porous silicon-polymer composites for cell culture and tissue engineering applications

© 2014 Woodhead Publishing Limited

420

18 Porous silicon–polymer composites for cell culture and tissue engineering applications

S. J. P. McINNES and N. H. VOELCKER ,

University of South Australia, Australia

DOI : 10.1533/9780857097156.3.420

Abstract : Porous silicon (PSi) is a promising biomaterial for a wide range of biomedical applications. There are several reasons for this accolade. Tunable porous structures can be fabricated with relative ease in large quantities. Degradation of PSi in vivo causes little to no adverse effect on surrounding tissue and the degradation product, orthosilicic acid, is rapidly cleared from the body. PSi has a large surface area that can be modifi ed with a wide variety of readily available chemistries, including polymers, and permits therapeutic molecules to be loaded. Hence, PSi is currently being intensely investigated for drug delivery, cell culture and tissue engineering applications. This chapter briefl y introduces PSi, its functionalization and use in polymeric composites. The challenges and requirements of PSi–based biomaterials are also discussed, along with reviews of specifi c applications of PSi–polymer composites in cell culture and tissue engineering. We conclude with our future vision for these biomaterials.

Key words: porous silicon, polymer, biomaterials, tissue engineering, cell culture.

18.1 Introduction

18.1.1 Fundamentals of cell culture, tissue engineering and biomaterials

Cell culture is a fundamental aspect of tissue engineering, as it establishes

the basic protocols involved with growing and maintaining cells ex vivo .

Tissue engineering utilizes aspects of cell culture to generate functional tis-

sues ex vivo or in situ with the intent to use them as functional replacements

for damaged or diseased tissue (Lanza et al. , 2007).

The need for tissue engineering technology is becoming more apparent as

the world’s populations continue to age. The replacement of failing organs

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Porous silicon–polymer composites 421

with artifi cial or donated organs is currently stymied by organ donation

shortages and immuno-rejection (Ikada, 2006a). In this context, repair of

tissues or organs to a healthy state using the potential of cells to regenerate,

ideally without the need for immuno-suppression, becomes an ideal solu-

tion to this problem (Kretlow et al. , 2007).

The creation of autologous engineered tissue grafts involves harvesting

cells from a patient and then seeding them onto a scaffold for expansion

in a laboratory environment. Once the cells are expanded and conditioned

into a suitable graft (a process which may take weeks) the completed graft

is re-implanted into the patient. While this method provides compatible

tissue for the patient, there are drawbacks. These include risk associated

with initial surgical harvest of the cells, elaborate culture procedures and

cell manipulation (Fioretta et al. , 2012). There is also a high risk of infection

from repeated surgery and tissue handling. If the graft is urgently required,

for instance in the case of skin grafts for burn victims, then the time needed

for autologous cell expansion may be prohibitive (Iwai et al. , 2005; Torikai

et al. , 2008). To date, grafts produced in this manner have rarely managed to

recover the tissue properties of the original tissue, in terms of mechanical

strength, structure and function (Ikada, 2006b).

A more attractive approach to tissue engineering in a clinical setting is in situ regeneration for tissue. This approach minimizes the risks involved in

tissue engineering, as the ‘bioreactor’ for the tissue engineering is the host’s

body (Fioretta et al ., 2012). This takes advantage of the body’s own regen-

erative capabilities, which provide the ideal environment for reintegration

(Mol et al. , 2009; Torikai et al ., 2008). In situ tissue engineering, however,

is problematic: the process of repair may take weeks or months. Scaffold

material must be able to replace the tissue in the interim, allow appropriate

tissue development, and then subsequently degrade at an appropriate rate

(Fioretta et al ., 2012).

Tissue construct materials need to be developed so they can support cells

and promote differentiation and proliferation in such a way that a functional

new tissue is formed (Mano et al. , 2007). Scaffold materials, cells and bio-

active molecules are combined to stimulate various cellular effects (Mano

et al ., 2007). The ideal scaffold biomaterial should be non-toxic, biocompat-

ible, promote favourable cellular interactions and tissue development, and

provide suitable mechanical and physical cues. It should also be bioresorb-

able and support tissue reconstruction without infl ammation. The degra-

dation rate of the material is dependent on the particular application. For

example, scaffold materials for bone-tissue engineering must dissolve rela-

tively slowly (months to years) (Middleton and Tipton, 2000), so that they

maintain mechanical strength for as long as possible. In contrast, skin tissue

scaffolds need to stay in place for no more than 1 month. The fi ne-tuning of

the materials’ degradation rate is important as remaining biomaterials can

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422 Porous silicon for biomedical applications

hinder the development of the newly forming tissue rather than promote it

(Ikada, 2006a). Vascularization is also a signifi cant issue in in vitro/ex vivo

tissue engineering, as the lack of vascularization will eventually lead to a

lack of oxygen and necrosis of the cells in the inner regions of the tissue

scaffold (Fig. 18.1). However, vascularization can be enhanced by the pres-

ence of the surrounding tissue and the newly forming tissue is less likely to

suffer necrosis (Ikada, 2006a).

A major challenge currently facing tissue engineering is the development

of materials which induce the desired cellular/tissue behaviour (Davis and

Leach, 2008). First and foremost, any successful tissue engineering scaffold

Ex vivo tissue engineeringPorous scaffold material

Cell culture and ECM formation

Lack of vascularization andlimited medium perfusion

NecrosisDue to lack of oxygen and nutrients

Implantation and ECM formationSurrounding tissue Blood vessels

Vascularization and cell proliferarion

In situ tissue engineeringCell seeding

18.1 The difference in vascularization for ex vivo and in situ tissue

engineering in terms of vascularization. ( Source : Adapted from Ikada

et al . (2006a).)

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must provide a surface upon which the cells can attach and proliferate.

Conventional cell culture is typically performed on sterile irradiated polysty-

rene culture plates and fl asks. Those surfaces do not capture the three-dimen-

sional (3D) character of cell environments in vivo using a two-dimensional

(2D) plate system. Typically, in situ , cells are surrounded by other cells, a

complex extra cellular matrix (ECM) and a wide range of different phys-

ical and chemical signals (Zhang, 2004). Due to the lack of a natural 3D

environment in typical cell culture, some particular cell phenotypes are not

expressed (Timmins et al. , 2004). To rectify this, 3D scaffolds are being devel-

oped from biologically-derived matrices such as collagen, chitosan and fi brin

(Guan et al. , 2011; Liu et al. , 2012; Moraes et al. , 2012), or synthetic hydrogels

such as polyacrylamide (Bayliss et al. , 1997b; Sailor, 2012) or poly(ethylene

glycol) (PEG) (Park et al. , 2005; Temenoff et al. , 2003). Recent reviews on this

topic are available by Drury and Mooney (2003) and Kretlow et al . (2007).

Within this chapter we will describe how PSi can be combined with various

polymers in a range of formats to generate PSi–polymer composite materials

that address some of these issues.

18.2 Fundamentals of porous silicon (PSi) and PSi/polymer composite fabrication and functionalization

PSi is a relatively new player in the biomaterials fi eld, but it is being

extensively investigated. This comes down to four particularly attractive

properties:

High surface area (up to 800 m • 2 g − 1 ) (Loni, 1997).

Photonic properties (exploitable for biosensing (Alvarez • et al. , 2009;

Coffer et al. , 2003; Palestino et al. , 2008; Worsfold et al. , 2006), self-

reporting drug delivery (Anglin et al. , 2008; Koh et al. , 2008) and in situ

monitoring cell response (Schwartz et al. , 2006)).

Ability to be processed into a variety of shapes including membranes •

and microparticles (MP) (Lehto et al. , 2005; McInnes et al. , 2006; Meade

and Sailor, 2007; Salonen et al. , 2005a, b; Wu et al. , 2008) or nanoparticles

(NP) (Bimbo et al. , 2010; Park et al. , 2009; Russo et al. , 2011; Secret et al. , 2012) suitable for applications including implantation or injection.

Degradation into non-toxic orthosilicic acid in aqueous environments, •

such as exist in vivo (Canham et al. , 1996; Loni, 1997).

PSi undergoes degradation in biological fl uids (Anderson et al. , 2003;

Canham, 1995; Loni, 1997) and is broken down into orthosilicic acid

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(Si(OH) 4 ), which is non-toxic and is the main form of Si in the human body

(Canham and Aston, 2001). This makes PSi a suitable material for a wide

range of biomedical applications, such as drug delivery systems and tissue

engineering scaffolds. The porosity of PSi can be tuned over a wide range

(up to 90%), and this parameter has a signifi cant infl uence on the degrada-

tion kinetics (Canham, 1997).

To be an effective biomaterial, PSi must be able to integrate into living

tissue, that is, it must be bioactive (Angelescu et al. , 2003). Additionally, it

is desirable for the material to be able to infl uence the response of cells via

tunable surface topology, chemistry and electrical properties (Angelescu

et al ., 2003). An additional benefi t of PSi is that its optical properties

can allow sensing of changes in the biological environment, for example

related to changes in cell behaviour such as cell adhesion and proliferation

(Angelescu et al ., 2003).

18.2.1 PSi fabrication

The conversion of bulk Si into its high surface area, biocompatible porous

counterpart is commonly achieved by etching in hydrofl uoric acid (HF).

By simply altering wafer resistivity, HF concentrations and current densi-

ties, different porous structures can be generated. Pore sizes can range in

diameter from a few nanometres to a few microns (Loni, 1997). This wide

range of pore sizes allows PSi to be generated with high surface areas from

400 − 800 m 2 /g (Loni, 1997). In general, highly doped p- and n-type Si are

needed for the preparation of mesopores of 20 − 50 nm (Ouyang et al. , 2005),

while macropores greater than 500 nm are usually formed on low-doped

n-type Si with backside illumination. Pores up to 2000 nm in diameter can

be obtained from p-type Si-wafers (Janshoff et al. , 1998). The chemistry

and mechanisms underlying the electrochemical etching of PSi has been

reviewed elsewhere (Canham, 1997). The tunability of pore dimensions

combined with the ability of PSi to degrade in aqueous solutions (Canham,

1995; Canham and Aston, 2001; Loni, 1997) has seen PSi become the focus

of several biomedical research groups.

18.2.2 PSi functionalization

PSi can be functionalized using a range of chemical reactions including oxi-

dation, silanization, hydrosilylation, electrografting, nitridization and thermal

carbonization/hydrocarbonization. This section will outline some of the most

common surface functionalization methods. For more in-depth reviews on PSi

functionalization, please see some of the other chapters of this book or the

following reviews (Kilian et al. , 2009; Schmeltzer and Buriak, 2000, 2004).

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Ozone and thermal modifi cations

The surface chemistry of PSi immediately after HF etching is hydride-

terminated. This surface chemistry is unstable in air and aqueous media.

In air, an oxide layer forms rapidly. In turn, in water oxidative hydrolysis

leads to surface degradation. This Si–H functionality has been modifi ed

in many different ways, ranging from simple oxidation (Gao et al. , 2002b)

through to patterned modifi cation with light-induced reactions (Lee et al. , 1996). When PSi is deliberately oxidized under a fl ow of ozone, the sur-

face Si–H groups are primarily converted to Si–OH groups (Letant and

Sailor, 2000). These Si–OH groups are stable in air and cause the PSi

surface to be hydrophilic in nature (Gao et al. , 2002a, b). Thermal oxida-

tion produces a thicker oxide layer (Gao et al ., 2002b), possessing fewer

Si–OH bonds, and is more glass-like in nature with a range of Si–O–Si

bonds (Gao et al ., 2002b).

Nitridization (Anderson et al. , 1993; James et al. , 2010; Kimoto and Arai,

2000), thermal carbonization (M ä kil ä et al. , 2012; Salonen et al. , 2000,

2002) and thermal hydrocarbonization (Bimbo et al ., 2010; Jalkanen et al. , 2012a, b) of PSi have been described previously. Surface chemistry and topo-

logical stability can be manipulated by selecting the specifi c modifi cation

techniques. Furthermore, nitridized surfaces are stable in HF acid solution,

which enables further micromachining fabrication steps to be performed

and presents the possibility of PSi inclusion in microelectrical mechanical

systems (James et al ., 2010). The benefi ts of thermal carbonization of PSi

includes increased thermal and electrical conductivity, as well as enhanced

mechanical strength and chemical stability of the PSi substrate (Salonen

et al ., 2002). Nitridization of PSi has a slight advantage over both oxidation

and carbonization techniques in that it leads to a decrease in the refractive

index of the PSi layer and can lead to bandwidth reduction in optical fi lter

devices (James et al ., 2010).

Other types of oxidations may be performed to stabilize PSi, including

chemical (Frotscher et al. , 1996; Mattei et al. , 2000; Salonen et al. , 1997; Song

and Sailor, 1998), anodic (Halimaoui et al. , 1991; Vial et al. , 1992) and photo-

induced (Salonen et al. , 1999; Tischler et al. , 1992) oxidations.

The treatment of freshly etched PSi with the techniques listed above

stabilizes the surface so that it is compatible with cell culture and tissue

engineering applications. This is because cells are unable to adhere to a PSi

surface that is rapidly degrading during cell attachment (i.e. over a time-

frame of hours). However, fi nding the balance between stability and the

controlled resorption of the PSi structure is one of the challenges currently

facing the development of these biomaterials.

A further advantageous characteristic of PSi is the wide range of possible

surface modifi cations with readily available chemicals under mild reaction

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conditions (Janshoff et al ., 1998; Khung et al. , 2006; Stewart and Buriak,

2000; Xu et al. , 2003).

Chemical functionalization

The presence of Si–OH groups on oxidized PSi surfaces allows for the func-

tionalization with alkoxy- or chloro-silanes. Ambient conditions or moder-

ate temperature elevation are suffi cient for common silanization reactions.

Silanization can be used to introduce a wide range of functional groups

including amines (Low et al. , 2006), isocyanates (Lowe et al. , 2010), meth-

acrylate (Cole et al. , 2006), PEG (Khung et al ., 2006; Sweetman et al. , 2011b)

and customized functionalities (Vasani et al. , 2011).

Silanes have been widely used to functionalize Si and PSi surfaces. The

most popular silanes are alkoxy silanes, which are produced commercially

with one, two or three alkoxy groups that can be linked to oxidized PSi

surfaces presenting –OH groups (Tinsley-Bown et al. , 2000). The use of

trialkoxy silanes can lead to pore blockage due to polymerization of the

silane (Strother et al. , 2000; Ulman, 1996). This can be avoided by the use of

monoalkoxy silanes. However, these silanes produce lower stability coatings

than the corresponding trialkoxy silanes (Strother et al ., 2000; Tinsley-Bown

et al ., 2000; Waddell et al. , 1981). The plethora of silanes available makes it

possible to convert the oxidized PSi surface into just about any functionality

imaginable. Not only can silanes be purchased or readily synthesized, they

can also be easily modifi ed further via the use of a myriad of commercially

available cross-linking agents.

Hydrosilylation of hydride-terminated PSi can be achieved via thermal

(Boukherroub et al. , 2000, 2001), white-light (Buriak, 1999; Stewart and

Buriak, 1998, 2001), Lewis acid (Buriak and Allen, 1998; Buriak et al. , 1999;

Holland et al. , 1999) and microwave-assisted (Boukherroub et al. , 2003a)

methods. Thermal hydrosilylation has been used to covalently link PSi with

alkenes and alkynes in order to generate functional interfaces (Anglin et al. , 2004; Bateman et al. , 1998; Lie et al. , 2004; Park et al. , 2006; Schwartz et al. , 2005). Hydrosilylation using alkenes and alkynes greatly increases the sta-

bility of PSi in aqueous medium (Anglin et al ., 2004). Both cathodic and

anodic electrografting methods have also been performed on freshly etched

PSi, and are reviewed elsewhere (Buriak, 2002; Stewart and Buriak, 2000).

Many other methods of Si–C bond creation have been performed, the most

popular of which is electrografting (Hurley et al. , 2003; Robins et al. , 1999).

These techniques are expertly reviewed by Buriak et al. (Schmeltzer and

Buriak, 2000, 2004).

PSi functionalization via silanization or hydrosilylation has been utilized

for a variety of applications, including spatially controlled DNA growth

(Hurley et al ., 2003), vapour sensor chips (Bakker et al. , 2003; Gao et al .,

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2002b), and the generation of interfaces between biological and semicon-

ductor surfaces (Wojtyk et al. , 2002) among others (Boukherroub et al ., 2000, 2003a, b; Buriak, 1999; Lees et al. , 2003; Robins et al ., 1999; Singh and

Lakshmikumar, 2002).

The functionalization of PSi with various surface chemistries can assist the

adsorption of proteins onto the surface, a phenomenon that precedes cell

adhesion. However, the functionalization of PSi with PEG groups reduces

protein and cell adhesion. This can be exploited to pattern cells (Sweetman

et al. , 2011a, 2012).

18.3 PSi/polymer composites

The combination of a pliable and soft polymeric material with a hard inor-

ganic porous material of high drug-loading capacity such as PSi improves

control over degradation and drug release profi les and is benefi cial for the

preparation of advanced drug delivery devices and biodegradable implants

or scaffolds. Polymers can be easily combined with PSi. This can either be

done directly, in a bulk fashion without any covalent bonding of the two

materials, or via utilizing the surface functionalization techniques discussed

above to create anchoring points. Since the polymer component is often

formed at similar or even higher mass fraction to the PSi, we refer to the

resulting materials as composites.

18.3.1 Surface-initiated polymerization techniques

PSi is easily functionalized with various polymer chemistries using polymer

grafting techniques. The covalent grafting of polymers to surfaces can be

performed in three (Moad et al. , 2006; Wang et al. , 2006) general ways (see

Plate XIII, in the colour section between pages 240 and 241):

‘Grafting from’: Involves surface modifi cation with a radical-forming •

species followed by surface-initiated polymerization.

‘Grafting to’: Involves the covalent or non-covalent attachment or pre-•

formed macromolecules, often via reactive end groups.

‘Grafting through’: Involves the copolymerization of surfaces already •

modifi ed with a polymer or monomer functionality.

The ‘grafting to’ approach usually leads to a lower surface coverage when

compared to ‘grafting from’ due to the steric hindrance of the macromol-

ecules already attached to the surface (Wang et al ., 2006). Meanwhile, the

‘grafting from’ method can cause the broadening of the molecular weight

distributions (Wang et al ., 2006). Subsequently, the choice between ‘grafting

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from’ or ‘grafting to’ must be based on the required application, considering

whether uniform polymer molecular weights or greater surface coverage

is necessary. For more information on various surface-initiated polymeri-

zation techniques please see the expert reviews provided by Edmondson

(Edmondson et al. , 2004) and Harabagiu (Harabagiu et al. , 2011).

18.3.2 Examples of surface-initiated PSi–polymer composites

One of the earliest examples of surface-initiated polymerization from

PSi was demonstrated by Yoon et al. (2003), who impregnated a fresh PSi

membrane with a ring-opening metathesis polymerization (ROMP) pre-

cursor and subsequently cross-linked the PSi scaffold via the ROMP of

poly(norbornene). This process generated a composite material that was

chemically and mechanically very stable. Errien et al. demonstrated the

‘grafting from’ approach via the electropolymerization of poly(3-dodecyl

thiophene) onto PSi. They have also used the thermal polymerization of

poly(diacetylene-bis-toluene-sulfonide) in acetone at 65 ° C for 48 h (Errien

et al. , 2005). Both methods of polymerization gave a homogeneous polymer

layer. Pike et al. covalently attached polypyrrole (PP) to hydrogen-termi-

nated PSi surfaces by conveying monomer functionality to the surface via

hydrosilylation with alkenyl-pyrrole before performing the light-induced

electropolymerization of pyrrole (Pike et al. , 2003). Xia et al. have used the

grafting approach to covalently graft polydimethylsiloxane (PDMS) poly-

mer monolayers to PSi layers to improve their stability towards oxidation.

To this end, they used thermal hydrosilylation with vinyl-PDMS at 80 ° C for

2 h to introduce a PDMS monolayer. Stringent washing assured that phys-

isorbed polymer chains were removed from the surface (Xia et al. , 2005).

In 2008, Gorman et al . studied the application of atomic transfer radical

polymerization (ATRP) for the surface grafting of poly(methyl methacry-

late) (PMMA) (Gorman et al. , 2008). The resulting composites have been

studied to determine the effect the porosity exerted on the molecular weight

and polydispersity during surface-initiated polymerization. It was found the

molecular weight of the grafted polymer decreased from polymerization

in solution, to fl at substrates and fi nally porous substrates (50 nm pores).

This was attributed to the increase of confi nement of the growing polymer

chains.

In 2010, Chiboub et al . covalently grafted polyaniline (PANi) onto ani-

line-terminated PSi substrates (Chiboub et al ., 2010a, b).. The oxidized PSi

surface was reacted with 3-bromopropyltrichlorosilane to yield a bromi-

nated surface. The aniline-terminated PSi surface was formed by reacting

the brominated PSi layer with aniline molecules at 60 ° C for 24 h. They then

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performed the graft polymerization of aniline to the surface using either oxi-

dative polymerization or electrochemical polymerization using cyclic volta-

mmetry. These composites have potential applications in sensing devices.

In an effort to harness the properties of poly(L-lactide) (PLLA) and

PSi together, PLLA-coated composite material was created from PSi

fi lms and PSi MPs (McInnes et al. , 2009, 2012a). Both PSi and PLLA

show good biocompatibility and tunable degradation behaviour, sug-

gesting composites of these materials to be suitable to support localized

drug delivery into the human body. Three different PSi and PLLA com-

posite formats were prepared: grafting PLLA from PSi fi lms via surface-

initiated ring-opening polymerization (ROP) (PSi-PLLA; grafted); spin

coating PLLA solution onto oxidized PSi fi lms (PSi-PLLA; spin-coated);

and melt casting a PLLA monolith containing dispersed PSi MPs (PLLA-

PSi; monoliths). Composite materials were loaded with a model cytotoxic

drug, camptothecin (CPT). Release profi les of CPT showed distinct char-

acteristics for each of the composites studied. The limited linear phase

release exhibited by PSi-PLLA composites was adequate for short-term

release application, such as that required for the delivery of antibiotics,

but clearly unsuitable to support long-term controlled drug release. Spin-

coated PSi-PLLA composites were less susceptible to hydrolytic degra-

dation due to the thicker PLLA layer protecting the PSi scaffold, and

released CPT for up to 400 h. Comparison of PLLA monoliths with and

without CPT pre-loaded into the pores of PSi MPs revealed the release to

be slower if CPT was pre-loaded into the PSi pores. Inclusion of PSi MP

into PLLA monoliths was calculated to extended CPT release from 50

to 200 days. The cytotoxicity of the composites was confi rmed by contact

with human lens epithelial (SRA) cells. Monolithic materials were deter-

mined to induce maximum cell death over the 5 day monitoring period.

We believe that biodegradable hybrid materials such as these will fi nd use

in tissue engineering and drug delivery, for example in applications where

complex degradation profi les are required that cannot be achieved with

one type of material alone.

18.3.3 Non-surface-initiated polymerization techniques

Besides surface-initiated polymerizations, other techniques can be used to

generate composite materials with pre-polymerized polymers, without nec-

essarily requiring covalent attachment to the surface (Plate XIII). These

techniques include layer-by-layer deposition, spin coating, electrospinning,

electrostatic assembly and encapsulation. These techniques will not be

reviewed in detail here; however, relevant examples of these can be found

in Table 18.1.

Cop

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St

even

McI

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(20

9-12

-940

)

Wed

nesd

ay, F

ebru

ary

19, 2

014

4:10

:01

PM

IP A

ddre

ss: 1

30.2

20.8

.10

Tab

le 1

8.1

P

Si–

po

lym

er

co

mp

osit

e m

ate

rials

(2

00

8–2

012

)

Su

rface(s

) P

oly

meri

zati

on

meth

od

P

oly

me

r P

ossib

le b

iom

ed

ica

l

ap

pli

ca

tio

ns

Re

fere

nce

s

PS

i fi

lms

‘Gra

ftin

g f

rom

’ u

sin

g

AT

RP

PM

MA

C

ell

sca

ffo

lds

Go

rman

et

al ., 2

00

8

PS

i fi

lms (

DB

R)

Casti

ng

P

MM

A

Dru

g r

ele

ase

mo

nit

ori

ng

K

oh

et

al ., 2

00

8

PS

i fi

lms

Ele

ctr

op

oly

meri

zati

on

P

P

Ele

ctr

ica

lly

acti

ve

ce

ll

sca

ffo

lds

Fu

ka

mi et

al. , 2

00

8

PS

i ru

gate

fi lm

s a

nd

mem

bra

nes

Casti

ng

P

oly

sty

ren

e (

PS

) C

ell

mo

nit

ori

ng

, ce

ll

sca

ffo

ld

Kim

et

al. , 2

00

8

PS

i la

tera

l

gra

die

nt

fi lm

Pla

sm

a

po

lym

eri

zati

on

Po

ly a

lly

lam

ine

(PA

) C

ell

sca

ffo

lds

Cle

men

ts e

t a

l. , 2

00

7

PS

i fi

lms

Ele

ctr

och

em

ical

oxid

ati

ve

po

lym

eri

zati

on

tech

niq

ue

PP

E

lectr

ica

lly

acti

ve

ce

ll

sca

ffo

lds

Ha

rra

z et

al. , 2

00

8

PS

i fi

lms

Sp

in c

oati

ng

P

oly

[2

-me

tho

xy-5

(2-

eth

ylh

ex

ylo

xy-p

-

ph

en

yle

ne

vin

yle

ne

)]

(ME

H-P

PV

)

Ph

oto

lum

ine

sce

nt

ce

ll

sca

ffo

lds

Mis

hra

et

al.

, 2

00

8

PS

i fi

lms

Sp

in c

oati

ng

P

oly

[2

-me

tho

xy-5

-(3

0,7

0-

dim

eth

ylo

cty

lox

y)-

1,4

-

ph

en

yle

ne

vin

yle

ne

]

Ph

oto

lum

ine

sce

nt

ce

ll

sca

ffo

lds

La

ho

z et

al.

, 2

00

8

PS

i m

em

bra

ne

C

asti

ng

Fo

rmv

ar ®

(o

r p

oly

vin

yl

form

al)

Bio

se

nsin

g

Jia

o a

nd

We

iss, 2

010

a, b

;

Jia

o e

t a

l ., 2

00

9; R

on

g e

t

al., 2

00

8

PS

i fi

lms

Infi

ltra

tio

n

PM

MA

or

PD

MS

C

ell

sca

ffo

lds

Ma

rsal et

al. , 2

00

8

PS

i fi

lms

Dip

co

ati

ng

, sp

in

co

ati

ng

, sp

read

ing

an

d c

asti

ng

PS

C

ell

sca

ffo

lds

Kim

et

al ., 2

010

; Tig

hilt

et

al., 2

00

8

Cop

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Mat

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l dow

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http

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dpub

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line.

com

St

even

McI

nnes

(20

9-12

-940

)

Wed

nesd

ay, F

ebru

ary

19, 2

014

4:10

:01

PM

IP A

ddre

ss: 1

30.2

20.8

.10

PS

i fi

lms

Sp

in c

oati

ng

an

d

e-b

eam

lit

ho

gra

ph

y

ZE

P 5

20

A p

ho

tore

sis

t C

he

mic

al

an

d b

iose

nsin

g

We

i et

al. , 2

00

8

PS

i fi

lms

‘Gra

ftin

g t

o’

via

carb

od

iim

ide

co

up

lin

g

Po

ly N

-iso

pro

py

l-

acry

lam

ide

(P

NIP

AM

)

Ce

ll s

he

et

rele

ase

S

eg

al e

t a

l. , 2

00

9

PS

i fi

lms a

nd

MP

s

Su

rface-i

nit

iate

d

rin

g-o

pen

ing

po

lym

eri

zati

on

(SI-

RO

P)

PL

LA

B

iod

eg

rad

ab

le s

up

po

rts,

dru

g d

eli

ve

ry

McIn

ne

s e

t a

l ., 2

00

9, 2

012

a

PS

i fi

lms

Dip

co

ati

ng

C

ross-l

inke

d c

hit

osa

n

hy

dro

ge

l

Dru

g d

eli

ve

ry

Wu

an

d S

ailo

r, 2

00

9

PS

i N

Ps

Casti

ng

P

oly

vin

ylb

uty

ral

Ele

ctr

olu

min

esce

nt

de

vic

es f

or

op

tica

l

da

ta i

ma

gin

g

Da

vid

en

ko

et

al. , 2

00

9

PS

i N

Ps

Sp

in c

oati

ng

M

EH

-PP

V

Ph

oto

lum

ine

sce

nt

ce

ll

sca

ffo

lds

Sv

rce

k e

t a

l. , 2

00

9

PS

i fi

lms

Ele

ctr

op

oly

meri

zati

on

P

P

Ele

ctr

ica

lly

acti

ve

ce

ll

sca

ffo

lds

Fu

ka

mi et

al. , 2

00

9

PS

i fi

lms

Sp

in c

oati

ng

P

MM

A

Ce

ll s

ca

ffo

lds

Na

va

rro

-Urr

ios e

t a

l. , 2

00

9

PS

i fi

lms a

nd

PS

i p

ho

ton

ic

cry

sta

ls

Sp

in c

oati

ng

A

min

o f

un

cti

on

ali

zed

po

ly ε

-ca

pro

lacto

ne

(PC

L)

Ce

ll s

ca

ffo

lds a

nd

bio

-

se

nsin

g

Ste

fan

o e

t a

l. , 2

00

9, 2

010

PS

i fi

lms a

nd

mem

bra

nes

(DB

Rs)

Sp

in c

oati

ng

P

MM

A

Op

toe

lectr

on

ic b

iose

nsin

g

Sy

che

v e

t a

l. , 2

00

9

PS

i N

Ps a

nd

PS

i

fi lm

s

Sp

in c

oati

ng

an

d

casti

ng

ME

H-P

PV

P

ho

tolu

min

esce

nt

ce

ll

sca

ffo

lds

Jla

ssi e

t a

l. , 2

00

9

PS

i fi

lms

Ele

ctr

op

oly

meri

zati

on

P

PV

P

ho

tolu

min

esce

nt

an

d

ele

ctr

olu

min

esce

nt

ce

ll

sca

ffo

lds

Dje

niz

ian

et

al.

, 2

010

;

Ge

llo

z et

al. , 2

010

(Co

nti

nu

ed

)

Cop

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Mat

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St

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McI

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(20

9-12

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)

Wed

nesd

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ebru

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19, 2

014

4:10

:01

PM

IP A

ddre

ss: 1

30.2

20.8

.10

PS

i fi

lms

Infi

ltra

tio

n

Cro

ss-l

inke

d p

oly

acry

lam

ide

(PA

Am

)

Ch

em

ica

l a

nd

bio

se

nsin

g

Bo

na

nn

o e

t a

l. , 2

010

PS

i fi

lms

‘Gra

ftin

g f

rom

’ (u

sin

g

oxid

ati

ve a

nd

ele

ctr

och

em

ical

po

lym

eri

zati

on

)

PA

Ni

Ph

oto

lum

ine

sce

nt

ce

ll

sca

ffo

lds

Ch

ibo

ub

et

al ., 2

010

a,

2010b

PS

i M

Ps

Infi

ltra

tio

n

Sp

urr

’s r

esin

D

rug

de

liv

ery

an

d c

ell

pe

ne

tra

tio

n a

ge

nts

Se

rda e

t a

l. , 2

010

PS

i M

Ps

Ele

ctr

osp

inn

ing

P

CL

B

iod

eg

rad

ab

le c

ell

sca

ffo

lds

Ka

sh

an

ian

et

al. , 2

010

PS

i m

em

bra

ne

s

Melt

casti

ng

P

oly

(e

thy

len

e o

xid

e)

Lo

w f

ou

lin

g s

urf

ace

s

Ku

sm

in e

t a

l. , 2

010

PS

i fi

lms

Casti

ng

P

oly

[2

,1,3

-be

nzo

-

se

len

ad

iazo

le-(

2,5

-

did

od

ecy

lox

y-1

,4-

ph

en

yle

ne

)eth

yn

yle

ne

]

Op

toe

lectr

on

ic s

en

so

r

su

rfa

ce

s

Xia

ng

et

al. , 2

010

PS

i fi

lms

AT

RP

P

NIP

AM

C

ell

sh

ee

t re

lea

se

V

asa

ni et

al .

, 2

011

PS

i fi

lter

(ru

ga

te)

Sp

in c

oati

ng

C

ross-l

inke

d c

hit

osa

n

Ch

em

ica

l a

nd

bio

se

nsin

g

Sh

an

g e

t a

l. , 2

011

PS

i fi

lter

(ru

ga

te)

Casti

ng

P

S

Ch

em

ica

l a

nd

bio

se

nsin

g

Ru

min

ski et

al.

, 2

011

PS

i fi

lms

Sp

in c

oati

ng

P

roL

IFT

™ (B

rew

er

Scie

nti

fi c R

oll

a,

MO

,

US

A)

Ce

ll p

att

eri

ng

or

gu

ida

nce

La

i et

al. , 2

011

PS

i M

Ps

Infu

sio

n

PS

O

pti

ca

l se

nsin

g

Ch

en

g e

t a

l. , 2

011

PS

i fi

lms

Po

ten

tio

sta

tic a

nd

galv

an

osta

tic

po

lym

eri

zati

on

PP,

PA

Ni

an

d p

oly

thio

ph

en

e

Ph

oto

vo

lta

ic d

ev

ice

an

d

se

nsin

g a

pp

lica

tio

ns

Ha

rra

z, 2

011

Tab

le 1

8.1

C

on

tin

ued

Su

rface(s

)P

oly

meri

zati

on

meth

od

Po

lym

er

Po

ssib

le b

iom

ed

ica

l

ap

pli

ca

tio

ns

Re

fere

nce

s

Cop

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hted

Mat

eria

l dow

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ded

from

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http

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line.

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St

even

McI

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(20

9-12

-940

)

Wed

nesd

ay, F

ebru

ary

19, 2

014

4:10

:01

PM

IP A

ddre

ss: 1

30.2

20.8

.10

PS

i fi

lms

Ele

ctr

och

em

ical

po

lym

eri

zati

on

Po

ly v

iny

l-ca

rba

zole

P

ho

tovo

lta

ic a

pp

lica

tio

ns

Na

ho

r et

al.

, 2

011

PS

i fi

lms

AT

RP

P

oly

acry

lic a

cid

(PA

A)

an

d p

oly

me

tha

cry

lic

acid

(P

MA

A)

Ce

ll s

ca

ffo

lds,

dru

g

de

liv

ery

de

vic

es

Wa

ng

et

al.

, 2

011

PS

i an

iso

tro

pic

DB

R

mem

bra

ne

Casti

ng

P

S

Ce

ll s

ca

ffo

lds

Ch

o e

t a

l. , 2

011

PS

i la

tera

l

gra

die

nt

fi lm

Pla

sm

a

po

lym

eri

zati

on

1,7

-octa

die

ne

an

d

pro

pio

nic

acid

Ce

ll s

ca

ffo

lds

Mic

he

lmo

re e

t a

l. , 2

012

PS

i fi

lms

Acti

vato

rs g

en

era

ted

by e

lectr

on

tran

sfe

r – A

TR

P

(AG

ET-

AT

RP

)

Po

ly h

yd

rox

ye

thy

l

me

tha

cry

late

Bio

se

nsin

g

Ho

lth

au

se

n e

t a

l. , 2

012

PS

i fi

lms

Infi

ltra

tio

n a

nd

sp

in

co

ati

ng

PA

Am

an

d P

NIP

AM

C

ell

sca

ffo

lds,

ce

ll s

he

et

rele

ase

Ma

ssa

d-I

va

nir

et

al. , 2

012

PS

i fi

lms

AT

RP

PA

A

Bio

se

nsin

g,

ce

ll s

ca

ffo

lds

Wa

ng

et

al.

, 2

012

a

PS

i fi

lms

Ele

ctr

och

em

ical

po

lym

eri

zati

on

PA

Ni

Ph

oto

lum

ine

sce

nt

ce

ll

sca

ffo

lds

Ole

ny

ch e

t a

l. , 2

012

PS

i fi

lms

Sp

in c

oati

ng

M

EH

-PP

V a

nd

la

dd

er

typ

e p

oly

pa

ra-

ph

en

yle

ne

Ce

ll s

ca

ffo

lds

Pra

ncu

lis e

t a

l. , 2

012

PS

i fi

lms

iCV

D

p(M

AA

-co

- E

DM

A)

Dru

g d

eli

ve

ry,

ce

ll

sca

ffo

lds

McIn

ne

s e

t a

l ., 2

012

b

PS

i m

icro

cavit

y

Infi

ltra

tio

n a

nd

sp

in

casti

ng

ME

H-P

PV

an

d p

oly

fl u

ore

ne

(P

FO

)

Ch

em

ica

l se

nso

rs

On

g a

nd

Le

vit

sky, 2

011

No

tes :

iCV

D, in

itia

ted

ch

em

ical

vap

ou

r d

ep

osit

ion

; p

(MA

A-c

o-E

DM

A),

po

ly (

me

tha

cry

lic a

cid

-co

-eth

yle

ne

dim

eth

acry

late

).

Cop

yrig

hted

Mat

eria

l dow

nloa

ded

from

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D

eliv

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by

http

://w

ww

.woo

dhea

dpub

lishi

ngon

line.

com

St

even

McI

nnes

(20

9-12

-940

)

Wed

nesd

ay, F

ebru

ary

19, 2

014

4:10

:01

PM

IP A

ddre

ss: 1

30.2

20.8

.10


One newly emerging polymerization technique that should be mentioned

in detail is initiated chemical vapour deposition (iCVD). Chemical vapour

deposition (CVD) is commonly used in the semiconductor industry to pro-

duce high-purity, high-performance thin fi lm coatings on solid substrates.

During CVD, the substrate is exposed to volatile precursors, which react

and/or decompose on the substrate surface to produce the desired coating.

In contrast to CVD, which is primarily used to deposit inorganic materials,

iCVD is able to produce polymeric thin fi lm coatings on any substrate in

a one-step, solvent-free process (Gupta et al. , 2008). Benefi ts of this pro-

cess include very low power usage, the production of highly conformal coat-

ings and, most importantly, the fact that fully functional coating with intact

pendent moieties remain after the deposition process (Gupta et al ., 2008).

There is now a wide library of different functional moieties that can be pro-

duced via this technique on Si (Baxamusa et al. , 2010; Chan and Gleason,

2005; Coclite et al. , 2009, 2010; Gupta et al ., 2008; Im et al. , 2009; Lee et al. , 2009; Mar í -Buy é et al. , 2009; O’Shaughnessy et al. , 2007; Ozaydin-Ince and

Gleason, 2010; Tenhaeff et al. , 2010; Trujillo et al. , 2010; Xu and Gleason,

2010) and other materials (Coclite et al ., 2010; Lau et al. , 2003; Ma et al. , 2005; Tekin et al. , 2011; Xu and Gleason, 2010; Yang et al. , 2011).

In iCVD of polymeric materials, an initiating species and monomer are

simultaneously introduced and fi laments heat the initiator (e.g., tert -butyl

peroxide) to generate free radicals without cracking the monomer and

retaining its functionality (Asatekin et al. , 2010; Baxamusa et al. , 2009;

Tenhaeff and Gleason, 2008). The polymerization mechanism followed in

the iCVD chamber is proposed to be similar to the classical free radical

polymerization of vinyl monomers. Extensive reviews of the iCVD process

and its applications can be found in the literature (Alf et al. , 2009; Asatekin

et al ., 2010; Baxamusa et al ., 2009; Tenhaeff and Gleason, 2008). iCVD dem-

onstrates an exceptional versatility in terms of substrates, polymers and

applications (Lau and Gleason, 2007a, b).

McInnes et al . used iCVD for the fi rst time to generate poly methacrylic

acid-co-ethylene dimethacrylate (p(MAA-co-EDMA)) coatings on drug-

loaded PSi (McInnes et al. , 2012b). The results indicated that the iCVD pro-

cess was able to coat drug-loaded PSi fi lms with a homogeneous 340 nm

polymeric layer. The deposition of this layer successfully occluded the pores

and encapsulated the drug. This is of much interest as the ability to coat sub-

strates in a solvent-free environment, without the loss of the loaded drug,

has benefi ts over conventional techniques. More importantly, iCVD allows

for the pre-loading of sensitive drug payloads (such as proteins and other

biomolecules), which helps ensure the drug loading is high, evenly distrib-

uted and embedded beneath the drug release controlling polymer layer and

not damaged by post-polymerization loading techniques.

Cop

yrig

hted

Mat

eria

l dow

nloa

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In summary, it is conceivable that the studies of PSi–polymer composite

materials discussed above will soon lead to the development of polymer-

coated porous materials that are capable of supporting the growth of cells

and their differentiation by the release of growth factors from the PSi layer,

possibly even in a patterned manner. Once grown to confl uence the cell

sheets/tissues could then be released from the scaffold and then implanted

or grafted into patients.

18.4 Polymers for tissue engineering

Polymeric scaffolds can be derived from natural or synthetic sources, or a

combination of the two. The natural polymers often mimic the ECM and

therefore promote cell attachment and growth (Mano et al ., 2007).

Benefi ts of natural polymers include their similarity to biological macro-

molecules and ECM components, which leads to lower chronic infl amma-

tion and immunological reactions. Naturally derived polymers include ECM

components (such as collagen (Wallace and Rosenblatt, 2003), fi bronectin,

glycosaminoglycans and fi brin (Drury and Mooney, 2003)) and protein-

based polymers (such as gelatin and casein) (Dhandayuthapani et al. , 2011).

There are also a wide variety of polysaccharides from sources such as plants

(i.e. starch and cellulose), animal (i.e. chitin, chitosan and hyaluronic acid)

(Gutowska et al. , 2001), algae (i.e. galactans and carrageenan), exudate

gums (gum arabic) and microbial (gellan gum, pullulan, xanthan gum and

cellulose) (Mano et al ., 2007) that may be incorporated into biomaterials

synthesis. Synthetic polymers may not always meet biocompatibility criteria,

but they are often cheaper, available in larger quantities and have longer

shelf lives (Dhandayuthapani et al ., 2011).

There are a myriad of synthetic polymers (and copolymers) used in tis-

sue engineering and other biomedical applications. Some of the typical syn-

thetic polymers used for tissue engineering include poly( α -hydroxyacid)

(Morita and Ikada, 2002), polyphosphazenes (Deng et al. , 2010), polycarbon-

ate (Feng et al. , 2012), polyanhydrides (Seppala et al. , 2011), polyurethanes

(Khan and Dahman, 2012), hydrogels (such as poly(N-isopropylacrylamide))

(Gutowska et al ., 2001; Schwall and Banerjee, 2009), polyethylene/propyl-

ene glycol block copolymers (Sousa-Herves et al. , 2012; Zhang and Easteal,

2008) and polyfumarates (Cortizo et al. , 2008; Fernandez et al. , 2010) to men-

tion a few. There are many comprehensive reviews on the use of both natural

(Huang and Fu, 2010; Mano et al ., 2007) and synthetic polymeric scaffolds in

tissue engineering and other biomedical applications (Gutowska et al ., 2001;

Middleton and Tipton, 2000; Sokolsky-Papkov et al. , 2007). There are also

reviews available on the polymer/inorganic composite materials for biomed-

ical applications available (Habraken et al. , 2007; Rezwan et al. , 2006) and

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it is important to note that many polymers can now be deposited with both

chemical (Harding et al. , 2012, 2013) and stiffness (Wang et al. , 2012d) gradi-

ents, both of which can infl uence cell adhesion and proliferation.

One of the main advantages of the use of polymers for biomedical appli-

cations has been the development of injectable matrices. These polymers

can be injected and subsequently cured in situ . This brings benefi ts in terms

of minimal invasiveness of the implantation (Kretlow et al ., 2007), and

brings other advantages including the ability of the matrix to conform to

the specifi c shape of the actual defect or trauma site. This means that no

patient-specifi c prefabrication is required. Injectable systems can be used

for both parenteral and localized applications, and the degradation/release

kinetics can be varied by changing the size and hydrophobicity of the matrix

(Kretlow et al ., 2007). Additionally, it is also possible to include both cells

and biomolecules (including growth factors) into the injection mixture and

overcome issues commonly associated with cell adhesion, localization and

biomolecule delivery. These injectable matrices are polymerized by chemi-

cal, photo-initiated, ionic, and cross-linking and thermo-gelling techniques

in situ and this can be done in a mild manner, which allows for the inclu-

sion of pre-loaded cells or biomolecules (Kretlow et al ., 2007). Some inject-

able systems have even been designed with particles serving as the bulk

of the material; this could lead to crossover in the use of porous inorganic

materials as injectable drug or cell reservoirs (da Silva et al. , 2007; Salem

et al. , 2003). The mechanical properties of the injectable polymers can also

be varied by either controlling the synthetic process to promote more cross-

linking or adding particles such as ceramics (Frazier et al. , 1997; Habraken

et al ., 2007; He et al. , 2000; Kretlow et al ., 2007) and carbon nanotubes (Shi

et al. , 2006). These approaches can be easily attempted with PSi. For exam-

ple, PSi particles could be functionalized with multiple cross-linking points

to promote a high cross-linking ratio, and the particle size, thickness, poros-

ity and concentration can be varied to reinforce the polymeric network. The

biodegradation is also affected by the same two properties, cross-linking

density and particulate size.

Due to the limitations of using polymers alone in tissue engineering, the

combination of typically brittle inorganic porous scaffolds and fl exible poly-

meric scaffolds has been attempted to yield more robust hybrid scaffolds

(Dhandayuthapani et al ., 2011). The following sections of this chapter will

endeavour to give an overview of the current PSi–polymer composites, their

benefi ts and application to biomedicine, with particular emphasis on cell

culture and tissue engineering. An overview of all the PSi–polymer com-

posite materials generated for a myriad of applications over the past 5 years

(2008–2012) can be found in Table 18.1. We have also given insight into the

possible future biomedical applications of these materials.

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18.5 The grafting of biopolymers to PSi

Biopolymers such as peptides and oligonucleotides are commonly attached

to PSi via ‘grafting to’ methods. For example, our laboratory has generated

PSi MPs that are capable of the direct solid-phase synthesis of oligonucle-

otides from the PSi surface using chemical linkers that are cleavable in phys-

iological conditions (McInnes and Voelcker, 2012). These supports are able

to withstand the cleavage and deprotection of the oligonucleotides post-

synthesis and subsequently dissolve at physiological conditions (pH = 7.4,

37 ° C), slowly releasing the oligonucleotides. This type of composite could

allow for the future design, synthesis and release of therapeutic DNA for

the manipulation of various cellular effects. While we have not yet tested

these materials for their ability to induce cellular effects, it is possible that

the released DNA could be used to transfect mammalian cells and promote

the expression of proteins (Zhang et al. , 2004), or the particles could be

designed to be taken up by the target cells where they could release their

payload (Serda et al. , 2011). The approach we have taken with the design of

these MPs could also be adapted for solid-phase synthesis of other biomol-

ecules, such as proteins and RNA.

Others in our laboratory (Clements et al. , 2011) have also combined

biopolymers to PSi surfaces via the use of electrochemical modifi ca-

tion with ethyl-6-bromohexanoate in a gradient fashion. The hydrolysis

of the ester moieties was used to create an activated gradient of func-

tional carboxylic acid groups, which were, subsequently used to couple

cyclic RGD peptide via the use of (N-(3-Dimethylaminopropyl)-N ′ -ethylcarbodiimide hydrochloride/N-Hydroxysuccinimide) EDC/NHS.

The functionalized surfaces were used to screen the extent of rat mes-

enchymal stem cell (MSC) attachment. The cell culture results showed

that the cells responded to the gradient surface with MSC attachment

on these surfaces increasing with increasing cyclic RGD density. This

high-throughput technique was then extended to include lateral poros-

ity gradients when Clements et al . (2012) demonstrated the fabrication

of an orthogonal gradient platform combining a PSi pore size gradient

with the gradient of cyclic RGD. In this case, it was found that the MSC

responded to both the topographical and chemical cues arising from the

orthogonal gradient. However, the MSCs responded more strongly to

changes in RGD density than to changes in pore size during short-term

culture. Gradients of RGD could fi nd applications in high-throughput

screening of optimal cell growth conditions, as Le Saux et al . (2011) have

found that the average RGD spacing infl uences lipid raft accumulation,

which enhances sensitivity to vascular epithelial growth factor (VEGF)

stimulation, and controls migration in endothelial cells.

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Other recent work by Guan et al . (2011) has shown the ability to selec-

tively functionalize the inside and outside of the PSi structure, with two

completely different chemistries. To do this they used click chemistry to com-

bine different moieties such as tetra(ethylene glycol) and GRGDS peptides

to the outer surface and alkane moieties to the internal pore surface. They

found that the PSi rugate fi lters modifi ed externally with the cell-adhesive

peptide GRGDS allowed for cell adhesion via formation of focal adhesion

points. The authors proposed to use such structures for applications in cell-

based biosensing.

18.6 PSi and tissue engineering

18.6.1 Cell culture on oxidized and small molecule modifi ed PSi

Most of the research in the literature to date has focussed on the use of

PSi modifi ed by oxidation, silanization or other depositions for cell culture.

Some studies have focussed on the use of lateral porosity gradients for the

optimization of cellular response. These studies/materials, while not being

polymeric in nature, have formed the basis for future work with polymeric/

PSi composite materials. We will briefl y discuss some of these studies in this

section.

The earliest report of cell culture on PSi comes from Bayliss et al ., who

published two papers in 1997 (Bayliss et al. , 1997a, b). In these papers the

authors used both PSi and porous germanium materials. The materials were

produced with a range of different porosities, and were tested for their expo-

sure to body salts and proteins and Chinese hamster ovary (CHO) cells. The

CHO cells were found to adhere to the porous surfaces, even after several

rinses and 24 h incubation. This was in direct contrast to crystalline control

substrates, which were not conducive to cell growth.

In 1999, Bayliss et al . also pursued the deposition of hydroxyapatite to

improve the biocompatibility of various Si substrates. They then assessed

the viability of immortalized B50 (rat neuronal cells) and CHO cells. The

study tested three surfaces, poly-Si (grown via plasma enhanced CVD), bulk

Si and PSi (prepared via electrochemical anodization), and compared them

to a glass standard. After 4 days in culture, the authors used two viability

assays and found that the optimal surface for the B50 cells was in fact the

hydroxyapatite-modifi ed PSi, while the CHO cells preferred the modifi ed

poly-Si. These fi ndings showed that nanoporous Si is a good candidate for

cell culture and possible development of bio-interfaced devices.

In 2003, Angelescu et al . published a large study on different surface modi-

fi cations of two types of PSi. In this study they produced PSi with an average

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pore size of 10 nm on both p+ <111> and p+ <100> substrates. These sur-

faces were then treated via either thermal treatment or carbon deposition.

These surfaces were then incubated with a variety of cells including human

retinal endothelial cells, mouse aortic endothelial cells, murine melanomas,

B50 and CHO cells. The authors found that the PSi substrates were biocom-

patible, non-toxic and appropriate for the culture of adherent cells without

further coatings of adhesion factors such as poly-lysine or collagen.

Other surface modifi cations on PSi have also been attempted. For exam-

ple, Low et al . (2006) have used amino-silanized PSi to help promote

mammalian cell adhesion. In this study, the authors compare PSi surfaces

functionalized by ozone oxidation, silanization and collagen or serum coat-

ing. Rat pheochromocytoma cells (PC12) and human lens epithelial cells

(SRA 01/04) were incubated on the surfaces for 4 and 24 h, and assayed

using cell counts and fl uorescent vital stain. Cells preferred to adhere to

surfaces that were collagen coated or amino-silanized, rather than those

that were oxidized or PEG-silanized. The silanization of PSi was found to

stabilize the surface and reduce the degradation rate. The study also found

that the PSi is capable of acting as a reducing agent, and this caused inter-

ferences with cell-based assays relying on redox enzymes. The high surface

area of PSi also facilitated the adsorption of certain dyes including Neutral

Red, compromising assays based on these dyes as well. In a follow-up work

published in 2009, Low et al . chose thermally oxidized and amino-silanized

PSi membranes with 40 − 60 nm pores for implantation under the rat con-

junctiva. The PSi was slowly eroding but was still visible for 8 weeks. It was

also noted that the surrounding tissue of the eye was not damaged in any

way. These PSi implants were able to support the attachment of human lens

epithelial cells, and when membranes carrying these cells were implanted

into rats, it was observed that the donor cells were able to move off the

membrane and into the ocular tissue space. The same authors have also

studied the generation of reactive oxygen species (ROS) from PSi in cell

culture medium (Low et al. , 2009b). In this work, it was found that untreated

PSi MPs in cell culture medium produced toxic levels of ROS, causing cell

death, while those MPs that were thermally oxidized did not reduce cell

viability. The authors point out that while high levels of ROS are toxic and

cause cell death it may be benefi cial to exploit the PSi to generate low levels

of ROS, which have been known to induce mitogenic response and increase

mammalian cell growth.

Sun et al . (2007) have shown that PSi can be used to promote osteoblast

growth, protein–matrix synthesis and mineralization. The study also con-

fi rmed that the PSi topology could also control the regeneration of bone

(osteoconductivity). This study was carried out on PSi that had nanoscale

(<15 nm) pores, mesoscale (c.50 nm) pores and macroscale (c.1 μ m) pores.

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The study found that the mesoporous PSi degraded faster than the macrop-

orous PSi and that this PSi may have some potential as a biodegradable can-

didate for bone-tissue engineering. It was found that the PSi surfaces bound

slightly fewer cells than the tissue culture plates, and that the macroporous

PSi anchored the most cells, with cell binding to the mesoporous PSi the

lowest of test surfaces. The cells tended to spread on the macroporous and

nanoporous PSi, while the mesoporous PSi did not allow spreading. After

seven days in culture, the macroporous PSi showed calcifi ed ECM, but it

took nearly 2 weeks for the other surfaces to show similar results. As the

macroporous PSi showed good viability, mineralization and maintained the

expression of biomarkers for bone formation, it was considered to be the

best candidate for osteoblast growth promotion.

The ability to change cell response by inducing changes in the surface

topology was further explored by Khung et al . (2008), who used continuous

PSi gradients to study the infl uence on neuroblastoma behaviour. In this

work, the authors created a pore size gradient from 3 μ m to 4 nm across the

surface by placing the electrode perpendicular to the Si substrate during

etching, generating an electrical fi eld density gradient. The study found that

these oxidized surfaces were able to exert different effects on cells depen-

dent on the pore size region to which the cells adhered. It was observed

that on the 1 − 3 μ m pore size region, the cells could not use their fi lopodia

to fi nd anchorage points, and cells in this region were notably elevated from

the surface. However, as the pore size decreased to 100 nm the cells were

able to produce more protrusions and the main processes were shorter.

Once the pores, sizes were below 50 nm, the normal neuroblastoma mor-

phology was recovered. These gradient surfaces could be used as a screen-

ing method to determine the best topology needed to support the correct

cellular morphology for future tissue culture applications. Similar gradi-

ent work looking at the attachment and differentiation of MSCs to n-type

PSi lateral gradients was performed by Wang et al ., (2012b, c). Meanwhile,

some gradient work that did incorporate polymeric coatings has also been

performed by Michelmore et al . (2012) and will be discussed later in this

chapter.

Since the discovery that PSi could induce the growth of hydroxyapatite

by Canham in the mid 1990s (Canham, 1995; Canham et al. , 1997), the induc-

tion of hydroxyapatite formation by PSi has also been widely investigated

(Hern á ndez-Montelongo et al. , 2010; Pastor et al. , 2007, 2009). Recently,

Sanchez et al . (2011) have deposited nanoscale hydroxyapatite using a

cathodic bias. Hydroxyapatite is often combined with various materials as it

displays good bioactivity (Laranjeira et al. , 2010) and biocompatibility (He

et al. , 2011; Xiao et al. , 2010), and it has a similar composition to that of

mammalian bone (Meneghini et al. , 2003). The materials used in the study

by Sanchez et al. possessed pores in the size regime of 5 − 7 μ m, larger than

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those found to be optimal by Sun et al . (2007). The modifi ed surface did,

however, prove to be a better support for cell culture than the equivalent

unfunctionalized PSi.

Since the seminal work of Bayliss et al . there have been many attempts to

interface PSi with cells for applications in areas such as in vitro biosensors,

implantable devices and interfaced neuronal networks. For example, Mayne

et al . (2000) used confocal microscopy to image patterned B50 cells on PSi

substrates. This study was vital in demonstrating that the by-products of PSi

manufacture are indeed non-toxic, even at levels that exceed those found

in in vivo or in vitro environments. The study also used scanning electron

microscopy to study cell morphology, which was found to be normal, and

it was found that the B50 cells rested on a 100 nm thick layer of ECM pro-

teins and that there were points along the axons and dendrites where the

cells connected to the PSi surface, opening the possibility of direct signal

transduction.

In a similar fashion, de-Leon et al . (2005) used PSi as a substrate for cul-

turing of Aplysia neurons with the intention to use the PSi as a biosensor

for the activity of neural networks. The authors investigated both the pho-

toluminescence (PL) response of the PSi to changes in voltages and the PSi

refl ectivity response to chemical changes. The study found that the cultured

neurons were able to survive for at least 1 week on the PSi substrate and

that they showed normal passive membrane properties and were able to

generate action potentials. The PSi used was passivated via ozone oxidation,

which stabilized the PL against the acidic environments used for sensing.

These surfaces were able to sense neuronal activity using both changes in

the PL intensity and the refl ectivity of the surface upon sensing of acetyl-

choline. In a later study, Unal (2011) has found that the PL of PSi was blue-

shifted and decreased in intensity in cell culture.

Over the years, there have been many attempts to optimize the properties

of PSi to help enhance or guide the adhesion, proliferation and differen-

tiation of a variety of mammalian cells. Examples of this include work by

Chin et al . (2001) who used PSi of approximately 70% porosity with 2 − 5 nm

pores to culture primary rat hepatocytes. The authors employed ozone oxi-

dation to create a thin (5 nm) oxide surface that was silanol-terminated. The

authors argued that this produces a surface similar to bioactive glass and,

although it renders the surface less electro-active, tunnelling of electrons

can still occur through the thin oxide layer, allowing for electrical control to

be retained. The rat hepatocytes used in this study were found to preferen-

tially adhere and spread on PSi in the presence of serum and type I collagen.

Adhesion and spreading did occur when serum only was used but to a lesser

extent. It was also shown that the nanotopology of the PSi surface did not

change with the addition of the matrix and adhesion molecules, suggesting

that surface chemistry is the key variable here. The cells cultured on PSi in

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this study remained viable after 5 days and the viability was comparable to

tissue culture PSi (TCPS). This work also found that the liver function (urea

production) of the cells on PSi was comparable to that of cell culture on

TCPS and crystalline Si.

In 2006 Khung et al . (2006) showed that SK–N–SH cells could be guided

to grow on PSi surfaces. To create the cell guidance channels the PSi was

ablated with laser from a commercial Matrix-assisted laser desorption ion-

ization (MALDI) mass spectrometer; this could be done without the use

of a mask. In this work, the PSi was silanized with a terminal PEG silane

that prevented cells from adhering; instead, the cells grew on the ablated

PSi (Plate XIV). These patterns could be used to study the neuronal pro-

cesses and communication in an environment that supports cell growth

and contains PSi region, which could provide growth factors to support

the growing cells.

Sapelkin et al . (2006) have also investigated the growth of B50 cells on

PSi. In this study, they used stain-etched PSi with pore sizes ranging from

50 to 100 nm that was fabricated from both crystalline and polycrystalline

substrates. These surfaces were patterned with 100 μ m square pads or 100

μ m stripes. The authors found that the cells showed preference for the PSi

surface rather than the untreated surface. The authors found that the gap

between the cells and the PSi was only 20 nm. The study shows that the cell

growth pattern can be controlled by the surface topography alone.

Recently, our group has developed a simple method to pattern bio-elements

onto PSi surfaces and has subsequently used these surfaces to support

mammalian cells (Sweetman et al ., 2012). In this work, an NHS-alkene was

patterned onto the PSi surface using UV-initiated hydrosilylation. The un-

reacted PSi areas were subsequently backfi lled with a PEG-alkene to pre-

vent cell attachment. The NHS was further reacted to functionalize these

areas with fi bronectin. Fibronectin was chosen as it has the ability to medi-

ate cell adhesion (Shi et al. , 2012) and it was found that on these surfaces

99% of the cells adhered to the fi bronectin regions. The SK–N–SH cells

used in this study were found to begin spreading after just 6 h of incuba-

tion. This work highlights the ability of biocompatible PSi to be used with

conventional photolithographic techniques for patterning. The ability to use

this process with a wide variety of organic molecules and to subsequently

conjugate many different chemical and biological entities will prove very

useful for a myriad of biomedical applications.

The combination of PSi with polymers will also open up more complex

cell culture and tissue engineering prospects that could also allow for the

delivery of complex drug mixtures and the subsequent release of the tissue

from the surface. Works that are currently using PSi and polymer compos-

ites will be discussed in the next section.

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18.6.2 PSi as cell and tissue penetrating agents

PSi can be used in two different ways to assist the penetration of molecules,

such as DNA, proteins, drugs or dyes, across the cell membrane. The fi rst

method utilizes the physical presence of large PSi structures as a perme-

ation enhancer. The second method processes PSi into MPs or NPs in order

to permeate the cells.

The fi rst method was demonstrated in 2003 by Foraker et al . (2003), who

developed large (100 × 200 × 25 μ m 3 ) porous plate-like PSi particles by com-

bining thin fi lm deposition, photolithography and selective electrochemical

etching. These particles were then incubated on a Caco-2 human epithelial

cell monolayer, and did not penetrate or disrupt the cell monolayer integ-

rity (Fig. 18.2). The fl uorescein isothiocyanate (FITC)-loaded particles were

shown to enhance the traffi c of the fl uorophore into the cells by more than

10- and 50-fold when compared to FITC and permeation enhancer (sodium

laurate or sodium caprate) and FITC alone, respectively.

In 2007, Kaukonen et al . (2007) used thermally carbonized PSi MPs to

enhance the permeation of furosemide across Caco-2 monolayers. Again,

PSi particles

(a)

(b)

Caco-2 cell monolayer

Polycarbonate membrane

Collagen coating

18.2 (a) Scanning electron microscopy image, particle demonstrating

the thickness, of the PSi particles and (b) schematic of the cells lying on

Caco-2 monolayers. ( Source : Adapted from Foraker et al . 2003.)

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the authors found permeation was enhanced with no decrease in cell mono-

layer integrity. This is likely due to the PSi nanostructure not allowing the

drug to form a crystalline structure and, thereby, remaining in a highly dis-

ordered, more soluble form. This effect is also enhanced by the high surface

area, high localized concentration and increased wettability of the PSi MPs.

This work was further extended in 2010 by Bimbo et al . (2010), who showed

that the thermally hydrocarbonized MPs could be sized into the NP regime

and labelled with 18 F, allowing their bio-distribution to be monitored after

enteral and parenteral administration in rats, and demonstrating their ver-

satility as a fl exible platform for drug delivery and imaging applications.

In 2011, Bimbo et al . (2011) also performed similar work with thermally

oxidized PSi MPs and NPs. Enhanced permeation of griseofulvin was again

demonstrated with Caco-2 monolayers with no cellular internalization of

particles. In contrast, immune response cells (RAW 264.7 macrophages)

incubated with the 1 − 10 μ m particles demonstrated rapid internalization

of the MPs.

Serda et al . (2009a, b, 2011) have published extensively on a PSi MP sys-

tem that is capable of permeating cells. The particles were fabricated with

the assistance of photolithography to produce particles with a diameter of

1.6 or 3.2 μ m and pore sizes of approximately 6 nm or 26 nm. Additionally, it

was shown that these particles could be loaded with a secondary NP payload

such as iron oxide NPs. These particles were internalized by human umbili-

cal vein endothelial cells by phagocytosis. It was observed that cell prolif-

eration was unaffected over several days and cationic (amino-silanized)

particles were taken up more readily than oxidized particles (Fig. 18.3). This

illustrates the potential of these particles in biomedical applications, such as

drug delivery and imaging.

In much the same way, NPs can be used to penetrate the cellular mem-

brane. Park et al . (2009) used luminescent PSi NPs with near-infrared

luminescence to observe penetration into HeLa (immortalized epithe-

lial cervical carcinoma) cells and also monitor their in vivo distribution

in a mouse model. This was the fi rst demonstration of organ and tumour

imaging in live animals with fully biodegradable PSi NPs. Importantly,

as this imaging was achieved with the intrinsic luminescence of PSi, very

low toxicity was observed. Alsharif et al . (2009) have also used non-toxic

alkyl modifi ed PSi NPs for cellular uptake. These particles were fabricated

by etching PL PSi fi lms followed by hydrosilylation with undecene. The

resulting particles were found to lack cytotoxicity and resulted in different

accumulation rates in different cell lines, including HeLa, A172 (glioblas-

toma tumour), MCF7 (breast adenocarcinoma), pancreatic epithelial-like

carcinoma (PANC1), SW1353 (chondrosarcoma), and MDA231 (breast

cancer) and non-malignant primary human cells isolated from two dif-

ferent healthy subjects (skin fi broblast, (Human skin fi broblasts) HSF1

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and HSF2), along with human myoblast and renal cells (A5UG). The rate

and extent of accumulation of alkyl-SiNCs showed only minor differences

between the neoplastic cell lines. However, the accumulation rate was

observed to be higher in the malignant cells compared to normal human

primary cells. The authors propose to use these particles, with modifi ca-

tions such as active proteins or oligonucleotides, for the evaluation of

18.3 Internalization of oxidized (top), APTES (middle) and PEG (bottom)

functionalized PSi microparticles into endothelial cells. ( Source : From

Serda et al ., 2009b.)

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cellular processes, such as signalling pathways, in vivo due to the low tox-

icity and rapid accumulation.

At present, the enhancement of cell permeation using PSi–polymer

composites remains largely unexamined. In future work, polymers could

be more highly utilized in conjunction with PSi structure on particles to

enhance their longevity in vivo as a protective barrier, aid in the more

effective attachment of targeting ligands, or a combination of both

(Gabizon et al. , 2004; Serda et al ., 2009b). In an early example of the poten-

tial of this approach, PSi NPs have also been reported as a permeation

enhancer to improve molecular penetration through the cellular mem-

brane. One early example of an Si NP (non-porous) and polymer com-

posite for biological staining was provided in 2004 by Li and Ruckenstein

(Li and Ruckenstein, 2004). They generated UV-induced PL in crystalline

Si using HNO 3 and HF to reduce the particle size to <3 − 5 nm. The par-

ticles were then polymerized with PAAc and utilized for the bioimaging

of CHO cells. It was found that the PAAc coating passivated the sur-

face of Si particles and stabilized the PL, making them more resistant

to photobleaching than conventional cellular organic dyes. Very recently

Secret et al . (2012) have demonstrated the vectorization of PSi NPs, by

coating the CPT loaded NPs with antibodies to target neuroblastoma,

glioblastoma and B lymphoma cells. The authors demonstrated the suc-

cessful targeted uptake of the PSi NPs using fl ow cytometry and the drug

delivery abilities were assayed via trypan blue and MTS viability assays.

In all cases, the drug-loaded PSi NPs did not affect the cells which did not

express the specifi c receptor for the antibody-functionalized PSi NP. This

work holds great promise for cancer therapy and will hopefully soon be

extended into animal models.

18.7 Applications of PSi-polymer composites in tissue culture and bioengineering

18.7.1 PSi-polymer composites as a cell growth support

In 2006, Sailor’s group reported an in situ sensor for cell behaviour which

they termed ‘Smart Petri Dish’ (Schwartz et al ., 2006). This was achieved

by fi rst generating a PSi photonic crystal functionalized with undecylenic

acid prior to spin coating with 45 kDa PSi–polymer. The surface was then

treated at 180 ° C for 30 min to melt the polymer into the porous structure.

Subsequently, the PSi-coated surface was treated with an oxygen plasma to

generate a hydrophilic surface suitable for cell growth, similar to the tech-

nique employed in TCPS. The optical scattering signal from the PSi was

used to measure changes in cell viability in a label-free, non-invasive and

real-time manner.

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As mentioned earlier in this chapter, PSi has been used as a screening

tool for cell culture conditions such as chemistry and topology. Recently,

this work has been extended by Michelmore et al. (2012), who combined

PSi pore size gradients with plasma polymerization. The authors used a

PSi topology gradient of pores ranging from 10 to 100 nm combined with

a plasma polymer gradient ranging from 0.7 − 3.0% carboxylic acid groups.

The surfaces were then used to culture osteoblast-like (MG63) cells. It was

found that the greatest cell spreading and adhesion occurred on PSi with

2 − 3% carboxylic acid groups. This work has the potential for the develop-

ment into a high-throughput screening for cell–material interactions, and

the rapid analysis of the best combinations for the culture of cells on PSi

in vivo .

18.7.2 PSi-polymer composites for in vivo applications (transplantable scaffolds)

Coffer’s group produced composites of PSi and PCL via salt leaching and

microemulsion techniques (Coffer et al. , 2005). The fi nal composites con-

tained either 1% or 5% of PSi by mass. These composites were subsequently

tested to determine their cytotoxicity and cell proliferation and found that

the composites did not impede the proliferation of fi broblast cells (Coffer

et al ., 2005). Coffer’s group then extended this work in 2007 by the addi-

tion of PANi to the PSi–PCL composites already prepared (Whitehead

et al. , 2007). PANi is an electrically conductive polymer and was used by

the authors to electrically connect the PSi particles throughout the network

so that an external bias could be applied. The presence of the PSi in the

network was then used to induce calcium phosphate formation within the

porous polymer network. This inductive effect of the PSi was found to occur

within just 7 h under an electrical bias; however, under no bias this induc-

tion did not occur for up to a month. The composite of PSi, PCL and PANi

had no deleterious effect on cell growth. These materials if developed fur-

ther could prove very benefi cial in orthopaedics as they are non-toxic, able

to induce calcination, support cell proliferation and with small refi nements

could lead to fully degradable implantable scaffolds.

In 2010, Kashanian et al . (2010) encapsulated hydride-terminated and

oxidized PSi MPs (ca 55 μ m) into PCL fi bres produced by electrospinning

from a chloroform solution to form a non-woven fabric. These composite

materials were investigated for ophthalmic applications, by in vivo testing

in rat subconjunctiva (Fig 18.4). It was found that these materials actively

supported epithelial cell attachment and only produced a mild histiocytic

foreign-body reaction after 8–9 weeks in the rat subconjunctival space.

The authors rightly point out that the composite material generated here

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is advantageous due to the possibility of loading therapeutic or bioactive

factors into the PSi component, while the PCL component makes the scaf-

fold more fl exible than the PSi alone. Interestingly, the authors propose

that these materials could possibly be exploited as a two-stage drug deliv-

ery vehicle and for the transfer of primary cells for regenerative medicine,

either individually or simultaneously.

18.7.3 PSi-polymer composites as a reservoir for the release of cell infl uencing factors

Growth factors are a range of proteins that play a key role in the prolifera-

tion and differentiation of cells. Typically, these factors are excreted endog-

enously either by the cells themselves (autocrine) or upon communication

with surrounding cells (paracrine). Typical growth factors used in tissue

engineering include bone morphogenic proteins (BMP), fi broblast growth

Highvoltage

Polymer suppplyneedle

(b)

(a)

(c)

Taylor cone

Fibrous polymermat

18.4 (a) Schematic representation of the electrospinning process using

PSi particles, (b) SEM image of the fi brous PCL mat incorporating the

PSi microparticles and (c) 20 wt.% as-prepared PSi in PCL at 1 week

(left panel) and 8 weeks (middle panel) after implantation beneath the

conjunctiva; end-point histology (right panel). ( Source : Adapted from

Kashanian et al . (2010).)

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factors (bFGF), VEGF and transforming growth factor- β (TGF- β ) (Ikada,

2006a).

Growth factors can be delivered by three different methods. The fi rst

two rely on genetic manipulation with either plasmid DNA or transfec-

tion of the encoding gene with vectors to stimulate the cells to produce

the growth factor at the desired site. The third technique is to use the

implanted biomaterial to deliver a sustained amount of growth factor to

the cells (Ikada, 2006a). This once again highlights the overlap that exists

between drug delivery and tissue engineering. Of course, any carrier used

must prevent the growth factor from denaturing and must have its degra-

dation rate tuned to facilitate a constant release rather than a burst release

profi le. Currently, most studies investigating the release of growth factors

from biomaterials have only analysed the use of one growth factor at a

time; however, it is more plausible to deploy multiple growth factors to

promote tissue regeneration. For example, it would be benefi cial to induce

both the differentiation and neovascularization into newly forming tissue

to help induce the supply of nutrients to the newly forming tissue (Ikada,

2006a). When growth factors become more readily available and cheaper

(possibly by using only active peptide fragments), the controlled release of

growth factor cocktails will surely feature prominently in the next phase in

tissue engineering.

There are many examples of PSi and polymer composite materials that

have been used to deliver various other drugs, and these have the poten-

tial to be extended for the delivery of growth factors and other biologi-

cals. Some of the most interesting and novel developments in recent years

include the work of Batra et al. (2006) who have covered dye-loaded PSi

substrates via repetitive coating with a PCL layer of varying thickness. The

PCL layer effectively slowed the release rate of the dyes, depending on the

layer thickness used. Similarly, Coffer et al. have constructed composites of

high porosity PSi particles with PCL to generate scaffolds that are highly

biocompatible and could potentially be used for the sustained release of

drugs (Coffer et al ., 2005). Demonstrating increased control over drug

(insulin) release, Wu and Sailor prepared a PSi scaffold containing a cross-

linked chitosan hydrogel capping layer. The insulin release was stimulated

by immersion in buffer at a pH 6, avoiding the burst release observed for

insulin loaded onto uncapped PSi (Wu and Sailor, 2009).

The modifi cation of PSi with thermoresponsive polymers, such as poly

N-isopropylacrylamide (PNIPAM), is also quite popular. PNIPAM has also

been extensively combined with fl at Si, and there are several examples of

this type of work in the literature. It is important, as the switching of these

surfaces can cause the delamination of cell sheets for harvesting and sub-

sequent implantation (Plate XV). This effect was fi rst described in 1990

by both Yamada et al . (1990) and Takezawa et al . (1990). These techniques

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are advantageous as harvesting cell sheets from TCPS (involving trypsin

or Ethylenediaminetetraacetic acid (EDTA)) disaggregates the ECM and

destroys any tissue-like structures that may be present. Examples of cur-

rent work with PNIPAM for composite generation includes papers from

Cole et al . (2006) LeMieux et al. (2007) and Xu et al . (2004).

Examples of this work on PSi have been reported by Sailor’s group who

produced thermoresponsive PSi-PNIPAM composite materials in two dif-

ferent ways. First, by polymerizing a NIPAM hydrogel into the pores of PSi

(Segal et al. , 2007) and, second, with two-step grafting of pre-polymerized

PNIPAM chains to undecylenic acid modifi ed PSi surfaces via EDC cou-

pling (Segal et al ., 2009). In the fi rst case, the temperature-dependent swell-

ing of the polymer led to refractive index changes that were monitored by

interferometric refl ectance spectroscopy. It is conceivable that this compos-

ite material could be infused with drugs and used for controlled release trig-

gered by a temperature change, which is sustained for a period of time. This

could potentially give greater control over the release of drugs, particularly

if multiple delivery cycles are required.

In a similar fashion to the work by Segal et al. (2009), our laboratory has

produced PSi fi lms coated with PNIPAM. However, we used an SI-ATRP

technique as opposed to Segal’s ‘grafting to’ approach (Vasani et al ., 2011).

In our work, we were able to show that the PNIPAM switches in a thermo-

responsive manner and that the thermoresponsive nature of the PNIPAM

was affected by the thickness of the polymer. We could easily control the

thickness of the PNIPAM layer by controlling the polymerization time. The

infl uence of PSi pore size on the PNIPAM layer formation was also inves-

tigated using lateral PSi pore size gradients with pores ranging from 20

to 200 nm. It was found that the polymer loading decreased as the pore

size decreased along the gradient. Most importantly, we were able to show

that the release of the CPT payload was able to be switched on (above the

lower critical solution temperature (LCST) of the polymer) and off (below

the LCST of the polymer); this was controlled by the temperature of the

release solution. In the future, adaptation of this system could allow for the

release of a drug payload above the LCST of the polymer at high tempera-

ture, for example, when a wound becomes infected (Dargaville et al. , 2013;

Pace et al. , 2013).

Koh et al. (2008) investigated composites of photonic PSi fi lms distrib-

uted Bragg refl ector (DBR) and PMMA. Refl ectance peaks in the region

of low optical absorption for human tissue (500 − 600 nm in this case) were

chosen for the PSi DBR. The free-standing PSi DBR was cast in a PMMA

solution containing the drug to be released. The resulting composite was

about 140 μ m thick, fl exible and chemically robust. Release of caffeine from

the composite was monitored in vitro by UV-vis, by monitoring changes in

the refl ection peak. Different refl ection peaks could be designed/controlled

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by varying the number of etching repeats used to generate the initial DBR.

This work opens the possibility for the development of ‘smart-patches’,

which enable the delivery of drugs to be monitored in real-time.

Mukherjee et al. prepared drug-loaded blocks of various shapes and sizes,

approximately 3 mm 3 in volume, from PCL. PSi MPs were subsequently

embedded in selected faces (enrichment sites) of the blocks by localized

melting of the PCL. The blocks could be self-assembled into networks by

exploiting the hydrophilic attraction of the oxide terminated PSi regions

and the use of starch to help drive the coupling of the enrichment sites.

The composites released the model compound Ru(bpy) 3 Cl 2 at varying rates

depending on the physical structure, spatial organization, PSi enrichment

site and composition of the polymer blocks (Mukherjee et al. , 2006).

18.8 Conclusions and future trends

Single crystalline Si has been modifi ed by various techniques and used in

a range of fundamental studies, mostly with cell culture and drug delivery

applications in mind. However, the in vivo use of these devices is unlikely,

as Si is not biodegradable and its low surface area limits drug-loading capa-

bilities. Conversely, PSi is biocompatible, shares common functionalization

techniques with crystalline Si, and benefi ts from a much larger drug-loading

capacity. Additionally, PSi is biodegradable, so no future surgical interven-

tion is required to remove implanted devices. Further benefi ts include the

ability to be processed into MPs and NPs, which increases the range of deliv-

ery options. PSi is a versatile material with a variety of tailorable properties,

including pore sizes, porosity and optical properties.

The combination of PSi with polymers to create hybrid materials for

implant applications can allow for the tuned release of therapeutic agents

(such as anti-infl ammatories or growth factors). This can be achieved by

modifying properties, such as hydrophobicity and porosity, which in turn

can infl uence the degradation and drug release rate. The combination of

PSi and stimuli-responsive polymers has also allowed for environmental

stimuli to be incorporated with the potential to be utilized to trigger drug

release or cell detachment. The main advantage of these stimuli-responsive

systems is the ability to switch from a passive to accelerated release of drugs

upon application of the stimulus. Release rates can be tuned by variations

in the polymer composition, loading technique and stimulation conditions.

The use of external stimuli such as heat, magnetism or light to induce drug

release could lead to the development of drug delivery devices that can be

surgically implanted subcutaneously and activated as required, eliminating

the need for multiple injections or frequent medical consultations.

The ability to easily functionalize with readily available techniques,

including oxidation, silanization, hydrosilylation and polymerization, results

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in the possibility of creating a myriad of materials with unique properties

ranging from biocompatibility to high mechanical strength and controlled

growth factor/drug delivery properties.

We suggest that future work with PSi–polymer composites will be

directed towards the generation of new injectable matrices that take full

advantage of the ability to fabricate PSi particles in the micro- and nano-

size range. This would allow PSi to operate as either a drug/cell carrier or as

a degradable porosifi cation agent. The ability to tune the in vivo degrada-

tion of PSi could also lead to its use as an agent to tune the biodegradation

of injectable matrices. This could be done not only by tuning the oxidation

state of the PSi from readily hydrolysable Si–H terminated to porous (SiO 2 –

terminated) glass, but also by changing the PSi particle’s surface function-

ality via silanization or hydrosilylation. In this manner, the degradation of

the matrix could be tuned from days to years. This line of research allows

for the incorporation of various growth factors and other drugs, and utilizes

truly 3D cell culture for the eventual replacement of damaged tissue; and

it does so in an in situ environment, which bypasses the known downfalls

of in vivo tissue engineering. Another potential stream of future work with

PSi–polymer composites will be to use the optical and electronic properties

of PSi to monitor cell behaviour in situ , and perhaps even in vivo .

The diverse PSi–polymer hybrid architectures that can be achieved, along

with the ability to tailor the functionality and properties of these hybrid

materials including their thickness, optical properties, release profi les, degra-

dation rates and mechanical strength, will lead to the generation of hybrids

that are well suited for different aspects of cell culture and tissue engineer-

ing and are likely to become more widely available in various applications

in the future as this technology is explored further.

18.9 Sources of further information and advice

Other sources of information for the avid reader include the following:

Books

Properties• of porous silicon, edited by L. T. Canham (Canham, 1997).

Porous silicon in practice: Preparation, Characterisation and Applications •

(Yamada et al ., 1990).

Tissue engineering and biodegradable equivalents: scientifi c and clinical applica-•

tions, edited by K. U. Lewandrowski, D. L. Wise, D. J. Trantolo, J. D. Gresser, M. J.

Yaszemski and D. E. Altobelli (Morita and Ikada, 2002).

Tissue Engineering: Fundamentals and applications, edited by Y. Ikada (Ikada, •

2006b).

Advances in Tissue Engineering, edited by Julia M. Polak (Polak, 2008). •

Topics in Multifunctional Biomaterials and Devices, edited by N. Ashammakhi •

(Davis and Leach, 2008).

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Principles of Tissue Engineering, 3 • rd Edition, Edited by R. Lanza, R. Langer and

J. Vacanti (Lanza et al. , 2007).

Professional and trade bodies

Int• ernational Society for Stem Cell Research (ISSCR)

Website available at http://www.isscr.org//AM/Template.cfm?Section=Home American Chemical Society •

Website available at http://portal.acs.org/portal/acs/corg/content

American Institute of Chemical Engineers (AIChE) •

Website available at http://www.aiche.org/

Institute of Chemical Engineers (IChemE, Worldwide) •

Website available at http://www.icheme.org/

Society for Biological Engineering (US) •

Website available at http://www.aiche.org/sbe

Biomedical Engineering Society (BMES) •

Website available at http://www.bmes.org/aws/BMES/pt/sp/home_page

The European Alliance for Medical and Biological Engineering & Science •

(EAMBES)

Website available at http://www.eambes.org/

National tissue engineering societies such as:

The Australasian Society for Biomaterials and Tissue Engineering (ASBTE) •

Website available at http://www.biomaterials.org.au/

and European Society for Biomaterials (ESB) •

Website available at http://www.esbiomaterials.eu/

Research and interest groups

The S• ailor Research Group, University of California, San Diego – Department

of Chemistry and Biochemistry.

Website can be found at http://sailorgroup.ucsd.edu/

The Voelcker Laboratories, Mawson Institute, University of South Australia, •

Adelaide. Website can be found at http://www.unisa.edu.au/research/mawson-institute/research/surface-engineering/

The Coffer Research Group, Texas Christian University, College of Science and •

Engineering.

Website can be found at http://www.chm.tcu.edu/faculty/coffer/

Society for Biomaterials, featuring many specifi c research interest groups •

Website can be found at http://www.biomaterials.org/special_interest_group.cfm

Tissue Engineering and Regenerative Medicine International Society •

(TERMIS)

Website can be found at http://www.termis.org/

18.10 Acknowledgement

The authors would like to acknowledge Dr Frances Harding for her

helpful and insightful discussions in the fi elds of cell culture and tissue

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engineering and we greatly appreciate the time-spent proof reading this

manuscript.

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Grafting to

PSi films

PSi membranes

Spin coating, impregnationand dip coating

Grafting from

Electropolymerisation Plasma polymerisation

iCVD

ElectrospinningGrafting through

Casting

PSi microparticles and nanoparticles

Plate XIII (Chapter 18) Schematics of common ‘grafting from’, ‘grafting

to’ and ‘grafting through’ polymerisation reactions of PSi as well as

non-covalent composite fabrication techniques such as spin coating,

electrospinning and casting.

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100 µm

(a) (b)

(c) (d)

100 µm 100 µm

200 µm

100 µm

PSi sufacePlain siliconsurface

Plate XIV (Chapter 18) Fluorescence microscopy images of cell

guidance provided by the ablated materials developed by Khung

et al . (2006). (a) Confl uent cell growth was observed on nonporous

PEG-functionalized Si surface and not the porous PEG-functionalized

surface. Selective cell adhesion to (b) ablated lines and (c) multiple

parallel lines (inset) crossed-line pattern and (d) squares. The cells were

stained with 3,3ʹ-dioctadecyloxacarbocyanine and propidium iodide.

PSi with PNIPAM coatingT < LCST

Protein resistance

PSi with PNIPAM coatingT > LCST

Protein adsorption(from cell culture media)


Cell adhesion


ECM formation

PSi with PNIPAM coatingT < LCST

Cell sheet lift-off

Plate XV (Chapter 18) Schematic representation of the harvesting of

cell sheets from thermoresponsive PNIPAM functionalised PSi surfaces.

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porous silicon-polymer composites for cell culture and tissue engineering applications

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