extruded collagen-polyethylene glycol fibers for tissue engineering applications

Extruded Collagen-Polyethylene Glycol Fibers forTissue Engineering Applications

D. I. Zeugolis,1,2,3 R. G. Paul,4 G. Attenburrow5

1 Tissue Modulation Laboratory, National University of Singapore Tissue Engineering Programme,National University of Singapore, 117510, Singapore

2 Division of Bioengineering, Faculty of Engineering, National University of Singapore, 117576, Singapore

3 Immunology Programme, Department of Microbiology, Yong Loo Lin School of Medicine,National University of Singapore, 117456, Singapore

4 Devro Plc, Glasgow, Scotland, G69 0JE, UK

5 School of Applied Sciences, The University of Northampton, NN2 7AL, UK

Received 28 April 2007; revised 5 July 2007; accepted 26 July 2007Published online 23 October 2007 in Wiley InterScience (www.interscience.wiley.com). DOI: 10.1002/jbm.b.30952

Abstract: The repair of anterior cruciate ligament, skin, tendon and cartilage remains a

challenging clinical problem. Extruded collagen fibers comprise a promising scaffold for tissue

engineering applications; however the engineering of these fibers has still to be improved to

bring this material to clinical practice. Herein we investigate the influence of collagen

concentration, the amount of PEG Mw 8K and the extrusion tube internal diameter on the

properties of these fibers. Ultrastructural evaluation revealed packed intra-fibrillar structure.

The thermal properties were found to be independent of the collagen concentration, the

amount of PEG or the extrusion tube internal diameter (p > 0.05). An inversely proportional

relationship between dry fiber diameter and stress at break was found. The 20% PEG was

identified as the optimal amount required for the production of reproducible fibers. Increasing

the collagen concentration resulted in fibers with higher diameter (p < 0.001), force (p < 0.001)

and strain at break (p < 0.02) values, whilst the stress at break (p < 0.001) and the modulus (p <0.007) values were decreased. Increasing the extrusion tube internal diameter influence

significantly (p < 0.001) all the investigated mechanical properties. Overall, extruded collagen

fibers were produced with properties similar to those of native or synthetic fibers to suit a wide

range of tissue engineering applications. ' 2007 Wiley Periodicals, Inc. J Biomed Mater Res Part B:

Appl Biomater 85B: 343–352, 2008

Keywords: collagen; scaffolds; tissue engineering; biomimetic

INTRODUCTION

Tissue engineering aims to design and use materials that

would interact with the body to encourage tissue repair. A

number of synthetic and natural biomaterials are currently

in use as tissue scaffolds,1 with the ideal ones to be those

that most closely mimic the naturally occurring environ-

ment in the host tissue matrix.2 Between them, the natu-

rally derived collagen has demonstrated a number of

desirable features such as high tensile strength, biodegrad-

ability and low immunogenicity,3,4 favoring it for tissue

engineering applications.

Previous studies on extruded collagen fibers have dem-

onstrated that such materials comprise an excellent scaffold

for soft and hard tissue replacement with structural, me-

chanical, and thermal properties similar to the native tis-

sue.5–7 Moreover their structural characteristics facilitate

fibroblast migration and neotissue formation.8 Although

momentous achievements have been made, many chal-

lenges still exist in the engineering of these scaffolds. For

example, extruded collagen fibers have been manufactured

up to now using either acid soluble rat tail tendon9,10 or

acid soluble bovine Achilles tendon.11,12 However, in both

cases the nonhelical peptides remain intact and a less bio-

compatible scaffold could be produced.4,13 Thus, herein we

utilised different concentrations of pepsin soluble bovine

Achilles tendon collagen, a raw material that has been

favored for biomimetic scaffold manufacturing.

Correspondence to: Dr. Dimitrios I. Zeugolis (e-mail: [email protected])Contract grant sponsors: The University of Northampton, EPSRC

� 2007 Wiley Periodicals, Inc.

343

Recently, many water-soluble polymers have become an

attractive field of study, due to their good processing char-

acteristics and variable degradation rates.14 It has been sup-

ported that collagen, which is a hydrogen donor, forms

hydrogen bonds with the hydroxyl group from the water-

soluble polymers.15 Polyethylene glycol (PEG), a low toxic

and low antigenic poly-ether-diol of general structure

HO��(CH2CH2O)n��H, has been FDA approved for sev-

eral medical and food industry applications.16,17 Addition-

ally, it has been shown to facilitate cell infiltration, tissue

in-growth, significant increase in the mechanical stability

and enzyme degradation with improved blood compatibility

and ability to resist protein adsorption.18,19 Furthermore, it

has been shown that the use of PEG for precipitating native

collagen has a number of advantages over conventional salt

precipitating methods.20,21 Despite these advantages, PEG

has been used sporadically in the fabrication of extruded

collagen fibers,12,22 whilst NaCl has been used

broadly,9,10,23,24 Therefore, herein we investigated the influ-

ence of different amounts of PEG on the properties of

extruded collagen fibers.

Finally, using extrusion tubes of different internal diam-

eter, we aimed to produce collagen fibers that would not

only imitate extracellular matrix assemblies with diameter

range between 1 and 300 lm,25–27 but also compete with

synthetic materials with stress and strain at break values

ranging from 32 to 840 MPa and 3 to 62%, respectively.28–30

Overall, the structural, mechanical, and thermal properties of

the produced fibers were found to match closely those of

native or synthetic fibers.

MATERIAL AND METHODS

Materials

All chemicals, unless otherwise stated, were purchased

from Sigma-Aldrich, UK. The bovine Achilles tendons

(BAT) were kindly provided by the BLC Research Centre

(Northampton, UK), whilst the rat tail tendons were

donated from the Department of Biomedical Sciences, Uni-

versity of Sheffield (Sheffield, UK).

Preparation of Collagen

The tendons were manually dissected out from the sur-

rounding fascia, minced in ice, and extensively washed

with distilled water and neutral salt solutions. Subsequently

were suspended in 0.5M ethanoic acid for 72 h at 48C in

the presence or absence of pepsin (porcine gastric mucosa

pepsin; 2500 U/mg, Roche Diagnostics, UK); enzyme to

tendon wet weight ratio 1:100. The suspension was filtered

and purified by repeated salt precipitation (0.9M NaCl),

centrifugation (12,000g at 48C for 45 min) (Gr20.22 Jouan

refrigerated centrifuge, Thermo Electron Corporation, Bath,

UK) and re-suspension in 1M ethanoic acid. The final col-

lagen solutions were dialysed (8000 Mw cut off) against

0.01M ethanoic acid and kept refrigerated until used. The

collagen concentration for the pepsin soluble bovine Achil-

les tendon was adjusted to 3, 6 and 7 mg/mL, whilst for

the acid soluble rat tail tendon at 3 and 4 mg/mL and for

the acid soluble bovine Achilles tendon at 3 and 5 mg/mL.

Viscosity Measurements

Viscosity measurements were carried out using a DV-III

Brookfield Viscometer (Brookfield Viscometers Ltd, Essex,

UK). The temperature (9.6–9.98C) of the collagen solutions

was maintained during the measurements by means of a

thermostatically controlled tank. A fixed volume of colla-

gen solution was measured each time. Readings were taken,

using the spindle No CP42, at rotational speeds of 1.2, 2.4,

3.6, 4.8, 6.0, and 7.2 rpm. Scale values were read every 5s

under shear. For each collagen solution, the measurements

were an average of two replicates.

Fiber Formation

The procedure for fiber formation has been described in

detail previously.9–12,23 Briefly, a 5 mL syringe (Terumo

Medical Corporation UK, Merseyside, UK) containing ei-

ther of the collagen solutions was loaded into a syringe

pump system (KD-Scientific 200, KD-Scientific, MA) con-

nected to silicone tubing (Samco Silicone Products, Ltd.,

Warwickshire, UK) of 30 cm in length and either 1.0, or

1.5 or 2.5 mm in internal diameter. The pump was set to

infuse at 0.4 mL/min. One end of the tube was connected

to the syringe pump with the other end placed at the bot-

tom of a container. Collagen solution was extruded into a

‘‘Fiber Formation Buffer’’ (FFB) comprising of 118 mMphosphate buffer and 5, 20, or 40% of polyethylene glycol

(PEG), Mw 8000 at pH 7.55 and 378C. Fibers were allowed

to remain in this buffer for a maximum period of 10 min,

followed by further incubation for additional 10 min in a

‘‘Fiber Incubation Buffer’’ (FIB) comprising of 6.0 mMphosphate buffer and 75 mM sodium chloride at pH 7.10

and 378C. Thereafter the fibers were transferred into a dis-

tilled water bath for further 10 min and finally air-dried

under the tension of their own weight at room temperature.

Mechanical Testing and Ultrastructural Matrix Analysis

Stress–strain curves of dry reconstituted collagen fibers

were determined in uniaxial tension using an Instron 1122

Universal testing machine (Instron Ltd, Buckinghamshire,

UK) operated at an extension rate of 10 mm/min. The

gauge length was fixed at 5 cm and soft rubber was used to

cover the inside area of the grips to avoid damaging the

fibers at the contact points. Results obtained with fibers

that broke at contact points with the grips were rejected.

The cross sectional area of each fiber was calculated by

measuring the diameter at five places (every 1 cm) along

its longitudinal axis using a Nikon Eclipse E600 optical

microscope with a calibrated eyepiece (Nikon Instruments,

Surrey, UK). It was assumed that the fibers were circular

344 ZEUGOLIS, PAUL, AND ATTENBURROW

Journal of Biomedical Materials Research Part B: Applied Biomaterials

for the cross-sectional area determinations. Fracture surfa-

ces of collagen fibers that had been extended to failure

were examined using a Hitachi S3000N Variable Pressure

Scanning Electron Microscope (Hitachi, Berkshire, UK).

The following definitions were used to calculate the me-

chanical data: stress at break was defined as the load at

failure divided by the original cross-sectional area (engi-

neering-stress); strain at break was defined as the increase

in fiber length required to cause failure divided by the orig-

inal length and modulus was defined as the stress at 0.02

strain divided by 0.02.

Thermal Properties

The shrinkage temperature was determined by Differential

Scanning Calorimetry (DSC) using the 822e Mettler-Toledo

differential scanning calorimeter (Mettler-Toledo Interna-

tional, Leicester, UK). The reconstituted collagen fibers

were hydrated overnight at room temperature in 0.01MPBS at pH 7.4. The wet fibers were removed and quickly

blotted with filter paper to remove excess surface water

and hermetically sealed in aluminium pans. Heating was

carried out at a constant temperature ramp of 58C/min in

the temperature range from 15 to 958C, with an empty alu-

minium pan as the reference probe. Thermal denaturation,

an endothermic transition, was recorded as a typical peak,

and two characteristic temperatures were measured corre-

sponding to the peak (the temperature of maximum power

absorption during denaturation) and onset (the temperature

at which the tangent to the initial power versus temperature

line crosses the baseline) temperatures.

Statistical Analysis

Numerical data is expressed as mean 6 SD. Analysis was

performed using statistical software (MINITABTM version

13.1, Minitab, Inc.). One way analysis of variance

(ANOVA) for multiple comparisons and 2-sample t-test forpair wise comparisons were employed after confirming the

following assumptions: (a) the distribution from which

each of the samples was derived was normal (Anderson-

Darling normality test); and (b) the variances of the popula-

tion of the samples were equal to one another (Bartlett’s

and Levene’s tests for homogenicity of variance). Nonpara-

metric statistics were utilised when either or both of the

above assumptions were violated and consequently Krus-

kal-Wallis test for multiple comparisons or Mann-Whitney

test for 2-samples were carried out. Statistical significance

was accepted at p\ 0.05.

RESULTS

Matrix Morphology

Scanning electron microscopy evaluation revealed homolo-

gous surface independent of the collagen concentration, or

the relative amount of PEG in the FFB. The fibers were

characterised by ridges and crevices running roughly paral-

lel to the axis of the fibers, with a ‘‘cavity’’ lying along

their longitudinal axis [Figure 1(a)]. After microscopic ex-

amination of extended to failure fibers, it became apparent

that their inter-fibrillar space was filled [Figure 1(b)]. Typi-

cal fracture modes were obtained classified as (a) smooth

fracture; (b) rough fracture with the internal structure of

the fiber appeared drawn out; (c) split fracture; and (d) split

fibrillation fracture. Fibers derived from the 2.5 mm inter-

nal diameter tube appeared to deviate from the above on

the grounds of having air-bubbles trapped within their

structure and consequently when these fibers extended to

failure, the fracture tended to occur at the side of the bub-

ble (results not shown).

Thermal Stability

Table I summarises the thermal characteristics of the

PSBAT collagen produced fibers. The hydrothermal stabil-

ity of the fibers was independent of the collagen concentra-

tion (p [ 0.05) and the amount of PEG in the FFB.

Similarly, the thermal properties of extruded collagen fibers

derived from increasing concentrations of acid soluble rat

tail tendon and acid soluble bovine Achilles tendon were

found to be independent of the collagen solution concentra-

Figure 1. Typical morphology of extruded collagen fibers: (a) sur-

face structure; (b) packed interfibrillar space.

TABLE I. Denaturation Temperatures of Rehydrated PepsinSoluble Bovine Achilles Tendon Collagen Derived Fibers

Treatment Onset 6 SD (8C) Peak 6 SD (8C)

20% PEG

1.5 mm id

3 mg/mL 44.64 6 0.68 46.83 6 0.50

20% PEG

1.5 mm id

6 mg/mL 45.00 6 0.52 47.93 6 0.51

20% PEG

1.5 mm id

7 mg/mL 44.64 6 0.65 48.78 6 1.19

6 mg/mL

1.5 mm id

5% PEG 47.34 6 0.03 50.24 6 0.20

20% PEG 45.00 6 0.52 47.93 6 0.51

40% PEG 43.81 6 1.93 47.09 6 1.90

20% PEG

6 mg/mL

1.0 mm id 44.84 6 0.58 47.12 6 0.02

1.5 mm id 45.00 6 0.52 47.93 6 0.51

2.5 mm id 48.50 6 0.04 52.79 6 0.58

Sample number: 3; SD, standard deviation.

345TISSUE ENGINEERING APPLICATIONS


tion (p [ 0.05; Table II). Although no significant differ-

ence was observed in thermal stability of the fibers derived

from the 1.0 and 1.5 mm tube (p [ 0.05), the utilisation of

the 2.5 mm tube produced fibers of significant higher dena-

turation temperature (p \ 0.008 and 0.007 for the 1.00 and

1.5 mm tube, respectively).

Biomechanical Analysis

All dry collagen fibers tested herein exhibited a typical s-

shape stress-strain curve, characterised by a small toe

region, followed by a region of steeply rising stress up to

the knee point, where the gradient of the curve reduced,

and by a long region of constant gradient up to the point of

failure. The stress-strain curves demonstrated a diameter

dependent variation; fibers of low diameter showed small

toe regions followed by high stress/low strain graphs,

whilst fibers of high diameter displayed longer toe region

and gave rise to low stress/high strain graphs. To demon-

strate visually this variation on Figure 2 we show the

thickest, the thinnest and an intermediate in diameter fiber

derived from the PSBAT collagen 6 mg/mL; 1.5 mm

extrusion tube; and 20% PEG Mw 8 K as a representative

example.

Effect of PEG Mw 8K. In this set of experiments the

influence of the amount of PEG Mw 8 K in the FFB on the

properties of extruded collagen fibers was investigated. The

mechanical properties of the fibers produced are summar-

ised in Table III. At 20% co-agent content, the smallest in

dry diameter (p \ 0.001) fibers were obtained and they

exhibited the highest stress (p \ 0.002) and strain (p \0.002) at break values. At 40% PEG, the largest in diame-

ter fibers were obtained (p \ 0.001) with the lowest stress

at break (p \ 0.002) and modulus (p \ 0.006) values.

Fibers produced at 5% PEG required the highest forces to

break (p \ 0.002), and they were found to be the stiffest

(p\ 0.006) and the least extendable (p\ 0.002).

Fitting a linear regression model between stress at break

and dry fiber diameter for every treatment (results not

shown), the strongest correlation was obtained for the 20%

PEG content (R2 value of 0.89).

Effect of Collagen Concentration. Figure 3 illustrates

the nonlinear relationship between shear stress and shear

rate for the collagen solutions of different concentration.

Moreover, all collagen solutions were found to be non-

Newtonian shear thickening fluids. Significant difference (p\ 0.001) was observed in viscosity between the 3 and 6

mg/mL and 3 and 7 mg/mL collagen solutions, but not

between the 6 and 7 mg/mL collagen solutions (p[ 0.05).

Table IV summarises the mechanical properties of

extruded collagen fibers derived from collagen solutions of

3, 6, and 7 mg/mL concentration. By increasing the colla-

gen concentration, the fiber diameter (p \ 0.001), the force

at break (p\ 0.001) and the strain at break (p\ 0.02) val-

ues were increased, whilst the stress at break (p \ 0.001)

and the modulus (p\ 0.007) values were decreased.

Supplementary results, comparing the mechanical prop-

erties of extruded fibers (Table II) that produced from

increasing concentration of ASRTT and ASBAT, demon-

strated proportional findings. Collagen solutions of high

concentration yielded fibers of high diameter (p \ 0.001)

that were characterised by high force at break values (p \0.001) and low stress at break (p \ 0.05) and modulus at

TABLE II. Thermal, Physical, and Mechanical Properties of Extruded Collagen Fibers Derived From Acid Soluble Rat Tail Tendon(ASRTT) and Bovine Achilles Tendon (ASBAT) Collagen Solutions

Treatment

Peak 6 SD

(8C)Dry Diameter 6 SD

(lm)

Stress at

Break 6 SD

(MPa)

Strain at

Break 6 SD

Force at

Break 6SD (N)

Modulus at 2%

Strain 6 SD (GPa)

ASRTT

3 mg/mL

(n 5 8)

45.38 6 0.49 53 6 11 255 6 89 0.19 6 0.02 0.50 6 0.05 3.31 6 1.09

4 mg/mL

(n 5 7)

44.75 6 0.27 126 6 21 134 6 43 0.21 6 0.05 1.56 6 0.10 1.59 6 0.61

ASBAT

3 mg/mL

(n 5 6)

47.34 6 0.32 123 6 21 168 6 62 0.30 6 0.06 1.84 6 0.15 1.49 6 0.57

5 mg/mL

(n 5 7)

46.7 6 0.82 194 6 23 108 6 19 0.36 6 0.06 3.12 6 0.26 0.56 6 0.46

Sample number n in parentheses; SD, standard deviation.

Figure 2. The influence of dry fiber diameter on the stress-strain

graph. Thin fibers exhibit small toe regions and high stress/lowstrain graphs; thick fibers display long toe region and low stress/

high strain graphs; and the intermediate in thickness fibers reveal

stress–strain curves laying in between the above described graphs.



2% strain (p\ 0.05) values. In this case, no significant dif-

ference was observed for the strain at break values (p [0.05).



shown), the strongest correlation was obtained for the

6 mg/mL collagen concentration (R2 value of 0.89).

Effect of Extrusion Tube Internal Diameter. Table V

summarises the mechanical properties of the dry reconsti-

tuted collagen fibers derived from different internal diame-

ter tubing. It was found that the smallest internal diameter

tubing (1.0 mm) produced fibers with the smallest diameter

(p \ 0.001) and with the highest stress at break (p \0.001), whilst the largest internal diameter (2.5 mm) extru-

sion tube produced dry fibers with the highest diameter (p\ 0.001) and lowest stress at break (p \ 0.001) values.

Furthermore, it was found that by increasing the extrusion

tube diameter, a great decrease in modulus values and an

increase in force required to break the fibers was observed

(p\ 0.001). Fibers derived from the 1.5 mm tube were the

most extendable (p\ 0.001).



shown), the strongest correlation was obtained for the 1.5

mm tube internal diameter (R2 value of 0.89).

DISCUSSION

Structural Evaluation

The fibers exhibited a rough surface with undulations along

their length and ridges running roughly parallel to the

longitudinal axis. Similar observations have been attributed

to the handling28 and/or to the substructure of the fibers.31

The nano-texture surface of the fibers has been shown to

facilitate cell attachment and migration.8 A consistent filled

sub-fibrillar structure was observed for the fibers studied in

this work and was independent of the collagen concentra-

tion, PEG amount or the tube internal diameter. This result

is in agreement with previous observations, where it was

mentioned that very little free inter-fibrillar space occurs in

fibers formed in vitro.32 Furthermore, the inter-fibrillar

space has been reported to be filled when polymers were

utilised for the production of composite fibers29 and elas-

tin-based materials.33 The fracture mechanisms identified

are in agreement with previous work and have been attrib-

uted to the handling of the fibers.34,35

The utilisation of the 2.5 mm extrusion tube derived

fibers with air-bubbles trapped within their structure and

even increasing the extrusion rate (from 0.4 mL/min to 1.2

mL/min) did not improve the situation. It can be therefore

concluded that there is a limit for the extrusion tube inter-

nal diameter above which inconsistent in shape fibers are

produced.

Thermal Analysis

The fibers produced in this study exhibited denaturation

temperatures ranging from 47 to 538C, which were higher

than any other non-cross-linked collagenous matrix

reported in the literature. We would attribute the improved

thermal stability of the fibers to the closer packing of the

molecules in the fiber form, whilst in collagen gels, for

example, the molecules are more scattered.36,37 We would

further identify PEG as a structure enhancer, which conse-

quently lead to improved thermal properties. It has been

claimed before that the use of PEG may serve to promote

the formation of a liquid crystalline phase by removing

free water, which is essential for the formation of highly

ordered arrays of fibrils.38 Moreover, it has been shown

that polymers may increase the shrinkage temperature by

shifting the equilibrium between native and denaturated

forms of collagen towards a more compact native form by

steric exclusions.39

Neither the increase of collagen concentration nor the

increase of the amount of PEG in the FFB affected signifi-

cantly the thermal properties of the fibers, although it has

been shown that the thermal stability is dependent on both

TABLE III. Physical and Mechanical Properties of Extruded Collagen Fibers as a Factor of % PEG Mw 8K

Treatment

Dry Diameter

6 SD (lm)

Stress at

Break 6 SD

(MPa)

Strain at

Break 6 SD

Force at

Break 6 SD (N)

Modulus at 2%

Strain 6 SD

(GPa)

5% PEG (n 5 13) 132 6 17 139 6 33 0.22 6 0.04 1.84 6 0.20 1.2 6 0.5

20% PEG (n 5 18) 123 6 26 136 6 57 0.36 6 0.13 1.41 6 0.17 0.9 6 0.7

40% PEG (n 5 10) 216 6 73 57 6 40 0.30 6 0.13 1.44 6 0.33 0.4 6 0.5


Figure 3. Non-Newtonian shear thickening pepsin soluble bovine

Achilles tendon collagen solutions of (a) 6 mg/mL, (b) 7 mg/mL, and (c)

3 mg/mL plotted as shear stress (mPa 3 s) against shear rate (1/s).Sample number: 2.



amount of hydroxyproline40,41 and solute concentration.42,43

In both cases we would correlate the thermal properties of

the fibers to the collagen-solute interaction, which conse-

quently altered the fiber diameter. Thus, under the condi-

tions utilised herein, we obtained fibers that, relatively

speaking, were comprised of the same amount of collagen-

solute and thus they exhibited no significant different dena-

turation temperatures. However, the excessive increase in

tube internal diameter resulted in fibers with diameter of

almost 6- and 4-fold difference between the 2.5 mm tube

and the 1.0 mm and the 1.5 mm tube, respectively, and

altered that way the balance collagen-solute and thus higher

denaturation temperatures were obtained.

Stress–Strain Curves

Uniaxial tensile tests of dry reconstituted collagen fibers,

produced stress–strain curves similar to those reported for

semi-crystalline polymers which yield and undergo plastic

flow.44 Similar curves have been reported previously for

dry native rat tail tendon fibers45 and ligaments,29 reconsti-

tuted collagen fibers,12,24,46 even synthetic nano-fibrous

meshes of poly(L-lactic acid)-co-poly(e-caprolactone)47 or

collagen-poly(L-lactide-co-e-caprolactone).48 The yielding

mechanism would imply some form of flow that takes

place within the fiber, possibly inter-fibrillar slippage. Since

collagen molecular and fibrillar slippage plays an important

role in the tensile deformation of aligned connective tissue

such as tendon, it is more likely that PEG plays a role in

promoting either molecular or fibrillar slippage by

‘‘reptating’’ parallel to the collagen molecules backbone as

has been suggested previously for PEG and collagen matri-

ces,49 PEG and clay films,50 even hydroxyapatite nano-

crystals and collagen.51 Moreover, it has been suggested

that the co-secretion of glycosaminoglycans with collagen

in living systems may be analogous to the use of PEG in invitro systems.24,38 We would attribute the longer toe region

of the thicker fibers to the fibrillar packing that is tighter

for thin fibers (see below). The lower modulus of the non-

linear toe region in native tissues such as tendons or liga-

ments has been attributed to the un-crimping of the

collagen fibrils, as well as the initiation of stretching of the

triple helix, the nonhelical ends and the cross-links.52

Effect of PEG Mw 8K

Although PEG has been employed previously for the fabri-

cation of extruded collagen fibers,12,22 its influence on the

properties of these fibers has not been explained

adequately. It is of importance to mention that at the total

absence of PEG, although fibers were produced, they were

too brittle to handle. Thus the utilisation of polyethylene

glycol or other solutes is essential in order to produce col-

lagen fibers of sufficient mechanical strength, as has been

shown previously for collagen gels38 and collagen micro-

fibrillar networks.53

The fiber diameter was decreased as the amount of PEG

was increased from 5 to 20%, whilst the diameter was

increased as the amount of PEG was increased from 20 to

40% in the FFB. During precipitation experiments it has

been shown that the mean particle size of the precipitates

was generally reduced as the ionic strength was increased

due to the high level of counter-ions around the protein in

the solution.54 In low ions concentration, due to the salting

in effect, the protein molecules are in the right configura-

tion to build up a rigid and more viscoelastic material.55 At

higher ions concentrations, due to passing the optimal

amount of water removable for fiber stability, this right

configuration might be lost56 and the particle size could be

remarkably increased,54 leading to a decrease in the effec-

tiveness of the precipitation.20 These observations are in

accord with other work, where it was mentioned that flocs

formed under polymer overdosing conditions are larger

than flocs formed under polymer under-dosing conditions.57

TABLE IV. Physical and Mechanical Properties of Extruded Collagen Fibers as a Factor of Pepsin Soluble BovineAchilles Tendon Collagen Concentration

Treatment

Dry Diameter 6SD (lm)

Stress at Break

6 SD (MPa)

Strain at Break

6 SD

Force at Break

6 SD (N)

Modulus at 2%

Strain 6 SD (GPa)

3 mg/mL (n 5 5) 52 6 8 169 6 53 0.27 6 0.04 0.35 6 0.04 1.66 6 0.61

6 mg/mL (n 5 18) 123 6 26 136 6 57 0.36 6 0.13 1.41 6 0.17 0.86 6 0.75

7 mg/mL (n 5 14) 190 6 10 55 6 5 0.42 6 0.10 2.88 6 0.28 0.32 6 0.20


TABLE V. Physical and Mechanical Properties of Extruded Collagen Fibers as a Factor of Extrusion Tube Internal Diameter

Treatment

Dry Diameter 6SD (lm)

Stress at Break 6SD (MPa)

Strain at

Break 6 SD

Force at

Break 6 SD (N)

Modulus at 2%

Strain 6 SD

(GPa)

1 mm (n 5 21) 84 6 13 200 6 58 0.24 6 0.04 1.03 6 0.12 1.8 6 0.5

1.5 mm (n 5 18) 123 6 26 136 6 57 0.36 6 0.13 1.41 6 0.17 0.9 6 0.7

2.5 mm (n 5 5) 459 6 43 13 6 3 0.34 6 0.07 2.16 6 0.55 0.02 6 0.01




Bearing in mind that (a) PEG20,21,58 can precipitate colla-

gen; (b) the in vitro self-assembly of collagen molecules

into fibrils with periodic patterns59 is analogous to the pre-

cipitation of collagen56; (c) precipitation (bulky result

occurring at low temperatures) and self-assembly (struc-

tural, highly organised result occurring at around 378C) areprocesses driven by forces, which are mainly electrostatic

attraction plus hydrogen and hydrophobic bonding57; and

(d) taking as particle size the diameter of the fiber pro-

duced, the same results obtained from these experiments;

the fiber diameter was decreased as the amount of PEG

was increased from 5 to 20% (under-dosing conditions),

and the fiber diameter was increased as the amount of PEG

was increased from 20 to 40% (overdose conditions) in the

FFB. Furthermore, since at 20% PEG the highest correla-

tion between stress at break and fiber diameter was

obtained, we tend to believe that the 20% PEG in the FFB

comprises the optimum amount of PEG required for the

production of reproducible extruded collagen fibers.

At 5% PEG the least extendable fibers were produced.

We would speculate that at low PEG content the elongation

derives only from the collagen, whilst at higher amounts of

PEG, both collagen and PEG contribute to the elongation;

thus increased strain at break values were obtained. Simi-

larly, it has been shown before for composite fibers that

their elongation derives from the contribution of the poly-

mer and the filler used.60 The modulus values appeared to

decrease, as the amount of PEG in the FFB was increased.

This observation appears to be in agreement with previous

work, where it was shown that by increasing the concentra-

tion of electrolytes, a decrease in tensile modulus was

observed.61,62 Moreover, the observed decrease in tensile

strength of the fibers, especially in the higher concentration

of PEG, is in agreement with previous results, where it was

shown that increases in dermatan sulphate concentration

has an inhibitory effect on the mechanical strength of

reconstituted collagen fibrils.63

Effect of Collagen Concentration

The collagen concentration was found to be an important

variable for the fabrication process; by increasing it, the

handling of the fibers while in the wet state was improved.

When as low as 1 mg/mL collagen solutions were

employed, although fibers were produced, they were too

fragile to handle. Similarly, it has been reported that there

is a critical concentration for the fibrillogenesis in vitro to

occur64,65 and for collagen sponges a minimum effective

concentration has been identified.66,67 Even during the pro-

duction of nano-fibers using electro-spinning, concentration

has been identified as an important variable for the produc-

tion of uniform and reproducible scaffolds.68,69

Fibers derived from the high in concentration collagen

solutions were of significant higher diameter than those

derived from the low concentration solutions. Likewise, it

has been shown that the fiber diameter depends on the

polymer utilized.22,68,69 Such differences in diameter may

be attributed to differences in the solution viscosity, which,

as has been shown in the literature, can influence the diam-

eter of extruded fibers.70,71 By increasing the fiber diame-

ter, consequent decreases in stress at break values were

become apparent. Similarly, it has been shown that the ten-

sile strength of rat skin depends on the total hydroxyproline

content.72 Moreover, it has been shown that collagen con-

tent not only can influence the mechanical properties of

collagen fibers,73 but also it has a negative correlation on

the tensile strength of rat skin.74 Even the tensile strength

of co-electro-spun collagen-poly(L-lactide-co-e-caprolac-tone) has been shown to decrease with increasing collagen

content.48 Also the strength of gelatin gels has

been shown to depend on the gelatin concentration.75 The

high correlations obtained between stress at break and

dry fiber diameter can been attributed to (a) the tensile

strength increases as cross-sectional area decreases because

there is less chance for defects in thinner sections76,77;

or (b) as the fiber diameter is decreased, an increased lon-

gitudinal alignment and improved packing density occurs

leading to strong interactions within or between collagen

fibrils.9,11,24,46 These strong interactions are illustrated in

the stress-strain curves with the short toe region of the thin

fibers, whilst the loose interactions are represented by the

longer toe regions in the thick fibers. Moreover, the strong

interactions would have an inhibitory effect on the extensi-

bility of the fibers, thus fibers of low diameter would ex-

hibit reduced strain at break values. Thin fibers comprised

of aligned fibrils exhibited lower forces at break, high ten-

sile strength and modulus values. Modulus values reflect

the stiffness or rigidity of the material; the higher its value,

the greater the load would be required to produce a given

extension.78,79 These results are in agreement with previous

work on collagen fibers,35 composite fibers60 and synthetic

sutures.80 Overall, from the strong correlations between the

collagen concentration and the dry fiber diameter and the

dry fiber diameter and the stress at break, it can be con-

cluded that by controlling the collagen concentration, tai-

lored made collagen fibers can be produced to suit a wide

range of tissue engineering applications.

Effect of Different Extrusion Tubing

The utilisation of different internal diameter tubing

appeared to influence all the measured mechanical proper-

ties of the dried reconstituted collagen fibers produced. It

can be observed that by increasing the fiber diameter, the

stiffness and the stress at break values of the fibers were

decreased, as was observed when the collagen concentra-

tion was increased. This result further enhances previous

observations, where it was shown that lower diameter fibers

exhibit a more packed structure and higher fibril align-

ment.9,11,24,46 The lowest diameter fibers, however, failed

at lower loads than the other two groups, which is in

accord with previous observations.11 Finally, the utilisation



of the 1 mm and 1.5 mm internal diameter tubing produced

fibers of 0.084 mm and 0.123 mm diameter, respectively.

These values are approximately one tenth smaller than

the tube utilised and demonstrates that the dry fiber diam-

eter can be controlled by varying the inner diameter of

the extrusion tube as has been shown before.11,46 However

herein we demonstrate that there is a critical tube diame-

ter (somewhere between 1.5 and 2.5 mm) over which

the extrusion becomes disrupted and the fiber becomes

ultimately nonuniform, at least under the conditions used

here.

CONCLUSIONS

Natural occurring collagen preferred for scaffold manufac-

turing, since such biomaterials closely imitate the extracel-

lular matrix. Herein, we manufactured collagen-PEG nano-

textured micro-fibers from reduced immunogenicity atelo-

collagen that were characterised by thermal, physical, and

mechanical properties similar to those of native tissues or

synthetic fibers. Moreover, we identified the 20% PEG Mw

8K amount as the optimum solute content for fabricating

reproducible fibers. By controlling the collagen concentra-

tion and/or the extrusion tube internal diameter, tailored-

made materials were produced to suit a wide range of tis-

sue engineering applications.

The authors thank Mrs. P. Potter, Ms. S. Lee, Mrs. T. Hayes,and Mr. L. Stathopoulos for excellent technical assistance andDr. S. Jeyapalina and Dr. P. Antunes for their useful discussion.

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extruded collagen-polyethylene glycol fibers for tissue engineering applications

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