7

CHAPTER 2

LITERATURE REVIEW

This chapter includes literature related to tissue engineering,

materials and structure of scaffold, and method of production of scaffold used

for tissue engineering applications. The literature related to the application of

silk for tissue engineering applications, characteristics of different varieties of

silks are included. Different forms of scaffolds, especially the nano fibrous

scaffolds, produced by the electrospinning method are discussed in detail.

2.1 TISSUE ENGINEERING

2.1.1 Current Organ Repairing Therapies

Tissue loss or organ failure, resulting from traumatic or non-

traumatic destruction, gives rise to a major health problem that directly affects

the quality and length of patient’s life. These circumstances often call for

surgical treatments to repair, replace, maintain, or augment the functions of

the affected tissue or organ using some additional functional component that

facilitates an improved life to the patients. They have been traditionally

treated with the help of tissues or organs procured from the donors.

Depending on the location of implantation, the procured tissue or organ (also

called graft) is termed as autograft, allograft, or xenograft as shown in

Figure 2.1 If the graft is implanted in the same patient, it is termed as

autograft. Autografts (tissues removed from the patient) are typically

considered to be the gold standard in treating injuries. Autografts possess the

necessary amount of initial mechanical strength and promote cell proliferation

8

and new tissue growth. There is no risk of rejection or disease transmission

associated with autografts, since the tissue comes from the patient. However,

autografts have disadvantages as well (Murugan and Ramakrishna 2007).

Autografts require additional surgery for tissue harvest, which may cause

donor site morbidity (Freeman and Kwansa 2008). In the case of allograft and

xenograft, there have been concerns about infection and second site morbidity

of such tissues, and this type of graft may be rejected by host body due to the

immune response to tissue (Jackson et al 1990).The procurement of living

tissue/organ is complex and expensive, and requires additional surgery. This

clearly indicates the need for a potential solution to overcome the limitations

of traditional therapies and, at the same time, to increase the accessibility and

long-term survivability of the tissue implants (Yarlagadda et al 2005).

Figure 2.1 Conventional medical therapies

9

2.1.2 Tissue Engineering Concepts

Tissue Engineering is the application of principles and methods of

engineering and life sciences, towards fundamental understanding of the

structure/function relationships in normal and pathological mammalian tissues

and the development of biological substitutes to restore, maintain or improve

body functions. The key challenges in Tissue Engineering are synthesis of

new cell adhesion-specific materials and development of fabrication methods

to produce three-dimensional synthetic or natural biodegradable polymer

scaffolds with tailored properties. In tissue engineering, the scaffold serves as

a three-dimensional (3D) template for cell adhesion, proliferation and

formation of an extracellular matrix (ECM), as well as a carrier of the growth

factors or other bio- molecular signals. The scaffold fundamentally needs

some properties, such as porosity, pore size distribution, mechanical strength,

and required rate of degradation and bio-compatibility (Chen et al 2004).

2.2 FUNDAMENTAL NEEDS OF SCAFFOLDS

The fundamental needs of scaffolds are:

Biocompatibility: acceptance within the body without causing

bio-fouling, in which the body attacks the implant, or the cells

do not grow on the material.

Biodegradability: ability to degrade in the body into

compatible by-products without causing inflammatory

responses.

Mechanical integrity: ability to maintain the original structure

and mechanical properties upon exposure to the body’s

environment, i.e., 37°C, pH 7.4, phosphate buffer saline

solution

10

High porosity: ability to allow the transfer of nutrients/oxygen

and removal of wastes via diffusion.

Bioactivity: ability to transform or conform, depending upon

the influence from the internal milieu that surrounds the

scaffold seeded with cells (McCullen 2006).

2.3 BIOPOLYMERS FOR SCAFFOLD PREPARATION

2.3.1 Bio-degradable Polymer

Biodegradable polymers are solid polymeric materials and devices

which break down due to macromolecular degradation with dispersion in

vivo, however no proof exists for their elimination from the body (this

definition excludes environmental, fungi or bacterial degradation).

Biodegradable polymeric systems or devices can be attacked by biological

elements, so that the integrity of the system and in some cases but not

necessarily, of the macromolecules themselves, is affected and produces

fragments or other degradation by-products. Such fragments can move away

from their site of action but not necessarily from the body (Vert et al 1992,

Hutmacher 2000).

2.3.2 Bio-resorbable Polymer

Bio-resorbable polymers are solid polymeric materials and devices

which show bulk degradation and further resorb in vivo; i.e., polymers which

are eliminated through natural pathways, either because of simple filtration of

degradation by-products or after their metabolization. Bio-resorption is thus a

concept, which reflects the total elimination of the initial foreign material and

of bulk degradation by-products (low molecular weight compounds) with no

residual effects (Vert et al 1992).

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2.3.3 Bio-erodible Polymer

Bio-erodible polymers are solid polymeric materials or devices,

which show surface degradation and further, resorb in vivo. Bio-erosion too is

thus a concept, which reflects total elimination of the initial foreign material

and of surface degradation by-products (low molecular weight compounds)

with no residual side effects (Vert et al 1992).

2.3.4 Bio-absorbable polymer

Bio-absorbable polymers are solid polymeric materials or devices,

which can dissolve in body fluids without any polymer chain cleavage or

molecular mass decrease. For example, it is the case of slow dissolution of

water-soluble implants in body fluids. A bio-absorbable polymer can be bio-

resorbable, if the dispersed macromolecules are excreted (Vert et al 1992).

Both synthetic polymers and biologically derived (or natural) polymers have

been extensively investigated as biodegradable polymeric biomaterials. The

biodegradation of polymeric biomaterials involves the cleavage of

hydrolytically or enzymatically sensitive bonds in the polymer leading to

polymer erosion (Katti et al 2002). Depending on the mode of degradation,

polymeric biomaterials can be further classified into hydrolytically degradable

polymers and enzymatically degradable polymers. Natural polymers can be

considered as the first biodegradable biomaterials used clinically. The rate of

in vivo degradation of enzymatically degradable polymers however, varies

significantly with the site of implantation, depending on the availability and

concentration of the enzymes. The chemical modification of these polymers

can also significantly affect their rate of degradation. Natural polymers

possess several inherent advantages, such as bioactivity, the ability to present

receptor-binding ligands to cells, susceptibility to cell-triggered proteolytic

degradation and natural remodeling. The inherent bioactivity of these natural

polymers has its own downsides. These include a strong immunogenic

12

response associated with most of the polymers, the complexities associated

with their purification and the possibility of disease transmission (Nair and

Laurencin 2007).

2.4 HYDROLYTICALLY DEGRADABLE SYNTHETIC

POLYMERS

Hydrolytically degradable polymers are polymers that have

hydrolytically labile chemical bonds in their back-bone. The functional

groups susceptible to hydrolysis include esters, orthoesters, anhydrides,

carbonates, amides, urethanes, ureas etc.

2.4.1 Aliphatic Polyesters

Poly ( -esters) are thermoplastic polymers with hydrolytically

labile aliphatic ester linkages in their backbone. Although all polyesters are

theoretically degradable, as esterification is a chemically reversible process,

only aliphatic polyesters with reasonably short aliphatic chains between ester

bonds, can degrade within the required time for most of the biomedical

applications. Poly ( -esters) comprises the earliest and most extensively

investigated class of biodegradable polymers. The uniqueness of this class of

polymers lies in its immense diversity and synthetic versatility. Poly ( -ester)

can be developed from a variety of monomers via ring opening and

condensation polymerization routes, depending on the monomeric units.

Bacterial bioprocess routes can also be used to develop some poly ( -esters).

Various synthetic routes for developing polyesters have been recently

reviewed by Okada (2007) and Davachi et al (2011).

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2.4.2 Polyglycolide

Polyglycolide is the simplest linear aliphatic polyester, which is one

of the first biodegradable synthetic polymers investigated for biomedical

applications. It is prepared by the ring opening polymerization of a cyclic

lactone, glycolide. Polyglycolide is a highly crystalline polymer, with a

crystallinity of 45–55%, and therefore, exhibits a high tensile modulus with

very low solubility in organic solvents. The glass transition temperature of the

polymer ranges from 35 to 40 C, and the melting point is higher than 200 C.

Polyglycolide is not soluble in most organic solvents; the exceptions are

highly fluorinated organic solvents such as hexafluoro-isopropanol

(Morentet al 2011). Common processing techniques such as extrusion,

injection and compression molding can be used to fabricate polyglycolide into

various forms; its high sensitivity to hydrolytic degradation requires careful

control of processing conditions (Sabir et al 2009). Porous scaffolds,

electrospun nanofibrous scaffolds and foams can also be fabricated from

polyglycolide.

An excellent fibre forming ability of polyglycolide was initially

investigated for developing resorbable sutures. The first biodegradable

synthetic suture called DEXON that was approved by the United States Food

and Drug Administration (USFDA) in 1969 was based on polyglycolide.

Non-woven polyglycolide fabrics have been extensively used as scaffolding

matrices for tissue regeneration, due to their excellent degradability, good

initial mechanical properties and cell viability on the matrices, and their

ability to help regenerate biological tissue. The polymer is known to lose its

strength in 1–2 months when hydrolyzed, and loses mass within 6–12 months.

In the body, polyglycolides are broken down into glycine, which can be

excreted in the urine or converted into carbon dioxide and water via the citric

acid cycle (Terasaka et al 2006).

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2.4.3 Polycaprolactone

Polycaprolactone is an aliphatic polyester and hydrophobic

polymer, prepared by the ring opening polymerization of caprolactone. It is

readily degraded by a variety of bacteria and moth. The ring-opening

polymerization of e-caprolactone yields a semi crystalline polymer with a

melting point of 58–63°C and a glass transition temperature of -60°C. The

repeating molecular structure of PCL homopolymer consists of five non polar

methylene groups and a single relatively polar ester group (Vroman and

Tighzer 2009). This structure gives PCL, the unique properties that are similar

to polyolefin, because of its high olefinic content, while the presence of

hydrolytically unstable aliphatic-ester linkage causes the polymer to be

biodegradable. Polycaprolactone has a slower biodegradation in in vivo than

PLA due to its higher crystallinity and hydrophobic property. It is used for

soft and hard tissue engineering and drug delivery. The main drawback of the

PCL polymer is that it has low cell attachment, as it is a hydrophobic

polymer. After blending starch, cellulose and PLA polymers with the PCL,

enhanced cell attachment could be achieved (Yu et al 2006).

2.4.4 Poly (dioxanone)

Poly(dioxanone) is a polyether – ester aliphatic polyester; it is

synthesized by the ring opening polymerization of p-dioxanone, resulting in

the first clinically tested mono filament synthetic suture, that is known as

PDS' marketed by Ethicon. P-Dioxanone or 1,4-dioxan-2-one, abbreviated as

PDO, is a colorless crystal or liquid. Poly (dioxanone) exhibits a crystalline

fraction of 55% and a glass-transition temperature between 10°C and 0°C

(Middleton and Tipton 2000). PDS is used for a suture for a long period, and

provides higher flexibility, higher strength retention, slow absorption rates

and lower inflammatory response rates, when compared to Vicryl (poly

(glycolic-co-lactic acid)) and Dexon (poly (glycolic acid)). It is mostly

15

preferred for application as sutures, because of the flexibility and easier

knotting capacity. In addition, vascular prostheses made of PDO have been

shown to be less thrombogenic than both PGA and Dacron® synthetic grafts

(Pillai and Sharma 2010). PDO is a synthetic bio-resorbable polymer, that

offers several advantages over the more traditional bio-resorbable polymers

like poly(glycolic acid), poly(lactic acid) and poly(lactic-co-glycolic acid),

because it has a slower resorption rate, and induces a lower inflammatory

response (Smith et al 2008).

2.4.5 Poly (trimethylene carbonate)

Poly(trimethylene carbonate) (PTMC)) aliphatic polycarbonates are

employed in many biomedical applications, owing to their high

biocompatibility, facile bio-degradation, low toxicity, and superior

mechanical properties as compared to those of structurally similar polyesters

(Nederberg et al 2007). It is manufactured by the ring opening polymerization

of trimethylene carbonate. The polycarbonate derived from trimethylene

carbonate (TMC or 1,3-dioxan-2-one) has been investigated quite extensively

for its potential utilization as a biodegradable polymer in important

biomedical and pharmaceutical applications, such as sutures, drug delivery

systems and tissue engineering (Gangly 2006). High-molecular-weight PTMC

an amorphous, rubbery polymer at room temperature shows good mechanical

performance, combining high flexibility with high tensile strength (Pego et al

2002). The advantage of PTMC is slow degradation; hence it is suitable for

long term medical applications.

2.4.6 Poly-3-hydroxybutyrate

Poly-3-hydroxybutyrate (PHB) is a bacterial polyester and of a

hydrophobic nature; it is obtained from many types of bacteria, and was

discovered in the year 1920. The bacteria (Bacillus megaterium) have the

16

ability to synthesize and polymerize the monomer of hydroxybutyric acid.

Polyhydroxybutyrate is a naturally occurring -hydroxyacid (linear polyester).

Its ability to degrade and resorb in the human body environment makes it a

suitable matrix for bioactive and biodegradable composite implants that will

guide tissue growth and be replaced eventually by newly formed tissue. The

PHB can also be used as surgical sutures, bone implant material, and drug

release system (Chen and Wang 2002). The brittleness of the PHB is largely

due to the presence of large crystallinity in the form of spherulites, which

form upon cooling from the melt. PHB can be injection molded or extruded,

provided care is taken to lower the melting temperature, and minimize the

residence time. Injection molded PHB bars often show high crystallinity and

higher melting temperature, especially, below the glass transition temperature.

It suffers from some disadvantages, including a narrow processability

window, relatively low impact strength, higher brittleness, and a hydrophobic

nature (Park et al 2001).

2.4.7 Polyurethanes

Polyurethane (PUR) elastomers are multi-block copolymers with an

alternating sequence of hard and soft-block locks. Polyurethane (PU)

elastomers have been used extensively in biomedical applications because of

their excellent biocompatibility and mechanical properties (Hung et al 2009).

Segmented polyurethanes are elastomeric block copolymers that generally

exhibit a phase-segregated morphology made up of soft rubbery segments,

and hard glassy or semi crystalline segments (Lligadas et al 2007). The

Melting temperature of polyurethane is in the range of 125-138°C. Segmented

polyurethanes can also be designed to have chemical linkages that are

degradable in a biological environment, and there has been some interest in

developing degradable polyurethanes for medical applications, such as

scaffolds for tissue engineering. Polyurethane semi crystalline polymer

17

possesses many useful properties; but it also exhibits drawbacks such as

hydrophobicity, low resiliency, and swelling and deformation upon

degradation. It can also be blended with amorphous lactic acid to overcome

above drawbacks (Vainio et al 1997).

2.4.8 Poly (ester amide)

Aliphatic segmented poly (ester amide)s, comprising a

crystallizable amide phase and a flexible amorphous ester phase, were

investigated for potential use in biomedical applications. Segmented poly

(ester amide) s are prepared by the melt condensation of preformed bisamide-

diols, 1,4-butanediol, and dimethyl adipate. Poly (ester amide)s are an

emerging group of biodegradable polymers that may cover both commodity

and specialty applications. These polymers have ester and amide groups in

their chemical structure, which are of a degradable character and provide

good thermal and mechanical properties (Galan et al 2011).

2.4.9 Pseudo Poly (amino acid)

As part of the continuing efforts to develop improved biomaterials,

a method for synthesizing a new class of poly (amino acids) was recently

proposed. These polymers, named “pseudopoly (amino acids)”, are different

from conventional poly (amino acids)s in that the polymer backbone is

formed by utilizing the side-chain functional groups on the monomeric -L-

amino acids or dipeptides. Such an approach offers the opportunity to create

polymers from naturally occurring metabolites, but without some of the

potential disadvantages of conventional poly (amino acids)s resulting from

the repeating amide bonds, e.g. poor mechanical strength and enzymatic

degradation. When degradable polymers are used as implant materials in

patients, the potential toxicity of the polymer degradation products and their

subsequent metabolites become a major anxiety. For this reason, poly amino

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acids (PAA) were particularly attractive candidates for biomedical

applications (Pulapura et al 1992). Synthetic PAAs that contain one or at most

two different amino acid residues were virtually non-immunogenic and were

found to degrade in vitro and in vivo to their respective amino acid building

blocks, which are non-toxic and natural metabolites. They provide good

structural stability and biocompatibility for biomedical applications

(Mallakpour et al 2011).

2.5 NATURAL BIODEGRADABLE POLYMERS

2.5.1 Chitin and Chitosan

Chitin is the most abundant natural amino polysaccharide, and is

estimated to be produced annually almost as much as cellulose. Chitin and its

de-acetylated derivative chitosan, are natural polymers composed of randomly

distributed -(1-4)-linked D-glucosamine (de-acetylated unit) and N-acetyl-D-

glucosamine (acetylated unit) (Ravikumar 2000). Chitin functions naturally as

a structural polysaccharide, like cellulose, but differs from cellulose in its

properties. Chitin is highly hydrophobic and is insoluble in water and most of

the organic solvents. It is soluble in hexafluoro iso-propanol, hexafluoro-

acetone, chloro alcohols in conjugation with aqueous solutions of mineral

acids. Chitin is a white, hard, inelastic, nitrogenous polysaccharide found in

the outer skeleton of insects, crabs, shrimps, and lobsters, and in the internal

structure of other invertebrates. Chitin is obtained basically from prawn/crab

shells; the chemical treatment of chitin produces chitosan. Chitin and chitosan

are recommended as suitable materials for wound dressing, since these natural

polymers have excellent properties such as biodegradability, biocompatibility,

non-toxicity, and adsorption (Dutta et al 2004). The reaction of chitosan is

considerably more versatile than cellulose due to the presence of NH2 groups.

The natural polysaccharide chitosan and its quaternized derivatives possess

high intrinsic activity against bacteria and fungi, low toxicity,

19

biodegradability and the ability to affect macrophage functions. Chitosan-

containing materials might contribute to faster wound healing processes,

making them good candidates for wound dressing applications. However,

these scaffolds have a low mechanical strength under physiological

conditions, thus limiting their applicability (Ignatova et al 2009).

2.5.2 Collagen

Collagen is an important biomaterial in medical applications due to

its special characteristics, such as biodegradability and weak antigenecity

(Zhang et al 2005). Collagen is found to make up one quarter of all the total

protein in the body. It possesses a fibrous structure; collagen is able to impart

structure and strength to body tissues, such as tendons, ligaments and skin.

Collagen proteins are comprised of polypeptide chains ( -chains) that form a

unique triple-helical structure that is 300 nanometers long and 1.5 nanometers

in diameter. There are more than twenty disparate collagen types that exist in

animal tissue, five of which are known to form fibres; Types I, II, III, V, and

XI. These types tend to self-assemble into periodic, cross-striated fibres,

which can reach centimeters in length and tens of microns in diameter. Type I

collagen is the predominant fibre-forming collagen type and is found in

bones, skin, teeth, and tendons. Type II collagen, considered the second most

abundant, is found in cartilaginous tissue, developing cornea, and vitreous humor

(Kadler et al 1996). The major disadvantage of collagen nanofibrous matrix may

be the loss of its structural integrity in an aqueous environment such as in the

human body. However, the combination of collagen with biocompatible

polymers that enhance biological interactions with cells and speed up tissue

regeneration, could improve the dimensional stability of scaffolds (Yeo et al

2008).

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2.5.3 Silk

Silk is popularly known in the textile industry for its lustre and

mechanical properties (Kaplan et al 1998). Silk was discovered in China

around 2700 B.C. Silk is traditionally manufactured by sericulture. This

ancient art was practiced in China, Korea and Japan since the fourth century,

and in the sixth century this technique reached Europe via the Silk Route

(Hyde 1984). Silk is now produced across Asia and Europe, although the

main sources are Japan, China and India. Silk has been of interest for over

5000 years not only for its properties of texture, tenacity and dyeing, but also

for its use in cosmetics, creams, lotions, makeup, powders, bath preparations

and pharmaceuticals (Brooks 1989).

Silks are generally defined as protein polymers that are spun into

fibres by some Lepidoptera larvae such as silkworms, spiders, scorpions,

mites and flies (Altman et al 2003). Silk is a natural filament produced by the

silkworm, Bombyx mori, which has been used traditionally in the form of

filaments in textiles for thousands of years. This silk contains a fibrous

protein termed fibroin (both heavy and light chains) that forms the thread core

and glue-like proteins termed sericin that surround the fibroin filament to

reinforce them together. The fibroin is a highly insoluble protein containing

up to 90% of the amino acids glycine, alanine, and serine, leading to anti-

parallel -pleated sheet formation in the filament (Asakura et al 1994). Silk

has been used as a textile material for a long period. After the removal of the

sericin from silk fibroin, it has been considered as the starting raw material for

non-textile applications, especially in the biomedical, cosmetic and

biotechnological fields, such as surgical sutures, wound cover materials,

controlled drug release carriers, tissue engineering scaffolds and repair

materials for skins, bones, ligaments etc. (Altman et al 2003). Silk fibroin has

more mechanical strength than other synthetic bio materials and it has higher

combined strength and toughness due to the presence of an anti-parallel -

21

sheet structure. The most extensively used silk is obtained from different types

of silk worm such as mulberry silk (Bombyx mori), and non- mulberry silks such

as eri (Attacus ricinii), muga (Antherae assama) and tassar (Antherae perni). The

mulberry silk belongs to the Bombycidae family, whereas the non –mulberry silk

belongs to Saturniidae.

2.5.3.1 Life cycle of mulberry silk

The life cycle of the mulberry silk worm (Bombyx mori) is shown

in Figure 2.2. In about 50 days, it completes its four-step metamorphosis; egg

or embryo, larva, pupa and adult (moth). The worms consume food (mulberry

leaves) only at the larval stage. Pupation occurs at the end of spinning (cocoon

formation). Silkworm silk is produced basically at one stage in the life cycle,

during the fifth larval instars just before the molt to pupa (Asakura et al 1997).

Figure 2.2 Life cycle of mulberry silk

2.5.3.2 Composition of silk gland

The silk worm extrudes two proteins, namely, fibroin and sericin.

The fibroin forms the core while the sericin is deposited as a coating. The

22

glands of the silk worm in which the proteins are secreted are shown in Figure

2.3.The silk gland consists of three sections, the rear, middle and front. The

rear section is generally 200 to 250mm long and has a diameter of 0.4 to

0.8mm. The fibroin is secreted in this section. The middle section is about 60

to 65 mm long and has a diameter of 1.2 to 2.5mm. Depending on the

products of secretion and other characteristics, the middle section of the silk

gland is in turn divided in to the rear portion, the central portion with a distal

and proximal section, and the front portion. The frontal section is 35 to 40

mm long and has a diameter of 0.05 to 0.3 mm, and has the function of

conducting the silk proteins. The fibroin is secreted in the posterior section of

the silk gland, and transferred to the middle section in which it is stored as a

viscous liquid. The bulk of the sericin is produced in the middle section

together with the pigments, which impart colour to coloured silks. The rear

section of the silk gland, which is narrow and highly convoluted, secretes the

fibroin as a 12- 15% polymer solution. The fibroin then passes in to the broader

middle section, where it is enveloped by sericin, which is synthesized in this

portion of the silk gland. The silk proteins get concentrated in to 30% in this gland.

The water present in the fibroin solution passes in to the sericin layer (Glurajani

1993).

Figure 2.3 Silkworm gland

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2.5.3.3 Composition of the mulberry silk filament

Gulrajani et al (1988) reported that the silk fibre is almost a pure

protein fibre composed of two types of proteins, viz., sericin and fibroin.

Sericin is chemically a non-filamentous protein. Besides sericin, raw silk also

contains other natural impurities, namely, fat and waxes, inorganic salts and

colouring matter as shown in Table 2.1.

Table 2.1 Composition of mulberry silk

Component Percentage

Fibroin 70-80

Sericin 20-30

Wax matter 0.4-0.8

Carbohydrate 1.2-1.6

In organic matter 0.7

Pigment 0.2

Total 100

2.5.3.4 Structure of mulberry silk filament

Mulberry silk has been shown to be composed of two protein-

monofilaments (named brins) embedded in a glue like sericin coating. A

similar structure has been observed in other silkworms’ silk. The brins are

fibroin filaments made up of bundles of nanofibrils, approximately 5 nm in

diameter, with a bundle diameter of around 100 nm. The nanofibrils are

oriented parallel to the axis and are thought to interact strongly with each

other (Hakimi et al 2007). A schematic representation of the structure of

mulberry thread is shown in Figure 2.4.

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Figure 2.4 Structure of mulberry silk filament

2.5.3.5 Physical properties of mulberry silk fibroin

Mulberry silk fibres are generally 10-20 m in thickness (size) and

each fibre is actually a duplet of two individual fibres, each with its own silk

coating (sericin) and an inner core (fibroin). The fibroin consists of thousands

of parallel fibrils (100-400 nm), which after ion-etching can be viewed by a

scanning electron microscope. The fibrils give the microfilament its grainy

structure. The fibres also contain small quantities of carbohydrate, wax and

inorganic components, which play significant roles as structural elements

during fibre formation. Mulberry silk fibres are not circular in the cross

section, but appear triangular. The fine structure of the silk fibres gives them

the dynamic qualities of excellent lustre, colour, exquisite texture and

excellent temperature retainability (Ayutsede 2005).

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2.5.3.6 Chemical properties of mulberry silk fibroin

Mulberry has two silk glands that constitute approximately

one-quarter of the worm’s mass, and produce liquid silk. This polymer is

composed of a 350 kDa (kilo Dalton) fibroin heavy chain (H-fibroin), a

25 kDa fibroin light chain (L-fibroin), and a family of proteins called sericin,

that bind the two threads together as they emerge from the glands, and harden

in contact with the air. The silk thread is pulled from the gland and can attain

a length of more than a kilometer (Wurm 2003). The elements of the

supramolecular structure of silk fibres are macro fibrils with a width of up to

6.5×105 nm, which in turn, consist of helically packed nanofibrils of 90 -170

nm diameter. Nanofibrils play an important role to enhance tensile strength.

Silk fibres produced by cultivated Bombyx mori silkworm consist

mainly of two proteins, sericin and fibroin; they also contain minor amounts

of residues of other amino acids and various impurities: fats, waxes, dyes, and

mineral salts. Depending on the cocoon strain, the fibroin content is

66.5- 73.5, and the sericin content, 26.5- 33.5 wt % (Sashina et al 2006). The

primary structure of mulberry silk consists of 12 repetitive regions called

crystalline regions and 11 non repetitive interspaced regions called

amorphous regions (Zhou et al 2001). The remarkable properties of silk fibres

are attributed to the distribution of microcrystalline and amorphous domains,

which are formed in the process of spinning by protein–protein interactions.

The overall composition of the silk fibroin in mol% consists of glycine

(42.9%), alanine (30.0%), serine (12.2%), tyrosine (4.8%), and valine (2.5%)

(Asakura and Yao 2002). The mole fraction of glycine, alanine, serine, and

tyrosine residues combined is 90%; their sequence is represented by the

general formula: Gly-Ala-Gly-Ala-Gly-Ser-Gly-Ala-Ala-Gly- [-Ser-Gly-(Ala-

Gly)n]8-Tyr- (Shimura et al 1982, Shimura et al 1976). The pendant fragments

26

of the fibroin macromolecules are nonpolar hydrophobic aliphatic

hydrocarbon (alanine, leucine, isoleucine, valine, proline) and aromatic

(phenylalanine) substituents, polar hydrophilic hydroxyl-containing residues

of serine, threonine, and tyrosine, carboxyl groups of aspartic and glutamic

acids, amino groups of lysine, and guanidine groups of arginine. The

secondary structure of fibroin is stabilized by various kinds of interactions.

Hydrogen bonds arise between functional groups of peptide macro chains and

between side fragments of macromolecules (Baker et al 1984). Three main

kinds of secondary structures of natural silk fibroin are distinguished: -

helical and -folded structures (silk I and silk II respectively) in crystalline

areas and as disordered conformation of random globules in amorphous areas.

The fibroin of natural mulberry fibres contains 56±5% macromolecules in the

-folded form and 13±5% macromolecules in the -helical form. Thus, the

fraction of highly ordered (crystalline) areas of the polymer reaches 60 -70%

in the silk fibroin (Trabbic and Yager 1998).

2.5.4 Eri Silk

Samia cynthia ricini (eri silk) (Family:Lepidoptera:Saturniidae),

the Indian eri silkworm, contributes significantly to the production of

commercial silk, and is widely distributed in the Brahmaputra river valley in

North-Eastern India (Vijayan et al 2006). Eri silkworm (Samia ricini) is a

traditional source of food in northeast India, where it is grown primarily for

silk and food uses (Longvah et al 2011). The eri silkworm is polyphagous in

nature and feeds on leaves of several food plants (Rajesh Kumar and Gangwar

2010). The eri silk worm feeds on the leaves of a variety of plants, but ‘Castor

plant’ (Ricinuscommunis) is the most important host plant. The leaves of

plants like kesseroo (Heteropanaxfragrans), ‘Gomari’ (Gmelinaarborea),

‘Gulancha’ (Tinosporacordifolia) etc., are also used (Mishra et al 2003). The

27

study shows that the castor plant feed yields more silk than other plant feeds.

Eri silkworm is multivoltine in character, and can be reared indoors. It is

generally hardy and not susceptible to diseases. It belongs to the family

Saturnidae and species ricini. The most widely domesticated eri silk worm is

Samia cynthia ricini (Kulkarni 2007).

2.5.4.1 Life cycle of eri silk

Eri silk worm has four stages; the egg, larva, pupa and the adult or

moth shown in Figure 2.5. The moth lays white eggs, which turn grey then

black just before hatching. Eggs hatch in seven days in hot weather, but may

take as long as 24 days in cold weather. Female moths lay eggs in clusters that

may contain as many as 100 or more eggs. During the first to the third larva

stage, the head is black and shiny but eventually will turn greenish-yellow or

yellow with a brownish patch on each cheek, when they reach the fourth and

fifth stage. In the fifth stage, the larvae eat enormously and grow very quickly

to their maximum stage of development. The well fed, full grown larvae are

cylindrical and about 90 -100 mm long. The general body colour is white,

which turns yellow before spinning (Capinera 2008).

28

Figure 2.5 Life cycle of eri silk worm

Figure 2.6 Composition of the eri silk gland

29

2.5.4.2 Composition of eri silk gland

Similar to the mulberry silk, the eri silk filament is extruded from

the silkworm; it consists of two portions, the fibroin and the sericin. The

sericin is glue like gum that covers the core filament. Both fibroin and sericin

are produced by very large flattened cells lining a pair of long tubular silk

glands as shown in Figure 2.6. The silk gland is divided into three regions, the

thin and flexuous posterior part, and the wider middle and anterior parts. The

silk fibroins are synthesized in the posterior part of the silk gland, and then

transported down the lumen into the middle part of the silk gland, in which it

is stored in a concentrated state as a weak gel before spinning. It has been

pointed out that the silk press part is important in the process of fibre

formation from the liquid silk fibroin (Asakura et al 2007).

Table 2.2 Chemical composition of eri silk bave (Manuals on sericulture 1987)

Component Percentage %

Fibroin 72.2

Sericin 5-11.9%

Fat 1.3

Moisture 14.6

Eri silk is mainly made up of fibroin, sericin and fat. Sericin is a

non filament which is dissolved in water, whereas the fibroin is a filament

covered by the sericin. The percentage of chemical composition of eri silk is

listed in Table 2.2. Eri silk contains low amount of sericin of 4.96% as

compared to 10% in mulberry silk, 8.62% in tassar silk and 7.88% in muga

silk (Mishra 2000).

30

2.5.4.3 Physical properties of eri silk

Eri cocoons are usually white in colour; however, brick red colour

cocoons are also available. Eri silk cocoons cannot be reeled as they are made

up of entangled layers, and are therefore spun like cotton into yarn. The

reeling process does not involve the killing of silk moth, as it is an open-

mouthed cocoon (Sarkar 1980). The fineness of eri silk ranges from 14 to

16µ. The average filament length of eri silk is approximately 450 meters. Eri

silk is durable and strong with a typical texture. Eri silk is similar to cotton

and has a unique aesthetic appeal. It appears like wool mixed with cotton and

the softness of silk. Each eri cocoon weighs about 1 to 5 g with a shell weight

of 0.2 to 0.7 g. The denier (d) of the filament is 2.2 to 2.5 d, with a tenacity of

3 to 3.5 g/d. Eri silk has an elongation percentage of up to 20-22. It has

excellent thermal properties, and can be substituted for wool. The moisture

retention capacity is 11 %. Eri silk is more crystalline than any other

non-mulberry silks (Sreenivasa et al 2005); also, eri silk has a higher

elasticity, strong durability, and immunity against disease and insects.

2.5.4.4 Chemical properties of eri silk fibroin

Eri silk fibroin consists mainly of repeated similar sequences (about

100 times) of alternative appearances of the polyalanine (Ala) 12–13

region and

the Gly-rich region (Nakazawa et al 2009). Eri silk is a protein, which is

containing two major amino acid residues alanine and glycine. Samia cynthia

ricini is a wild silkworm and the amino acid composition of the silk fibroin is

different from that of the silk fibroin of the domesticated silkworm Bombyx

mori. The sum of Gly and Ala residues in eri silk is 82% which is similar to

mulberry silk (71%), but the relative composition of Ala and Gly is reversed

(Asakura et al 1999). The primary structure of eri silk fibroin is composed of

alternate blocks of polyalanine regions and glycine-rich regions. The alanine

is dominant in the crystalline region of the silk fibroin, whereas the glycine

31

(Gly) motif provides elasticity to the amorphous region of the silk fibroin.

The eri silk fibroin is in the form of - helix and random coil structure in the

silkworm gland as (silk I structure) shown in Figure 2.7 and the silk II

structure is present in the spun silk filament after spinning, which is attributed

to the - sheet in the silk fibroin (Rousseau et al 2006, Asakura et al 1999).

Three major polypeptides are found in the eri silk with different molecular

weights of 97 KDa, 45KDa and 66 KDa. The molecules of 66 KDa represent

sericin, where as 97KDa and 45KDa indicate the presence of polypeptides,

which are connected by a disulfide bond in the silk fibroin. The eri silk fibroin

aqueous solution contains 70% alanine in a helix structure, and the rest of

alanine in the form of a random coil structure. The silk liquid transition from

- helix to structure by a thermal or mechanical method was observed by

Asakura et al (1988) and Nakazawa et al (1999).

Figure 2.7 – Helix structure of silk fibroin (Nakazawa and Asakura 2003)

32

2.5.4.5 Difference between mulberry silk and eri silk fibroin properties

2.5.4.5.1 Amino acid composition

Mulberry and eri silk fibroins contain three major amino acids viz.,

glycine, alanine and serine as shown in Table 2.3. The total composition of

the three amino acids of the eri silk fibroin (84.26%) is higher than that of the

mulberry fibroins (82.8%). The eri silk fibroin has higher hydrophilic to

hydrophobic ratio than the mulberry silk fibroin, which indicates that the eri

silk fibroin contains less amount of hydrophobic amino acid than the

mulberry silk fibroin. The eri silk possesses higher amount of sulfur content

amino acids (cystine and methionine) than the mulberry silk fibroin, which

helps to connect the heavy and light weight chain in the silk fibroin. The eri

silk fibroin possesses a higher ratio of basic to acidic amino acids, as well as a

substantially greater proportion of positive arginine and negative aspartic acid

residues. The positive charged amino acid supports cell growth and

attachment on the eri silk fibroin scaffold better than on mulberry silk (Sen

and Babu 2004, Mai-ngam et al 2011).

2.5.4.5.2 Moisture regain

The moisture regain percentage of eri silk fibroin is higher than that

of the mulberry silk fibroin due to the higher proportion of amino acids with

bulky side groups, and also the higher hydrophilic to hydrophobic amino acid

ratio (9.06–9.85) for eri silk fibroin, compared to that of the mulberry silk

(5.29–6.22). Moisture regain is crucial for cell attachment and spreading on

the silk fibroin scaffold (Sen and Babu 2004).

33

Table 2.3 Amino acid composition of eri and mulberry silk fibroin

Amino acids Eri silk (mol %) Mulberry silk (mol

%)

Aspartic acid 3.89 1.64

Glutamic acid 1.31 1.77

Serine 8.89 10.38

Glycine 29.35 43.45

Hystidine 0.75 0.13

Arginine 4.12 1.13

Threonine 0.18 0.92

Alanine 36.33 27.56

Proline 2.07 0.79

Tyrosine 5.84 5.58

Valine 1.32 2.37

Methionine 0.34 0.19

Cystine 0.11 0.13

Isoleusine 0.45 0.75

Leucine 0.69 0.73

Phenylalanine 0.23 0.14

Tryptophan 1.68 0.73

Lysine 0.23 0.23

2.5.4.5.3 Thermal behavior

Thermal analysis is a useful tool for monitoring the important

processing parameters and properties of textiles. In addition to other factors,

the response of a fibrous polymer to thermal treatment depends on its

chemical architecture and microstructure. Thermal treatments bring about

34

morphological changes in silks that will have a bearing on the mechanical and

other properties. The thermal stability of silk varies from mulberry to wild

silk. The thermal stability of wild silk fibroin ( tassar, muga and eri) is higher

than that of the mulberry silk fibroin, because of the higher ratio of

bulky/non-bulky amino acids (0.24-0.32) in wild silk fibroins compared to

that of mulberry (0.17-0.18), and also the presence of higher amount of

(Ala)n sequences in the crystalline regions (Babu and Sen 2007). Thermal

stability is essential for sterilizing purposes in biomedical applications.

2.5.4.6 Mulberry and non – mulberry silk fibroin for tissue

engineering applications

The silk protein fibroin, isolated from the cocoon of the

domesticated mulberry silkworm, Bombyx mori, is used extensively in

biomaterial design and tissue culture. The study by Acharya et al (2009) on in

vitro cell culture on Antheraea mylitta (A.mylitta) as a substrate, showed the

higher mechanical strength of A. mylitta , better adherence, growth and

proliferation patterns of feline fibroblast cells on antheraea mylitta fibroin

films compared to that of mulberry fibroin films. The antheraea mylitta silk

fibroin scaffolds are used for tissue engineering applications in different

forms, such as film, foam, salt leach, sponge and nanofibre. The Antheraea

pernyi silk (Oak silkworm) fibroin supported the attachment and growth of

human bone marrow mesenchymal stem cells (hBMSCs). Compared to the

mulberry silk fibroin, the Antheraea pernyi (A. pernyi) silk fibroin contains a

special [Arg - Gly - Asp] (RGD) tripeptide sequence, which favours the cells

to attach. It also contains a certain amount of amino acid with positive

charges, thus the A. pernyi silk fibroin is more beneficial for cells of human

beings and many kinds of mammals to adhere and proliferate than the of

mulberry silk fibroin (Tao et al 2009). The Antheraea pernyi silk film made of

nanofibre was similar to the extracellular matrix (ECM) on the nanoscale,

35

which promoted cell migration and proliferation (Mao et al 2012). Tassar silk

fibroin (TSF) nanofibre is the most favourable silk fibroin material for

supporting the attachment and growth of neurons (Qu et al 2010), and the TSF

is more compatible for the development of neurons than mulberry silk fibroin,

suggesting the potential use of tassar silk fibroin for preparing the tissue-

engineered nerve guides to treat nerve injuries or diseases. Tassar electrospun

nanofibre is very suitable for developing human embryonic stem cells

(hESCs)-derived neural precursors (NPs) (Wang et al 2012). The Antheraea

assama (muga) fibroin based micro / nanofibrous nonwoven scaffold

possessed good bio-compatibility, and blood compatibility and the scaffold

was found to be nontoxic and efficient in supporting cell adhesion and growth

(NareshKasoju et al 2009). Eri silk fibroin films exhibited greater adhesion,

proliferation rate and spreading of cells than mulberry fibroin films. The

better cell supporting properties of eri silk fibroin may be mainly governed by

the initial non-specific binding between the serum proteins and the substrate.

The greater compositions of hydrophilic and positively charged amino acids

of the non-mulberry fibroin molecules may result in appropriate hydrophobic

and electrostatic interactions, allowing the adsorption of protein layer with

proper composition and conformation for cell adhesion and spreading. Eri silk

fibroin provides better biomaterial scaffold design than the more commonly

used mulberry fibroin (Mai-ngam et al 2011). When compared with mulberry

fibroin, wild silk materials are used very limitedly in biomaterial applications,

because most of the wild silk varieties are rare in the world, apart from the

Asian region. Among the wild silks, the tassar silk fibroin is used for tissue

engineering in different forms, but the muga silk and eri silk fibroins are

hardly used for tissue engineering. Mai-ngam et al (2011) carried out research

on eri silk fibroin scaffold for tissue engineering in the form of films.

36

2.6 MINERALS AND ANTIBIOTICS USED AS BIO

MATERIALS

2.6.1 Hydroxyapatite

Hydroxyapatite (Hap) (Ca10 (PO4)6 (OH), is an important inorganic

biomaterial which has attracted the attention of researchers related to the

biomaterial field in recent years. Due to its chemical and structural similarity

with the mineral phase of bones and teeth, hydroxyapatite (Hap) has been

used clinically for many years. It has good biocompatibility in bone contact as

its chemical composition is similar to that of bone material (Porous

hydroxyapatite for artificial bone applications). Hydroxyapatite exhibits

excellent biocompatibility with soft tissues such as skin, muscle and gums

making it an ideal candidate for orthopedic and dental implants or

components of implants (Zhou and Lee 2011). The Hap has higher

crystallinity and higher chemical stability (Agrawal et al 2011). The

advantage of the electro spun biopolymer with Hap is higher strength and

enhanced hydrophilicity of the scaffold (Ito et al 2005). Hydroxyapatite

ceramics have been used as tissue engineering material for cell culture and

tissue repair due to their biocompatibility and osteo-conductivity. However,

brittleness and fatigue failure in the body of Hap ceramics limit their clinical

applications only for repair and substituting purposes. It has been found that

each category of bio-ceramic and polymer cannot fulfill well the demand of

bone repair application by itself. Therefore, hydroxyapatite/polymer bio-

composite offers a possible combination of the advantages of the two

biomaterials (Jie and Yubao 2004).

37

2.6.2 Amoxicillin Drug

Amoxicillin is a broad spectrum antibiotic effective against various

types of microorganisms, but it possesses a short biological half life of about

60 minutes. Hence, repeated administration is needed to maintain the blood

plasma concentration of amoxicillin (Ramesh et al 2010). In order to reduce

the adverse effects due to frequent dosing, there is a need of a controlled

release formulation. Hence, the nanoparticles of amoxicillin were prepared by

using naturally available, nontoxic, low cost polymer obtained from natural

biodegradable polymers. Amoxicillin has been used for various infections,

including septic absorptions, urinary tract infections, upper and lower

respiratory infections, skin and soft tissue, and gastro-intestinal tract

infections (Patel et al 2007). The amoxicillin is a poorly soluble broad-

spectrum antibiotic; hence, this drug is added with bio degradable polymers to

maintain a sustainable drug release.

2.7 DIFFERENT FORMS OF SILK BIOMATERIAL

Silk biomaterials can be prepared directly from silk, or regenerated

from silk fibroin solutions. An alternative way is to convert silk into ultrafine

particles through milling. Figure 2.8 shows a schematic diagram of processing

silk into various forms of diverse morphologies, which are used for tissue

engineering, wound dressing and drug delivery (Rajkhowa et al 2010).

38

Figure 2.8 Different forms of silk biomaterial

2.8 TYPES OF SCAFFOLDS AND PROCESSING TECHNIQUES

2.8.1 Felts or Meshes

Polyglycolic acid (PGA) in the form of tassels and felts were

utilized as scaffolds to demonstrate the feasibility of organ regeneration.

Meshes consist of either woven or knitted three-dimensional patterns of

39

variable pore sizes are used as scaffolds. The advantageous characteristic

features of meshes are large surface area for cell attachment and rapid

diffusion of nutrients in favour of cell survival and growth. However, they

lacked the structural stability necessary for in vivo use, which led to the

development of bonding techniques (Yang et al 2001, Mikos et al 1993).

2.8.2 Bonding

The bonding was prepared by different techniques such as chemical

bonding and composite. PGA fibres were aligned in the shape of the desired

scaffold and then embedded in PLLA- methylene chloride solution. After the

evaporation of the solvent at room temperature, the PLLA-PGA composite

was heated above its melting temperature and then cooled in a room

atmosphere to produce the desired scaffold. This technique is not suitable for

producing fine porosity control (Salgado et al 2004).

2.8.3 Phase Separation

The polymer is dissolved in solvents such as molten phenol,

naphthalene or dioxane at a low temperature (Lo et al 1996). Liquid – liquid

or solid – liquid phase separation is induced by lowering the solution

temperature. The subsequent removal of the solidified solvent rich phase by

sublimation leaves a porous polymer scaffold. One major advantage is to

incorporate the bioactive molecules into the matrix, without decreasing the

activity of the molecules, due to the harsh, chemical and thermal

environments. A slight change in the parameters, such as types of polymer,

polymer concentration, solvent/non solvent ratio, and most importantly, the

thermal quenching strategy, significantly affects the resultant porous scaffold

morphology (Nam and Park 1999).

40

2.8.4 Solvent Casting and Particulate Leaching

This is one of the methods of the porous scaffold preparing

technique. This method involved mixing of water-soluble salt particles into a

biodegradable polymer solvent solution. The mixture was then casted into the

desired shape mould, and the solvent was removed by vacuum drying and

lyophilisation. The water-soluble salt particles were then leached out with

water to leave a porous structure. This method is characterised by its simple

operation and adequate control of the pore size, and the porosity of material is

determined by the amount of salt, and the size of the particle added to the

solution. In this method, the distribution of salt is not uniform within polymer

solution (Mikos et al 1994). The cast solvent and particulate method is used to

produce a two dimensional structure, with a thickness between 500 and

2000µm. In this method, it is very difficult to produce a scaffold of a

thickness of more than 3000µm (Yang et al 2001).

2.8.5 Melt Molding

This process involves physically mixing a polymer with the defined

amount of calibrated leachable particles and loading this powder into a mould.

This is followed by the application of heat and pressures that result in the

melting of the polymer. The compression maximizes the packing of the

mixture. The heating process causes the fusion of the polymer, and promotes

the formation of a continuous polymeric network that gives mechanical

stability to the structure. The last stage consists of immersing the moulded

polymer-porogen composite in a solvent that selectively dissolves the porogen

agent. This methodology has been successfully applied in the production of

natural origin starch-based scaffolds. The drawback of the melt moulding

scaffold processing is the requirement of high temperature (Gomes et al 2002,

Correlo et al 2009).

41

2.8.6 Polymer- Ceramic Composite Foam

In solvent-casting technique the hydroxyapatite short fibres and the

porogen are dispersed in a PLGA/ methylene chloride solution. After the

solvent evaporation, leaching of the porogen leaves the open-cell porous

composite foam of PLGA reinforced with hydroxyapatite short fibres. With a

certain range of fibre content, these scaffolds have superior compressive

strength compared to non -reinforced materials of the same porosity (Yang et

al 2001).

2.8.7 Gas Foaming

Highly open porous biodegradable poly(L-lactic acid) (PLLA)

scaffolds for tissue regeneration were fabricated, by using ammonium

bicarbonate as an efficient gas foaming agent, as well as a particulate porogen

salt. A binary mixture of PLLA-solvent gel containing dispersed ammonium

bicarbonate salt particles, in a paste form, was cast in a mould, and

subsequently immersed in a hot water solution, to permit the evolution of

ammonia and carbon dioxide within the solidifying polymer matrix. This

resulted in the expansion of pores within the polymer matrix to a great extent,

leading to well interconnected macro porous scaffolds, having mean pore

diameters of around 300–400 µm. The major disadvantage of the gas foaming

technique is the failure to obtain uniform porosity in the scaffold (Nam et al

2000, Salerno et al 2009).

2.8.8 Membrane Lamination

Membrane lamination is another solid freeform fabrication

technique (SFF) used for constructing three-dimensional biodegradable

polymeric foam scaffolds with precise anatomical shapes. Membrane

lamination is prepared by solvent casting and particle leaching, and

42

introducing peptides and proteins layer by layer during the fabrication

process. The membranes with appropriate shapes are soaked in the solvent,

and then stacked up in three-dimensional assemblies with a continuous pore

structure and morphology (Maquet and Jerome 1997). The bulk properties of the

final 3D scaffolds are identical to those of the individual membranes. This method

generates the porous 3D polymer foams with defined anatomical shapes, since it is

possible to use computer assisted modeling to design a template with the desired

implant shape. The disadvantages of this technique are the layering of porous

sheets, resulting in lesser pore interconnectivity (Hutmacher et al 2000, Hutmacher

et al 2001), and the fact that it is a time consuming process since, only thin

membrane can be used in this process (Subia et al 2010).

2.8.9 Freeze Drying Method

The freeze drying technique is also used for the fabrication of

porous scaffolds (Whang et al 1995, Schoof et al 2001). This technique is

based upon the principle of sublimation. Polymer is first dissolved in a

solvent to form a solution of the desired concentration. The solution is frozen

and the solvent is removed by lyophilisation under the high vacuum, that

fabricates the scaffold with high porosity and inter connectivity (Mandal and

Kundu 2009a, 2009b). This technique is applied to a number of different

polymers including silk proteins (Vepari and Kaplan 2007, Altman et al 2003),

PGA, PLLA, PLGA and PLGA/PPF blends. The pore size can be controlled by

the freezing rate and pH; a fast freezing rate produces smaller pores. Controlled

solidification in a single direction has been used to create a homogenous 3D-pore

structure (Schoof et al 2001). The main advantage of this technique is that, it

neither requires high temperature nor a separate leaching step. The drawback of

this technique is the smaller pore size, long processing time and poor homogeneity

of the pore structure (Boland et al 2004, Hou et al 2003). The freeze drying

technique is used to make silk protein porous scaffolds for tissue engineering

applications (Mandal and Kundu 2008a, 2008b, 2009, Kundu et al 2008).

43

2.8.10 Knitted Scaffold

A knitted scaffold is a structure made from the interlacing of yarn.

The knitted scaffold was made from silk filament, PLA, PCL and silk etc.

(Chen et al 2010) for tissue engineering applications. The knitted scaffolds

provide a suitable architecture and mechanical properties to withstand the

stresses faced by tissue. The knitted micro fibrous scaffold possesses an

interconnected porous structure allowing better tissue growth and nutrient

supply. The knitted structure is most adequate scaffolding for ligament tissue

engineering (Sahoo et al 2007). The major drawback of the knitted structure is

that controlling and homogeneous cell seeding could be very difficult to

achieve, due to high porosity and larger hole size in the knitted scaffold

(Vaquette et al 2009).

2.8.11 Braided Scaffold

Three-dimensional braiding is defined as a system, where three or

more braiding yarns are used to form an integral braided structure, with a

network of continuous filaments and yarn bundles with fibrous architecture,

oriented in various directions. Three dimensional braiding systems can

produce thin and thick structures in a wide variety of shapes through the

selection of the yarn bundle size. The braided scaffolds are mostly fabricated

for ligament growth due to their tensile strength. The advantage of the braided

scaffold is the interconnected network of porous structure, which supports the

transportation of oxygen and nutrients to the implant site (Cooper et al 2005).

Drawback of the braided structure is poor cell seeding due to the limited

internal space (Ouyang et al 2003).

44

2.8.12 Electrospun Fibrous Scaffold

Electrospinning is a unique process to produce polymeric fibres in

the diameter range of 100 nm -5 m. Fibres produced by this approach are at

least one or two orders of magnitude smaller in diameter, than those produced

by conventional production methods like melt or solution spinning

(Srinivasan and Reneker 1995, Subbiah et al 2005). Based on earlier research

results, it is evident that the average diameter of the electrospun fibres ranges

from 100 nm–500 nm. In textile and fibre science related literature, fibres

having diameters in the range of 100 nm–500 nm are generally referred to as

nanofibres. Electrospun fibres have a small pore size and high surface area to

volume ratio. There is also evidence of sizable static charges in electrospun

materials that could be effectively handled to produce three dimensional

structures.

The stable electrospinning jet was described in detail by Reneker

and Chun (1996) as being composed of four regions: the base, the jet, the

splay and the collection. Electrospinning is aided by the application of the

high electric potential of a few kV magnitudes to a pendant droplet of

polymer solution/melt from a syringe or capillary tube as shown in Figure 2.9

(Kumbar et al 2008). A polymer jet is ejected from the surface of a charged

polymer solution when the applied electric potential overcomes the surface

tension. The ejected jet under the influence of applied electrical field, travels

rapidly to the collector and collects in the form of a non-woven web as the jet

dries (Reneker et al 2000). Before reaching the collector, the jet undergoes a

series of electrically driven bending instabilities in the base region and

emerges from the needle to form a cone known as the Taylor cone (Hsu and

Shivkumar 2004). The shape of the base depends upon the surface tension of

the liquid and the force of the electric field; jets can be ejected from surfaces

that are essentially flat if the electric field is strong enough. The charging of

45

the jet occurs at the base with solutions of higher conductivity being more

conducive to jet formation (Pham et al 2006). The diameter of the electrically

charged jet decreases under electro-hydrodynamic forces, and under certain

operating conditions this jet undergoes a series of electrically induced bending

instabilities during passage to the collection plate, which results in extensive

stretching. The stretching process is accompanied by a rapid evaporation of

the solvent, which leads to a reduction in the diameter of the jet (Sukigara et

al 2003).

Doshi and Reneker (1995) classified the parameters (shown in

Table 2.4) that control the process in terms of solution properties, controlled

variables and ambient parameters. The solution properties include viscosity,

conductivity, surface tension, polymer molecular weight, dipole moment, and

dielectric constant. The fibre diameter can be controlled by varying the

processing parameters, such as polymer solution concentration, viscosity,

applied charge and electric field, type of solvent employed, distance from the

tip of the capillary to the collection plate, flow rate, diameter and angle of

spin of the spinneret (Pham et al 2006). In 1969, Taylor derived the condition

for the critical electric potential needed to transform the droplet of liquid into

a cone (commonly referred to as the Taylor cone), and to exist in equilibrium

under the presence of both electric and surface tension forces as given in

Equation (2.1).

2 2 2

c

2L 3v 4H L 0 117 R

R 2/ . (2.1)

where Vc is the critical voltage, H is the distance between the capillary tip and

the ground, L is the capillary length, R is the capillary radius and is the

surface tension of the liquid (Lyons 2004).

46

Table 2.4 Process parameters of electrospinning (Pham et al 2006)

Process parameters Effect on morphology of fibre diameter

Viscosity/concentration Low concentrations/viscosities yielded defects in the

form of beads and junctions; increasing

concentration/viscosity reduced the defects

Conductivity/solution charge

density

Fibre diameters increased with increasing

concentration/viscosity

Increasing the conductivity aided in the production of

uniform bead-free fibres

Higher conductivities yielded smaller fibres in general

Surface tension No conclusive link established between surface tension

and fibre morphology

Polymer molecular weight Increasing molecular weight reduced the number of

beads and droplets

Dipole moment and

dielectric

Successful spinning occurred in solvents with a high

dielectric constant

Flow rate Lower flow rates yielded fibres with smaller diameters

High flow rates produced fibres that were not dry upon

reaching the collector

Field strength/voltage At too high voltage, beading was observed

Correlation between voltage and fibre diameter was

ambiguous

Distance between tip and

collector

A minimum distance was required to obtain dried fibres

At distances either too close or too far, beading was

observed

Needle tip design Using a coaxial, 2-capillary spinneret, hollow fibres

were produced

Multiple needle tips were employed to increase the

throughput

Collector composition and

geometry

Smoother fibres resulted from metal collectors; more

porous structure was obtained using porous collectors

Aligned fibres were obtained using a conductive frame,

rotating drum, or a wheel-like bobbin collector

Yarns and braided fibres were also obtained

Ambient parameters Increased temperature caused a decrease in the solution

viscosity, resulting in smaller fibres

Increasing humidity resulted in the appearance of

circular pores on the fibres

47

Figure 2.9 Electrospinning set up

2.8.12.1 Nanofibre scaffolds for tissue engineering

The nanofibres are extremely thin fibres with diameters ranging

from microns down to a few nanometers. Such small-size fibres could

physically mimic the structural dimension of the extracellular matrix of a

great variety of native tissues and organs, which are characterized by well

organized hierarchical fibrous structures realigning from nanometer to

millimeter scale. The scaffolds produced provide a highly porous

microstructure with interconnected pores and extremely large surface area to

volume ratios, which are conducive to tissue growth. They are very versatile

and allow the use of a variety of polymers, blends of different polymers, and

inorganic materials as well as the integration of additives, biomolecules and

living cells for tailoring different application requirements. The

electrospinning process is a simple, straightforward and cost-effective method

to make various types of scaffolds (Zhang et al 2005).

Tissue engineering has emerged as a promising alternative

approach to treat the loss or malfunction of a tissue or organ without the

limitations of current therapies. Tissue engineering involves the expansion of

48

cells from a small biopsy, followed by the culturing of the cells in temporary

three-dimensional scaffolds to form the new organ or tissue by using the

patient's own cells as shown in Figure 2.10.This approach has the advantages

of autografts, but without their associated problems of inadequate supply. The

porous three-dimensional temporary scaffolds play an important role in

manipulating cell function and guidance of new organ formation (Chen et al

2002). Tissue engineering holds great promise as an alternative strategy to

current treatment modalities of diseased or otherwise failed tissues. Most

strategies of tissue engineering rely on three-dimensional porous scaffolds to

mimic the natural extracellular matrix (ECM) as templates, onto which cells

attach, multiply, migrate and function. When cells are harvested from a donor

and seeded, scaffolds facilitate the organization of these cells into a three-

dimensional architecture, control cell behavior and subsequently direct the

formation of organ-specific tissue (Johnson et al 2010).

For tissue engineering, various forms of scaffolds such as sponge,

foam, film, woven, knitted and non-woven materials are being used. The

nanofibrous matrix has a much higher surface-to-volume ratio than those of

fibrous non-woven fabrics fabricated with the textile technology, or foam

fabricated with other techniques (Smith et al 2009). As biomaterials for tissue

engineering, electrospun meshes exhibit important advantages when

compared with other scaffolds. First, the interconnectivity of the voids

available for tissue ingrowths is beneficial compared to foams and sponges.

Second, ultra thin fibres produced by electrospinning offer an unsurpassed

surface- to- volume ratio among established tissue scaffolds. The latter is

expected to have important advantages on the availability and activity of

immobilized molecules (e.g. peptides, lectines, enzymes etc.). Thus,

electrospun fibres have been explored as better extra cellular matrix for tissue

engineering, novel carriers for bioactive drugs and filtrations for bimolecules.

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Besides the three-dimensional structure, biocompatibility and biological

activity is required, and this remains a major challenge for tissue engineering

applications (Burger et al 2006). Several methods have been developed to

fabricate highly porous biodegradable scaffolds including fibre bonding,

braiding, solvent casting, particle leaching, phase separation, emulsion freeze

drying, gas foaming and 3D- printing techniques. However, the simplicity of

the electrospinning process to generate nanofibres makes it an ideal process

for scaffold fabrication (Gandhi 2006).

Figure 2.10 Tissue culture on electrospun nanofibrous scaffolds

2.8.12.2 Other application of electrospun fibres

Electrospun nanofibrous mats are used for various applications as

shown in Figure 2.11 (Huang et al 2003)

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Figure 2.11 Electrospun nanofibre applications

2.8.12.2.1 Filtration

The nanofibrous filter had been widely used in both households and

industries for removing substances from air or liquids. Filters for environment

protection are used to remove pollutants from air or water. In the armed force,

they are used in uniform garments and isolating bags to decontaminate

aerosol dust, bacteria and even virus. The respirator is another example that

requires an efficient filtration function. A similar function is also needed for

some fabrics used in the medical area (Fang et al 2008). Conventional

mechanical fibrous filters made of microfibres exhibit minimum fractional

collection efficiency for the aerosol particle size ranging between 100 and 500

nm, which is called the most penetrating particle size (MPPS). Simple

theoretical calculations predict that this efficiency may be significantly

increased using nanofibrous media (Podgorsk et al 2006). Electrospun

nanofibres have the diameters that are 5-10 times smaller than the smallest

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melt blown filter. The nanofibrous filter possesses some special properties

such as smaller fibre size and higher pressure drop, which cause interception

and inertial impaction efficiencies higher than that of conventional melt

blown nonwoven filters (Grafe and Graham 2003).

2.8.12.2.2 Wound dressing

Electrospun nanofibre based wound dressing material potentially

offers many advantages than conventional wound dressing. Nanofibrous mat

provides intrinsic properties such as high surface area and micro porous

structure, which affords quick start signaling pathway and attracts fibroblasts

to the derma layer that can excrete important extracellular matrix components,

such as collagen and several cytokines (e.g., growth factors and antigenic

factors). The electrospun membrane enhances the cell attachment and

proliferation in wound healing (Chen et al 2008). The electrospun nanofibre

mats usually have pores which are small enough to prevent bacterial

penetration. The high surface area is of importance for fluid absorption and

dermal drug delivery. Although plenty of polymers have been successfully

electrospun into nanofibres, reports on electrospun nanofibrous mats suitable

for wound dressings are still scarce (Kanani and Bahrami 2010). Electrospun

nanofibrous membrane wound dressings can also meet the requirement of

high gas permeation, apart from providing effective protection of the wound

against infection and dehydration. Electrospun nanofibre can be easily

incorporated with pharmaceutical compounds, such as antiseptics, antifungal,

vasodilators and cell growth factor etc. But conventional wound dressing is

not provided with the feasibility for drug incorporation; and another major

advantage over conventional wound dressing is that the nano fibres hold a

promise of healing wounds without leaving scars (Zahedi et al 2010).

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2.8.12.2.3 Drug delivery

Many therapeutic compounds can be conveniently incorporated into

the electrospun fibres through the electrospinning process, which is unlike

common encapsulation methods involving some complicated preparation

processes. At present, both degradable and non-degradable polymers are

under investigation to be developed as drug carriers for local delivery of

antibiotics and anticancer drugs, and electrospun fibrous materials with

different structures are the preferred selection. The advantages of the

electrospinning technique are, maintaining of molecular structure and

bioactivity of the incorporated drugs or bioactive molecules due to the mild

process conditions, and reducing the burst release of drugs in vitro. Ignatious

et al (2010) defined that the release of pharmaceutical dosage from nanofibres

can be designed as rapid, immediate, delayed or modified dissolution

depending on the polymer carrier used for drug delivery. The biocompatible

and biodegradable polymer fibre mats produced by electrospinning methods,

are able to serve both as drug encapsulation vehicles and tissue engineering

scaffolds (Venugopal and Ramakrishna 2005). The release profile can be

finely controlled by the modulation of nanofibre morphology, porosity and

composition. Nanofibres for drug release systems are obtained mainly from

biodegradable polymers and hydrophilic polymers.

2.8.12.2.4 Other applications of nanofibres

Nanofibres fabricated through electrospinning have a specific

surface, approximately one to two orders of magnitude larger than flat films,

making them excellent candidates for potential applications in sensors. The

sensitivity of chemical gas sensors is strongly affected by the specific surface

of the sensing materials. A higher specific surface of sensing material leads to

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a higher sensor sensitivity; therefore, many techniques have been adopted to

increase the specific surface of sensing films with fine structures, especially to

form nanostructures, taking advantage of the large specific surface of nano

structured materials. Electrospun nanofibrous materials have attracted

considerable interest in the food industry for their utilization as highly

functional ingredients, and high-performance packaging materials. The

nanofibres are used as filtration membranes for environmental remediation,

which minimize the pressure drop and provide better efficiency than

conventional fibre mats. The large surface area-to-volume ratio of the

nanofibre membrane allows greater surface adsorption of contaminants from

air and water, and increases the life-time of the filtration media. It is also used

in solar and fuel cell applications (Ding et al 2009).