

the scaffolds on upregulated expression of collagen I post TGFβ1 stimulation. It was

observed that cells cultured on scaffolds containing drug (ES-PGMA + analogue 2)

showed a tendency for a reduced increase in expression of collagen I after TGFβ1

stimulation. However, this was not statistically significant. Therefore, it is likely that

the drug is restricted within the fibers and is not able to be released. It has previously

55

been reported that DMSO can facilitate drug release. It facilitates drug release by

providing a congenial environment for drug dissolution in an otherwise hydrophilic

physiological environment. When DMSO was incorporated in the culture, a

significant reduction in collagen I gene expression was attained as anticipated. No

significant effect was observed of DMSO alone or DMSO and control scaffold.

Results are presented in Agarwal, V., Wood, F. M., Fear, M. and Iyer, K. S.,

Inhibiting the activation of transforming growth factor-β using a polymeric

nanofiber scaffold, Nanoscale (Submitted)

56

Chapter 3

Series of papers

The results of this thesis are presented in the following series of published papers or

submitted manuscripts, the citations of which are listed below. Supporting

information for these manuscripts can be found in Appendix A, where applicable.

1. Agarwal, V., Ho, D., Ho, D., Galabura, Y., Yasin, F. M.D., Gong, P., Ye, W.,

Singh, R., Munshi, A., Saunders, M., Woodward, R. C., St. Pierre, T., Wood, F.M.,

Fear, M., Lorenser, D., Sampson, D. D., Zdyrko, B., Smith, N.M., Luzinov, I., Iyer,

K.S., A Functional Reactive Polymer Nanofiber Matrix, RSC Advances (Submitted)

2. Agarwal, V., Toshniwal, P., Smith, N. E., Smith, N. M., Li, B., Clemons, T.

D., Byrne, L. T., Hassiotou, F., Wood, F. M., Fear, M., Corry, B., and Iyer, K. S.,

Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor

Inhibitors by Intracellular Delivery, Angewandte Chemie International Edition

(Submitted)

3. Agarwal, V., Wood, F. M., Fear, M. and Iyer, K. S., Inhibiting the activation

of transforming growth factor-β using a polymeric nanofiber scaffold, Nanoscale

(Submitted)

Journal Name RSCPublishing

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Cite this: DOI: 10.1039/x0xx00000x

Received 00th January 2012,

Accepted 00th January 2012

DOI: 10.1039/x0xx00000x

www.rsc.org/

A Functional Reactive Polymer Nanofiber Matrix

Vipul Agarwal,a Dominic Ho,a Diwei Ho,a Yuriy Galabura,b Faizah M.D. Yasin,a Peijun Gong,c Weike Ye,d Ruhani Singh,a Alaa Munshi,a Martin Saunders,e Robert C. Woodward,f Timothy St. Pierre,f Fiona M. Wood,g Mark Fear,g Dirk Lorenser,c David D. Sampson,c,e Bogdan Zdyrkob, Nicole M. Smith,a Igor Luzinovb,*, and K. Swaminathan Iyera,*

Synthetic fractal materials have been regarded as a new class of hybrid materials with many potential applications. However, the lack of an efficient, reactive large-area fractal substrate has been one of the major limitations in the development of these materials as advanced functional platforms. Herein, we demonstrate the utility of electrospun polyglycidyl methacrylate (PGMA) fractal-like films as a highly versatile platform for the development of functional nanostructured fractal-like materials anchored to a surface. The utility of this platform as a reactive substrate is demonstrated by grafting poly (N-isopropyl acrylamide) to incorporate stimuli-responsive properties. Additionally, we demonstrate that functional fractal-like nanocomposites can be fabricated using this platform with properties for sensing, fluorescence imaging and magneto-responsiveness.

The development of nanostructured polymeric matrices to obtain organic-inorganic nanocomposites has been actively researched to produce hybrid materials for applications in electronics, optics, medical devices, sensors and catalysis.1-4 Of the various techniques developed to produce large area nanoscale polymeric matrices, one of the most researched, cost effective and facile method is electrospinning. It has been adapted to cover a wide range of polymers and optimized to regulate fiber diameter, alignment and shape.5-7 There have been numerous reports using this technique to develop matrices with enhanced mechanical strength,8 matrices with selective filtration/permeability,9 fire retarding material, optoelectronic devices10 and substrates for catalysis.11, 12 One of the key steps involved in the development of organic-inorganic nanocomposites is grafting to achieve excellent integration by minimizing interfacial tension of the nanoparticles in the

organic nanofiber matrix. Additionally, the ability to modify the surface of the nanofiber to alter the adhesion, lubrication, wettability and biocompatibility is pivotal in its customisation for end-use applications. The achievement of a certain degree of grafting universality requires the establishment of a controlled method of introducing the desired functional groups on a substrate.5 Currently, this is achieved by physisorption.13,

14 In contrast, chemisorption, which is difficult to achieve on a polymeric nanofibers, would result in permanent irreversible surface modification. Polymers containing epoxy groups are examples of functional polymers that are able to react with a wide range of substrates through ‘‘grafting to’’ interactions mediated by the epoxy groups.15, 16 The versatile chemistry of epoxy groups renders a polymer that is exceptionally suitable as a universal electrospun nanofiber matrix to provide reactive groups for further grafting reactions. To this end, poly(glycidyl methacrylate) (PGMA), which contains an epoxy group in every repeating unit, has been used extensively as a macromolecular anchoring layer for grafting of polymers to the surfaces.17-20 Upon electrospinning, epoxy groups in the polymer will undergo self-crosslinking upon heating, providing mechanical integrity to the matrix.21 Approximately 40% of the epoxy groups are still available for surface modification following a 12 hour treatment at 120°C.

The main advantage of using PGMA as a matrix for electrospinning, as opposed to modifying the surface using monolayers, is the high mobility of the epoxy groups located in the “loops” and “tails” of the polymer. The mobility of the free groups results in the formation of a highly effective interpenetrating anchoring zone.22 In this article, we report that PGMA can be directly electrospun (ES-PGMA) to form large area nanofibers. We demonstrate that this polymer nanofiber matrix can be used as an effective platform to graft polymers to impart switchability, and can be used to produce

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nanocomposites with upconverting properties, with hydrogen sensing capability or with magneto-responsive properties.

The ES-PGMA nanofibers (see Supporting Information for method of synthesis) were uniform over a large area and had an average diameter of 0.69 ± 0.04 µm (average ± standard error mean) (Fig. 1A). The average thickness of the 1 x 1 cm2 ES-PGMA generated was 127 ± 3 µm in 7 hours (Fig. 1B). In order to test the efficacy of the ES-PGMA nanofiber matrix as an anchoring platform, carboxylic acid-terminated poly (N-isopropyl acrylamide) (PNIPAM-COOH) was end grafted to the nanofibers via a ring opening reaction with the epoxy groups to yield ES-PGMA-g-PNIPAM-COOH.23 PNIPAM is a thermo-responsive polymer, which has been utilized in various

forms, such as thermo-responsive hydrogels, particles, brushes, spheres and micelles.24-27 Importantly, PNIPAM exhibits a

temperature-sensitive phase transition in water at what is known as a lower critical solution temperature (LCST), 32 °C.28

The transition is due to the coil-to-globule transition at the critical temperature resulting in switching from hydrophilic to hydrophobic behavior.29 At temperatures below the LCST, PNIPAM chains arrange into an expanded and hydrated conformation. Conversely, at temperatures above the LCST, PNIPAM chains collapse and arrange into a compact, dehydrated conformation.29 This thermo-responsive behavior is retained post-grafting and post-end group functionalization. The ES-PGMA-g-PNIPAM-COOH nanofibers demonstrated thermo-responsive behavior as monitored using contact angle measurements. The contact angle changed from 60 ± 2° at 70°C (Fig. 1C) to 15 ± 2° at room temperature (Fig. 1D). The contact angle of the unmodified ES-PGMA remained unchanged at 100 ± 2° both at room temperature and at 70°C. The ability of the ES-PGMA nanofiber matrix to produce nanocomposites was further evaluated using three distinct classes of nanoparticles: upconverting fluorescent particles of NaGdF4:Yb, Er (UCNP), palladium (Pd) and magnetite (Fe3O4). The nanoparticles synthesized (see Supporting Information for methods) had a narrow size distribution of 7.4 ± 1.4 nm (average ± standard error mean) for UCNP, 19.3 ± 0.2 nm for Pd and 6.7 ± 1.4 nm for Fe3O4 respectively (Fig. 2A, C and E). One of the major hurdles in developing functional materials by electrospinning nanocomposites is the lack of control in attaining a homogeneous distribution of nanoparticles throughout the polymer matrix. In the present case, electrospinning PGMA with the aforementioned nanoparticles resulted in relatively uniform distributions of the nanoparticles throughout the fiber matrix (Fig. 2B, D and F) which was observed through various images obtained at similar fields of view. It has been reported that variations in solution properties such as surface tension and solution conductivity in the presence of nanoparticles result in changes in the nanofiber diameter.30-32 In the present case, the electrospun fiber diameter increased in the presence of nanoparticles to 2.56 ± 0.16 µm (average ± standard error mean) for UCNP, 1.75 ± 0.07 µm for Pd and 4.37 ± 0.44 µm for Fe3O4 (Insets in Fig. 2G, H and I, respectively). Small fibers were chosen for TEM analysis because of the contrast problems related to thicker samples. The ability of the nanocomposites to be used as functional materials was evaluated by testing the upconverting properties, hydrogen sensing properties and magnetic properties of the, UCNP/ES-PGMA, Pd/ES-PGMA and Fe3O4/ES-PGMA fibers, respectively.

In the case of UCNP/ES-PGMA fibers, the ability to convert near-infrared excitation into visible emission was evaluated (see supporting information for methods). UCNP have been successfully used as ultrasensitive magnetic/upconversion fluorescent dual-modal molecular probes for MRI and upconversion fluorescence imaging.33-35 In the present case the UCNP/ES-PGMA fibers demonstrated excellent upconversion properties upon 974 nm laser excitation (Fig. 3A). The three major emissions were located at 521, 541, and 655 nm. Green emission from 500- 600 nm was attributed to 2H11/2 → 4I15/2 and 4S3/2 → 4I15/2 transitions, respectively, and the red emission from 635-670 nm was attributed to the 4F9/2 → 4I15/2 transition (Fig. 3A). The green to red (G/R) ratio for the fibers was 1.35:1. The UCNP/ES-PGMA composite fibers retained the upconversion signal levels to the pure NaGdF4:Yb,Er nanoparticle samples (G/R ratio = 1.37:1).

Fig. 1: (A) SEM secondary electron image of the electrospun PGMA (ES-PGMA) fibers, (B) cross-sectional image of ES-PGMA, (C) Water contact angle θ = 60° at 70°C, (D) Water contact angle θ = 15° at room temperature respectively measured on ES-PGMA -g-PNIPAM-COOH.

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The Pd/ES-PGMA nanofibers were evaluated for sensing hydrogen (see Supporting Information for methods). Palladium has emerged as an important candidate for hydrogen gas sensing because of its ability to absorb high quantities of hydrogen and its highly selective response.36 Sensing herein is based on the well-established principle that palladium spontaneously absorbs H2 gas as atomic hydrogen which diffuses into the lattice to form palladium hydride, PdHx, resulting in an to phase transition and a corresponding change in the lattice spacing.37 The change in phase and lattice spacing leads to a measurable resistance change of the palladium material. However, either replacing precious noble metals with cheaper materials or alternatively development of methods that result in the reduction of material used by several orders of magnitude, especially in applications that require large amounts of material, would be beneficial. Currently, hydrogen sensing platforms are based on all-palladium constructs or hybrids with high Pd loading to stimulate an effective sensing response. Herein, using the electrospun polymer/nanoparticle nanocomposite material we demonstrate a

response is obtainable for as low as 0.6 ng of Pd dispersed across a 650 m x 900 m area over interdigitated electrodes (IDE). The ability of Pd/ES-PGMA nanofibers to sense different hydrogen concentrations (between 1 and 10% in N2 as a carrier gas) was tested38 (Fig. 3B). An increase in resistance with hydrogen gas-flow and a return to the original state in the absence of a hydrogen gas flow was observed for hydrogen concentrations (1-10%) with a response time 90 of ~14 seconds (Fig. 3B).

Finally, the magnetization properties of the Fe3O4/ES-PGMA fibers were measured by SQUID magnetometry

Fig. 2: TEM images of the (A) Upconverting nanoparticles (UCNP), (B) UCNP/ES-PGMA composite fibers, (C) Pd nanoparticles, (D) Pd/ES-PGMA composite fibers, (E) magnetite (Fe3O4) nanoparticles, (F) Fe3O4/ES-PGMA composite fibers. X-ray microanalysis spectrum obtained on: (G) UCNP/ES-PGMA composite fibers showing the presence of Gd and Yb among other elements (Inset: SEM micrograph of UCNP/ES-PGMA composite fibers): (H) Pd/ES-PGMA composite fibers showing the presence of Pd (Inset: SEM micrograph of Pd/ES-PGMA composite fibers) and (I) Fe3O4/ES-PGMA composite fibers (Inset SEM micrographs of Fe3O4/ES-PGMA composite fibers). Scale bars for images A, C and E 10 nm, for images B, D and F 1 µm and for inset images G, H and I 20 µm

Fig. 3: (A) Upconversion fluorescence spectrum of both UCNP’s (blue) and UCNP/ES-PGMA fiber composite (black) showing three main emissions green at 521 and 541 nm and red between 635 and 670 nm upon 974 nm laser excitation, (B) Current response of the Pd/ES-PGMA matrix sensor to 1-10% hydrogen gas, with alternating 4 min hydrogen and 20 min. nitrogen exposure, (C) Zero-field cooled (orange) and field cooled (blue) curves for Fe3O4/ES-PGMA composite (Inset: hysteresis loop at 5 K (pink) and 300 K (green) for Fe3O4/ES-PGMA composite) as measured by SQUID magnetometry.

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4 | J. Name., 2012, 00, 1-3 This journal is © The Royal Society of Chemistry 2012

(Superconducting Quantum Interference Device) (see Supporting Information). The Fe3O4/ES-PGMA fibers are superparamagnetic at room temperature with the zero field cooled/field cooled curves showing a maximum blocking temperature of 30 K (where the two curves merge) and the absence of hysteresis at 300K (Fig. 3C). The mass specific saturation magnetization, Ms of the fibers was 4.0 emu g−1. Particle loading was estimated to be ~7% by weight as determined from the Ms values of the Fe3O4 nanoparticles and Fe3O4/ES-PGMA fibers.

Conclusions In summary, we have developed a robust polymeric platform for the large scale production of electrospun nanofibers based on poly(glycidyl methacrylate) (PGMA). We have demonstrated that the epoxy groups of the polymeric matrix can be effectively used as a grafting platform for surface modifications and the polymer serves as an excellent platform to fabricate functional nanocomposites. We believe our findings presented herein will aid in the design of novel electrospun materials with tailorable surfaces for applications as scaffolds in regenerative medicine, optoelectronics, magnetic filtration and catalysis. Notes and references a School of Chemistry and Biochemistry, The University of Western Australia, WA 6009, Australia b Department of Materials Science and Engineering, Clemson University, Clemson, SC 29634, USA c School of Electrical, Electronic and Computer Engineering, The University of Western Australia, WA 6009 Australia d School of Chemistry and Chemical Engineering, Nanjing University, China e Centre for Microscopy, Characterisation and Analysis, The University of Western Australia, WA 6009 Australia f School of Physics, The University of Western Australia, WA 6009 Australia g Burn Injury Research Unit, School of Surgery, The University of Western Australia, WA 6009, Australia E-mail: [email protected], [email protected] The authors would like to acknowledge the Australian Microscopy & Microanalysis Research Facility at the Centre for Microscopy, Characterization & Analysis, The University of Western Australia, funded by the University, State and Commonwealth Governments. Authors would also like to thank Dr C.W.Evans and Dr Peter R. T. Munro for assistance with valuable experimental and result discussions. Peijun Gong is supported by The University of Western Australia and the China Scholarship Council. † Footnotes should appear here. These might include comments relevant to but not central to the matter under discussion, limited experimental and spectral data, and crystallographic data. Electronic Supplementary Information (ESI) available: [details of materials and methods]. See DOI: 10.1039/c000000x/ 1. S. Komarneni, J. Mater. Chem., 1992, 2, 1219-1230. 2. D. R. Paul and L. M. Robeson, Polymer, 2008, 49, 3187-3204.

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1

Intracellular Drug Delivery

DOI: 10.1002/anie.201((will be filled in by the editorial staff))

Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor Inhibitors by Intracellular Delivery** Vipul Agarwal, Priyanka Toshniwal,# Natalie E Smith,# Nicole M Smith, Binbin Li, Tristan D Clemons, Lindsay T Byrne, Foteini Hassiotou, Fiona M Wood, Mark Fear, Ben Corry* and K Swaminathan Iyer*

Abstract: Extracellular targeting of the cation-independent mannose 6-phosphate/insulin-like growth factor II (M6P/IGFII) receptors has been an attractive approach for the development of antifibrotic drugs. Several M6P analogues have been developed to regulate the activity of transforming growth factor-β1 (TGFβ1) by inhibiting its conversion from the latent to the active form. Herein, we adopt a combinatorial approach using an in vitro wound healing model and molecular dynamic simulations, to reveal that the efficacy of M6P/IGFII inhibitors can be significantly enhanced by adopting an intracellular approach. We demonstate this using systematic analysis of a bioisosteric M6P analogue, a lipophilic prodrug which upon cellular internalization undergoes ester hydrolysis to yield an

active M6P analogue to effectively downregulate collagen 1 expression in primary human dermal skin fibroblasts. In mammalian cells, the cation-independent mannose 6-phosphate/insulin-like growth factor II (M6P/IGFII) and cation-dependent mannose 6-phosphate (CD-MPR) receptors, have been identified as pivotal targets that modulate cellular response because of their role in protein trafficking. Both these receptors are functionally complimentary and can partially compensate for the absence of the other.[1] These sorting receptors play an important role of transporting M6P-bearing glycoproteins from the trans-Golgi network (TGN) to lysosomes mediated through their M6P binding sites.[2] Both receptors transport important enzymes to the intracellular acidic pre-lysosomal compartments where low pH leads to the release of the enzymes from the complex. The receptor then gets recycled into the Golgi apparatus.[3] However, only the M6P/IGFII receptor is anchored to the cell surface membrane and has been implicated in the internalization of M6P bearing compounds.[4] Importantly, it modulates the activity of a variety of extracellular M6P bearing glycoproteins including latent transforming growth factor-β (LTGFβ) precursor, urokinase-type plasminogen activator receptor, glycoprotein D of the herpes virus, granzyme B an essential factor for T cell-mediated apoptosis and proliferin.[4] This has resulted in an enormous interest in the design of M6P bearing compounds that target the M6P/IGFII receptor as it offers an efficient means for internalization of high specificity therapeutics.[5] This approach has been used to deliver therapeutic compounds in enzyme replacement therapies in lysosomal diseases like Fabry disease, aid wound healing, as a treatment for breast cancer, and to combat viral infections.[4] However, the approach suffers a major drawback as the phosphomonoester bond of M6P is prone to hydrolysis by various phosphatase enzymes.[6] This dramatically reduces its binding efficiency to the receptor thereby compromising its potency. This problem has been circumvented by

Figure 1: a) Chemical structure of mannose-6-phosphate (M6P and the two analogues, b) Schematic representation of the cLogP of three compounds

[] V. Agarwal, P. Toshniwal, Dr. N.M. Smith, Dr. T.D. Clemons, Dr. F. Hassiotou, Prof. Dr. K. S. Iyer School of Chemistry and Biochemistry, The University of Western Australia, Crawley, Western Australia, Australia E-mail: [email protected]

B. Li State Key Laboratory of Advanced Technology for Materials

Synthesis and Processing, Wuhan University of Technology, Wuhan, PR China

Dr. L. T. Byrne Centre for Microscopy, Characterization & Analysis, The University of Western Australia, Australia

Dr. N. E. Smith, Dr. B. Corry Research School of Biology, Australian National University, Canberra, Australian Capital Territory, Australia

Email: [email protected] Dr. M. Fear, Prof. Dr. F. M. Wood

Burn Injury Research Unit, School of Surgery, The University of Western Australia, Crawley, Western Australia, Australia

[] The authors would like to thank the Australian Research Council for funding the work, Pharmaxis Pvt Ltd for kindly providing analogues 1 and 2. The authors would also like to acknowledge the Australian Microscopy & Microanalysis Research Facility at the Centre for Microscopy, Characterization & Analysis, The University of Western Australia, funded by the University, State and Commonwealth Governments. Authors would like to thank Dr Bernard Callus, Dr Cameron Evans and Dr Megan Finch for their assistance with experimental analysis.

[#] These authors contributed equally

Supporting information for this article is available on the WWW under http://dx.doi.org/10.1002/anie.201xxxxxx.

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the design of several isosteric M6P analogues with phosphonate, carboxylate or malonate groups, which have higher affinity to the receptor and a stronger stability in human serum than M6P.[7] This approach is successful in overcoming the issues with hydrolysis of the phosphomonoester bond, yet falls short as these analogues can only target the receptors present on the cell surface. In the steady state, ~90 % of the M6P/IGFII receptors are localized in the transmembrane compartments while the remainder stays on the cell surface.[8] The receptor has a relatively long half-life (t1/2 ~ 20 hours) and recycles between the trans-Golgi network, endosomes and the plasma membrane.[9] In this communication, we report a novel approach to improve ligand-receptor protein interaction in cells whilst overcoming stability issues associated with M6P. We demonstrate this by exploring a prodrug (analogue 2) that undergoes intracellular chemical modification by esterases to yield an active M6P analogue (analogue 1) (Figure 1a),[10] resulting in a sustained and focused therapeutic strategy in an in vitro model of wound healing.

Table 1. Ligand-Receptor Protein interaction energies obtained for M6P and each of the two analogues in domain 3 and domain 5 as determined from 100 ns of molecular dynamics simulation. Two ligands were placed into the dimer binding pocket, because the receptor is secreted as a dimer.

Domain 3 Ligand-Receptor Protein interaction Energy (kcal/mol)

M6P Analogue 1 Analogue 2

Ligand 1 -368.4 -309.3 -81.6

Ligand 2 -347.4 -304.3 -79.6

Domain 5 Ligand-Receptor Protein interaction Energy (kcal/mol)

M6P Analogue 1 Analogue 2

Ligand 1 -128.2 -44.1 -43.6

Ligand 2 -118.3 -74.0 -47.8

The design of phosphonate analogue 1 is based on established principles of bioisosteric M6P analogues by replacing the P-O bond at C6 by a methylene bridge. Moreover, the replacement of the hydroxyl group at the anomeric position by an aromatic subtituent slightly improves recognition by the M6P/IGFII receptor.[11] This could be due to the hydrophobic interactions between the aromatic moiety of analogue 1 and the binding pocket of the M6P/IGFII receptor. Previous studies have demonstrated that neutral ester prodrugs are relatively benign towards enzymatic degradation, thereby altering their apparent elimination and half-life.[12] Hence analogue 2 was designed by masking analogue 1 via esterification of the phosphate group to yield a non-charged bis(pivaloyloxymethyl) (POM) derivative. Importantly, derivatization of phosphates decreases the polarity of the parent drug thereby promoting its cellular internalization and altering the elimination/distribution mechanism.[12-13] Notably, the clogP (calculated-logP evaluated using Chem Draw) for M6P, analogue 1 and 2 are -3.28, 0.10 and 3.29 respectively (Figure 1b). LogP is an estimate of a compound's overall lipophilicity, a value that influences its physiological properties such as solubility, permeability through biological membranes, hepatic clearance, and non-specific toxicity.[14] Polar compounds with low logP have very low cellular permeability due to their low affinity for the lipid bilayers. Alternatively, lipophilic compounds with high logP have high affinity for the phospholipid phase facilitating their internalization and prohibiting their escape into the aqueous basolateral side.[14] Herein the lipophilic prodrug,

analogue 2, will have improved cellular internalization compared to its charged parent analogue, 1. Once internalized the bis(pivaloyloxymethyl) linkers of analogue 2 will be gradually prone to ester hydrolysis by microsomal esterases present within the intracellular compartments,[15] resulting in the conversion to the charged parent analogue, analogue 1. Analogue 1 on the contrary would only target extracellular M6P/IGFII receptors, when administered directly, due to its low cellular permeability deemed to its low logP value. The extracellular region of the M6P/IGFII receptor is comprised of 15 repetitive domains and contains three distinct M6P binding sites located in domains 3, 5, and 9, with only domain 5 exhibiting preference for phosphodiesters.[16] In order to assess our strategy to use the intracellular conversion of the produg analogue 2 to a high receptor binding phosphonate analogue 1, it is pivotal to examine the ligand-receptor interactions to validate the hypothesis that analogue 2 will have minimal interaction with the extracellular receptors. In the current study, we used six independent molecular dynamics simulations to study the ligand-receptor protein interactions of M6P, analogues 1 and 2 with domains 3 and 5 of the extracellular M6P/IGFII receptor (see Supporting Information for experimental details, section S8.1). Domain structures were adopted from previously reported studies and two ligands were placed into the dimer binding pocket, because the receptor is secreted as a dimer.[17] Analogue 1 showed similar ligand-receptor protein interaction energies to M6P in domain 3 (Table 1). Importantly, the m-xylene ring of analogue 1 was positioned in the middle of the binding pocket further stabilizing the binding of this compound in comparison to M6P (see Supporting Information, Figures S1 and S2). This is in accordance with the previous studies of other phosphonate analogues of M6P, which are reported to display higher affinity and stronger stability in human serum than M6P.[6, 7c] The domain 5 binding pocket is larger than in domain 3, hence all the

Figure 2: Cell viability assays showing percentage of live cells in the culture post incubation with M6P, analogue 1 and 2. First and second column in each condition is representing 24 h and 72 h respectively. Data presented as average ± SEM (n=4). Significance was set at * p < 0.05 using bonferroni post-hoc test in one way ANOVA analysis.

Figure 3: Cell body area showing change in cell area post TGFβ1 stimulation and subsequent analogues treatment. Cell area was measured from the fluorescent images of live cells taken for viability assay (cells from minimum 40 images per group were measured). Significant increase in cell body area was observed for cells treated with TGFβ1 (2 ng/mL), however no such increase was observed in cells treated with analogues +/- TGFβ1 (2 ng/mL). Data presented as average ± SEM (n > 40). Significance was set at * p < 0.05 using bonferroni test in one way ANOVA.

3

compounds displayed weaker interactions with the receptor and occupied more diverse positions in domain 5 due to the increased space (see Supporting Information, Figure S2). Furthermore, in the case of analogue 1 in domain 5, the simulations suggested that one of the two analogue 1 ligands (ligand 1) bound to the protein dimer has weaker interactions with the protein as it primarily interacts with the second molecule of analogue 1 (ligand 2). Overall, the simulations suggested that analogue 1 has high affinity towards domain 3 similar to M6P whilst the prodrug 2 has weak interactions with both domains of the receptor (Table 1 and see Supporting Information, Figure S3). The molecular dynamics simulations further validated our aforementioned hypothesis that the prodrug will be internalized with minimal extracellular receptor-ligand interactions. We next validated our hypothesis in a well-established in vitro model for wound healing using primary human dermal skin fibroblasts (HDF). In mammals, wound healing is not a regenerative process that restores normal tissue architecture, but a reparative process that results in scar formation.[18] This process occurs in all tissues of the body in response to physical, chemical and biological stressors. Scar tissue is functionally and aesthetically inferior to normal tissue. It is a result of the excessive production of extracellular matrix (ECM) that occurs after injury.[19] One of the most important proteins influencing the ECM architecture during wound healing is collagen I. Collagen I is synthesized predominantly by fibroblasts and its synthesis is largely regulated by cytokine transforming growth factor β1 (TGFβ1).[20] TGFβ1 is secreted in an inactive form (LTGFβ1), requiring enzymatic conversion to active TGFβ1 to effect a change in cell function. One of the methods of TGFβ1 activation involves binding of M6P residues within the N-linked oligosaccharides on latent TGFβ1 to the M6P/IGFII receptor.[21] Since the M6P binding sites are involved in various steps of TGFβ1 activation and inactivation, it is believed that small molecule inhibitors that block the binding of M6P residues could present an opportunity to block the activity of TGFβ thereby reducing overproduction of an important profibrotic extracellular matrix protein collagen I. Cytotoxicity and cell viability of the analogues were initially assessed using MTS and live/dead assays (see experimental details in Supporting Information, sections S2.1, S3.1 and S4.1). Previous studies characterizing M6P binding affinity towards the M6P/IGFII receptor reported significant binding affinity at a concentration of 10 µM.[7b, 22] This concentration was therefore selected for our in vitro studies. All compounds showed no effect on cell viability and proliferation both in the presence and absence of TGFβ1 (Figure 2 and see Supporting Information, Figure S4 respectively). Exposure to TGFβ1 in the absence of analogues 1 and 2 resulted in a reduction in HDF proliferation (see Supporting Information, Figure S4). This growth suppressive response has been previously reported in many cell types.[23] The observed change in cell proliferation upon exposure to TGFβ1 influenced HDF cell morphology (and see Supporting Information, Figure S5b). Fibroblasts alter their morphology from dendritic to stellate upon exposure to various external cues caused by changes in actin polarisation and focal adhesion.[24] TGFβ1 has been shown to alter the morphology of many cell types including fibroblasts, potentially by inducing polymerisation of the actin cytoskeleton from globular to filamentous.[24a] Different factors such as cell motility and mechanical strain have also been reported to cause this alteration.[25] In the present case, we observed a reversal of HDF cell morphology back to initial cell morphology without TGFβ1 stimulation when treated with the analogues 1 and 2 (Figure 3 for quantification of cell body area and see Supporting Information, Figure S5c-e for

images). We next assessed if the observed change in morphology is correlated to collagen I gene expression using qRT-PCR, and if changes in collagen I gene expression could be altered by inhibition of TGFβ1 activity by targeting the M6P/IGFII receptor in the presence of the analogues (refer to Supporting Information for method, section S5.1). Indeed as previously reported, exposure of HDF to TGFβ1 (2 ng/mL) resulted in a significant increase in collagen I mRNA expression at 48 hours post-stimulation.[20, 26] It is noteworthy that although collagen I gene expression was upregulated throughout the study period (72 hours), the optimal response was observed after 48 hours exposure to TGFβ1 (see Supporting Information, Figure S6). Therefore, the efficacy of the aformentioned compounds was assessed in the presence of TGFβ1 at 48 hours. TGFβ1 induced collagen I mRNA expression was downregulated significantly (p < 0.05) with the addition of prodrug analogue 2 (10 µM) with levels returning to that of normal untreated cells (Figure 4a). Downregulation was also observed for M6P however, the change did not reach stastistical significance (Figure 4a). This suggests that the variable responses that have been reported in the use of hydrolytically unstable M6P may be due to its realtive instability and that the development of stable analogues may resolve this issue. Importantly, in the present case we observed no significant change in collagen I gene expression in HDF cells treated with analogue 1 (Figure 4a). This was expected given the low cellular permeability which is believed to affect the ligand-receptor protein interactions in cells. Next, we investigated if the observed change at transcription level would have a corresponding influence on protein translation. Changes in collagen I protein expression were quantified using immunoblotting (refer to Supporting Information for method, section S6.1). All protein expression studies were carried out at 72 h post-stimulation. Significant upregulation in collagen I protein expression was observed post TGFβ1 stimulation (Figure 4b; column 2; p < 0.05) which is consistent with previous reports.[27] Analogue 2 (10 µM) was observed to reduce TGFβ1

Figure 4: Change in Collagen 1 a) mRNA levels and b) protein levels post TGFβ1 stimulation in the presence and absennce of M6P, analogues 1 and 2 compared to untreated (negative) control. Collagen I protein expression was normalised against β-actin levels. Data are presented as average ± SEM (n = 3). Significance was set at * p < 0.05 using bonferroni post-hoc test in one way ANOVA analysis

4

mediated upregulation of collagen I protein to non-stimulated levels (Figure 4b; column 5; p < 0.05). No significant changes were observed in the case of M6P or analogue 1. This further confirms that analogue 2 is a potent repressor of TGFβ1 induced collagen I synthesis and thus can ameliorate the profibrotic effects of TGFβ1 in human skin dermal fibroblasts. In summary, we have developed a novel approach using an intracellular prodrug of M6P, analogue 2, to target M6P receptors. This approach overcomes the physiological problems associated with the hydrolysis of M6P whilst successfully targeting the receptors using an intracellular coversion of the analogue. We believe that this approach of intracellular drug coversion for receptor targeting will have far reaching implications in the design of highly potent drug candidates for enzyme replacement therapies of lysosomal storage diseases, to aid wound healing and in cancer therapy.

Received: ((will be filled in by the editorial staff)) Published online on ((will be filled in by the editorial staff))

Keywords: drug development • intracellular drug delivery • Mannose-6-Phosphate analogue

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Entry for the Table of Contents (Please choose one layout) Layout 2:

Intracellular Drug Delivery

Vipul Agarwal, Priyanka Toshniwal,# Natalie E Smith,# Nicole M Smith, Binbin Li, Tristan D Clemons, Lindsay T Byrne, Foteini Hassiotou, Fiona M Wood, Mark Fear, Ben Corry* and K Swaminathan Iyer* __________ Page – Page

Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor Inhibitors by Intracellular Delivery Intracellular delivery of M6P/IGFII receptor inhibitors exhibits better efficacy than

extracellular inhibitors to regulate TGFβ1 mediated upregulation of profibrotic marker, collagen I.

Journal Name RSCPublishing

COMMUNICATION

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Cite this: DOI: 10.1039/x0xx00000x

Received 00th January 2012,

Accepted 00th January 2012

DOI: 10.1039/x0xx00000x

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Inhibiting the activation of transforming growth factor-β using a polymeric nanofiber scaffold

Vipul Agarwal,a Fiona M. Wood,b,c Mark Fearb and K. Swaminathan Iyer*a

Electrospun poly(glycidyl-methacrylate) (PGMA) nanofibers were fabricated in the presence of a hydrophobic analogue of mannose-6-phosphate, (PXS64). The nanofibers were tested for biocompatibility as a tissue engineering scaffold and for their efficacy to inhibit transforming growth factor β1 (TGFβ1) activation in human dermal skin fibroblasts.

The principle of tissue engineering involves the development of advanced functional biomaterials that incorporate biochemical cues to aid in the regeneration of tissues to restore and maintain normal organ function.1 This approach has resulted in the fabrication of several advanced biomaterial platforms that have been used to repair damaged or diseased tissues and to create therapeutic approaches for entire tissue replacement.2-4 While tissue regeneration has successfully been demonstrated in the presence of biocompatible scaffolds, scarring remains one of the unresolved issues. In mammals postnatal wound healing results in scar formation, characterised by excessive collagen deposition and dysfunctional extracellular matrix formation. This is also a hallmark of fibrotic disease which occurs in many tissues. Central to wound repair is transforming growth factor β1 (TGFβ1), a cytokine secreted by several different cell types involved in wound healing.5 In the case of skin, scarring during the wound healing process is a result of TGFβ1 mediated imbalance in the fibroblast activity resulting in an architecturally disorganised extracellular matrix.6 One of the most important proteins influencing the ECM architecture during wound healing is collagen I. Collagen I is synthesized predominantly by fibroblasts and its synthesis is largely regulated by TGFβ1 signalling.7 Following wound healing in skin, a scar is not only aesthetically and psychologically detrimental but can also cause functional disability and pain.8 Importantly, inhibition of TGFβ1 activity has been documented to reduce scar formation9 whereas subcutaneous delivery of exogenous TGFβ1 in newborn mice was shown to promote fibrosis and angiogenesis at the site of injection.10 During wound healing TGFβ1 activation from its latent state involves binding of the mannose 6-phosphate (M6P) residues within

the N-linked oligosaccharides on the latent TGFβ to the M6P/IGFII receptor.11 Since the M6P binding sites are involved in various steps of the TGFβ1 activation and inactivation route, it is believed that the presence of small molecule analogues of M6P that competitively bind to the receptor could present an opportunity to block the activation of TGFβ1 thereby reducing overproduction of extracellular matrix protein collagen and potentially reducing scarring.12 The major drawback with using M6P as a drug in the reduction of scarring is that the phosphomonoester bond of M6P is prone to hydrolysis by various phosphatase enzymes.13 Isosteric M6P analogues with phosphonate, carboxylate or malonate groups have been shown to circumvent the aforementioned issues and have been reported to have greater stability in human serum than M6P.14-17 The analogue PXS64 reported in the present study is a lipophilic bioisosteric phosphonate analogue developed by [(bis(pivaloyloxymethyl)) (POM)] ester derivatization of M6P. Importantly, the high lipophilicity of PXS64 limits its solubility in aqueous solutions thereby reducing its bioavailability.18 In this communication we report that PXS64 can be incorporated in an electrospun poly(glycidyl methacrylate) (PGMA) nanofibrous scaffold. Furthermore, we demonstrate the utility of this scaffold as a biomaterial platform for wound healing using human dermal skin fibroblasts (HDF). We demonstrate its effectiveness as a drug delivery platform to mitigate TGFβ1 mediated upregulation of collagen I. Electrospinning is a widely used technique to fabricate large area nanofibrous scaffolds19 mimicking the architecture of extracellular matrix (ECM).20 They have been used as tissue engineering scaffolds to promote cell growth and migration and to achieve controlled delivery of drugs and growth factors.21 They have been widely used in skin, bone, cartilage, vascular and neural tissue engineering.22 For example, Yang et al. employed emulsion electrospinning to fabricate ultrafine core sheath poly(ethylene glycol)-poly(D,L-lactide) fibers loaded with basic fibroblast growth factor (bFGF) and demonstrated its gradual release over 4 weeks.23 In vitro studies on mouse embryonic fibroblasts showed enhanced cell adhesion, proliferation and secretion of ECM when cultured on the nanofibrous scaffold.24

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In the in vivo studies on dorsal wound model in diabetic rats, bFGF/ poly(ethylene glycol)-poly(D,L-lactide) mats showed elevated healing with complete re-epithelialization and regeneration of skin appendages such as hair and sebum, while bFGF promoted collagen deposition and its remodelling similar to normal architecture.23 In the case of electrospun scaffolds the ability to modify the surface properties of the nanofibers such as adhesion, wettability and biocompatibility is pivotal in its integration as a tissue engineering platform. Currently one of the biggest challenges using this technique to produce scaffolds is that the polymers used thus far to develop nanofibers need specific approaches for surface modification that are complex and laborious.25 The achievement of a certain degree of grafting universality requires the establishment of a controlled method of introducing the desired functional groups. Currently, this is achieved by physisorption.26 In contrast, chemisorption which is difficult to achieve on polymeric nanofibers would result in covalent attachment. Polymers containing epoxy groups are examples of functional polymers that are able to react with a wide range of biologically relevant molecules through

‘‘grafting to’’ interactions mediated by the epoxy groups.27, 28 To this end, poly(glycidyl methacrylate) (PGMA) used in the current study, has an epoxy group in every repeating unit and has been used extensively as a macromolecular anchoring layer.29, 30 In the present study, PGMA was electrospun in the presence of PXS64 and in the absence of PXS64 as a control (see supporting information for experimental conditions). PXS64 loading in the electrospun PGMA fibers was measured using high pressure liquid chromatography (HPLC) to be ~76 % w/w. Electrospun fibrous scaffolds were characterized using scanning electron microscopy (SEM). Fiber diameter was measured from the SEM images to be 0.69 ± 0.31 µm (average ± standard deviation) for ES-PGMA and 2.24 ± 1.06 µm in the case of ES-PGMA + PXS64 (Figure 1). Biocompatibility of the scaffold was investigated using the colorimetric MTS assay and live/dead cell viability assay (see Supporting Information for method) and were imaged using fluorescence and scanning electron microscopy.31 Human dermal fibroblasts (HDF) cells were cultured on the scaffolds and incubated

Figure 1: SEM secondary electron image of electrospun fibers of a) PGMA, b) PGMA + PXS64, C) showing the chemical structure of PXS64. Scale bars: a) and b) 10 µm.

Figure 2: a) Cell viability assay showing percentage of live cells in culture post incubation on ES-PGMA and ES-PGMA + PXS64, in the presence and absence of TGFβ1 (2 ng/mL), b) and c) showing HDF cell morphology on ES-PGMA + PXS64 + TGFβ1 (2 ng/mL) using fluorescent microscopy and scanning electron microscopy respectively. Red arrows highlighting the cell attachment and adhere points on the fibers. Scale bars: b) and c) 2 µm.

Figure 3: Collagen I gene expression analysis showing the percentage change in gene expression as compared to non-treated negative control (column 1). HDF cells incubated on both scaffolds and plastic tissue culture plate both in presence and absence of TGFβ1 and release media. Data presented as average ± SEM (n = 3). Significance was set at * p < 0.05 using bonferroni post-hoc test in one way ANOVA analysis.

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both with and without TGFβ1 over a period of 72 hours. No significant changes were observed on cell proliferation (see Supporting Information; Figure S1) or cell viability (Figure 2a) on both scaffolds with or without TGFβ1. To study the cell-matrix interactions, scaffolds were incubated with HDF cells and imaged using fluorescence and electron microscopy (see Supporting Information for method). HDF cells were observed to adopt their characteristic spindle shape morphology on both ES-PGMA and ES-PGMA + PXS64 scaffolds in the presence and absence of TGFβ1 (Figure 2b, c and see Supporting Information; Figure S2 and S3). Finally, the ability of ES-PGMA + PXS64 scaffold in regulating the over expression of collagen I gene in HDF cells post TGFβ1 stimulation was evaluated using RT-qPCR (see Supporting Information). HDF cells cultured on the biocompatible ES-PGMA control scaffold showed significant increase in collagen I gene expression in the presence of TGFβ1 (Figure 3). Cells incubated on ES-PGMA + PXS64 scaffold under similar conditions showed reduced expression of collagen I gene compared to ES-PGMA scaffold, which reached significance in the presence of supplemented release media. The use of DMSO in release media was adopted from previously reported methodology32 to accelerate the release of hydrophobic drugs in vitro and was shown to have no effect on collagen expression alone (Figure 3). Here we developed a PGMA scaffold and encapsulated an anti-scarring drug, PXS64. The biocompatibility of the scaffold was demonstrated in human dermal skin fibroblasts. Finally, the efficacy of scaffold and drug in mitigating the increased expression of collagen I in response to TGFβ1 stimulation was also demonstrated. We believe that this potential proof of principle approach can easily be adapted in the design of scaffolds for tissue regeneration in the presence of other anti-fibrotic agents. The authors would like to thank the Australian Research Council for funding the work. The authors would also like to acknowledge the Australian Microscopy & Microanalysis Research Facility at the Centre for Microscopy, Characterization & Analysis, The University of Western Australia, funded by the University, State and Commonwealth Governments. The authors would like to thank Dr Foteini Hassiotou to kindly provide the RT-qPCR instrument and Pharmaxis Ltd, Sydney for providing PXS64. MF is supported by Chevron Australia. This work was partly funded by the Fiona Wood Foundation. Notes and references a School of Chemistry and Biochemistry, The University of Western Australia, Crawley, Western Australia, Australia. E-mail: [email protected] b Burn Injury Research Unit, School of Surgery, The University of Western Australia, Crawley, Western Australia, Australia. c Burns Service Western Australia, Department of Health, Perth, Western Australia, Australia † Footnotes should appear here. These might include comments relevant to but not central to the matter under discussion, limited experimental and spectral data, and crystallographic data. Electronic Supplementary Information (ESI) available: [details of any supplementary information available should be included here]. See DOI: 10.1039/c000000x/ 1. T. Dvir, B. P. Timko, D. S. Kohane and R. Langer, Nat.

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Chapter 4

Conclusion and Future Works

In this thesis two complementary approaches were investigated to address the

problem of scarring after injury. The first part of the work focused on a tissue

engineering approach towards promotion of skin regeneration and subsequent drug

delivery application using a nanoscaffold while the latter part focused on the use of

novel anti-scarring therapeutics. The tissue engineering approach involved the

development of a multifunctional scaffold which could have a number of different

applications in addition to a potential drug-delivery scaffold including upconversion

probes, magneto-responsive material and biosensor. Subsequently, two analogues of

the natural anti-scarring molecule mannose-6-phosphate (analogue 1 and 2) were

investigated as potential scar modulators. Finally, the two concepts were blended to

fabricate a nanofibrous scaffold for the delivery of analogue 2 in scar therapy. In this

section, a summary of the work detailed in the publications will be presented.

4.1 Tissue engineering of a nanoscaffold

Tissue engineering has evolved as an innovative approach towards tissue

regeneration which usually relies on a platform to drive the cellular response.

Electrospinning was employed to fabricate a stable nanofibrous universal grafting

substrate using a polyglycidyl methacrylate (ES-PGMA) polymer. PGMA is an

interesting polymer because of the presence of an ester linkage and highly reactive

epoxy ring in each monomer unit. Presence of the epoxy ring provides a thermally

induced self-crosslinking capability to the scaffold which eliminates the need for

external chemical or physical crosslinking methods.

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Multi-functionality of ES-PGMA was demonstrated first by end grafting it with

carboxy-terminated poly N-isopropyl acrylamide (PNIPAM-COOH) to design a

thermo-responsive scaffold. Contact angle measurements confirmed the surface

grafting of ES-PGMA with PNIPAM-COOH which was demonstrated to undergo

phase transition at the glass transition temperature of PNIPAM. Next, ES-

PGMA/nanoparticle composite scaffolds using three different types of nanoparticles

((NaGdF4:Yb, Er); Pd and Fe3O4) were generated using the electrospinning protocol

established above.

Nanoparticle loading in the case of hybrid scaffolds was confirmed by x-ray

microanalysis, TEM, and ICP and AAS analysis. SEM analysis however revealed

an increase in the fiber diameters in the hybrid materials. Changes in the solvent

conductivity and surface tension have been claimed to have major influence on fiber

diameter.

One of the major problems encountered during the electrospinning process caused

by the low vapour pressure of solvent used was the coagulation and drying of

Taylor’s cone at the tip of the needle which required periodic cleaning and resulted

in loss of material. Solvents with relatively higher vapour pressure such as THF can

be used to improve the electrospinning process and address the above mentioned

problem.

Applicability of these hybrid materials were examined in gas sensing for Pd,

upconversion imaging in the case of upconverting nanoparticles and magneto-

responsiveness for magnetite nanoparticles. Optimisation of the hydrogen gas

sensing measurement was difficult as the Pd concentration on the surface of the

matrix was unknown. Fiber porosity also limited the sensitivity to high

concentrations of hydrogen. Potentially, surface functionalization of the scaffold

with Pd nanoparticles, in addition to encapsulation, could be attempted to elevate the

sensitivity of the system towards lower concentrations of hydrogen gas. SQUID

analysis on the magnetite/polymer composite showed superparamagnetic behaviour

of the matrix making it ideal for applications in protective fabric material and

currency notes.304

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4.2 Scar therapy and mannose-6-phosphate analogues

The M6P/IGFII receptor is important in trafficking of cargoes from the cell surface

and trans-golgi network to lysosomes.305-307 It plays an important role in various

diseases including fibrosis and lysosomal storage disorders.307 The trafficking

function of the receptor is exploited mainly for enzyme replacement therapy but has

not been a focus of anti-fibrotic therapeutic approaches 308

M6P is a natural occurring ligand of the M6P/IGFII receptor found in a number of

glycoproteins. The M6P/IGFII receptor consists of 15 repeating domains out of

which domains 3, 5 and 9 are specific towards M6P binding. With the emergence of

the anti-scarring properties of M6P309 and the structural determination of the

M6P/IGFII receptor,305 a great deal of research has focused on the development of

more efficient analogues of M6P.155, 156 This is because the therapeutic potential of

M6P is limited because of its low metabolic stability against phosphatases and

difficulty in retaining high concentrations at the injury site. Research has mainly

focussed on manipulating the structure of M6P in order to enhance its binding

efficiency towards the receptor while improving metabolic stability.157, 310, 311

Bioisosteric phosphonate analogues have been shown to have greater serum stability

than M6P. Taking advantage of this, two bioisosteric phosphonate analogues, 1 and

2, were explored for their potential as anti-scarring agents in this thesis. Analogue 1

was rationalised to have higher metabolic stability than M6P, thereby enhancing its

bioavailability, without lowering its affinity for the receptor. Molecular dynamics

simulation studies confirmed that analogue 1 has a similar interaction towards the

receptor as M6P (carried out in collaboration with Corry group). However,

intercellular delivery of analogue 1 showed no significant response in impeding the

expression of collagen I gene, the model used to study anti-scarring potential of the

analogues.

Neutral lipophilic phosphate compounds have been reported to undergo rapid

internalisation increasing their local concentration within the cellular

microenvironment.312, 313 Therefore, analogue 2 was designed as a neutral molecule

with high logP to achieve intracellular drug delivery. Analogue 1 was masked by

esterification of the phosphate group to yield a non-charged bis(pivaloyloxymethyl)

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derivative, analogue 2. Upon internalisation, phosphoester linkers undergo ester

hydrolysis by esterases present within the cellular microenvironment and release

analogue 1, allowing intracellular targeting of the M6P/IGFII receptor. Gene

expression studies showed significant reduction in collagen I expression, similar to

treatment with M6P, upon analogue 2 treatment. Therefore, it was concluded from

this part of the study that intracellular targeting of the receptor is effective using

analogue 2 as a prodrug of analogue 1.

One of the major problems encountered was in the analysis of protein expression.

Analysis of collagen I and fibronectin expression was attempted to supplement the

data on RNA expression levels. Immunoblotting was unreliable and had low

reproducibility, most likely in part due to the large molecular weights of the target

proteins and ineffective protein transfer. Alternative techniques such as HPLC and a

recently developed ‘scar-in-jar’ protocol may be more effective methods for the

quantitation of protein expression, especially for collagen I.314, 315

Despite the positive response in reducing the effects of TGFβ1, analogue 2 has

limitations. It is vulnerable to metabolic degradation mediated by esterases present

in growth serum. In addition, one of the hydrolysis products is formaldehyde, which

is toxic to cells at high concentrations limiting the dose that can be administered.

4.3 Scaffold based delivery of analogue 2

To achieve sustained delivery of the effective analogue 2, the electrospun PGMA

nanofibrous platform was manipulated for a drug delivery application. It was

anticipated that enzymatic vulnerability of analogue 2 could be subdued by

encapsulating it within the polymer scaffold by limiting its direct exposure to serum.

Using the parameters established above, analogue 2 was electrospun with PGMA to

yield a fibrous drug delivery platform with loading efficiency of ~76 % as measured

by HPLC. Despite the high drug loading its subsequent release from the fibers could

not be quantified. It was postulated that either the rate of drug release was below the

limit of detection of the instrument or there is no drug release from the fibers. In an

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attempt to see if any cellular response could be detected, loaded scaffold was

incubated with HDF cells. Cyto-compatibility of the scaffold was confirmed by

MTS and a fluorescent staining live/dead assay. Interestingly, efficacy of the

released drug from the fibers was evident, in the presence of DMSO (facilitating the

release of drug from the fibers), from the collagen I gene expression studies which

showed a similar significant response as free drug in curtailing the upregulated

expression of collagen I post TGFβ1 stimulation. Therefore, it can be concluded that

slow sustained release of analogue 2 can be accomplished using electrospun PGMA

scaffold.

4.4 Future recommendations

There are still several questions that need to be addressed to translate this

preliminary work into the clinic. For example, the lack of biodegradability of the

PGMA scaffold, the serum stability of the drug and its release kinetics all need

further development.

The general strategy adopted to induce biodegradability is by copolymerising non-

degradable polymers with another biodegradable polymer. Polymers have also been

copolymerised with peptides and proteins to not only improve biodegradability but

also to provide biologically active functional cues for their integration into the

matrix. For example, the RGD (arg-gly-asp) sequence peptide motif can be

copolymerised on PGMA using atom transfer living radical polymerisation (ATRP).

The RGD motif is strongly recognised by a subset of integrins which are protein

receptors necessary for cell adhesion and migration.316-318 Wound healing is a

combination of cellular proliferation and migration working in tandem. Copolymers

with RGD motifs are shown to promote cell migration. In addition, PGMA scaffolds

can be covalently functionalised with growth factors such as insulin like growth

factor I (IGF I) and epidermal growth factor (EGF) via an SN2 substitution ring

opening reaction. The ester groups present in the PGMA backbone can undergo

esterase mediated ester hydrolysis to sustainably deliver growth factors. Both IGF I

and EGF are growth inducing proteins which play pivotal roles in wound healing.319

Chemical conjugation of the growth factor to the scaffold allows their passive

76

diffusion which is continued over the life-time of the scaffold.320 Incorporation of

growth factors and the RGD peptide sequence in the electrospun PGMA matrix

could elevate the rate of healing by simultaneously promoting cell growth and

migration. Electrospun scaffolds have found clinical applications because of their

propensity to mimic extracellular matrix thereby providing a suitable environment to

promote regeneration. Matrix based delivery of growth factors are preferred over

their exogenous delivery because it protects the growth factor from degradation by

limiting their exposure to the proteolytic wound environment.321

Despite the efficacy of analogue 2 ([(bis(pivaloyloxymethyl) (POM)] derivative of

analogue 1) in reducing the expression of collagen I gene after TGFβ1 treatment, its

structural integrity in serum limits it’s in vivo translational potential. [Bis(S-acyl-2-

thioethyl) (SATE)] derivatives of phosphonate analogues have been shown to

possess higher enzymatic stability against serum and gastric juices as compared to

POM derivatives.313, 322 Benzaria et al. synthesised SATE derivatives of 9-[2-

(Phosphonomethoxy)ethyl]adenine (PMEA) and studied their stability in vitro in

comparison to bis[(pivaloyloxymethyl) (POM)]- and bis[dithiodiethyl (DTE)]PMEA

analogues.322 They concluded that although bis(POM)- and bis(tBu-SATE)PMEA

had similar anti-viral activity, bis(tBu-SATE)PMEA had greater stability than

bis(POM)PMEA in human gastric juice and human serum. It would potentially be

beneficial if the replacement of the POM functionality by a SATE derivative in

analogue 2 could improve its serum stability whilst maintaining its functional

efficacy. This could lead to oral or intravenous delivery of the drug. However, SATE

as a linker has its own limitations, with one of the degradation products, ethylene

sulphide, inducing toxicity at high concentrations in cells.

Analogue 2 can be modified at the anomeric centre by replacing the lipophilic group

with a fluorophore which would be beneficial in visualising its internalisation within

the cell microenvironment which can be tracked using confocal microscopy. This

will also allow another platform to study drug release profile from the scaffold.

Further, in a more combinatorial approach, analogue 2 could be combined with a

lysyl oxidase inhibitor. While analogue 2 works predominantly on controlling the

expression of collagen, lysyl oxidase inhibitors reduce collagen crosslinking, making

collagen more prone to degradation. A combined approach could be more effective

77

in reducing fibrosis than the use of a single compound targeting one part of the

fibrotic pathway.

Incorporation of such approaches within the functionalised scaffold would address

the shortfalls with both the scaffold and analogue 2 and could lead to pre-clinical

assessments. The efficacy of the system can be studied on an in vivo porcine wound

healing model. Porcine is a well-established wound healing model which has similar

characteristics to humans i.e. porcine heals with scars unlike rodents which heal

more via contraction instead of granulation tissue formation. The developed scaffold

could potentially be used as a dressing, incorporating growth factors and/or other

anti-scarring agents to be applied directly at the wound site. Scaffold mediated

delivery of the anti-scarring agents would need to be compared against free drug to

confirm the efficacy of this combinatorial approach.

4.5 Final remarks

The three overarching aims of the work compiled in this thesis are set out below:

1. Optimise the parameters for electrospinning PGMA and study its efficacy as a

biocompatible universal polymeric scaffold

2. Investigate the biocompatibility and potential of M6P analogues as anti-

scarring agents in an in vitro model using human dermal skin fibroblasts

3. Investigate the potential of a combinatorial scaffold:drug approach to limit

fibrosis

The data presented within this thesis provides evidence of successful incorporation

of tissue engineering and drug delivery approaches to tackle the problem of skin

scarring. The scaffold developed by electrospinning PGMA can be used in various

other applications, some of which has been demonstrated in this thesis.

Two bioisosteric phosphonate analogues were investigated as potential anti-fibrotic

treatments. Analogue 2, a POM derivative of analogue 1, was shown to function as a

prodrug. Its mode of action was substantiated to be intracellular where its POM

78

linkers get hydrolysed to produce the active analogue 1. Therefore, it was

demonstrated that intracellular targeting of M6P/IGFII receptor using this analogue

is a potentially effective approach to inhibit TGFβ1 signalling and therefore reduce

fibrosis. In a tissue engineering approach, analogue 2 was co-electrospun with

PGMA to fabricate a fibrous scaffold. The biocompatibility of the scaffold was

confirmed in a human dermal skin fibroblast in vitro model. Scaffold mediated

delivery of analogue 2 was shown to regulate collagen I expression. This is the first

example of using this tissue engineering approach in this context and has the

potential with further work to address some of the major limitations with current

treatments aimed at reducing fibrosis.

79

Appendix A

Supporting information for papers

1. Agarwal, V., Ho, D., Ho, D., Galabura, Y., Yasin, F. M.D., Gong, P., Ye, W., Singh,

R., Munshi, A., Saunders, M., Woodward, R. C., St. Pierre, T., Wood, F.M., Fear, M.,

Lorenser, D., Sampson, D. D., Zdyrko, B., Smith, N.M., Luzinov, I., Iyer, K.S., A

Functional Reactive Polymer Nanofiber Matrix, RSC Advances (Submitted)

2. Agarwal, V., Toshniwal, P., Smith N. E., Smith, N. M., Li, B., Clemons, T. D.,

Byrne, L. T., Hassiotou, F., Wood, F. M., Fear, M., Corry, B., Iyer, K.S., Intracellular

Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor Inhibitors

by Intracellular Delivery, Angewandte Chemie International Edition (submitted)

3. Agarwal, V., Wood, F. M., Fear, M. and Iyer, K. S., Inhibiting the activation of

transforming growth factor-β using a polymeric nanofiber scaffold, Nanoscale (Submitted)

80

Supporting Information

A Functional Reactive Polymer Nanofiber Matrix

Vipul Agarwal,a Dominic Ho,a Diwei Ho,a Yuriy Galabura,b Faizah M.D. Yasin,a Peijun Gong,c Weike Ye,d Ruhani Singh,a Alaa Munshi,a Martin Saunders,e Robert C. Woodward,f Timothy St. Pierre,f Fiona M. Wood,g Mark Fear,g Dirk Lorenser,c David D. Sampson,c, e Bogdan Zdyrkob, Nicole M. Smith,a Igor Luzinovb, *, and K. Swaminathan Iyera,*

Materials

Polyglycidyl methacrylate (PGMA) with Mn = 220515 and Mw = 433730 was synthesized

by radical polymerisation as reported previously,1 Carboxy terminated N-isopropyl

acrylamide (PNIPAM-COOH) (Polymer source, Cat. # P5589, Mn = 42000). Methyl ethyl

ketone (MEK, Merck), iron (II) acetylacetonate, tetradecanediol, oleic acid, oleylamine,

dibenzyl ether, 1- octadecene and palladium (II) acetylacetonate purchased from Sigma

Aldrich. Gadolinium chloride, ytterbium chloride and erbium chloride purchased from

GFS chemicals. All other chemicals were purchased from Sigma-Aldrich. All chemicals

used were of analytical grade purity.

Methods

Electrospinning Procedure:

PGMA (15 w%) was dissolved in MEK (10 mL) overnight at room temperature with

constant stirring. In the case of composites, metal nanoparticles resuspended in MEK were

added (magnetite – 35.7 mg, palladium - 40 mg, Upconverting particles - 23.6 mg) to

maintain PGMA concentration at 15 w% the next day and further stirred for 1h to

homogenise the polymer solution.

81

PGMA polymer and polymer composite fibers were obtained via electrospinning (ES-

PGMA) (Nanofiber Electrospinning Unit, Cat. # NEU-010, Kato Tech, Japan). The

electrospinning parameters after optimisation were a voltage of 9.1 kV, working distance

of 9 cm, syringe pump speed of 0.04 mm/min (1 mL/h). Fibers were annealed at 80 °C for

5 hours post electrospinning. Nanoparticle loading in the ES-PGMA was determined by

ICP-MS and AAS analysis (in the case of the magnetite/ES-PGMA composite).

Magnetite Synthesis:

Magnetite nanoparticles were synthesised using the thermal decomposition method as

described previously.2 Briefly, iron (II) acetylacetonate {Fe(acac)3} (1 mmol),

tetradecanediol (5 mmol), oleic acid (3 mmol) and oleylamine (3 mmol) were dissolved in

dibenzyl ether. The reaction mixture was heated under a N2 atmosphere at 100 °C for 30

min, 200 °C for 2h and further refluxed at 300 °C for another 1h. The black coloured

product was collected via precipitation and washed with 3 x ethanol by centrifugation at

4000 rpm and dispersed in hexane. Nanoparticles were stored under an inert atmosphere.

Palladium Nanoparticles Synthesis:

Palladium nanoparticles were synthesised as per the method described.3 Briefly,

palladium(II) acetylacetonate {Pd(acac)3} (150 mg) was mixed with oleylamine (246 µL)

in toluene (61.5 mL) and stirred vigorously for 10 min at room temperature yielding a

yellow coloured reaction mixture. To this formaldehyde (300 µL) was added and further

reacted for 10 min. The reaction was carried out for a further 8h at 100 °C and the colour

changed to black, indicating the completion of the reaction. The product was brought to

82

room temperature, washed with 3 x ethanol and resuspended in chloroform. Nanoparticles

were stored under an inert atmosphere.

NaGdF4: Yb, Er Upconverting Nanoparticle Synthesis:

NaGdF4:Yb, Er upconverting nanoparticles were synthesised as per the method described.4

Briefly, GdCl3.6H2O (0.80 mmol), YbCl3.6H2O (0.18 mmol) and ErCl3.6H2O (0.02 mmol)

were added to the solution mixture of oleic acid (14 mL) and 1- octadecene (16 mL) and

homogenized under N2 while heated to 150 °C. The solution was then cooled to 50 °C,

methanol (10 mL) containing NaOH (2.5 mmol) and NH4F (4 mmol) was then added

slowly and reacted for another 30 min. Next, methanol was removed by heating the

reaction mixture at 100 °C under vacuum for 10 min. Under atmospheric pressure, the

reaction temperature was raised to 300 °C and the reaction was carried out for 1h under a

N2 atmosphere. The reaction was terminated by cooling to room temperature. The product

was precipitated in ethanol, collected by centrifugation and washed with ethanol. Finally,

the product was resuspended in THF.

PNIPAM- COOH attachment on ES-PGMA:

The PGMA polymer was electrospun onto 12 mm diameter glass coverslips (Cat. # G401-

12, ProSciTech) and annealed at 80 °C for 5h. Dried polymer was then exposed to carboxy

terminated poly N-isopropyl acrylamide (PNIPAM- COOH) (0.6 w% in water) and

reacted in the oven for 2h at 120 °C. The polymer mat on the coverslip was washed twice

with THF to remove any excess PNIPAM. Finally, the coverslip was dried in the oven

above 40 °C to remove the THF. The contact angle was measured on this coverslip at

room temperature and again at 40 °C.5

83

Scanning Electron Microscopy:

Dried electrospun PGMA fibers both with and without nanoparticles were coated with 4

nm of platinum and imaged using scanning electron microscope (Zeiss 1555, VP-FESEM)

at an accelerating voltage of 4-5 kV. The fiber diameters and width of the matrix were

calculated using the image analysis software ImageJ (NIH). A minimum of 50 random

fibers were measured. The data is reported as an average ± standard error mean.

X-ray Microanalysis:

Elemental analysis was carried out using the Oxford x-ray microanalysis system (Oxford

instrument X-MAX (80mm2)) set up on the SEM (Zeiss 1555, VP-FESEM). Elemental

data was collected at higher accelerating voltage of 15 kV at 10 mm working distance

required for x-ray microanalysis. Data was analysed using the Aztec version 2.1a analysis

instrument software.

Gas Sensing:

The hydrogen sensor setup consists of a test chamber with inlet gas and outlet gas, a

potentiostat and an electronic recorder. See reference for complete detail of the

experimental setup.6 Prior to entering the test chamber, hydrogen and nitrogen gas, which

are tested and carrier gas respectively, were mixed in a cyclonic mixer. Two silver epoxy

electrodes were painted to the Pd/ES-PGMA composite fibers mounted on an IDE and the

whole integration was mounted into the gas chamber subject to current-voltage (I-V)

sweeps. The procedure involved alternating nitrogen gas (20 min) and varying

concentrations of hydrogen gas (4 min). The change in the current was monitored

simultaneously. The total gas flow rate was 1000 mL/min. The voltage applied between

84

the electrodes was 100 mV dc. An IDE consists of 15 fingers, each 15 μm in width and

550 μm in length, with a finger gap of 10 μm. Each individual electrode is connected to a

boding pad (200 μm×250 μm) to provide a sufficient area for subsequent wire bonding.

Transmission Electron Microscopy (TEM):

Nanoparticles were air dried onto carbon-coated copper grids and imaged using a JEOL

3000F TEM operating at 300 kV. The nanoparticle size was determined using Image J

software. A minimum of 200 particles were measured and the data reported as the average

± standard error mean.

Superconducting Quantum Interference Device (SQUID) Magnetometry:

The magnetic properties of the dried magnetite/ES-PGMA composite (88.2 mg) were

measured using a SQUID magnetometer (Quantum Design 7 Tesla MPMS) operating

between 5K and 300K. Magnetite nanoparticles (22.5 mg) were lyophilised prior to their

measurement and their saturation magnetisation was recorded at 70 kOe at 5K. The zero

field cooled and field cooled measurements were conducted in a field of 0.1 kOe.

Contact Angle:

Static contact angles of MilliQ water on the surface of the electrospun polymer matrix

were measured using a home-built goniometer with Rame- Hart scope attachment.7 The

polymer was electrospun on the 12 mm glass coverslips (Cat. # G401-12, ProSciTech)

both with and without nanoparticles (refer PNIPAM attachment on ES-PGMA). 5 µL of

water was pipetted onto the membrane and images were taken after the drop edges came to

rest (~2 min) using a Canon EOS450D and the angle was measured using the calibrated

85

Rame-Hart scope. Measurements were carried out at room temperature and again at 40 °C.

Images were processed using ImageJ software. Measurements were done in duplicates and

reported as average ± standard deviation.

NIR Room Temperature Emission Spectroscopy (λexc

= 974 nm):

Upconversion spectra measurements were carried out as previously described.8 Briefly,

upconverting particles were suspended in chloroform and were air dried on glass slides.

The upconversion spectra were obtained with an optical set up incorporating a 980 nm

laser diode. The peak wavelength of the laser diode is 974.5 nm. The optical excitation

intensity for obtaining the spectrum shown in Fig. 3A was 7665 W/mm2

References

1. V. Tsyalkovsky, R. Burtovyy, V. Klep, R. Lupitskyy, M. Motornov, S. Minko and I.

Luzinov, Langmuir, 2010, 26, 10684-10692.

2. S. Sun, H. Zeng, D. B. Robinson, S. Raoux, P. M. Rice, S. X. Wang and G. Li, J.

Am. Chem. Soc., 2003, 126, 273-279.

3. Z. Niu, Q. Peng, M. Gong, H. Rong and Y. Li, Angew. Chem. Int. Ed., 2011, 50,

6315-6319.

4. C. Liu, Z. Gao, J. Zeng, Y. Hou, F. Fang, Y. Li, R. Qiao, L. Shen, H. Lei, W. Yang

and M. Gao, ACS Nano, 2013, 7, 7227-7240.

5. G. D. Fu, L. Q. Xu, F. Yao, K. Zhang, X. F. Wang, M. F. Zhu and S. Z. Nie, ACS

Appl. Mater. Inter., 2009, 1, 239-243.

86

6. J. Zou, L. J. Hubble, K. S. Iyer and C. L. Raston, Sensor Actuat.B-Chem., 2010, 150,

291-295.

7. M. V. Baker and J. D. Watling, Langmuir, 1997, 13, 2027-2032.

8. M. Challenor, P. Gong, D. Lorenser, M. Fitzgerald, S. Dunlop, D. D. Sampson and

K. Swaminathan Iyer, ACS Appl. Mater. Inter., 2013, 5, 7875-7880.

87

Enhancing the Efficacy of Cation-Independent Mannose 6-Phosphate Receptor Inhibitors by Intracellular Delivery** Vipul Agarwal, Priyanka Toshniwal,# Natalie E Smith,# Nicole M Smith, Binbin Li, Tristan D Clemons, Lindsay T Byrne, Foteini Hassiotou, Fiona M Wood, Mark Fear, Ben Corry* and K Swaminathan Iyer*

Supporting Information

S1.1 Cell culture

Human primary dermal fibroblast cell cultures from normal skin were cultured in

Dulbecco’s Modified Eagle’s Medium (DMEM/F12 - GlutaMAX; Invitrogen Gibco)

supplemented with 10 % fetal bovine serum (FBS; Invitrogen Gibco) and 1%

penicillin/streptomycin (Invitrogen Gibco). Analogues 1 and 2 were provided by

Pharmaxis Ltd after stringent QC analysis. M6P and analogue 1 were dissolved in

PBS filter sterilized, aliquoted and stored at -20 °C. Each aliquot was used for

maximum 2 freeze thaw cycles. Analogue 2 was dissolved in DMSO and treated the

similar way as other analogues. The cells were incubated at 37 °C in a humidified

atmosphere of 5 % CO2. All experiments were carried out with cells between

passages 3-6.

S2.1 Cell Viability

Cell viability was determined using a LIVE/DEAD viability/cytotoxicity Kit

(Invitrogen, UK) which measures the membrane integrity of cells, as per

manufacturer’s protocol. In brief, 20000 cells were seeded in each well in a 24 well

plate and incubated with analogues at 10 µM concentration in cell culture media

(DMEM F-12 containing 10% FBS and 1% Penicillin/ Streptomycin) for 30 min

(media was added in case of controls). Following which human recombinant TGFβ1

88

(2 ng/mL) (Cat.# 14-8348-62, eBioscience) was added to the wells (fresh media was

added in controls) and plates were incubated for 24 h and 72 h in the humidified

incubator at 37 °C with 5 % CO2. At the stipulated time (24 h and 72 h), cells were

washed with PBS (3 times) and then stained with calcein (100 µL, 1 µM)/ ethidium

bromide (100 µL, 2 µM) in PBS and incubated in the humidified incubator at 37 °C

and 5 % CO2 for 30 min. Images were captured using an Olympus IX71 inverted

microscope with a 20 x objective with fixed exposure time. Both live and dead cells

were counted using Image J software (NIH) with cell counter plug in. Data presented

as mean ± standard error mean (n = 4).

S3.1 Cell Body Area

Cell size was measured using Image J software (NIH).[1] Minimum of 40 cells were

randomly selected from the fluorescence images and their area was measured.

Values reported as mean ± standard error mean.

S4.1 Cell Proliferation

Cell proliferation was measured using the MTS assay (Cell Titer 96 ®Aqueous,

Promega, Madison, USA) as per the manufacturer’s protocol. Briefly, 1500 cells

were seeded in each well of a 96 well plate and treated with analogues at 10 µM

concentration in cell culture media (DMEM F-12 containing 10 % FBS and 1 %

Penicillin/ Streptomycin) for 30 min (media was added in case of controls) prior to

the addition of human recombinant TGFβ1 (2 ng/mL) to each well (fresh media was

added in controls) and the plates (individual for each time point) were incubated for

72 h in the humidified incubator at 37 °C with 5 % CO2. MTS solution (40 µL) was

added in each well the next day, and was considered as 0 h time point, and incubated

for 3 h in the humidified incubator at 37 °C with 5 % CO2. Following which 80 µL

89

from each well was transferred into a new 96 well plates and read under a plate

reader at 490 nm excitation wavelength. Same protocol was followed at every time

point for next 72 h. Data presented as mean ± standard error mean (n = 5).

S5.1 Gene Expression

Gene expression was measured using real time quantitative polymerase chain

reaction. Cells (50000) were seeded in 24 well plates and incubated for 24 h in the

humidified incubator at 37 °C and 5% CO2. Next day culture media (10 % FBS) was

replaced with starve media (0.1 % FBS) and incubated for further 24 h in the

incubator to bring all the cells under same physiological cycle. Next day, cells were

treated with required concentration of M6P analogue 30 min prior to TGFβ1 (2

ng/mL in starve media) stimulation and further incubated for 48 h in the humidified

incubator at 37 °C and 5% CO2. mRNA was extracted using RNeasy mini kit®

according to manufacturers’ protocol (Qiagen GmbH). For reverse transcription 1.5

µg of total mRNA was converted to cDNA using Superscript VILO (Cat.# 11754,

Applied Biosystems) according to manufacturers’ protocol. 150 ng of cDNA was

analysed by ABI 7500 fast analysis real-time PCR system using TaqMan® master

mix and col1a1 probes (Hs01076777_m1, Life Technologies). GAPDH was used as

a reference gene (Cat.# 4326317E, Life Technologies). Analysis was carried out

using the instrument software. Data presented as mean ± standard error mean (n =

3).

S6.1 Protein Expression

90

Protein expression was measured using western blotting. Cells (1 x 105) were seeded

in each well of a 6 well plate and incubated for 24 h in the humidified incubator at

37 °C and 5% CO2. Next day culture media (10 % FBS) was replaced with starve

media (0.1 % FBS) and incubated for further 24 h in the incubator to bring all the

cells under same physiological cycle. Next day, cells were treated with required

concentration of M6P analogue 30 min prior to TGFβ1 (2 ng/mL in starve media)

stimulation and further incubated for 72 h in the humidified incubator at 37 °C and

5% CO2. Whole cell lysates were prepared from treated cells. 35 µg of protein was

denatured and subjected to SDS-PAGE, and transferred to nitrocellulose membrane

(Cat.# 10600007, Amersham, General Healthcare Lifesciences) by standard transfer.

Post transfer, membrane was blocked with 5% skim milk/0.1% TBST for 30 min at

room temperature, then incubated overnight with rabbit anti-human collagen I

antibody (1:2000 in 5% skim milk/0.1% TBST, Cat. # NB600-408, Novus) at 4 °C.

Membranes were washed with 0.1% TBST and incubated with peroxide conjugated

mouse anti-rabbit (1:5000 in 5% skim milk/0.1% TBST, Cat.# NA934VS, GE

Healthcare Lifesciences) for 1 h in 5% skim milk/0.1% TBST at room temperature.

Immunoreactivity was detected using the chemiluminescent HRP substrate (Cat.#

WBKLS0100, Millipore IMMobilon) and the signal was captured with the

Chemidoc (BioRad, Model #731BR02144) and analysed using ImageJ software

(NIH). To confirm equality of protein loading, all membranes were stripped and

reanalysed for β-actin expression using 1° antibody (1:50000 in 5% skim milk/0.1%

TBST, Cat.# A1978, Sigma) and 2° (1:5000 in 5% skim milk/0.1% TBST, Cat.#

NA9310V, GE Healthcare Lifesciences). Data presented as mean ± standard error

mean (n = 3).

91

S7.1 Statistics

The results for cell viability, cell body area, cell proliferation, gene and protein

expression are expressed as mean ± standard error mean (SEM) and analysed for

analysis of variance (ANOVA). Significance was evaluated using Bonferroni and

Turkey’s post-hoc analysis and set at 95% confidence (p < 0.05).

S8.1 Simulation Systems:

Experimental Methods:

6 simulation systems of the M6P/IGFII receptor were investigated; these included

the domain 3 and domain 5 dimers with a ligand in each binding site (2 binding sites

per dimer). These ligands were M6P, analogues 1 and 2. The coordinates for the

domain 3 dimer with M6P bound (pdb accession code 1SYO)[2] and the domain 5

monomer (pdb accession code 2KVB)[3] were obtained from the protein database. In

order to obtain the dimer for domain 5 and position M6P in the binding pocket, the

domain 5 beta sheet regions were aligned with the corresponding beta sheet regions

of domain 3. Analogue 1 and 2 were positioned in each binding pocket by aligning

the mannose ring and phosphate group to the M6P coordinates obtained for domain

3.[2] Each system was then solvated in a TIP3P water box of dimensions 65 x 60 x

114 Å and ionised with 150 mM KCl. All simulations were run with periodic

boundary conditions, constant temperature (310 K) maintained using Langevin

dynamics and constant pressure (1 atm) maintained with a Langevin piston, and the

particle mesh Ewald method was used to compute full system electrostatics.[4] The

CHARMM 36 force field was used for protein, water and M6P parameters.[5] The

ion parameters were obtained from Joung and Cheatham.[6] Missing parameters for

analogues 1 and 2 were obtained using ab-initio techniques [7] with the program

92

Gaussian 09.[8] All molecular dynamics simulations were run with the program

NAMD[9] using rigid bonds to hydrogen and 2 fs time steps. Molecular graphics

were generated using VMD.[10]

Water and ions were energy minimized for 5000 steps and equilibrated for 20 ps

with the protein and substrate held fixed. A harmonic restraint with a force constant

of 20 kcal/mol was then applied to the backbone atoms of the protein and on each

ligand and the system was minimized for a further 10,000 steps prior to 500 ps of

equilibration. This step was repeated with gradual reductions in the force constant

with values of 10, 5, 2.5 and 1 kcal/mol. Finally, to replicate the influence of the

surrounding protein domains on the individual domain being simulated, a harmonic

restraint with a force constant of 0.1 kcal/mol was placed on all of the protein Cα

atoms which were located more than 10 Å from the binding pocket in order to

ensure no loss of secondary structure throughout the simulation. The system was

then minimized for a further 10,000 steps prior to 10 ns of equilibration.

Subsequently 100 ns of equilibrium simulation were obtained for each of the 6

systems.

Data Analysis:

Cluster analysis was performed on the final 100 ns of equilibrium simulation for

each ligand in order to determine the most occupied binding positions. Each ligand

was clustered according to the RMSD of its coordinates with a cut-off of 3 Å.

Subsequently the NAMDEnergy plugin of VMD was utilized to determine the

interaction energy of each ligand with the protein for the entire time and for the most

populated clusters.

93

Figure S1: Most populated binding positions (clusters) of each drug in the domain 3 and domain 5 binding domains as determined from 100 ns of molecular dynamics simulation. In each case the protein is shown in grey, and the most to the least populated clusters are for ligand 1 (initially bound to dimer 1): cyan, pink, mauve, white and green and for ligand 2 (initially bound to dimer 2): dark blue, red, orange, yellow and green. a) M6P in domain 3: Ligands 1 and 2 both remain in the vicinity of the binding site determined in the crystal structure. However as there are multiple hydrogen bond acceptors and donors, both ligands sample multiple positions as the mannose ring hydroxyl groups form hydrogen bonds with various residues and these change over the 100 ns simulation. b) Analogue 1 in domain 3: Ligand 1 has only one binding position (cyan) and ligand 2 has only two clusters indicating that in both cases it binds stably to the protein and remains in the binding pocket throughout the simulation. This binding appears to be stabilized by the interaction between the m-xylene rings of the 2 ligands. c) Analogue 2 Domain 3: Both ligand 1 and 2 are oriented such that the benzyl groups associate in the center of the dimerization domain, hence while the ligands do have more than one binding position they remain in the vicinity of the predicted binding site. d) Domain 5 M6P: In this case the ligands occupy positions which are not asymmetrical, while ligand 1 favors a horizontal orientation with one major binding position (cyan) ligand 2 favors a vertical orientation where it has more flexibility to sample new positions. e) Analogue 1 in domain 5: Ligand 1 moves away from its initial position in the binding site and occupies a position where it is in close contact with ligand 2. f) Analogue 2 in domain 5: In this case both ligands remain in the vicinity of the original binding position. However, while ligand 1 has one major cluster implying its motion is restricted, ligand 2 has more flexibility to move occupying multiple clusters.

94

Figure S2: Snapshots from 100 ns molecular dynamics simulations representative of the most heavily occupied binding positions for each ligand in the domain 3 and domain 5 dimers. The two protein subunits are shown in pink and grey respectively and residues which form common hydrogen bonds are labelled. a) M6P and b) Analogue 1 in domain 3: Ligand 1 and 2 form the strongest energetic interactions with Lys350, Lys358, Ser386 and Arg391. c) Analogue 2 in domain 3: Ligand 1 and 2 form the strongest interactions with Gln348, Arg391 and Glu416. d) M6P in domain 5: Ligand 1 and 2 are oriented differently to each other in the binding pocket. M6P forms its strongest interactions with Arg687. e) Analogue 1 in domain 5: interacts most favorably with Asn680 andArg687. f) Analogue 2 in domain 5: Ligand 1 and 2 form the strongest interactions with Gln644, Trp653, Arg687 and Tyr714.

Figure S3: Average interaction energies obtained for the most occupied positions of each ligand in domains 3 and 5. R391 in domain 3 is equivalent to R687 in domain 5. M6P (a) and analogue 1 (b) in domain 3 both show similar interactions for each ligand and the protein subunit it is most closely associated with (for example Protein 1 with Ligand 1 as compared to Ligand 2 with Protein 2). This is because the binding mode is similar for each ligand in its respective binding site (Figure S2). Similarly, in each case the ligands also interact closely with Lys350 which is actually located on the alternate protein subunit (for example Ligand 1 to Protein 2). When compared to (a) and (b) analogue 2 (c) in domain 3 shows significant decreases in the interaction energies which is reflected by its lower overall interaction energy. M6P (d) and analogue 1 (e) in domain 5 have two major interactions the first one to Arg687 and the second to Glu709. Overall the interaction energies for domain 5 are much lower than those observed for domain 3. Analogue 2 (f) in domain 5 shows a similar pattern with a further reduction in the observed interaction energies.

95

Figure S4: Cell Proliferation assay showing cell growth over the period of 72 h post incubation with analogues both in the presence and absence of TGFβ1. First and second column in each condition is representing 24 h and 72 h respectively. Significant proliferation was observed in all groups despite the reduction in proliferation in TGFβ1 treated groups. Data presented as Mean ± SEM. Significance was set at * p < 0.05 using one way ANOVA with Bonferroni post-hoc analysis.

Figure S5: HDF cell morphology post calcein AM staining imaged using fluorescent microscopy. Cells were treated for 72h, stained and imaged: a) untreated (control), b) TGFβ1 (2 ng/mL) treatment, c) M6P (10 µM) + TGFβ1 (2 ng/mL), d) Analogue 1 (10 µM) + TGFβ1 (2 ng/mL) and e) Analogue 2 (10 µM) + TGFβ1 (2 ng/mL). Scale bar 2 µm.

Figure S6: Collagen I gene time response curve post TGFβ1 (2 ng/mL) stimulation. Collagen I gene expression was significantly upregulated at 24 and 48 h post TGFβ1 stimulation compared to non-treated control. The 48 h stimulation was selected for all further experiments as it yielded significantly higher response. Data presented as average ± SEM (n = 3). Significance was set at * p < 0.05 using bonferroni post-hoc test in one way ANOVA analysis.

96

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[1] T. J. Collins, BioTechniques 2007, 43, S25-S30.

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[3] L. J. Olson, F. C. Peterson, A. Castonguay, R. N. Bohnsack, M. Kudo, R. R.

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[4] U. Essmann, L. Perera, M. L. Berkowitz, T. Darden, H. Lee, L. G. Pedersen, J.

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[5] A. D. MacKerell, D. Bashford, M. Bellott, R. L. Dunbrack, J. D. Evanseck, M.

J. Field, S. Fischer, J. Gao, H. Guo, S. Ha, D. Joseph-McCarthy, L. Kuchnir, K.

Kuczera, F. T. K. Lau, C. Mattos, S. Michnick, T. Ngo, D. T. Nguyen, B. Prodhom,

W. E. Reiher, B. Roux, M. Schlenkrich, J. C. Smith, R. Stote, J. Straub, M.

Watanabe, J. Wiorkiewicz-Kuczera, D. Yin, M. Karplus, J Phys Chem B 1998, 102,

3586-3616.

[6] I. S. Joung, T. E. Cheatham, 3rd, J. Phys. Chem. B 2008, 112, 9020-9041.

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Cheeseman, G. Scalmani, V. Barone, B. Mennucci, G. A. Petersson, H. Nakatsuji,

M. Caricato, X. Li, H. P. Hratchian, A. F. Izmaylov, J. Bloino, G. Zheng, J. L.

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Tomasi, M. Cossi, N. Rega, J. M. Millam, M. Klene, J. E. Knox, J. B. Cross, V.

Bakken, C. Adamo, J. Jaramillo, R. Gomperts, R. E. Stratmann, O. Yazyev, A. J.

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98

Inhibiting the activation of transforming growth factor-β using a polymeric nanofiber scaffold Vipul Agarwal,a Fiona M. Wood,b,c Mark Fearb and K. Swaminathan Iyer*a

Supporting Information

S1. Materials

Polyglycidyl methacrylate (PGMA) with Mn = 220515 and Mw = 433730 was synthesized by radical

polymerisation as reported previously.1 MEK was purchased from Merck. PXS64 was kindly provided by

Pharmaxis Ltd after intensive QC testing.

S2. Electrospinning Procedure:

PGMA (15 wt%) was dissolved in MEK overnight at room temperature with constant stirring. In case of PGMA

+ PXS64, PXS64 (2.95 mg) was mixed with PGMA and dissolved in MEK overnight at room temperature.

PGMA polymer and PGMA + PXS64 fibrous scaffold were obtained via electrospinning (Nanofiber

Electrospinning Unit, Cat. # NEU-010, Kato Tech, Japan) onto 12 mm and 6 mm glass coverslips (Cat. # G401-

12, and G401-06 respectively, ProSciTech). The electrospinning parameters after optimisation were a voltage of

9.1 kV, working distance of 9 cm, syringe pump speed of 0.04 mm/min (1 mL/h), and the traverse and collection

drum speeds were set at 0 cm/min and 0 m/min respectively. Fibers were dried and crosslinked by annealing at

80 °C for 5 h and stored at room temperature for maximum of a month. Scaffolds were UV sterilised in the tissue

culture hood for 15 min and further washed with 2 x PBS prior to cell experiments.

S3. Scanning Electron Microscopy:

Dried electrospun PGMA fibres both with and without PXS64 were coated with 4 nm of platinum and imaged

using scanning electron microscope (Zeiss 1555, VP-FESEM) at an accelerating voltage of 4-5 kV. The fiber

diameter was calculated using the image analysis software ImageJ (NIH). A minimum of 50 random fibers were

measured. The data is reported as an average ± standard deviation. In the cell experiments, cells were incubated

as per cell viability protocol. At the stipulated time point, culture media was removed and coverslips loaded with

scaffold and cells were washed with 2 x PBS and fixed for 10 min using glutaraldehyde (2.5 % in PBS),

99

followed by 2 x PBS washes. Next serial dehydration steps were performed with increasing concentration of

ethanol to replace all the water in the samples with dry ethanol before critical point drying step. Post critical

point drying samples were coated with 4 nm of platinum and imaged using aforementioned parameters.

S4. High Pressure Liquid Chromatography

Drug loading analysis was carried out using HPLC (solvent A: water with 0.1 % TFA, solvent B: acetonitrile

with 0.1 % TFA, Phenomenex-C18 (2) 100A column (4.6 × 150 mm, 5 microns) at room temperature, 1 mL/ min,

λ = 280 nm, gradient: 100 % solvent B in 17 min) with Water 2695 pumping system and 2489 UV/Vis detector.

Drug loading was calculated by dissolving 35 mg of fibers (PGMA + PXS64) in acetonitrile (0.5 mL) and

calculated from the standard curve developed by dissolving free PXS64 in acetonitrile.

S5. Cell culture

Human primary dermal fibroblast cell cultures from normal skin were cultured in Dulbecco’s Modified Eagle’s

Medium (DMEM/F12 - GlutaMAX; Invitrogen Gibco) supplemented with 10% fetal bovine serum (FBS;

Invitrogen Gibco) and 1 % penicillin/streptomycin (Invitrogen Gibco). The cells were incubated at 37 °C in a

humidified atmosphere with 5 % CO2. All experiments were carried out with cells between passages 3-6.

S6. Cell Viability

Cell viability was determined using a LIVE/DEAD Viability/Cytotoxicity Kit (Invitrogen, UK) which measures

the membrane integrity of cells, as per manufacturer’s protocol. In brief, 20000 cells were seeded on the

sterilised electrospun scaffold on glass coverslips in a 24 well plate and supplemented with cell culture media

(DMEM F-12 containing 10 % FBS and 1 % Penicillin/ Streptomycin) for 30 min. Next, human recombinant

TGFβ1 (2 ng/mL, Cat. # 14-8348-62, eBioscience) was added to the wells (fresh media was added in controls)

and plates were incubated for 24 h and 72 h in the humidified incubator at 37 °C with 5 % CO2. At the stipulated

time (24 h and 72 h), cells were washed with PBS (3 times) and then stained with calcein AM (100 µL, 1

µM)/ethidium bromide I (100 µL, 2 µM) in PBS and further incubated for 30 min in the incubator. Images were

captured using an Olympus IX71 inverted microscope with a 20 x objective with fixed exposure time. Both live

and dead cells were counted using Image J software (NIH) with cell counter plug in. Data presented as mean ±

standard error mean (n = 4).

100

S7. Cell Proliferation

Cell proliferation was measured using the MTS assay (Cell Titer 96 ®Aqueous, Promega, Madison, USA) as per

the manufacturer’s protocol. Briefly, 1500 cells were seeded on the sterilised electrospun scaffold on 6 mm glass

coverslips in each well of a 96 well plate and supplemented with cell culture media (DMEM F-12 containing 10

% FBS and 1 % Penicillin/ Streptomycin) for 30 min followed by the addition of human recombinant TGFβ1 (2

ng/mL) to the wells (fresh media was added in controls) and the plates (individual for each time point) were

incubated for 72 h in the humidified incubator at 37 °C with 5 % CO2. MTS solution (40 µL) was added in each

well the next day, and was considered as 0 h time point, and incubated for 3 h in the humidified incubator at 37

°C with 5 % CO2. Following which 80 µL from each well was transferred into a new 96 well plates and read

under a plate reader at 490 nm excitation wavelength. Same protocol was followed at every time point for next

72 h. Data presented as mean ± standard error mean (n = 5).

S8. Gene Expression

Gene expression was measured using real time quantitative polymerase chain reaction. Cells (50000, in

DMEM/F12 media containing 10% FBS) were seeded on sterilised scaffolds (electrospun on coverslips) in 24

well plates and incubated for 24 h in the humidified incubator at 37 °C and 5% CO2. Next day culture media

(containing 10 % FBS) was replaced with starve media (containing 0.1 % FBS) and incubated for further 24 h in

the incubator to bring all the cells under same physiological cycle. Next day, TGFβ1 (2 ng/mL in starve media)

was added in the culture and further incubated for 48 h in the incubator. mRNA was extracted using RNeasy

mini kit® according to manufacturers’ protocol (Qiagen GmbH). For reverse transcription 1.5 µg of total mRNA

was converted to cDNA using Superscript VILO (Cat. # 11754, Applied Biosystems) according to

manufacturers’ protocol. 150 ng of cDNA was analysed by ABI 7500 fast analysis real-time PCR system using

TaqMan® master mix and col1a1 probes (Hs01076777_m1, Life Technologies). GAPDH was used as a reference

gene (Cat. # 4326317E, Life Technologies). Analysis was carried out using the instrument software. Release

media was prepared using DMEM/F-12 media (containing 10 % FBS and 1 % Penicillin/ Streptomycin)

supplemented with 0.2 % DMSO. Data presented as mean ± standard error mean (n = 3).

S9. Statistics

101

The results for cell viability, cell proliferation and gene expression are expressed as mean ± standard error mean

(SEM) and analysed for analysis of variance (ANOVA). Significance was evaluated using Bonferroni and

Turkey’s post-hoc analysis and set at 95% confidence (p < 0.05).

Figure S1: Cell proliferation assay showing cell growth over the period of 72 h post incubation with the two scaffolds both with and without TGFβ1 (2 ng/mL). Data presented as Mean ± SEM. Significance was set at * p < 0.05 using one way ANOVA and Bonferroni post-hoc analysis.

Figure S2: A representative florescent images showing cell morphology and the number of live and dead cells in culture. HDF cells were incubated on a) ES-PGMA scaffold, b) ES-PGMA + TGFβ1 (2 ng/mL) and c) ES-PGMA + PXS64 scaffold. Scale bar 2 µm. Cells were stained using Calcein AM/Ethidium bromide I staining where live cells fluoresce green while dead cells fluoresce red.

102

Figure S3: A representative image showing cell adherence on the two scaffolds. HDF cells were incubated a) ES-PGMA scaffold, b) ES-PGMA + TGFβ1 (2 ng/mL) and c) ES-PGMA + PXS64 scaffold. Scale bar: a) 2 µm, b) and c) 1 µm. Red arrows highlighting the cells.

References

1. K. S. Iyer, B. Zdyrko, H. Malz, J. Pionteck and I. Luzinov, Macromolecules, 2003, 36, 6519-6526.

103

Appendix B

Published work not directly included in the thesis

Peer reviewed publications contributed by the candidate are listed below. However, only

published articles are attached in the thesis.

1. Agarwal, V., Tjandra, E.S., Iyer, K. S., Humfrey, B., Fear, M., Wood, F. W.,

Dunlop, S. and Raston, C. L., Evaluating the effects of nacre on human skin and scar cells

in culture, Toxicology Research, 3, 223-227 (2014)

2. Eroglu, E., Chen, X., Bradshaw, M., Agarwal, V. , Zou, J., Stewart, S.G., Duan, X.,

Lamb, R.N., Smith, S.M., Raston, C. and Iyer, K. S., Biogenic production of palladium

nanocrystals using microalgae and their immobilization on chitosan nanofibers for

catalytic applications, RSC Advances, 3, 1009-1012 (2013)

3. Eroglu, E., Agarwal, V., Bradshaw, M., Chen, X., Smith, S. M., Raston, C. and Iyer,

K. S., Nitrate removal from liquid effluents using microalgae immobilized on chitosan

nanofiber mats, Green chemistry, 14 (10), 2682-2685 (2012)

4. Ho, D., Zou, J., Chen, X., Munshi, A., Smith, N. M., Agarwal, V., Hodgetts, S.I.,

Plant, G. W., Bakker, A., Harvey, A. R., Luzinov, I., Iyer, K. S., Hierarchical patterning of

multifunctional conducting polymer nanoparticles as a bionic platform for topographic

contact guidance, ACS nano, 9 (2), 1767-1774 (2015)

Toxicology Research

Chinese Society Of Toxicology

www.rsc.org/toxicology

ISSN 2045-452X

COMMUNICATION Colin L. Raston et al.Evaluating the eff ects of nacre on human skin and scar cells in culture

Volume 3 Number 4 July 2014 Pages 217–292

Toxicology Research

COMMUNICATION

Cite this: Toxicol. Res., 2014, 3, 223

Received 6th January 2014,Accepted 2nd May 2014

DOI: 10.1039/c4tx00004h

www.rsc.org/toxicology

Evaluating the effects of nacre on human skin andscar cells in culture†

Vipul Agarwal,a Edwin S. Tjandra,a K. Swaminathan Iyer,a Barry Humfrey,b Mark Fear,c

Fiona M. Wood,c Sarah Dunlopd and Colin L. Raston*e

Pearl nacre, a biomineralisation product of molluscs, has growing

applications in cosmetics, as well as dental and bone restoration,

yet a systematic evaluation of its biosafety is lacking. Here, we

assessed the biocompatibility of nacre with two human primary

dermal fibroblast cell cultures and an immortalised epidermal cell

line and found no adverse effects.

There are three main types of pearl oysters of the genusPinctada: the “Akoya” pearl oyster called Pinctada fucata, the“Golden lipped” oyster Pinctada maxima and the “Blacklipped” oyster named Pinctada margaritifera. Mollusc shells aremainly made up of two layers of calcium carbonate, compris-ing an outer layer of calcite and an inner layer of aragonite.Nacre (mother of pearl) in all oyster shells is a calcified struc-ture that forms the lustrous inner layer. It is mainly composedof aragonite (∼95–97%) tablets oriented in multiple layers,each surrounded by organic matrix.1,2 This organic matrixmakes up ∼5% of the nacre composition and is mainly com-prised of polysaccharides and proteins.3 According to a Euro-pean Commission report published in 2007 the cosmetic andtoiletries industry in the EU, Japan, China and the US had atotal market value of €136.2 billion.4 The cosmetics industrymaintains its edge by constantly developing novel topical skintreatments. A popular example is the use of all-natural ororganic ingredients, such as fruit and plant extracts to offerwrinkle relief that mimics the painful and potentially danger-ous side effects associated with invasive chemical remedies.5

Clinically, topical treatments containing, for example, aloevera, vitamin C, corticosteroids and tacrolimus are used with

the aim of minimizing scarring.6 Recently, there has beeninterest in the cosmetics industry in the use of nacre as a keyingredient.7 Most of the formulations are reported to eitheruse powdered pearl shell or powdered nacreous layer shell.Powdered shell and powdered nacre comprises of both organicand inorganic components. It is reported that nacre stores inits mineral-based organic structure a variety of bioactive mole-cules. Efficacy of this water soluble matrix (WSM) has beentested in a porcine burn injury model.8 WSM was obtained bysuspending powdered nacre in ultra-pure water and collectingthe supernatant via precipitation of insoluble components bycentrifugation. It was concluded that the active mineral basedorganic component has beneficial effects on the skin withenhanced wound healing.8,9

Nacre has also attracted attention for its potential in sup-porting bone grafting and bone regeneration. In culture underphysiological conditions, nacre can transform to hydroxy-apatite, the phosphorous rich main constituent of the mamma-lian bone framework.10,11 Nacre and its WSM can also aid inosteogenic regeneration.9,12–17 High phosphorous rich domainshave been described at the interface between bone andimplants made from Margaritifera shells which are biocompa-tible, biodegradable and osteoconductive and thus are thoughtto promote bone formation.18 Furthermore, nacre powder hasbeen used as an implantable material for reconstruction andregeneration of maxillary alveolar ridge bone in humans.19 Inthis example, the implanted nacre dissolves gradually and iseventually replaced by the mature lamellar bone suggestingthat the nacre acts as a biocompatible substrate for bone repla-cement.19 The water soluble components of the crushed nacrehave also been investigated for their potential in bone regener-ation in a similar vein.20,21 Lee et al., demonstrated the woundhealing potential of WSM component in a deep burn porcineskin model and showed enhanced collagen secretion anddeposition at the injury site resulting in enhanced healing.8 Inanother in vivo study using a rat skin incisional injury model,powdered nacre was implanted between the epidermis anddermis at the incisional site, with an aim of studying the effectof nacre on the synthesis of certain constituents of the dermal

†Electronic supplementary information (ESI) available. See DOI:10.1039/c4tx00004h

aSchool of Chemistry and Biochemistry, The University of Western Australia, Crawley,

AustraliabPearl Technology Pty Ltd, Geraldton, AustraliacBurn Injury Research Unit, School of Surgery, The University of Western Australia,

Crawley, AustraliadSchool of Animal Biology, The University of Western Australia, Crawley, AustraliaeSchool of Chemical and Physical Sciences, Flinders University, Bedford Park,

Australia. E-mail: [email protected]; Tel: +61 82017958

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extracellular matrix. It was concluded that implanted nacreincreased collagen synthesis by dermal fibroblasts.22 Whileextensive investigations have been carried out in bone, theevaluation of the biocompatibility of nacre with human skincells is lacking. Thus the growing number of cosmetic formu-lations in the market with nacre as a key ingredient7,23,24

clearly warrants a thorough assessment with human skin cells.Since scars are also common, and contain cells with a pheno-type distinct from normal skin,25 it is also important to testpotential cosmetic ingredients with both cell types.

In the present study, we use nacre from the inner calcifiedlayer of the shell of Pinctada margaritifera and report thein vitro toxicity assessment of the material on three cell typesrepresenting both epidermal and dermal layers of humanskin. These were HaCaT cells, a human derived immortalisedkeratinocytes cell line, primary human dermal skin fibroblasts(HDF) and primary human scar fibroblasts (HSF).

Nacre used in the study was gently scraped26,27 from theinner layer of the shell to avoid the post processing required inthe case of powdered shell. SEM images (Fig. 1) confirmedthat the nacre was composed of pseudo-hexagonal shaped ara-gonite tablets which have basal plane dimensions of 2–6 µm,and a thickness of 300–400 nm.28 This structure is character-istic of previously reported nacre, which is a compositematerial consisting of alternating layers of mineral tablets sepa-rated by thin layers of biomacromolecular “glue”.29,30

To test the cytotoxicity of nacre, a live/dead assay wascarried out (see ESI S1.4†). Cells were incubated with nacre inculture media at physiological conditions for 24 h to 72 h andwere then stained for viability using calcein AM/ethidiumbromide I solutions. Viable cells fluoresce green through thereaction of calcein AM with intracellular esterase, whereasnon-viable cells fluoresce red due to the diffusion of ethidiumhomodimer across damaged cell membranes and binding withnucleic acids.

Fig. 2 shows live cells as the percentage of the total cells inhuman primary dermal skin fibroblast (HDF), human primaryscar fibroblast (HSF) and human derived immortalised HaCaTcell cultures when exposed to various concentrations of nacrefor 24 h and 72 h. Cytotoxicity of nacre was not observed forany of the concentrations examined in HDF cells (Fig. 2a).However, interestingly at a concentration of 2.5 mg ml−1 ofnacre (highest concentration tested) there was a significantreduction in viability at both 24 and 72 hours in the HSF cells(Fig. 2b). This underlines the importance of testing both scarand normal skin cell types for cosmetic application. Toxicitywas also observed at a concentration of 0.5 mg ml−1 in HaCaTcells (Fig. 2c), although this was only observed at 24 hours and

Fig. 1 Top view of the scrapped nacre, imaged using scanning electronmicroscopy (SEM). Scale bar (a) 1 µM and (b) 2 µM respectively. Samplewas coated with 4 nm platinum prior to imaging.

Fig. 2 Cell viability assays showing percentage of live cells in theculture post incubation with nacre. (a) Human dermal skin fibroblastscells, (b) human scar fibroblasts cells and (c) human derived immorta-lised HaCaT cells were incubated with various concentrations ofscrapped nacre and treated with calcein AM/ethidium bromide I to stainfor live and dead cells. Both live and dead cells were counted using fluo-rescence microscopy. ‘None’ is the untreated control. Data presented asaverage ± SEM (n = 4). Significance was set at *p < 0.05 using bonferronitest in one way ANNOVA.

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was no longer present at the 72 hour time point. These resultsare in line with the results obtained previously where constitu-ents of nacre were shown to promote wound healing in a ratmodel22 and deep burn porcine skin.8 In both these studies,nacre has been shown to promote the recruitment of fibro-blasts for restoration and coverage of the injury site whileshowing no apparent signs of cytotoxicity. It has also beenshown to promote bone formation when implanted in thefemur of sheep with midshaft hemidiaphysis resection of theirfemur in vivo reiterating the non-cytotoxic advantage ofnacre.31

Fibroblasts have been reported to undergo morphologicalchanges from dendritic to stellate shapes upon exposure toexternal cues caused by changes in actin polarisation andadhesion.32,33 Cell morphology in fibroblasts is known to beinfluenced by cytokines such as transforming growth factor β

which can potentially induce polymerisation of globular tofilamentous actin.34 Fibroblast morphology can also be modu-lated by extracellular matrix architecture during woundhealing via cell–matrix interaction.32 Such morphologicalchange has been observed in cells undergoing oxidativestress.35,36 In our study, we found similar changes in fibroblastmorphology for both HDF and HSF cells at the highest concen-tration of nacre of 2.5 mg mL−1 (Fig. 3). Similar altered mor-

phology was also observed for HaCaT cells (see ESI Fig. S1†). Itcould be postulated that the high concentration of nacreinduces cellular stress, resulting in changes in the actin cyto-skeleton and a more stellate morphology (Fig. 3i, iii, ii and iv).Cell area was calculated from the fluorescent images shown inFig. 3 using Image J software.37 It was found that both HDFand HSF had significantly larger cell areas (p < 0.05) whentreated with 2.5 mg mL−1 of nacre (HDF: 2.79 ± 0.13 µm andHSF: 3.0 ± 0.19 µm respectively) as compared to the non-treated controls (HDF: 1.56 ± 0.08 µm and HSF: 1.54 ± 0.10 µmrespectively) (see ESI Fig. S2†).

Altered fibroblast morphology has been thought to occur inresponse to various factors including aging,38 strength of theextracellular matrix39 or other etiologies that induce mechan-ical stress on the cell. Changes in morphology also commonlyindicate oxidative as well as mechanical stress.39 Therefore, weexplored whether the morphological changes and increase in

Fig. 3 Cell morphology post calcein AM staining and imaged usingfluorescent microscopy. Cells were treated with various concentrationsof nacre for 24 h, stained and imaged. (i) untreated (control) primaryhuman dermal skin fibroblasts (HDF), (ii) HDF treated with 0.05 mg mL−1

nacre, (iii) HDF treated with 2.5 mg mL−1 nacre, and (iv), untreated(control) primary human dermal scar fibroblasts (HSF), (v) HSF treatedwith 0.05 mg mL−1 nacre and (vi) HSF treated with 2.5 mg mL−1 nacre.Scale bar 1 µm.

Fig. 4 Reactive oxygen species (ROS) assay showing ROS levels in cellsstressed with various concentrations of nacre for 24 h. No significantstress was observed as a result of calcium (from nacre) induced oxidativestress at the concentrations studied. (a) Human dermal skin fibroblastscells, (b) human scar fibroblasts cells and (c) human derived immorta-lised HaCaT cells were incubated with various concentrations ofscrapped nacre for the specified period of time. Cells were then incu-bated with 2’,7’-dichlorodihydrofluorescein diacetate (DCFH-DA) solu-tion which fluoresce in the presence of reactive oxygen species. ‘None’is the untreated control. Data presented as average ± SEM (n = 3). Sig-nificance was set at *p < 0.05 using bonferroni test in one way ANNOVA.

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cell area at the highest concentration of nacre in culture was aresult of, or induced oxidative stress in, the cells. Oxidativestress was tested using the cell permeable fluorogenic probe2′,7′-dichlorodihydrofluorescein diacetate (DCFH-DA).DCFH-DA is taken up by cells and is deacetylated by cellularesterases to non-fluorescent 2′,7′-dichlorodihydrofluorescein(DCFH) which is rapidly oxidised by reactive oxygen species(ROS) to highly fluorescent 2′,7′-dichlorodihydrofluorescein(DCF). The fluorescent intensity is proportional to the ROSlevels within the cytosol (see ESI S1.5†). Cell responsiveness tothe assay was carried out by stressing the cells with the H2O2

solution provided in the kit which was also used to generatethe calibration curve (see ESI Fig. S3†). No changes in levels ofreactive oxygen species were observed in any cell type at anyconcentration of nacre (Fig. 4). This is important as oxidativestress is known to be a significant contributor to skin damageand excessive scarring in previous studies.40 It has beenknown that cells alter their morphology depending on theirenvironment.41,42 It is therefore hypothesized that the alteredfibroblast morphology in the present case is mainly due to theregulation of cell motility through geometrical constraint inthe presence of nacre. Indeed, it has been previously reportedthat when cells probe their physical surroundings, they acquiremechanical information or signals that help to determine thedirection of migration, with a consequential change in cellmorphology.43

Conclusions

We have established the biocompatibility of nacre using threehuman cell types representing the two primary layers ofhuman skin, using immortalised keratinocytes from the epi-dermal layer and two primary human dermal cell cultures. Thenacre used in the present study showed limited cytotoxicity athigh concentrations in scar derived cells, with the morphologyof the cells significantly changed by exposure at such concen-trations of nacre. No apparent oxidative stress was evident inany of the cell types. Overall, the data support the use of lowconcentrations of nacre in aesthetic formulations, with thepotential for high concentrations to cause changes in skinand/or scar cells which may have impact on efficacy.

The authors would like to thank the Australian ResearchCouncil and Pearl Technologies for funding the work underthe grant number LP100100812. The authors would also liketo acknowledge the Australian Microscopy & MicroanalysisResearch Facility at the Centre for Microscopy, Characteri-zation & Analysis, The University of Western Australia, fundedby the University, State and Commonwealth Governments.

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42 K. Burridge and M. Chrzanowska-Wodnicka, Annu. Rev.Cell Dev. Biol., 1996, 12, 463–519.

43 R. J. Petrie, A. D. Doyle and K. M. Yamada, Nat. Rev. Mol.Cell Biol., 2009, 10, 538–549.

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Evaluating the effects of nacre on human skin and scar cells in culture

Vipul Agarwal,a Edwin S. Tjandra,a K. Swaminathan Iyer,a Barry Humfrey,b Mark Fear,c Fiona M. Wood,c Sarah Dunlopd and Colin L. Rastone, *

Supporting Information

S1. Experimental

S1.1 Materials

Pinctada margaritifera shells were provided by Pearl Technologies Pty Ltd, which were grown in the waters of the Abrolhos Islands, Western Australia. The decalcified organic conchiolin layer was removed by wet sand blasting of the shell followed by gentle brushing to remove any dust particles that might otherwise contaminate the samples. Inner nacreous layer was then scraped using a surgical scalpel and stored at room temperature for a maximum of 2 weeks.

S1.2 Scanning Electron Microscopy

Scraped nacre above from the inner layer of the shell was mounted on SEM stubs (ProSciTech, Cat.# G040). Samples were coated with 4 nm of platinum. Images were taken using scanning electron microscope (Zeiss 1555, VP-FESEM) at 4-5 kV at 30µm aperture. Images were analyzed with the image analysis software ImageJ (NIH).1

S1.3 Cell culture

A human derived immortalized keratinocyte cell line, HaCaT 2 and two human primary dermal (fibroblast) cell cultures from normal skin and normal scar were used. All three cell types were cultured in Dulbecco’s Modified Eagle’s Medium (DMEM/F12 - GlutaMAX; Invitrogen Gibco) supplemented with 10% fetal bovine serum (FBS; Invitrogen Gibco) and 1% penicillin/streptomycin (Invitrogen Gibco). The cells were incubated at 37°C in an atmosphere of 5% CO2. Primary cells used were between the passages 7-10. Nacre was sterilized using UV sterilization technique in the tissue culture hood for 15 min prior its dissolution in media. Fresh nacre solution was prepared before every experiment.

S1.4 Cell Viability

Cell viability was determined using a LIVE/ DEAD Viability/Cytotoxicity Kit (Invitrogen, UK) which measures the membrane integrity of cells,3, 4 as per manufacturer’s protocol. In brief, 20000 cells were seeded in each well in a 24 well plate and treated with scraped nacre at various concentrations in cell culture media (DMEM F-12 containing 10% FBS and 1% Penicillin/ Streptomycin) and incubated for 24 h or 72 h in the humidified incubator at 37°C with 5% CO2. At the stipulated time (24h and 72h), cells were washed with PBS (3 times) and then stained with calcein (100 µL, 1µM)/ ethidium bromide (100 µL, 2 µM) in PBS and incubated in the humidified incubator at 37°C and 5% CO2 for 30 min. Images were captured using an Olympus IX71 inverted microscope with a 20 x objective with fixed exposure time. Both live and dead cells were counted using Image J with cell counter plug in. Experiments were performed in triplicate. Minimum of fifty images were captured per condition.

S1.5 Reactive Oxygen Species (ROS)

ROS was measured using the ROS assay kit (Oxiselect ROS assay kit, Cat.# STA 342, Cell Biolabs) following manufacturer’s protocol. In brief, 6000 cells were seeded in a 96 well plate and incubated in the humidified incubator at 37°C with 5% CO2 for 24h. Next day, wells were washed with PBS (3 times) and incubated with 2’, 7’- dichlorodihydrofluorescin diacetate (DCFH-DA) solution (0.1 x, 100 µL/ per well) for 1 h in the humidified incubator at 37°C with 5% CO2. DCFH-DA solution was removed and the wells washed with 3 x PBS. Cells were then treated with scraped nacre solution in culture media at a specified concentration for 24 h, wells were washed with 3 x PBS and cells were lysed using the lysis buffer provided (1 x, 100µL/ per well, incubated for 20 min. at room temperature) before reading the plate at 480 nm excitation/ 530 nm emissions using the plate reader. Experiments were performed in triplicate.

S1.6 Cell Body Area

Electronic Supplementary Material (ESI) for Toxicology Research.This journal is © The Royal Society of Chemistry 2014

Cell size was measured using Image J software (NIH).1 A minimum of 25 cells were randomly selected from the fluorescence images and their area was measured. Values reported as mean ± standard error mean.

S1.7 Statistics

The results for cell viability, ROS experiments and cell area are expressed as mean ± standard error mean (SEM) and analysed by analysis of variance (ANOVA). Significance was evaluated using Bonferroni and Turkey’s post-hoc analysis and set at 95% confidence (p < 0.05).

Supporting Figures

Figure S1: Cell morphology post calcein AM/ ethidium bromide staining and imaged using fluorescent microscopy. HaCaT cells were treated with various concentrations of nacre for 24h, stained and imaged. a) Untreated (control), b) HaCaT cells treated with 2.5 mg/mL nacre. Scale bar 1µm.

Figure S2: Cell area size showing increase in the cell area post incubation with nacre. Cell area was measured from the fluorescent images of live cells taken for viability assay. ‘None’ is the untreated control. Data presented as average ± SEM (n>25). Significance was set at * p < 0.05 using bonferroni post hoc test in one way ANNOVA

Figure S3: Reactive oxygen species (ROS) assay showing increase in ROS levels in cells stressed with various concentrations of H2O2 in dose dependent manner. Human dermal skin fibroblasts cells were incubated with various concentrations of H2O2 for the specified period of time to generate the standard curve. Cells were then incubated with 2’, 7’- dichlorodihydrofluorescin diacetate (DCFH-DA) solution which fluoresce in the presence of reactive oxygen species. ‘None’ is the untreated control. Data presented as average ± SEM (n=3).

References

1. T. J. Collins, BioTechniques, 2007, 43, S25-S30.2. P. Boukamp, R. T. Petrussevska, D. Breitkreutz, J. Hornung, A. Markham and N. E. Fusenig, J. Cell Biol., 1988, 106, 761-771.3. L. W. Zhang and N. A. Monteiro-Riviere, Toxicol. Sci., 2009, 110, 138-155.4. I. Gotman and S. Fuchs, in Active Implants and Scaffolds for Tissue Regeneration, ed. M. Zilberman, Springer Berlin Heidelberg,

2011, vol. 8, pp. 225-258.

Cite this: RSC Advances, 2013, 3, 1009

Received 4th October 2012,Accepted 12th November 2012

Biogenic production of palladium nanocrystals usingmicroalgae and their immobilization on chitosannanofibers for catalytic applications3

DOI: 10.1039/c2ra22402j

www.rsc.org/advances

Ela Eroglu,ab Xianjue Chen,a Michael Bradshaw,a Vipul Agarwal,a Jianli Zou,a ScottG. Stewart,c Xiaofei Duan,d Robert N. Lamb,d Steven M. Smith,*b Colin L. Raston*a

and K. Swaminathan Iyera

Spherical palladium nanocrystals were generated from aqueous

Na2[PdCl4] via photosynthetic reactions within green microalgae

(Chlorella vulgaris). Electrospun chitosan mats were effective for

immobilizing these biogenic nanocrystals, as a material for

recycling as a catalyst for the Mizoroki–Heck cross-coupling

reaction. This photosynthetically-driven metal transformation

system can serve as a good candidate for an environmentally-

friendly method for the synthesis of metal nanocatalysts.

Biogenic systems can control the phase, structure, and topographyof inorganic nanocrystals with a level of precision similar tosynthetic approaches.1 Various organisms such as bacteria,2

yeast,3 fungi4 and algae5 are capable of processing a wide rangeof metals. Such bioprocessing has been effectively used in thereduction of environmental pollution and also for the recovery ofmetals from waste.6 The formation of metal nanoparticles in thepresence of microorganisms primarily involves the reduction ofmetal ions in solution by enzymes generated by microbial cellactivities, which can be intracellular or extracellular. This dependson the nanoparticle crystallization site, as being either inside thecell or on the cell surface.7,8 When considering palladium, thebiosynthesis of nanoparticles of the metal have been reported inbacteria (Desulfovibrio desulfuricans, Shewanella oneidensis, Bacillussphaericus),9 cyanobacteria (Plectonema boryanum UTEX 485),10,11

plants,12–14 and viruses (e.g., tobacco mosaic virus).15,16 Herein wereport a biogenic synthesis of palladium nanocrystals in thepresence of Chlorella vulgaris with the photoautotrophic micro-

algal metabolism most likely providing the necessary reducingagents. Studies on the biogenic reduction of palladium by variousorganisms are usually limited to the generation of Pd(0). We haveused a non-toxic and environmentally-available microorganism toachieve this, and then collected the particles from the liquidculture via immobilizing on electrospun chitosan nanofibers. Inaddition, we have established the utility of this composite materialas a recyclable catalyst which is functional even at very low loadingrates. The overall integrated approach is without precedent, as arethe individual components.

Algae are a large group of photosynthetic organisms that areubiquitous in various aquatic habitats including sea, freshwater,wastewater and also in moist solids.17 Microalgal remediation hasbeen reported for several metal ions,18,19 and involves bothintracellular (polyphosphates) and extracellular (polysaccharideson algal cell wall) metal binding groups.20–22 Based on theseobservations, we initially evaluated the ability of a common greenmicroalga, Chlorella vulgaris, to reduce palladium(II) to Pd(0)starting with a solution of Na2[PdCl4]. Various concentrations ofPd salt were investigated (100, 50, 25, 12.5, 0 mg L21) in order todetect their effects on cell viability. The cells were grown insolutions of Na2[PdCl4] added to algal freshwater media (MLAmedia)23 at 25 uC and a rotation speed of 120 rpm, under diurnalillumination of 16 h light/8 h dark cycles, with the total chlorophyllcontent used to validate the viability of the cell cultures.24

Accumulation of total chlorophyll pigment (Chl a + Chl b) as afunction of growth time was monitored (Fig. 1a), and thethreshold concentration for toxicity on the cells establishedamongst 25–50 mg L21. Within the first five days, the cultureflasks containing less than this concentration clearly showed anincrease in cellular growth with solution retaining the green colorwhich is associated with the chlorophyll content of the cells. Incontrast, higher concentrations of Na2[PdCl4] resulted in the colorfading and loss of cell viability (Fig. 1b). The solution containing25 mg L21 of Na2[PdCl4] had the highest amount of palladiumprecursor while showing little reduction in growth rates comparedto the control. Thus, it was chosen as the prototype for furtherexperiments.

aCentre for Strategic Nano-Fabrication, School of Chemistry and Biochemistry, The

University of Western Australia, M313, 35 Stirling Highway, Crawley, WA 6009,

Australia. E-mail: [email protected]; Fax: +61 8 6488 8683;

Tel: +61 8 6488 3045bARC Centre of Excellence in Plant Energy Biology, The University of Western

Australia, M313, 35 Stirling Highway, Crawley, WA 6009, Australia.

E-mail: [email protected]; Fax: +61 8 6488 4401; Tel: +61 8 6488 4403cSchool of Chemistry and Biochemistry, The University of Western Australia, 35

Stirling Highway, Crawley, WA 6009, AustraliadSurface Science & Technology Group, School of Chemistry, The University of

Melbourne, VIC 3010, Australia

3 Electronic supplementary information (ESI) available: Experimental details,characterization, and an additional TEM image. See DOI: 10.1039/c2ra22402j

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Transmission electron microscopy (TEM) of the liquid culturemedia revealed the presence of crystalline spherical palladiumnanoparticles with an average particle size of around 7 nm, with arange of 2 to 15 nm in diameter (Fig. 2a). The correspondingelectron diffraction pattern further confirmed the presence ofelemental Pd nanocrystals with a characteristic face-centered cubic(fcc) structure and average d-spacing values of 0.22, 0.19, 0.14, and0.12 nm for the (111), (200), (220) and (311) planes, respectively(Fig. 2b). High resolution TEM analysis further confirmed thepresence of single palladium nanocrystals (Fig. 2c and d).

Reducing agents produced by photosynthesis are the keycomponents for the reduction of Pd(II) into Pd(0) nanoparticles(Scheme 1). Photosynthetic processes in green microalgae can takeplace either under oxygenic or anoxic environments.17,25 Greenalgae have chlorophyll-a and chlorophyll-b pigments, and can

accomplish oxygenic (oxygen-evolving) photoautotrophic reactionswhile using H2O as an electron donor.17 Oxygenic photoauto-trophic processes include two sets of photochemical reactions: (i)light reactions conserving chemical energy in the form ofadenosine triphosphate (ATP), and the reduced form of nicotina-mide adenine dinucleotide phosphate (NADPH), and (ii) ‘dark’reactions in which CO2 is reduced to organic compounds usingATP and NADPH. Light reactions have two separate sets ofphotosystems (PS), namely PS I and PS II, in which the electrontransfer follows a Z-scheme between these two photosystems. PS IIis mainly responsible for the splitting of water (H2O A KO2 + 2e2

+ 2H+) as a first stage of cyclic electron flow. While electron flowfollows the Z-scheme between PS II and PS I, a proton motive forceis created and used for the generation of ATP. CO2 is fixed by theenzyme ribulose bisphosphate carboxylase (RuBisCO) and reducedin the Calvin cycle using NADPH.17 Reducing equivalents can beexported from the chloroplast in the form of carbon metabolites,particularly triose phosphates (triose-P).26 Their oxidation bydehydrogenases in the cytosol can also generate NADH andNADPH for other reduction reactions. We postulate that suchreducing agents promote the reduction of Pd(II) into crystallinePd(0) nanoparticles, which is further partially oxidized due to theaerobic culture conditions. An hypothesis for the reduction ofpalladium(II) by photoheterotrophic bacteria (Rhodobacter sphaer-oides) cultures involves the reduced electron carriers (such asferredoxin, NADH, and FADH) as the electron donors.27

Consistent with this hypothesis, we found that addition ofNADPH to the MLA growth medium23 resulted in reduction ofNa2[PdCl4] to produce Pd nanocrystals in solution which wereeffectively recovered on exposure to cross-linked electrospunchitosan nanofiber mats (see ESI3, Fig. S1).

Fig. 2 (a) TEM image of palladium nanoparticles precipitated in the four week oldmicroalgae solution (scale bar: 20 nm). (b) Selective area electron diffraction patternof the palladium nanoparticles in 2(a), giving d-spacings of 0.22, 0.19, 0.14, and 0.12nm which correspond to 111, 200, 220, and 311 reflections, respectively, for Pd(0).(c) High resolution TEM image of palladium nanoparticles. (d) Pd(0) nanocrystalswith (111) lattice spacings of around 0.22 nm. Inset shows the Fast FourierTransform (FFT) pattern corresponding to the area shown within the yellowrectangle in 2(c), indicating the crystal structure of palladium nanoparticles.

Scheme 1 Palladium nanoparticle synthesis by photosynthetic green microalgae,and their uptake on an electrospun chitosan mat for use as a catalyst in Mizoroki–Heck reactions. The left stage shows a combination of mechanisms taking placewithin the photosynthetic organisms, resulting in the production of reducingagents.17,25,26 (ADP: adenosine diphosphate, ATP: adenosine triphosphate, Fd:ferredoxin, NADP+: oxidized form of nicotinamide adenine dinucleotide phosphate,NADPH: reduced form of nicotinamide adenine dinucleotide phosphate, PGA:phosphoglycolic acid, RuBisCO: rubilose biphosphate carboxylase). NADPH is likelyone of the main reducing agents for the reduction of Na2[PdCl4], which is partiallyoxidised as a result of aerobic culture conditions.

Fig. 1 (a) Total amount of accumulated chlorophyll (Chl a + Chl b) measured as afunction of growth time for Chlorella vulgaris cultures in the presence of variousconcentrations of Na2[PdCl4] (100, 50, 25, 12.5 mg L21), and control cultures in theabsence of Na2[PdCl4] (0 mg L21, hollow circles), with the total volume of solution inall flasks at 40 mL during the growth experiments. (b) Chlorella vulgaris culture flaskscontaining various concentrations of Na2[PdCl4].

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Chitosan is derived from chitin present in the exoskeletonsof crustaceans,28 and is an attractive, renewable feedstockwhich is effective in binding a number of metal ions,29 andeffective as a support for Pd, Ni, Cu, Cr and Zn based catalysts.30

Cross-linked chitosan mats (3 6 2 cm rectangular shape)generated using electrospinning31,32 were placed into a fourweek old microalgal culture containing 25 mg L21 Na2[PdCl4]. Acontrol experiment involved exposing chitosan mats to 25 mgL21 Na2[PdCl4] solution in the absence of microalgae. Chitosanmats were kept in the solution for about two weeks as to achievesufficient particle adsorption from the liquid media. Scanningelectron microscopy (SEM) revealed the nanofiber structure ofthe chitosan mats with an average diameter of 100 nm, whichwas maintained after exposure to the growth media (Fig. 3a). Itis noteworthy that there was a detectable difference in color ofthe recovered chitosan mats from the microalgae and thecontrol solutions. The initial light yellow chitosan mat turneddark brown after being incubated with algae (Fig. 3b), whereasno color change was observed for the control experimentwithout algae (Fig. 3c). The darker color is considered to beconsistent with the immobilization of the reduced Pd nano-crystals from the growth media, and corresponds to approxi-mately 1.03 wt% Pd loading on each rectangular piece ofelectrospun chitosan mat (3 6 2 cm), which was establishedusing inductively coupled plasma-optical emission spectro-scopy (ICP-OES). X-ray photoelectron spectrometric (XPS)analysis of the Pd loaded chitosan mats showed a dominatingpeak in the PdO 3d5/2 region (binding energy of 337.9 eV) andalso the presence of the Pd(0) peak in the Pd 3d5/2 region

(binding energy of 336.0 eV) (Fig. 3d). Analysis of the samplefollowing argon ion etching primarily showed a dominant Pd(0)peak in the Pd 3d5/2 region (binding energy of 335.9 eV) with anadditional lower peak of PdO in the 3d5/2 region (binding energyof 337.7 eV) (Fig. 3e). Thus the XPS data establish that thesurface of the palladium nanoparticles is partially oxidizedunder the aerobic culture conditions. Slight oxidation was alsoreported when palladium containing samples were exposed toan oxygen containing atmosphere before XPS analysis.33

Chlorine was not detected during these XPS analyses, indicatingthe absence of palladium(II) chloride species. The bindingenergies of both Pd 3d5/2 and PdO 3d5/2 were found to be slightlyhigher than the values for pure Pd metal,34 and PdO samples.35

This slight shift is reported to be common when the Pd/PdOnanoparticles are embedded in an insulating substrate,36 whichis the chitosan mat in the present case. In their study,Schildenberger and his colleagues also reported similar peaksfor Pd 3d5/2 (336.0 eV) and PdO 3d5/2 (337.9 eV) due to theisolated arrangement of metal nanoparticles on the layers ofoxidized silicon wafers.36

The utility of these palladium loaded chitosan mats were testedas catalyst supports for the standard Mizoroki–Heck reaction.37,38

Six dried mats (a total of 0.23% mol Pd per mol of iodobenzene)were introduced into a solution containing iodobenzene, butylacrylate, triethylamine and dimethylformamide (DMF). Thereaction temperature was kept constant at 80 uC for 16 h, andafter each reaction, the supported catalyst was recovered forrecycling studies after washing with DMF under nitrogen gas toavoid any oxygen induced regrowth of Pd(0) nanoparticles.39 Thecatalyst can be recycled at least four times with quantitativeconversion for each cycle, with the conversion yields of: 68% (1stcycle), 62% (2nd cycle), 45% (3rd cycle) and 36% (4th cycle) byweight. For the control comparison, a mat exposed to a Na2[PdCl4]solution without algae resulted in a conversion yield of only 5%.We have previously reported that a chitosan mat containingpalladium nanoparticles generated by reduction of pre-absorbedNa2[PdCl4] is also an effective catalytic support with no appreciableleaching of the metal.31 Conventional palladium catalysts usuallyrequire a 1–5 mol% loading rate for effective Mizoroki–Heck cross-coupling reactions.40 Our biogenic palladium nanocatalysts havehigh catalytic activity (68%) with respect to commercial material(5%), even at low palladium loadings (0.23 mol%), which issignificant for applications in the fine chemical industries.

In conclusion we have established a biogenic synthesis ofpalladium nanocrystals in the presence of Chlorella vulgaris andtheir subsequent immobilization on an electrospun chitosan matas a novel support for application in catalysis. In addition, we havedemonstrated that NADPH involved in the photoautotrophicmicroalgal metabolism is likely to play a role in the biogenicsynthesis. Utilization of easy-to-grow, nontoxic and environmen-tally-available microalgae for the synthesis of palladium haspotential for the development of green chemistry processes forother metals.

Fig. 3 (a) SEM image of the as-prepared electrospun chitosan nanofibers beforebeing exposed to palladium solution (scale bar: 1 mm). Mats of this material placedinto 4 week old 25 mg L21 Na2[PdCl4] solutions with and without microalgae whichwere collected after two weeks are shown in (b) and (c) respectively. Curve-fitted Pd3d XPS spectra for (d) as collected and (e) etched internal surface of a chitosan matunder argon ions for 60 s, both showing Pd(0) (blue curve) and PdO (red curve).

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Acknowledgements

We kindly acknowledge support of this work by the AustralianResearch Council (ARC) and internal grants of The Universityof Western Australia. The microscopy analysis was carried outusing the facilities in the Centre for Microscopy,Characterization and Analysis (The University of WesternAustralia). The authors also would like to thank Dr L. Byrne forhis kind help during the NMR analysis.

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8 S. Mann, Biomimetic Materials Chemistry, VCH, Weinheim, NewYork, 1996.

9 P. Yong, N. A. Rowson, J. P. G. Farr, I. R. Harris and L.E. Macaskie, Biotechnol. Bioeng., 2002, 80, 369–379.

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Supplementary Information

Biogenic production of palladium nanocrystals using microalgae and

their immobilization on chitosan nanofibers for catalytic applications

Ela Eroglua,b, Xianjue Chena, Michael Bradshawa, Vipul Agarwala , Jianli Zoua, Scott G. Stewartc,

Xiaofei Duand, Robert N. Lambd, Steven M. Smithb*, Colin L. Rastona*, and K. Swaminathan Iyera

a Centre for Strategic Nano-Fabrication, School of Chemistry and Biochemistry, The University of Western

Australia, Crawley, WA 6009, Australia, E-mail: [email protected] b ARC Centre of Excellence in Plant Energy Biology, The University of Western Australia M313, 35 Stirling

Highway, Crawley, WA 6009, Australia, E-mail: [email protected] c School of Chemistry and Biochemistry, The University of Western Australia, 35 Stirling Highway, Crawley,

WA 6009, Australia d Surface Science & Technology Group, School of Chemistry, The University of Melbourne, VIC 3010, Australia

S1. Biosynthesis of Palladium Nanoparticles

Wild type Chlorella vulgaris cultures (from the Australian National Algae Culture Collection

at CSIRO, Tasmania) were used as the green microalgae source for the biosynthesis of

palladium. Sterile algal freshwater media (MLA media)[1] containing standard micronutrients,

nitrate, phosphate, carbonate buffer and vitamins was used as the substrate source. The

microalgae cultures were mixed with Na2[PdCl4] solution at various concentrations (100, 50,

25, 12.5, 0 mg/L), towards reaching an initial total-chlorophyll content of around 1.8 mg/L

(Figure 1a).

Na2[PdCl4] solution was prepared by dissolving Na2[PdCl4] powder (Sigma-Aldrich) in

distilled and sterilized water overnight, and subsequently filtered through a sterile Pall®

Acrodisc® 32 mm syringe filter (0.2 µm membrane) for further sterilization. Concentrations

of Na2[PdCl4] in microalgae solutions are reported here as the initial concentrations

measured by the gravimetric analysis before the filtration process. The experiments were

conducted under batch conditions and cyclic diurnal conditions (16 h light/8 h dark) at a

constant temperature (25 °C). Algae cultures (total liquid volume of 40 mL) were grown in

250 mL Erlenmeyer flasks, under continuous cool-white fluorescent illumination at an

incident intensity of around 200 µmol photons m-2s-1(PAR) upon orbital shaking (Thermoline

Electronic Supplementary Material (ESI) for RSC AdvancesThis journal is © The Royal Society of Chemistry 2012

Scientific) at 120 rpm.

Algal growth was investigated by measuring the total chlorophyll content (Chl a + Chl b)

with respect to time, following the spectrophotometric method involving methanol

extraction.[2] Chlorophyll content was also used to validate the viability of the cell cultures.

All experiments were conducted in triplicates, with the standard deviation of each value given

in the form of error bars within the related figure.

S2. Characterization

A JEOL 2100 TEM instrument operating at 80 kV was used for determining the size and

morphology of chitosan nanofibers and palladium nanoparticles. Samples were prepared by

inserting the solutions on top of carbon-coated 200 mesh copper grids and allowing them to

dry. High resolution images were obtained using a JEOL 3000F instrument operating at 300

kV. A Zeiss 1555 VP-FESEM with a 3 kV accelerating voltage was used to image samples

coated with platinum (~3nm). XPS data was acquired using a VG ESCALAB220i-XL X-ray

Photoelectron Spectrometer equipped with a hemispherical analyzer. The incident radiation

was monochromatic Al Kα X-rays (1486.6 eV) at 220 W (22 mA and 10kV). Survey (wide)

and high resolution (narrow) scans were taken at analyzer pass energies of 100 eV and 50 eV,

respectively. Survey scans were carried out over 1200-0 eV binding energy range with 1.0 eV

step size and 100 ms dwell time. Narrow high resolution scans were run over a 20 eV binding

energy range with 0.05 ev step size and 250 ms dwell time. Base pressure in the analysis

chamber was 4.0x10-9 mbar and during sample depth profile analysis 1.5x10-7 mbar. A low

energy flood gun (~6 eV) was used to compensate the surface charging effect. Argon ions at

3 keV beam energy were used to sputter off approximately 18 nm surface layers at a rate of

~3 Angstrom/second. The ion source gave a crater of approximately 3x3 mm. The energy

calibration was referenced to the C 1s peak at 284.7 eV.

S3. Cross Coupling Reactions

A rectangular piece of electrospun chitosan mat (3 x 2 cm per each piece) was placed into

four week old 25 mg/L Na2[PdCl4] containing microalgae solutions (total volume: 40 mL per

flask) for two weeks. The amount of palladium was established using inductively coupled

plasma - optical emission spectroscopy (ICP-OES), at ~1.03 % w/w per each rectangular mat

(3x2 cm). ICP-OES analysis was acquired with ARL-3520B sequential scanning ICP-OES,

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with a 20mm torch, 1200W incident power, and a MDSN nebuliser. For the Mizoroki-Heck

reactions iodobenzene (55 µL; 0.49 mmol), butyl acrylate (85 µL; 0.59 mmol), triethylamine

(171 µL; 1.23 mmol) in dimethylformamide (DMF, 0.5 mL) were added to 6 hybrid

palladium nanoparticle-chitosan mats with a total of 0.23% mol Pd per mol iodobenzene.

The reaction mixture was heated to 80°C for 16 hours, whereupon the catalyst mat was

filtered and recovered for its recycling after washing with DMF five times under nitrogen

gas. Quantitative conversion yields were assessed gravimetrically upon the confirmation of

the final product (butyl cinnamate) using 1H and 13C NMR spectroscopy (Varian® 400

NMR).

S4. Electrospinning of Chitosan

Pd was collected by electrospun chitosan mat following the electrospinning procedure

optimized by Bradshaw et al. (2011)[3], which was slightly modified from the original

protocol given by Ohkawa et al. (2004)[4]. The variables of electrospinning processes were as

follows: (i) syringe pump speed: 0.1 mm/min, (ii) voltage: 18 kV, (iii) distance between the

target and the tip of the syringe: 11 cm, (iv) target speed: 1 m/ min, (v) traverse speed: 0.5

cm/min.[3] Chitosan (2-amino-2-deoxy-(1-4)-β-D-glucopyranose), 75-85% deacetylated

(Sigma-Aldrich), powder (6% wt) was mixed with TFA (trifluoroacetic acid): DCM

(dichloromethane) solution (70:30 v/v), and stirred overnight for its complete dissolution.

These ratios were taken to be optimum for attaining ultrafine chitosan nanofibers (Bradshaw

et al. 2011).[3] TFA (trifluoroacetic acid) and DCM (dichloromethane) were obtained from

Chem Supply. Before the electrospinning process, the mixture was sonicated for 15 min and

5.4% v/v glutaraldehyde (25% in H2O, Sigma-Aldrich) was immediately added to the

chitosan/TFA/DCM solution for an effective crosslinking. Electrospinning of this cross-

linked solution created insoluble nanofibers. Two rectangular pieces of electrospun chitosan

mat (3 x 2 cm per each piece) were placed into one flask of (four weeks old) 25 mg/L

Na2[PdCl4] containing microalgae solutions (total volume of 40 mL per flask) and left inside

for two weeks. The amount of palladium was established by using ICP-OES, yielding ∼1.03

% w/w per each rectangular mat (3x2 cm).

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S5. Reduction of Palladium with NADPH

Figure S1. High resolution TEM images of Pd(0) nanoparticles after the inoculation of

Na2[PdCl4] precursor with an excess amount of NADPH, dissolved inside MLA algal

growth-media,[1] indicating high crystallinity with (111) lattice constant of around 0.22 nm

(inset: FFT pattern corresponding to the area shown with red rectangle).

References

1. R.A. Andersen, in Algal Culturing Techniques, Elsevier Academic Press, 2005.

2. H.K. Lichtenthaler and C. Buschmann, in Current protocols in food analytical chemistry, eds:

R.E. Wrolstad RE, John Wiley & Sons Inc., New York, 2001, pp. F4.3.1- 8.

3. M. Bradshaw, J. Zou, L. Byrne, K.S. Iyer, S.G. Stewart, C.L. Raston, Chem. Commun. 2011, 47,

12292-12294.

4. K. Ohkawa, D. I. Cha, H. Kim, A. Nishida, H. Yamamoto, Macromol. Rapid Commun. 2004, 25,

1600–1605.

Electronic Supplementary Material (ESI) for RSC AdvancesThis journal is © The Royal Society of Chemistry 2012

Green Chemistry Dynamic Article Links

Cite this: Green Chem., 2012, 14, 2682

www.rsc.org/greenchem COMMUNICATION

Nitrate removal from liquid effluents using microalgae immobilized onchitosan nanofiber mats

Ela Eroglu,a,b Vipul Agarwal,a Michael Bradshaw,a Xianjue Chen,a Steven M. Smith,*b Colin L. Raston*a

and K. Swaminathan Iyera

Received 25th June 2012, Accepted 8th August 2012DOI: 10.1039/c2gc35970g

Mats of electrospun chitosan nanofibers were found effectivein immobilizing microalgal cells. These immobilized micro-algal cells were also used as a durable model system forwastewater treatment which has been demonstrated by theremoval of around 87% nitrate from liquid effluents ([NO3

−-N]initial 30 mg L−1). Virtuous nitrate removal rates werederived by superimposing physicochemical adsorption andbiological nutrient consumption phenomena for chitosan andmicroalgae, respectively.

Wastewater treatment focuses on eliminating unwanted chemi-cals and/or biological impurities from contaminated water. Ingeneral the treatment methods are mainly based on the separationof pollutants from the wastewater with a requirement for afurther processing stage to eliminate these pollutants.1 Integratedwastewater treatment processes are important in eliminatingundesired species, ideally converting them into valuable pro-ducts. As relatively recent bioprocesses, algal cultivation inwastewaters has a combination of several advantages such aswastewater treatment and simultaneous algal biomass pro-duction, which can be further exploited for biofuel production,food additives, fertilizers, cosmetics, pharmaceuticals, and othervaluable chemicals.2 However, there are inherent difficultiesassociated with algae-based bioprocessing in the harvesting,dewatering and processing of the algal biomass.

Immobilization of the cells on solid surfaces confer advan-tages over free cells in suspension, namely the immobilized cel-lular matter occupy less space, require smaller volume of growthmedium, are easier to handle, and can be used repeatedly forproduct generation.3,4 Moreover, immobilization can alsoincrease the resistance of cell cultures to harsh environmentalconditions such as salinity, metal toxicity and variations inpH.3,4 Entrapment is one of the most common immobilizationmethods which involves capturing the cells in a three dimen-sional gel matrix, made of polymeric materials or inorganicspheres.4 Both synthetic polymers (e.g., acrylamide,

polyurethane, polyvinyl) and natural polymers (e.g., collagen,agar, cellulose, alginate, carrageenan) have been used for thispurpose.5 Several studies have been reported on wastewater treat-ment involving the entrapment of microalgae cultures insidealginate beads, porous glass, and several synthetic polymers.6–8

However, most attempts to immobilize viable algae cells insidesuch insoluble materials have limitations, with the encapsulatingmaterials having volume/surface ratios usually orders of magni-tude larger than thin films. As a consequence, algal viability ismostly reported to decrease which relates to the need for thenutrients or reactants to diffuse far into the material to reach thealgal cells.5 We have developed a new technique to overcomethese problems using electrospun nanofiber mats as the matrixfor immobilizing the algal cells, with an overall strategy tocombine wastewater treatment processing with algal harvestingin a single process.

Electrospun nanofibers of chitosan were employed as apolymer/matrix support for green microalgae in the currentstudy. Chitosan is composed of D-glucosamine and N-acetyl-D-glucosamine, and is formed by the deacetylation of chitin (β-N-acetyl-D-glucosamine polymer).9 Chitin is usually extracted fromthe exoskeletons of crustaceans (e.g., crab, lobster and shrimp)and even from the cell walls of fungi.9,10 Chitosan is non-toxicand biodegradable, and can be used as an animal feed.9 Further-more, it can be used as a coagulant for wastewater treatment andfor the recovery of waste sludge.11

The removal of nitrate ions is regulated by law which relatesto its hazardous effects on human health and the environment.Several methods have been reported for the removal of nitratefrom water bodies, including biological denitrification,12 chemi-cal reduction,13 electrodialysis,14 and a combined bioelectro-chemical/adsorption process.15 In this study, we aimed tosuperimpose the treatment efficiencies of microalgae in reducingnitrate and electrostatic binding of the nitrate ion by chitosan, i.e.combining biological and chemical processing.

Chlorella vulgaris cultures were used as the green microalgaewhich were originally obtained from the Australian NationalAlgae Culture Collection at CSIRO, Tasmania. Cell growth wascarried out under artificial diurnal-illumination (16 h light/8 hdark cycle) at around 22 °C. Electrospun chitosan mats werefabricated following the procedure optimized in previousstudies.16,17 This involved dissolving chitosan powder (6 wt%)in a mixture of trifluoroacetic acid (TFA) and dichloromethane(DCM) (70 : 30 v/v), with 5.4% (v/v) addition of glutaraldehyde

aCentre for Strategic Nano-Fabrication, School of Chemistry andBiochemistry, The University of Western Australia, M313, 35 StirlingHighway, Crawley, WA 6009, Australia. E-mail: [email protected];Tel: +61 8 6488 3045; Fax: +61 8 6488 8683bARC Centre of Excellence in Plant Energy Biology, The University ofWestern Australia M313, 35 Stirling Highway, Crawley, WA 6009,Australia. E-mail: [email protected]; Tel: +61 8 6488 4403;Fax: +61 8 6488 4401

2682 | Green Chem., 2012, 14, 2682–2685 This journal is © The Royal Society of Chemistry 2012

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solution (25% in H2O) immediately prior to electrospinning.17

This was deemed necessary to cross-link the chitosan to avoidpolymer breakup and dissolution. The experimental settings ofelectrospinning processes were as follows: (i) syringe pumpspeed: 0.1 mm min−1, (ii) voltage: 18 kV, (iii) distance betweenthe target and the tip of the syringe: 11 cm, (iv) target speed:1 m min−1, (v) traverse speed: 0.5 cm min−1. Fiber mats wereannealed overnight to remove any remaining solvent and storeduntil required. After electrospinning, 2 mL of algae solution inits exponential phase of growth with a total chlorophyll content(Chl a and b) of ∼2 mg L−1, was placed onto a cut out rectangu-lar chitosan-mat (3 × 2 cm) and left at room temperature forapproximately 48 hours to allow sufficient attachment of algaecells to the surface of the mat. Microalgae cell walls containvarious polysaccharides, which are compatible with the surfaceof the chitosan nanofibers.18,19 The presence of negative surfacecharge on the surface of Chlorella cells, arising from dissociationof uronic acid groups, and/or the presence of sulfate groups forexample,18 provide electrostatic attraction to the positivelycharged primary amine groups of chitosan not involved in theabove cross linking. Moreover, the negatively charged surface ofthe microalgae can also result in binding metal ions, thereby pro-viding an opportunity for biosorption applications, along withthe removal of nitrate ions as functional algal cells.18–20

This bionano-composite material was then placed into nitratecontaining artificial growth medium which contained mainlyphosphates, nitrates, carbonate buffer, micronutrients and vita-mins,21 with an initial nitrate-nitrogen concentration of around30 mg L−1. Nitrate-nitrogen (NO3

−-N) term refers to the amountof nitrogen (N) in liquid solutions coming from nitrate ions(NO3

−). This nitrate-nitrogen concentration is within the rangeof other algal nitrate removal studies.2,22 Furthermore, it simu-lates the range of nitrate content present in ground-waters(∼0.1–50 mg L−1 NO3

−-N)23 and sewage treatment planteffluents. The regulatory limit for the maximum contaminantlevels of [NO3

−-N] in public drinking water is 10 mg L−1, asestablished by the United States Environmental ProtectionAgency (EPA).24,25

During the algal growth, the pH of the medium was keptaround 6.5–7.0 by the addition of dilute hydrochloric acid (HCl)when necessary. The colorimetric “cadmium reduction method”was employed for the nitrate-nitrogen analysis, using chemical-kits in the form of powder pillows (HACH®, NitraVer NitrateReagent) and a colorimeter (HACH® DR/870).26 For compara-tive purposes, a control experiment with a chitosan nanofibermat devoid of algae culture was also treated under the same con-ditions. Scanning electron microscopy (SEM) analyses wereacquired using a Zeiss 1555 VP-FESEM, while the acceleratingvoltage was changed between 3 to 5 kV. The air-dried sampleswere coated with approximately 3 nm layer of platinum beforeimaging. A NanoMan AFM system (Veeco Instruments Inc.)was used for the atomic force microscopy (AFM) analysis, oper-ating under the tapping mode. Chlorophyll content of the cellswas analyzed using spectrophotometric measurements of metha-nol extracts obtained from the algal culture pellets.27

A challenge in the present work was to fabricate an insoluble,fibrous structure with sufficient porosity, which can facilitate thediffusion of nutrients and cellular products between the environ-ment and the algae. Fig. 1a and b show scanning electron

microscopic (SEM) images of the nanofiber structure of thechitosan mats after the electrospinning process. The diameter ofthe fibers was between 50 to 180 nm, with an average diameterof around 91 nm. After placing the nanofiber mat into aqueoussolution, the fibers gradually swelled with a significant increasein porosity of the material, Fig. 2a, and became an effectivesupport matrix for the C. vulgaris cells, which have the expecteddiameter from SEM images around 3–4 μm, Fig. 2b and c. Thisporous structure has an advantage for facilitating the diffusion ofmaterials such as nutrients and waste products between theenvironment and the algae, while the replication of algal cells isaccomplished on the surface of the nanofiber mat, Fig. 2c. Theheight profile measurements obtained by AFM analyses esta-blished the thickness of the chitosan mat at close to 400 nm,Fig. 3a, with the overall thickness for the algae attached chitosanmat at 4.3 μm, Fig. 3b. The difference in height is consistentwith the attachment of a single layer of individual C. vulgariscells on the nanofiber mats.

Fig. 4 shows the optical images of algae cells attached to thesurface of chitosan mats (a) initially, (b) after 3 days, and (c)after 10 days from the start of the growth experiments in 40 mLliquid media. Note the increasing green color on the surface isdue to the increased concentration of algal cells on the chitosanmat with respect to time. Detailed imaging is given as SEMimages in Fig. 5. The amount of algae cells on a chitosan matwith same dimensions yielded around 4 cells per 100 μm2 for a3 day old sample, Fig. 5c, whereas this increased to around

Fig. 2 (a) SEM images of porous and swollen chitosan nanofiber mats,after exposure to nitrate containing media for two days, (b) SEM imagesof chitosan nanofibers surrounding individual, and (c) multiple algaecells. Scale bars are given as 1 μm.

Fig. 1 SEM images of as prepared electrospun chitosan nanofibers atlow and high magnifications with scale bars of (a) 1 μm, and (b) 10 μm,respectively.

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20 cells per 100 μm2 by the 10th day of the treatment process,Fig. 5d.

Fig. 6 shows the nitrate-nitrogen concentration versus time inthe absence of algae (triangles) or algae attached (rectangles)nanofiber mats. After the insertion of this bionano-compositeinto the liquid media (Vtotal: 40 mL), around 30% of the initialnitrate value was decreased within the first 2 days. This reductionin nitrate is mainly caused through the uptake by the chitosannanofibers rather than the algal cultures, as a physicochemicaladsorption process. A similar pattern was also observed for themats devoid of algal cells where after the second day there wasno further nitrate removal. In contrast the algae containing matscontinued their nitrate uptake, being used in their cellular meta-bolism for replication, building more biomass and energyproducts.

Amino groups in chitosan are protonated at acidic to neutralpH conditions,28 which enhance the adhesive properties of chito-san by increasing its tendency to attach negatively charged enti-ties which in this case are algal cell walls and nitrates. Chlorellavulgaris cell walls are known to be highly negatively chargedwith a zeta potential of around −30 mV in neutral water.29 Onthe other hand, the zeta potential of the positively charged chito-san nanofibers is +20 mV at neutral pH. Matsumoto et al.30

reported the zeta potential values of chitosan nanofibers to behighly dependent on the pH of the media, with it increasingto +30 mV in pH around 5–6, whereas it drops to zero for pHvalues above 8.30 Algal growth tends to alkalify its medium, asthe cellular uptake of anions (such as nitrates, phosphates, car-bonates, etc.) is stabilized with equivalent amounts of hydroxyl(OH−) anion efflux.31 For this reason, we maintained the pH ofthe culture around 6.5–7 by the regular addition of dilute HClduring the current study. Due to the nature of HCl, several otheracidifying agents (such as CO2) can be considered for any futuredevelopments and advanced scale-up processes for municipaland/or industrial wastewater samples.

Clearly the presence of the nanofiber mat in the liquidenvironment is responsible for the initial removal of nitrate whilethe continued growth of algae subsequently consumes theremaining nitrate in further stages with a slower rate. Overallnitrate removal rates were calculated as 32 ± 3%, and 87 ± 4%,for the “microalgae-absent” and “microalgae-attached” chitosanmats, respectively. Several studies have already been reported onwastewater treatment with immobilized microorganisms. Fierroet al.22 investigated the effect of nitrate removal by Scenedesmusspp. cyanobacterial cells immobilized within spherical chitosanbeads. They achieved 70% nitrate removal for the immobilizedcultures, while 20% of the initial nitrate content was removed bythe blank chitosan beads. In another study, Mallick and Rai32

also achieved relatively higher nitrate removal rates (73%) byAnabaena doliolum and Chlorella vulgaris cells immobilized inchitosan beads compared to the cells immobilized in other typesof gels made of alginate, carrageenan or agar. De-Bashan et al.33

achieved only 15% nitrate removal for co-immobilized

Fig. 3 AFM topographic mapping of chitosan nanofiber mats(4 × 4 μm) without, (a) and with, (b) algal cells.

Fig. 4 Progress of the algal growth on the surface of chitosan mats: (a)initially, (b) after 3 days, (c) after 10 days of the growth experiment.

Fig. 5 SEM images of immobilized C. vulgaris cells on the surface ofchitosan nanofiber mats after different time intervals. (a and c) are formats after 3 days, and (b and d) are for mats 10 days old.

Fig. 6 Nitrate-nitrogen (NO3−-N) concentration (mg L−1) of algal

medium versus time. Chitosan nanofibers devoid of algal cells are rep-resented with triangles, whereas those with immobilized algal cells areshown as rectangles.

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microorganisms (Chlorella vulgaris with a growth-promotingbacterium Azospirillum brasilense) within alginate beads. At theother end of the scale, Tam and Yong34 reported complete nitrateremoval using immobilized C. vulgaris cells within calcium algi-nate beads. The treatment efficiency of our current method iscomparable with that of these aforementioned methods, althoughlarge variations among the experimental parameters; includingwastewater composition, microbial species, duration of theprocess, type of bioreactor, chemical composition and shape ofthe immobilization matrix, make direct comparison difficult.

In summary, we have established the use of cross-linked chito-san nanofiber mat as a water-insoluble and non-toxic support foralgal growth and nitrate removal from waters. Algal growth on asupport material can lead to combine algal harvesting, dewater-ing, and processing steps in a single stage. This bionano-compo-site material is potentially an attractive, simple and highlydurable polymer, with the mats still retaining their integrity aftersix months in contact with an aqueous solution, and has promisefor industrial and/or municipal wastewater treatment processes.

Acknowledgements

This work has been supported by The University of WesternAustralia and the Australian Research Council. We would liketo acknowledge the facilities of the Australian Microscopy &Microanalysis Research Facility at the Centre for Microscopy,Characterization & Analysis, The University of Western Austra-lia, which was funded by the University, State and Common-wealth Governments. AFM images were performed at CurtinUniversity.

Notes and references

1 Metcalf and Eddy, Inc, Wastewater Engineering: Treatment and Reuse,McGraw-Hill, New York, 4th edn, 2003.

2 N. Mallick, BioMetals, 2002, 15, 377.3 L. Hall-Stoodley, J. W. Costerton and P. Stoodley, Nat. Rev. Microbiol.,2004, 2(2), 95.

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6 K. Abe, E. Takahashi and M. Hirano, J. Appl. Phycol., 2008, 20, 283.7 G. Schumacher and I. Sekoulov, Water Sci. Technol., 2002, 46, 83.8 E. Delahaye, R. Boussahel, T. Petitgand, J. P. Duguet and A. Montiel,Desalination, 2005, 177, 273.

9 A. Lavoie and J. de La Noue, J. World Maricul. Soc., 1983, 14, 685.10 A. Zamani, L. Edebo, B. Sjöström and M. J. Taherzadeh, Biomacromole-

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E. Elabbassi, M. Hafsi and D. Largeteau, Desalination, 2001, 136, 325.15 Z. Feleke and Y. Sakakibara, Water Sci. Technol., 2001, 43, 25.16 K. Ohkawa, D. I. Cha, H. Kim, A. Nishida and H. Yamamoto, Macromol.

Rapid Commun., 2004, 25, 1600.17 M. Bradshaw, J. Zou, L. Byrne, K. S. Iyer, S. G. Stewart and

C. L. Raston, Chem. Commun., 2011, 47, 12292.18 D. Kaplan, D. Christiaen and S. M. Arad, Appl. Environ. Microbiol.,

1987, 53, 2953.19 R. H. Crist, K. Oberholser, N. Shank and M. Nguyen, Environ. Sci.

Technol., 1981, 15, 1212.20 B. Volesky and Z. R. Holan, Biotechnol. Prog., 1995, 11, 235.21 C. J. S. Bolch and S. I. Blackburn, J. Appl. Phycol., 1996, 8, 5.22 S. Fierro, M. del P. Sanchez-Saavedra and C. Copalcua, Bioresour.

Technol., 2008, 99, 1274.23 U. N. Dwivedi, S. Mishra, P. Singh and R. D. Tripathi, in Environmental

Bioremediation Technologies, ed. S. N. Singh and R. D. Tripathi,Springer, New York, 2007, ch. 16, pp. 353–389.

24 EPA, National Pesticide Survey: Project Summary, U.S. EnvironmentalProtection Agency, Washington DC, 1990.

25 S. Ghafari, M. Hasan and M. K. Aroua, Bioresour. Technol., 2008, 99,3965.

26 APHA, Standard Methods for the Examination of Water andWastewater, American Public Health Association, Washington, DC,18th edn, 1992.

27 H. K. Lichtenthaler and C. Buschmann, in Current Protocols in FoodAnalytical Chemistry, ed. R. E. Wrolstad, John Wiley & Sons Inc.,New York, 2001, pp. F4.3.1–F4.3.8.

28 L.-Q. Wu, P. Gadre Anand, H. Yi, M. J. Kastantin, W. Rubloff Gary,E. Bentley William, F. Gregory Payne and R. Ghodssi, Langmuir, 2002,18, 8620.

29 B.-M. Hsu, Parasitol. Res., 2006, 99, 357.30 H. Matsumoto, H. Yako, M. Minagawa and A. Tanioka, J. Colloid Inter-

face Sci., 2007, 310, 678.31 J. Naus and A. Melis, Plant Cell Physiol., 1991, 32, 569.32 N. Mallick and L. C. Rai, World J. Microbiol. Biotechnol., 1994, 10,

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Can. J. Microbiol., 2002, 48, 514.34 N. F. Y. Tam and Y. S. Wong, Environ. Pollut., 2000, 107, 145.

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HO ET AL . VOL. 9 ’ NO. 2 ’ 1767–1774 ’ 2015

www.acsnano.org

1767

January 26, 2015

C 2015 American Chemical Society

Hierarchical Patterning ofMultifunctional Conducting PolymerNanoparticles as a Bionic Platform forTopographic Contact GuidanceDominic Ho,†,‡ Jianli Zou,§ Xianjue Chen,†,0 Alaa Munshi,† Nicole M. Smith,†, ) Vipul Agarwal,†

Stuart I.Hodgetts,‡GilesW.Plant,^Anthony J.Bakker,‡AlanR.Harvey,‡ Igor Luzinov,# andK.Swaminathan Iyer*,†

†School of Chemistry and Biochemistry, The University of Western Australia, Crawley, Western Australia 6009, Australia, ‡School of Anatomy, Physiology and HumanBiology, The University of Western Australia, Crawley, Western Australia 6009, Australia, §Institute for Integrated Cell-Material Sciences (iCeMS), iCeMS Complex 2,Kyoto University, Yoshida-Honmachi, Sakyo-ku, Kyoto, 606-8501, Japan, )Experimental and Regenerative Neurosciences, School of Animal Biology, The University ofWestern Australia, Crawley, Western Australia 6009, Australia, ^Stanford Partnership for Spinal Cord Injury and Repair, Department of Neurosurgery,Stanford University School of Medicine, Stanford, California 94305, United States, and #School of Materials Science and Engineering, Clemson University, Clemson,South Carolina 29634, United States. 0Present address: Centre for NanoScale Science and Technology, School of Chemical and Physical Sciences, Flinders University,Bedford Park, Adelaide, SA 5042, Australia.

Exogenous electrical stimulation hasbeen effectively used both in clinicalpractice and in laboratory research

to regulate cell-type-dependent adhesion,differentiation, and growth.1 This phenom-enon of introducing programmed electricalsignals locally to influence biological eventshas resulted in the pursuit of sophisticatedmedical bionic devices.2 An important prop-erty that dictates the performance of mostbionic electrodes is the electrode/cellularinterface and its ability to transmit chargeacross the biointerface.3 Traditionally metal-lic electrodesmadeofplatinum, gold, iridiumoxide, tungsten, their alloys, and morerecently carbon fibers have been effectively

employed in bionic devices.4 They havebeen employed for deep brain stimulation,as cochlear implants, for vagus nerve stimu-lation to treat epilepsy, and for stimulatingregeneration in the central nervous sys-tem.2 However, stiff metal electrodes sufferamajor drawback of eliciting tissue damageover long-term implantation.2 Importantly,it is now recognized that nanoscale patternsprovide topographic guidance cues forcells. This has been widely exploited toengineer sophisticated regenerative plat-forms for nerves, muscles, skin, and bones.5

The need to incorporate large-area nano-scale patterns for bionic applications coupledwith the demand toward miniaturization of

* Address correspondence [email protected].

Received for review November 20, 2014and accepted January 26, 2015.

Published online10.1021/nn506607x

ABSTRACT The use of programmed electrical signals to influence

biological events has been a widely accepted clinical methodology for

neurostimulation. An optimal biocompatible platform for neural

activation efficiently transfers electrical signals across the electrode�cell interface and also incorporates large-area neural guidance

conduits. Inherently conducting polymers (ICPs) have emerged as

frontrunners as soft biocompatible alternatives to traditionally used

metal electrodes, which are highly invasive and elicit tissue damage

over long-term implantation. However, fabrication techniques for the

ICPs suffer a major bottleneck, which limits their usability and medical translation. Herein, we report that these limitations can be overcome using colloidal

chemistry to fabricate multimodal conducting polymer nanoparticles. Furthermore, we demonstrate that these polymer nanoparticles can be precisely

assembled into large-area linear conduits using surface chemistry. Finally, we validate that this platform can act as guidance conduits for neurostimulation,

whereby the presence of electrical current induces remarkable dendritic axonal sprouting of cells.

KEYWORDS: multimodal nanoparticles . conducting polymers . capillary force lithography . neurostimulation

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biocompatible implantable devices has resulted inthe emergence of inherently conducting polymers asfrontrunners for fabricating flexible organic electrodematerials. However, advances in the applicability ofpatterned surfaces of inherently conducting polymersin bionic devices have been limited due to the diffi-culties of transferring printing techniques and theirintegration under physiological conditions. In thisarticle, we report a transferable method to fabri-cate multifunctional poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonate) (PEDOT:PSS) nanoparticles anddirect their self-assembly by electrostatic interactionsinto large-area patterns. Using the rat pheochromocy-toma cell line (PC12), we demonstrate the suitabilityof the assembly as a bionic platform for exogenouselectrical stimulation.The three primary classes of conducting polymers

that have been studied are polyanilines, polypyrroles,and polythiophenes.6 The ease of functionalization ofpolythiophenes and maintenance of conductivity un-der physiological conditions has made them primarycandidates for multifunctional organic bionic devices.7

Themost widely explored processes for the fabricationof organic conducting polymer patterns are electro-polymerization, extrusion printing, inkjet printing,microcontact printing, electrospinning, and morerecently high-precision Dip Pen Nanolithography(DPN).4,6 Electropolymerization has been widely usedfor coating metal/carbon substrates, following whichpatterning is achieved by top-down lithography onpolymer thin films covering larger area electrodes. Thistechnique results in controlled, high-resolution nano-scale patterns but is limited by the ability to regulatepolymerization of monomers on nanoscale implanta-ble electrodes.8 Similarly, printing techniques haveachieved significant advances in recent years, reachinghigh-throughput patterns, but are limited in resolutionby the liquid dispensing techniques, which operatewithin the limit of tens of micrometers.9 Electrospin-ning techniques have offered simple processable solu-tions to generate 3D scaffolds at resolutionsmimickingthe extracellular matrix architecture but are limitedby the inability to generate patterned conductingconduits for the development of bionic guidancechannels.4 The aforementioned shortfalls have beenrecently overcome by the advances of DPN, whichenables precise deposition, patterning down to nano-scale resolution, and most importantly applicabilityover a wide range of substrates.10 However, advancesare limited by their cost, need for specialized equip-ment, and low throughput. In the present paper weadopt a bottom-up self-assembly process to preciselypattern conducting polymer nanoparticles into pat-terns as conduits for guidance. The approach is easilyadoptable over multiple substrates, needs no specia-lized equipment, and affords large-area patterns. Impor-tantly, this approach enables drug encapsulation and

sustained release from the nanoparticles once pat-terned and multimodal imaging of the nanoparticleconstructs once implanted.

RESULTS AND DISCUSSION

Patterned Multifunctional PEDOT:PSS Nanoparticle Arrays.In this study poly(glycidal methacrylate) (PGMA) isused as a reactive macromolecular anchoring platformboth on the substrate as a nanoscale layer and as acolloidal nanoparticle to enable multilayer assembly(Figure 1). A polymer with epoxy functionality waschosen, since the reactions of epoxy groups are uni-versal and easily transferable to various substrates,affording ease of attachment of functional molecules.Furthermore, the epoxy groups of the polymer cancross-link to provide structural integrity to the patternand nanoparticle constructs.11 The mobility of thereactive loops of PGMA ensures greater access toanchoring, resulting in a 2�3-fold greater graftingdensity when compared to a monolayer of epoxygroups on a nanoparticle surface of similar dimension,enabling high loading using a layer by layer approachthat is adopted in the current study.11 Polymer nano-spheres were initially prepared using an oil in wateremulsion methodology from PGMA modified with arhodamine-B (RhB) dye, encapsulated with magnetite(Fe3O4) nanoparticles to form the core platform(Figure 1a,b). Not only does the incorporation ofmagnetite and RhB render these constructs multi-modal for both MRI and fluorescence imaging, butimportantly in the present case magnetite provides ameans to separate, wash, and purify the nanoparticlesusing a magnetic fractionation column during eachstep of layered assembly. Polyethylenimine (PEI) wasthen covalently bound to the RhB-PGMA core to facil-itate a cationic layer for electrostatic conjugation of ananionic conducting polymer, PEDOT:PSS (Figure 1c,d).Capillary force lithography (CFL) was then usedto generate large-area nanoscale conduits in whichPEDOT:PSS nanoparticles are electrostatically directedto self-assemble as linear channels from solution(Figure e,f). Capillarity allows the polymer melt to fillup the void space between the polymer and theapplied mold when the temperature is above theglass-transition temperature (Tg), thereby generatinga large-area pattern that depends on the size of stamp.Importantly, the technique needs no specializedinstrumentation for generation of large-area patterns.Patterns can easily be generated using polydimethyl-siloxane (PDMS) stamps, which in turn can be fabri-cated using the ubiquitous optical storage discs as amaster. An optical data storage disc is typicallymade ofa polymer (polycarbonate) disc, onwhich a single spiraltrack is drilled. The typical width and depth of each linein the spiral track are 800 and 130 nm, respectively, andthe periodicity of the track is∼1.5 μm (Figure S1). In thepresent study, an indium tin oxide (ITO) substrate was

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modified first by spin coating a thin film of PGMAfollowed by a second spin-coated layer of polystyrene(PS) using previously reported conditions.12 The PSlayer acts as a chemical resist to selectively react theepoxy groups of PGMA following patterning. ThePS/PGMA bilayer was annealed with the PDMS maskat 130 �C (T> Tg of PS) to induce patterning via capillaryflow. The reusable PDMS stamp was peeled off fol-lowing heat treatment to obtain a patterned sur-face resulting in alternating PGMA and PS stripes.

Ethylenediamine (EA) was then grafted to PGMA toresult in cationic linear patterns. PEDOT:PSS nano-particles were then electrostatically assembled ontothe patterned surface, followed by washing steps toremove PS to obtain linear arrays of assembled PEDOT:PSS nanoparticles. A detailed schematic of the fabrica-tion process is shown in Figure S2. The nanoparticleand the patterns were characterized at each step of theassembly (Figure 2). The PEDOT:PSS nanoparticleswere an average size of 200 nm (Z-average) with a

Figure 1. Schematic illustration of the fabrication protocol to pattern multifunctional PEDOT:PSS nanoparticle arrays forexogenous electrical stimulation. (a�d) Multilayer assembly of conducting PEDOT:PSS nanoparticle fabrication via non-spontaneous emulsification. (a) An organic phase is initially formed by dissolving RhB-modified PGMA (yellow) and Fe3O4

(purple) in a 1:3mixture of CHCl3 andMEK. (b) Colloidal fluorescent PGMA-Fe3O4 nanoparticles are fabricated upon dropwiseaddition of the organic phase to an aqueous solution of Pluronic F-108. (c) Cationic second layer via covalent attachment ofPEI (green) to the PGMA-Fe3O4 core. (d) Anionic conducting polymer layer via electrostatic attachment of PEDOT:PSS (blue).(e�g) Patterning of the multilayered PEDOT:PSS nanoparticles for exogenous electrical stimulation of PC12 cells. (e) Linearnanoparticle conduits patternedon a substrate via capillary force lithography (CFL) using charge complementarity. A detailedschematic of the CFL procedure can be found in Figure S2. (f) PC12 cells (green) were cultured onto the biocompatibleplatform, followed by (g) exogenous electrical modulation.

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Figure 2. Characterization of the multilayered PEDOT:PSS conducting nanoparticles and their assembly as linear conduits.(a) TEMmicrograph of the multilayered PEDOT:PSS nanoparticles. Scale bar = 200 nm. (Inset: high-magnification TEM imageof PEDOT:PSS-coated nanoparticles showing encapsulated Fe3O4 nanoparticles. Scale bar = 10 nm.) (b) DLS particle sizedistributions of the PEDOT:PSS nanoparticles in solution. (c) Zeta potential distributions of the nanoparticles: PGMA-Fe3O4

core (black) with an average zeta potential of 3.9 ( 1.3 mV, cationic PEI-coated (red) with an average zeta potential of 37 (1.2 mV, and anionic PEDOT:PSS-coated (blue) with an average zeta potential of �29 ( 6.15 mV. (d) Current vs voltageresponse for the nonconducting PEI-coated nanoparticles (red) and conducting PEDOT:PSS-coated (black) nanoparticles.(e�g) Tappingmode AFM topography images of the nanoparticle patterns at each stage of fabrication: PGMA and PS stripes(e), EA-modified PGMA and PS stripes (f), PEDOT:PSS nanoparticle patterns (g). (h�j) Corresponding height profiles of thenanoparticle patterns at each stage of fabrication.: PGMA and PS stripes (h), EA-grafted PGMA and PS stripes (i), PEDOT:PSSnanoparticles patterns (j). The AFM line scans corresponding to the height profiles are indicated on the topography images in(e)�(g). (k, l) SEM micrographs of the nanoparticle patterns at a magnification of 25k� (k) and 11k� (l) indicating theformation of tightly packed and highly ordered nanoparticle arrays. (m) Confocal fluorescence image of the RhB-functionalized PEDOT:PSS nanoparticle arrays at 20� magnification.

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polydispersity index (PDI) of 0.07, a zeta potential of�29( 6.15mV, and a conductivity of 2.5� 10�12 S/cm.Themeasured conductivity is in accordance with othervalues reported in the literature for polymer blends.13

Importantly, this low conductivity is important underphysiological conditions to induce local cellular stimu-lation and avoid tissue damage due to toxic over-stimulation.14 The final self-assembled linear arraysof PEDOT:PSS nanoparticles were of large-area high-density packing, as confirmed at various length scalesusing AFM, SEM, and fluorescence imaging.

Biocompatibility Assessment of the PEDOT:PSS NanoparticleArrays. Topographic modulation of tissue response isone of the most important considerations in develop-ing bionic implants. Topographic contact guidanceusing micropatterns has been widely exploited toinfluence cell migration, adhesion, and prolif-eration.15,16 One of the pivotal first steps in the pres-ent study was to establish biocompatibility of thepatterned structures. PC12 cells were chosen in thepresent case, as they have been demonstrated to show

enhanced neurite outgrowth and spreading uponexogenous stimulation on a conducting polymersubstrate.17 MTS (3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazo-lium, inner salt) assays, cell viability assays, and SEMimaging were performed after exogenous electricalstimulation and in the absence of electrical stimulationto determine effects on cell viability and cell adhesion(Figure 3a,b and Figure S3). The stimulation conditionsused in the present study involved a monophasicpulsed current at a frequency of 250 Hz with a 2 mspulse width and an amplitude of 1mA for 2 h, similar toprotocols previously reported for similar cell lines.18,19

Importantly, we observed no changes in cell viabilityupon exogenous stimulation and observed preferen-tial adhesion of the PC12 cells to the patterned sur-face over a nonpatterned surface in both cases(( stimulation). High-magnification SEM imaging (nostimulation) further revealed preferential interaction ofthe PC12 cells to the PEDOT:PSS nanoparticle arrays,confirming not only biocompatibility with the large

Figure 3. Biocompatibility of the PEDOT:PSS nanoparticle arrays with PC12 cells. (a) Cell viability determined using MTScalorimetric assay obtained at 72 h after an initial exogenous electrical stimulation for 2 h and in the absence of stimulationshowing no significant changes. (b) SEM micrograph demonstrating preferential cell adhesion to the pattern area (yellowbox). Image acquired at 323�magnification 72 h after the addition of NGFwithout exogenous electrical stimulation. (c) High-magnification (12k� magnification) SEM images demonstrating specific and preferential interactions of neurites (whitearrows) with the PEDOT:PSS linear conduits (red arrows).

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area of the pattern but also potential applicabilityof the nanoscale linear arrays as guidance conduits(Figure 3c).

Exogenous Electrical Stimulation Induced Dendritic Sproutingof the PC12 Cells. Electrical stimulation has been effec-tively used to modulate growth and differentiation of

anchorage-dependent cells such as neurons, fibro-blasts, and epithelium cells.17,20,21 In the central ner-vous system, brief stimulation to the proximal end oftransected peripheral nerves has been shown to aug-ment preferential motor reinnervation,22 improve thespecificity of sensory reinnervation,23 and accelerate

Figure 4. Exogenous electrical stimulation induced dendritic sprouting of the PC12 cells guided by the PEDOT:PSS linearconduits. (a) Significant increase in the average cell area is observed 72 h after exogenous electrical stimulation on thenanoparticle platform in comparison to unstimulated and nonpatterned controls. (b) Corresponding decrease in PC12 cellproliferation observed 72 h after exogenous electrical stimulation on the nanoparticle platform in comparison tounstimulated and nonpatterned controls. (c�f) Representative confocal images (40� magnification) of β-III tubulinimmunohistochemically stained cells 72 h after the following treatments: {(þ) pattern, (�) stimulation} (c); {(þ) pattern,(þ) stimulation} (d); {(�) pattern, (�) stimulation} (e); {(�) pattern, (þ) stimulation} (f), demonstrating modulation of cellmorphology. (g) High-magnification SEM image (magnification 8k�) indicating the formationof extensive dendritic networks(white arrows) guided by the PEDOT:PSS arrays. Inset: The corresponding low-magnification image of the area (yellow box)analyzed (magnification 3k�).

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the reinnervation of distal target tissues.24 These havebeen reported to depend on depolarization of theneuronal soma and its axon, involvement of axonguidance factors such as polysilylated neural celladhesion molecule,25 the L2/HNK-1 carbohydrate,26

and brain-derived neurotrophic factor.27 Finally elec-trical stimulation induced neurite outgrowth wasrecently reported to be dependent on calcium influxthrough L- and N-type voltage-dependent calciumchannels and calcium mobilization from IP3R andRYR-sensitive calcium stores.28 In the present case,we analyzed the morphological modulation of PC12cells following electrical stimulation having deter-mined no change in cell viability using the MTS assay.Nerve growth factor (NGF) induces PC12 cells tochange their phenotype and acquire a number ofproperties that are similar to sympathetic neurons.Importantly, although they can acquire propertiessimilar to sympathetic neurons upon NGF treatment,they do not develop definitive dendritic axons or formtrue synapses with each other in the absence ofexogenous stimulation.29 This change in phenotypeupon NGF treatment is associated with a retardation inproliferation, the extension of neurites making themelectrically excitable. Monitoring the cell numbers andcell area can assess this change from the proliferationstate to a differentiation state. Furthermore, micro-tubule levels correlate precisely with the neurite

extension during NGF-induced PC12 cell differ-entiation.29,30 Using immunohistochemical stainingfor β-III tubulin it was determined that stimulation onthe patterned surface resulted in a significant increasein the cell area and lower number of cells per unit area,indicating exogenous electrical stimulation induceddifferentiation of PC12 cells (Figure 4a�f, Figure S4).High-magnification SEM (Figure 4g) also revealedthat stimulation resulted in an extensive dendriticnetwork guided by the linear conduits of PEDOT:PSSnanoparticles.

CONCLUSION

In summary, we have demonstrated a practical andtransferable protocol to fabricate self-assembled large-area patterns of conducting polymers from solution.This overcomes some of the shortfalls in the currentfabrication techniques in developing patterned organicbionic devices. The patterns generated have demon-strated excellent biocompatibility. At the same time,they have been shown to induce exogenous electricalstimulation under physiological conditions to elicit ameasurable and consistent cellular response. Impor-tantly the methodology permits the design of bionicdevices capable of inducing local electrical stimulationfor in vivo applications while integrating multimodalimaging and simultaneous drug delivery capabilities ofnanoparticles.

METHODS SUMMARY

Nanoparticle Synthesis. The conducting nanoparticles wereprepared via a nonspontaneous emulsification route. Briefly,rhodamine Bwas attached to PGMA inMEK at 80 �Cunder N2 for5 h. The modified PGMA was then precipitated in diethyl etherand dried under N2. This was dispersed in a 1:3 mixture of CHCl3and MEK along with 25 mg of Fe3O4 to form the organic phase.This organic phase was added dropwise into a rapidly stirringaqueous solution of Pluronic F-108. The emulsion was homo-genized with a probe-type ultrasonic wand for 1 min. Theorganic solvents were then evaporated off under N2. Largeaggregates of Fe3O4 and excess polymer were separated viacentrifugation. The nanoparticles in the supernatant were thenmixed with PEI and heated to 80 �C for 16 h to facilitateattachment. The PEI-coated nanoparticles were isolated andwashed on a magnetic separation column. Next, a dilutedsolution of PEDOT:PSS was added dropwise under rapid stirringto nanoparticles at a concentration of 0.5 mg/mL to facilitateelectrostatic attachment. This was followed by sonication for10min and stirring for 18 h. The nanoparticles were thenwashedmultiple times inwater before being stored at 4 �C for further use.

Platform Fabrication. To direct the self-assembly of the nano-particles, a template was fabricated by CFL. A 0.2%w/v PGMA inCHCl3 solution was spin coated on ITO coverslips and annealedat 120 �C for 20 min. Next, 1.3% w/v PS in toluene was spincoated onto the PGMA surface. A PDMS stamp was then placedonto the PS layer, followed by heat treatment in an oven at130 �C for 1 h. Once cooled, the stamp was peeled off. This wasfollowed by exposure to EA at room temperature for 5 h.The pattern was next washed multiple times with water toremove unreacted EA. A 50 μL amount of 4mg/mL nanoparticlesuspension was drop casted onto the patterned area of thecoverslip. The setup was then placed in a sealed vial, facilitating

controlled evaporation, which allowed for electrostatic nano-particle attachment onto the EA surface. The PS mask was thenremoved by washing with toluene. The resulting patternedPEDOT:PSS nanoparticle array was then used for furtherexperimentation.

Electrical Stimulation Protocol. For electrical stimulation experi-ments, two silver epoxy electrodes were painted onto the endsof the patterned nanoparticle arrays. Prior to cell culture, thewhole platform was UV and ethanol sterilized. Wells werecoated with poly(L-lysine) and 15 μg/mL of laminin followedby cell seeding at a density of 50 000 cells/well. Cells were left toadhere for 18 h. Immediately prior to stimulation, the prolifera-tion media was replaced with low-serum nerve growth factorcontaining differentiation media. For stimulation, the cells weresubjected to a monophasic pulsed current at a frequency of250 Hzwith a 2ms pulsewidth and an amplitude of 1mA for 2 h,after which they were left for an additional 72 h before analysis.

Cell Viability Assessment. Cell viability was measured using theMTS assay as per the manufacturer protocols (Invitrogen, UK).For measurements, 80 μL from each well was transferred into anew 96-well plate and read under a plate reader at 490 nmexcitation wavelength. To analyze cell morphology, cells wereimmunohistochemically stained for β-III tubulin.

Material Characterization. AFMwas performed on a Dimension3100 AFM systemwith a Nanoscope IV controller used to obtainthe AFM images in tapping mode, using Pt/Ir-coated contactmode probes with a spring constant of 0.2 N/m (type SCM-PIC,Bruker). TEM was performed on a JEOL 2100 transmissionelectron microscope at an accelerating voltage of 80 kV. SEMwas performed on a Zeiss 1555 VP-FESEM, and all samples werecoated with 5 nm of Pt. Biological samples were initially fixed in2.5% glutaraldehyde and dehydrated in increasing concen-trations of ethanol followed by critical point drying prior toPt coating. Immunohistochemically stained samples were

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analyzed using a Leica TCS SP2 AOBS multiphoton confocalmicroscope.

Conflict of Interest: The authors declare no competingfinancial interest.

Supporting Information Available: Detailed materials andmethods: synthesis, characterization (TEM, SEM, AFM), cellculture, and electrical stimulation experiments. This material isavailable free of charge via the Internet at http://pubs.acs.org.

Acknowledgment. D.H., I.L., and K.S.I. designed the experi-ments, developed the concept, and analyzed the data. D.H., J.Z.,and N.M.S. optimized the capillary force lithography experi-ments. D.H., X.C., V.A., and A.M. performed image acquisitionusing confocal microscopy, transmission electron microscopy,scanning electron microscopy, and atomic force microscopy.D.H., A.R.H., G.W.P., S.I.H., and A.B. optimized and designed theelectrical stimulation experiments. This work was funded bythe Australian Research Council (ARC), the National Health &Medical Research Council (NHMRC) of Australia, and the NationalScience Foundation (CBET-0756457). The authors acknowledgetheAustralianMicroscopy&Microanalysis Research Facility at theCentre for Microscopy, Characterization & Analysis, and TheUniversity of Western Australia, funded by the University, Stateand Commonwealth Governments. The authors also wish tothank Margaret Pollett and Chrisna LeVaillant for their invaluablecontribution in assisting with the PC12 cell cultures and immu-nohistochemistry, and Ella Marushchenko (www.scientific-illustrations.com) for assistance with Figure 1.

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ARTIC

LE

S1

Hierarchical Patterning of Multifunctional Conducting Polymer Nanoparticles as a Bionic

Platform for Topographic Contact Guidance

Dominic Ho,1,2 Jianli Zou,3 Xianjue Chen,1,‡, Alaa Munshi,1 Nicole M. Smith,1,4 Vipul Agarwal,1

Stuart I. Hodgetts,2 Giles W. Plant,5 Anthony J. Bakker,2 Alan R. Harvey,2 Igor Luzinov6 & K.

Swaminathan Iyer1*

1School of Chemistry and Biochemistry, The University of Western Australia, Crawley, WA

6009, Australia;

2School of Anatomy, Physiology and Human Biology, The University of Western Australia,

Crawley, WA 6009, Australia;

3Institute for Integrated Cell-Material Sciences (iCeMS), iCeMS Complex 2, Kyoto University,

Yoshida-Honmachi, Sakyo-ku, Kyoto, 606-8501, Japan;

4Experimental and Regenerative Neurosciences, School of Animal Biology, The University of

Western Australia, Crawley, WA 6009, Australia;

5Stanford Partnership for Spinal Cord Injury and Repair, Department of Neurosurgery, Stanford

University School of Medicine, Stanford, CA 94305, USA;

6School of Materials Science and Engineering, Clemson University, Clemson, South Carolina,

29634-0971, USA.

‡Present Address: Centre for NanoScale Science and Technology, School of Chemical and

Physical Sciences, Flinders University, Bedford Park, Adelaide, SA 5042, Australia.

* Correspondance: [email protected]

S2

Supplementary Information

Materials. All chemicals were purchased from Sigma-Aldrich unless otherwise stated: iron(III)

acetylacetonate (97 %), benzyl ether (98 %), oleic acid (90 %), oleyl amine (70 %), 1,2-

tetradecanediol (90 %), rhodamine B (Fluka), methyl ethyl ketone (99 %, Fisher), chloroform

(99 %, merck), toluene (99 %, Fisher), diethyl ether (90 %, Asia Pacific Speciality Chemicals),

polyethylenimine (50 % solution, Mn 1200, Mw 1300), Poly(3,4-ethylenedioxythiophene)

Polystyrene sulfonate (1.3% solution, Mw 10355), ethylenediamine (99.5 %, Fluka) and Pluronic

F-108. All tissue culture reagents were purchased from Gibco unless otherwise stated.

Dulbecco's Modified Eagle's medium (DMEM), PBS, foetal calf serum (Sigma), horse serum

(Sigma), penicillin/streptomycin, L-glutamine, non-essential amino acids (NEAA),

trypsin/EDTA (Sigma), laminin (#L2020, Sigma) and nerve growth factor (β-NGF, PeproTech).

Magnetite Synthesis. Fe3O4 was synthesized by the organic decomposition of Fe(acac)3 in

benzyl ether at 300 oC, in the presence of oleic acid, oleyl amine, and 1,2- tetradecanediol, as

previously described by Sun et al.1 The method to synthesise 6 nm Fe3O4 nanoparticles was

followed.

Synthesis of RhB-Modified PGMA: PGMA was synthesized by radical polymerization

according to a published procedure.2 Briefly, glycidyl methacrylate was polymerized in methyl

ethyl ketone (MEK) to give PGMA (Mn = 220515, Mw = 433730), using azobisisobutyronitrile

as initiator. The polymer was purified by multiple precipitations from MEK solution using

diethyl ether. To attach the dye to the polymer, a solution of rhodamine B (RhB, 20 mg) and

PGMA (100 mg) in MEK (20 mL) was heated to reflux under N2 for 18 h. The solution was

reduced in vacuo before the modified polymer was precipitated with diethyl ether (20 mL). The

S3

polymer was redissolved in MEK and precipitated with ether twice to remove ungrafted RhB.

PEDOT:PSS Multilayer Nanoparticle (NP) Synthesis. To prepare the organic phase of the

emulsion, the dried RhB-PGMA polymer was initially dissolved in 2 mL of CHCl3 and dried

under N2, leaving a sticky residue. This was redissolved in a 1:3 mixture of CHCl3 (2 mL) and

MEK (6 mL) along with 25 mg of Fe3O4. This organic phase was added drop wise to a rapidly

stirring aqueous solution of Pluronic F-108 (12.5 mg/mL, 30 mL). The emulsion was

homogenised with a probe-type ultrasonic wand at the lowest setting for 1 min. The organic

solvents were evaporated off overnight under a slow flow of N2. The suspension was purified via

centrifugation at 3000 g for 45 mins. The supernatant was transferred to a 50 mL flask

containing PEI (50 wt % solution, 100 mg) and heated to 80 oC for 16 h. The magnetic polymer

nanoparticles were collected on a magnetic separation column (LS, Miltenyi Biotec) in 3 mL

batches, washed with water (5 mL) and then flushed with water until the filtrate ran clear. This

purified product produced 10 mL of nanoparticle suspension at a concentration of 1 mg/mL.

Next, PEDOT:PSS was electrostatically attached to the nanoparticles. 60 µL of PEDOT:PSS (1.3

wt % dispersion in H2O) was diluted in 2 mL of water and added drop wise under rapid stirring

to NPs at a concentration of 0.5 mg/mL. The PEDOT:PSS was further dispersed under sonication

for 10 mins to ensure complete dispersion and then left to stir for 18 h. After 18 h, the mixture

was again sonicated for 2 mins. Excess PEDOT:PSS was then removed via centrifugation 16800

g for 20 mins). NPs were then washed twice in water before being stored at 4 oC at a

concentration of 4 mg/mL for further use.

Nanoparticle Conductivity Measurement. The nanoparticle conductivity was determined using

4-point probe measurements. 80 µL of nanoparticle solution with a concentration of 1 mg/mL

S4

was selectively dried on a square area 0.5 cm x 0.5 cm. Electrodes were placed at the 4 corners

of the square and subject to current-voltage sweeps.

Fabrication of PDMS Stamp. The metal layer of a blank compact disc (CD) was peeled off and

the CD washed with ethanol. The remaining polycarbonate structure was used as a master for the

PDMS stamp. The polymer base and curing agent from a Sylgard® 184 (Dow Corning) silicone

elastomer kit were mixed at a 10:1 ratio by weight in a glass vial. The glass vial was placed in a

vacuum desiccator to remove trapped bubbles from the mixture. Following vacuum treatment,

the elastomer was restored to atmospheric pressure slowly several times until it was free of

bubbles. The PDMS mixture was then cast onto the surface of the grooved side of CD and cured

at 80 ºC for 2 hours.

CFL Procedure. Prior to the CFL procedure, the indium tin oxide (ITO) coverslips were first

clean in acetone and isopopanol under sonication. 0.2 % w/v PGMA in CHCl3 was spin coated

onto the conducting surface of the ITO coverslips. Coverslips were then placed in an oven at 120

ºC for 20 mins to anneal the PGMA. Unreacted PGMA on the coverslip surface was removed by

washing in CHCl3. Next, 1.3 % w/v PS in toluene was spin coated onto the PGMA surface. A

PDMS stamp was then placed onto the PS layer, followed by heat treatment in an oven at 130 ºC

for 1 hr. The assembly was then cooled down at room temperature for another hour before the

PDMS stamp was peeled off. Next, the substrate was exposed to ethylenediamine (EA) and left

at room temperature for 5 h. The substrate was then wasted with water to remove unreacted EA.

Next, 50 µL of 4 mg/mL nanoparticle solution was drop casted onto the patterned area of the

coverslip. The setup was then placed in a sealed vial, facilitating controlled evaporation which

allowed for electrostatic nanoparticle attachment onto the EA surface. The PS mask was then

S5

removed by toluene, leaving the patterned nanoparticle array.

Cell Culture. The rat pheochromocytoma cells (PC12 cells) used here were obtained from

Flinders University (Adelaide, Australia) courtesy of Professor Jacqueline Phillips (Macquarie

University, Sydney, Australia). PC12 cells were cultured in P75 flasks in a humidified

atmosphere containing proliferation media: 5 % CO2 at 37 oC and maintained in DMEM medium

containing horse serum (10 % v/v), fetal calf serum (5 % v/v), penicillin/streptomycin (0.5 %

v/v), L-glutamine (1 % v/v) and nonessential amino acids (1 % v/v). For PC12 differentiation,

cells were cultured in differentiation media consisting of DMEM, L-glutamine (1 % v/v), horse

serum (1 % v/v) and nerve growth factor (50 ng/mL).

Electrical Stimulation Experiments. Prior to stimulation experiments, two silver epoxy

electrodes were painted onto the ends of the prepared NP array and platinum wires attached to

allow for connections with the stimulator. Next, a cell culture well was created by first cutting a

1.5 mL microcentrifuge tube in half and then sealing the capped end with silicon vacuum grease.

This was stuck onto the ITO glass with the patterned arrays in the centre of the well. This was

done to ensure that the electrodes did not come into direct contact with the cell culture media.

The array was then placed in a Petri dish to maintain sterility throughout the course of the

experiment (Fig S5a). Prior to culturing cells on the arrays, the coverslips were UV sterilised (20

mins) and then washed with 70 % ethanol three times. Wells were then coated with poly-(L-

lysine) and 15 µg/mL of laminin. Cells were then seeded at a density of 50 000 cells/well and left

to adhere for 18 h. Prior to stimulation, the proliferation media was replaced with differentiation

media. The cells were then stimulated according to protocols as listed below. Following

stimulation, the cells were left for a further 72 h with fresh differentiating media added every 48

S6

h. Photographs of the electrical stimulation setup are described in Figure S5.

Stimulation Protocol. The electrical signals were supplied by Grass S44 Stimulator (Quincy,

Massachusetts, USA). The stimulation regime is similar to that used by Wallace et al.3-5 Briefly,

the cells were subjected to a monophasic pulsed current at a frequency of 250 Hz with a 2 ms

pulse width and an amplitude of 1 mA for 2 h.

Cell Viability Assays. Cell viability was measured using the (3-(4,5-dimethylthiazol-2-yl)-5-(3-

carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium, inner salt) (MTS) assay as per the

manufacturer protocols (Invitrogen, UK). Cells were plated as per “electrical stimulation

protocol” stated above. Viability was to be determined at 3 time points: (i) 0 h (immediately

prior to electrical stimulation), (ii) 72 h after the addition of differentiation media and (iii) 72 h

after the addition of differentiation media and electrical stimulation. For measurements, 80 µL

from each well was transferred into a new 96 well plate and read under a plate reader at 490 nm

excitation wavelength. The same protocol was followed for every sample and each measurement

was carried out in triplicate.

Immunohistochemical Staining. The PC12 cells were immunohistochemically stained for ß-III

tubulin. The cells were fixed in 4 % paraformaldehyde for 10 mins. Cells were first incubated

with a primary antibody solution containing PBS, 10 % Normal Goat Serum, 0.1 % Triton X-100

and the anti-β-III tubulin antibody (1:1000, anti-rabbit, Covance) at room temperature for 30

mins. After 3 PBS washes, the antibody binding was visualised with anti-rabbit FITC (1:100,

Sigma) following incubation for 30 mins at room temperature. Coverslips were mounted on glass

slides covered with Dako Fluorescent Mounting Medium (Dako, USA). All experiments were

S7

performed in triplicate.

Confocal and Fluorescence Microscopy Analysis. Immunohistochemically stained samples

were analysed using confocal and fluorescence microscopy. Confocal microscopy was carried

out using a Leica TCS SP2 AOBS Multiphoton Confocal microscope and fluorescence

microscopy with a Diaplan fluorescence microscope.

Image and Statistical Analysis. To determine the effects both stimulation and the NP arrays had

on the PC12 cells, the average area of each cell was determined. 3 randomly selected areas on

each sample was visualised at 40 x magnification. The average area covered by each cell was

assessed using Image J analysis software (version 1.48a, NIH). All immunohistochemical

analyses were conducted by a single investigator, ensuring constant selection criteria, and results

expressed as means ± SD. Data were analysed using Origin data management software to

conduct ANOVA on groups of data. Statistically significant differences between each treatment

were determined using Bonferroni/Dunn post hoc tests (p≤0.05).

Scanning Electron Microscopy (SEM). Prior to SEM imaging, samples without cells were

coated with 5 nm of Pt. Samples with cells were fixed in 2.5 % glutaraldehyde for 2 h at 4 oC and

dehydrated. Samples were washed with deionized water and dehydrated in a microwave in serial

concentrations of ethanol (50 %, 70 % and 90 % once then 3x in absolute ethanol), before critical

point drying with carbon dioxide for 1h and then coating with 5 nm of Pt. Samples were imaged

using a Zeiss 1555 VP-FESEM.

Transmission Electron Microscopy (TEM). Synthesized polymer nanoparticles were drop-

casted on carbon coated TEM grids and imaged with an accelerating voltage of 100 kV on a

S8

JEOL 2100 transmission electron microscope.

Atomic Force Microscopy. A Dimension 3100 AFM system (Bruker) with a Nanoscope IV

controller (Bruker) was used to obtain the AFM images in Contact Mode, using Pt/Ir coated

contact mode probes with a spring constant of 0.2 N/m (type SCM-PIC, Bruker). The scan

parameters were adjusted to ensure reliable imaging with the smallest possible contact force

setpoint. Data analysis was performed using the SPM analysis freeware Gwyddion

(http://gwyddion.net).

Dynamic light scattering (DLS) and zeta potential measurements. DLS experiments were

performed using a Malvern Zetasizer Nano series. For measuring the size distribution, 5

measurements were taken and in each measurement there were 10 data acquisitions. Zeta

potential (ζ) measurements were performed using the same instrument. Measurements for each

sample were recorded in triplicate and 100 data acquisitions were recorded in each measurement.

All measurements were recorded at 25 oC in Malvern disposable clear Folded Capillary Cells.

S9

Supplementary Figures

Figure S1. (a) The PDMS stamp used in the study. SEM micrograph of the grooved structure of

the PDMS. Image taken at 6k x magnification; (b) Photograph of the polycarbonate disc peeled

from a CD. PDMS was cast on the grooved surface and stamps of the desired size were cut out.

(a) (b)

S10

Figure S2. Schematic of CFL procedure. Briefly, ITO substrate was modified with a thin layer

of PGMA followed by second layer of PS; a PDMS stamp was placed over the PS film and heat

treated at 130 oC; PDMS stamp was peeled off after cooling; EA was selective reacted to the

exposed PGMA stripes to produce cationic stripes to enable charge complementarity to assemble

the anionic PEDOT:PSS nanoparticles. The PS mask was removed by washing with toluene, to

obtain linear PEDOT:PSS conduits.

T > Tg (PS) (130 oC)

EA grafting onto PGMA Conducting NPs

PS removal

Peel off PDMS Stamp

Electrostatic attachment

S11

Figure S3. SEM micrograph of PC12 cells 18 h after plating and immediately prior to electrical

stimulation. PC12 cells on the patterned surface (yellow box) were evenly spread out, in

comparison to the rounded cells on non-patterned areas of the substrate demonstrating

preferential adhesion. Image taken at 1000 x magnification.

S12

Figure S4. Representative low magnification (magnification 400 x) SEM micrographs of PC12

cells 72 h after the following treatments: (a) (+) pattern, (-) stimulation and (b) (+) pattern, (+)

stimulation demonstrating lower coverage due to reduction in proliferation upon stimulation.

.

(b) (a)

S13

Figure S5. Photographs of the electrical stimulation setup: (a) A sterile Petri dish containing the

modified cell culture well (red arrow) and the Platinum wires which allow for connections to the

stimulator (green arrows), (b) The stimulator (yellow arrow) was placed next to an incubator and

the wires from the machine leading into the stimulator (blue arrows), (c) The wires from the

stimulator were connected to the platinum wires via alligator clips.

(a)

(c)

(b)

S14

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