In vitro biodegradation of designed tubular scaffolds of electrospun protein/polyglyconate blend...

In Vitro Biodegradation of Designed Tubular Scaffoldsof Electrospun Protein/Polyglyconate Blend Fibers

Xing Zhang,1y Vinoy Thomas,2y Yogesh K. Vohra2

1 Department of Biomedical Engineering, University of Alabama at Birmingham (UAB), Birmingham, Alabama 35294

2 Center for Nanoscale Materials and Bio-integration (CNMB), Department of Physics, University of Alabamaat Birmingham (UAB), Birmingham, Alabama 35294

Received 16 January 2008; revised 30 April 2008; accepted 3 June 2008Published online 8 September 2008 in Wiley InterScience (www.interscience.wiley.com). DOI: 10.1002/jbm.b.31196

Abstract: Electrospun polyglyconate (Maxon1) and its blends with proteins such as gelatin

and elastin, with a spatially designed layer structure, were prepared as potential scaffolds for

vascular tissue engineering. In vitro biodegradation of the electrospun tubular protein/Maxon

scaffolds in phosphate buffered saline (pH 5 7.3) was studied for the first time. The

biodegradation is manifested by uptake of the PBS medium by the hydrophilic proteins and

also by the mass loss due to the removal of degraded fragments and uncrosslinked proteins

from the matrices. The effect of degradation on the structure–property relations was evaluated

by IR, XRD, and DSC analyses of the aged scaffolds. The degradation of amorphous phase of

Maxon in the early stages of aging has resulted in an increase in the crystallinity of the

polymer. SEM analysis indicated a significant change in nanofiber morphology and fiber-

breaking. The mass loss and fiber breaking have negatively impacted the mechanical

properties and the effect was maximum at 15–20 days of aging. The scaffold containing low

molecular weight buffer soluble elastin revealed relatively better degradation properties

compared to that containing high molecular weight buffer insoluble elastin. ' 2008 Wiley

Periodicals, Inc. J Biomed Mater Res Part B: Appl Biomater 89B: 135–147, 2009

Keywords: biodegradation; electrospun scaffolds; nanofibers; vascular grafts; tissue

engineering

INTRODUCTION

In the past two decades, several research groups have

attempted to engineer arterial substitutes with biological

and mechanical properties as close as possible to native

blood vessels.1–5 Owing to the formation of ultrafine poly-

meric fibers closely resembling the nanofeatures of native

extracellular matrices,6–10 electrospinning technique has

gained much attention as a fabrication-method for scaffolds

in tissue engineering. Many biodegradable synthetic poly-

mers such as polyglycolide (PGA), poly(caprolactone)

(PCL), poly(lactide-co-glycolide) (PLGA), and poly(L-lactic

acid) (PLA) have been electrospun into fibrous meshes.6–9

Moreover, nanofibrous tubular grafts based on native pro-

teins and/or synthetic polymers for vascular tissue regener-

ation can be easily constructed by electrospinning.11–14 For

biomedical applications such as drug delivery and tissue

engineering, a suitable degradation rate is one of the impor-

tant requirements for the matrix-material used. For exam-

ple, biodegradation of polycaprolactone (PCL) is quite slow

in the form of films but, it is very fast in the form of nano

or microparticles.15–17 Because electrospun nanofibers pos-

sess very high surface-to-volume ratio than films or micro-

particles, a high rate of biodegradation is expected for

electrospun meshes. However, the chemical nature and

composition, hydrophilicity or hydrophobicity, and crystal-

linity of the fibers are also key factors for biodegradation.

Zone et al.18 have reported the degradation behavior of

electrospun poly(glycolide-co-lactide) ultrafine fibers in

phosphate buffered saline (PBS) and explained their struc-

tural and morphological changes because of in vitro hydro-

lytic degradation. You et al.19,20 have reported the

degradation behavior of electrospun PGA, PLA, and PLGA

by soaking them in PBS (pH 5 7.3) at 378C for various

periods. The degradation results demonstrated that PGA

fibers possessed the highest degradation rate with only 40%

residual mass left after 20 days, whereas the PLA fibers

remained relatively without mass change throughout the pe-

riod. Again, the initial rate of degradation of PLGA fibers

was low but became fast after 25 days of aging. The elec-

trospun fibers of PLA blended with PGA exhibited a

yBoth authors contributed equally to this work.Correspondence to: V. Thomas (e-mail: [email protected]) or Y. K. Vohra

([email protected])Contract grant sponsor: National Science Foundation (NSF-NIRT Program);

Contract grant number: DMR-0402891

' 2008 Wiley Periodicals, Inc.

135

decreased rate of degradation with the increase of PLA

content.20

Because tubular geometry is one of the easiest 3D

structures that can be fabricated by electrospinning, many

researchers have reported the fabrication of electrospun

natural proteins and/or synthetic polymers by depositing

the fibers onto a rotating mandrel to produce tubular

grafts.11–14,21,22 Matthews et al.23 and Boland et al.12 have

electrospun type I collagen and collagen/elastin nanofibers,

respectively, for vascular applications. Tubular scaffolds (6

mm diameter) of electrospun polydioxanone/elastin blends

having different blending ratios were prepared by Sell

et al.13 to match tensile properties to those of femoral artery.

Polydioxanone was chosen to provide the mechanical integ-

rity to the prosthetic, whereas elastin provides elasticity and

bioactivity. Stitzel et al.14 have used a rotating rod (500 rpm)

to collect the electrospin PLGA-type I collagen-elastin blend

fibers in the form of small diameter tubular scaffold (4.75

mm) for vascular graft. Incorporation of collagen in synthetic

polymers such as poly(ester urethane) urea (PEUU) has sig-

nificantly enhanced the smooth muscle cell adhesion onto

electrospun PEUU/collagen scaffolds.22 In addition to blend-

ing, multilayering of electrospun bilayered PLA (outer)/PCL

(inner) based tubular scaffolds has also been reported to

design artificial grafts.24 In this context, Thomas et al.25

have successfully constructed a small diameter (4 mm inner

diameter) tubular graft with a spatially designed layer struc-

ture via sequential electrospinning to mimic the structure of

native blood vessels. The scaffold possessed a trilayer struc-

ture: gelatin and elastin blend fibers as lumen layer; gelatin,

elastin, and a polyglyconate (Maxon) ternary-blend fibers as

middle layer; and gelatin and Maxon blend fibers as outer

layer.25 Maxon is a copolymer of 67.5 wt % glycolide and

32.5 wt % trimethylene carbonate (TMC) and is clinically

used as an absorbable surgical suture. The preliminary

mechano-morphological studies of the trilayer graft have

shown promising mechanical and structural properties invitro.25 The scaffolds (hydrated) under a cyclic pressure

range (80–120 mmHg) exhibited a mean compliance value

of 2% and a burst pressure of 1700 mmHg, which is compa-

rable to the burst pressure of saphenous vein (1680 6 307

mmHg).26 Although the degradation behavior of pure

Maxon sutures in macrofiber form under in vitro27–29 and

in vivo30,31 conditions has been extensively studied earlier,

the biodegradation behavior of electrospun nano/microfi-

brous Maxon or protein-blended-Maxon fibrous meshes has

not been reported yet. This article reports the biodegradation

of the sequentially spun trilayered protein/Maxon scaffolds

in PBS medium and the effect of in vitro degradation on the

structural, morphological, and mechanical properties.

MATERIALS AND METHODS

Materials

Poly(glycolide-co-trimethylene carbonate) was purchased in

the form of surgical sutures under the trade name Maxon1

from Advanced Inventory Management, Mokena, IL. The

cosmetic green dye distributed throughout the suture was

removed by soaking in methylene dichloride for 1 day and

then dried under vacuum. Porcine gelatin A (300 bloom

number) and 1,1,1,3,3,3-hexafluoro-2-propanol (HFP) were

purchased from Sigma-Aldrich, St Louis, MO. Two types

of elastins, sourced from bovine neck ligament, (ES60 and

ES12 from Elastin Products, Owensville, MO) were used to

make the multilayered vascular grafts.

Fabrication of Tubular Scaffolds

Two types of electrospun trilayer tubular scaffolds of 4

mm inner diameter were fabricated such that one of which

contained ‘‘buffer insoluble elastin’’ and the other contained

‘‘buffer soluble elastin.’’ HFP was used as the common sol-

vent for both proteins and Maxon. A multilayering sequen-

tial electrospinning process was employed to fabricate

protein/Maxon tubular scaffolds having a total composition

of 58 wt % Maxon, 28 wt % gelatin, and 14 wt % elastin.

The sequential spinning order was as follows: first gelatin/

elastin (4:1) as the lumen layer, then gelatin/elastin/Maxon

(1:2:8) as the middle layer, and then gelatin/Maxon (1:4) as

the outer layer. The solution for each layer was loaded into

a syringe with a 27-gauge needle and a high voltage of

30 kV (M826, Gamma High-Voltage Research, Ormond

Beach, FL) was applied. When the electric charge on poly-

mer solution at the needle tip overcomes the force due to

surface tension, a whipping polymer jet is ejected from the

needle. The jet within the electric field is directed toward

the grounded rotating collector, during which time the sol-

vent evaporates and randomly oriented fibers are deposited

at the collector. A grounded 303 stainless steel mandrel (4

mm diameter) rotating at 400 rpm was set at 30 cm apart

from the needle tip for collecting the fibers. A feeding rate

of 1–3 mL/h was used for spinning. The wall thickness of

the graft was controlled by the volume of spinning solu-

tion. Typically, the spinning of 1 mL of gelatin-elastin so-

lution (inner layer), 2 mL of gelatin-elastin-Maxon solution

(middle layer), and 1 mL of gelatin-Maxon solution (outer

layer) yielded a thickness of �0.5 mm for a 20 cm long

graft. To ensure the uniform distribution of the fibers

throughout the scaffold length, the collector was traversed

at a speed of 15 mm/min.24,25 The scaffolds were recovered

from the mandrel by pushing gently out without affecting

the tubular structure and fiber morphology. Optical and

scanning electron microscopic techniques were used to

visualize the tubular structure and fiber morphology.

In Vitro Biodegradation Studies

Electrospun tubular grafts were cut into specimens with the

length of 25 mm and grouped into a number of sets (n 53) in separate vials. The specimens were soaked in 15 mL

phosphate buffered saline (PBS) (pH 5 7.3) in closed vials

and incubated at 378C for up to 30 days. At the end of

each degradation period, three specimens were withdrawn

136 ZHANG, THOMAS, AND VOHRA

Journal of Biomedical Materials Research Part B: Applied Biomaterials

from the aging medium and blotted with Kim wipes. The

weight of specimen before aging (Mb) and immediately af-

ter aging under wet condition after blotting with Kim wipes

(Ma) were used to calculate the PBS uptake (%).

PBS uptake %ð Þ ¼ Ma �Mb

Mb

� �3 100 ð1Þ

The specimens were dried under vacuum at room tem-

perature until a constant weight was obtained and the

weight after complete dehydration (Md) was noted to calcu-

late the mass loss (%).

Mass loss %ð Þ ¼ Mb �Md

Mb

� �3 100 ð2Þ

The aged and dried specimens were characterized for

structural, morphological, and mechanical properties to

evaluate the biodegradation behavior of the scaffolds.

Morphological and Structural Characterization

The scanning electron micrographs (SEM) were captured

using an SEM instrument (Philips SEM 510) at an acceler-

ating voltage of 30 kV. The aged specimens were sputter

coated with gold to make them charge-conductive.

The Fourier-Transform infrared (FTIR) spectra were

recorded with 32 scans per sample ranging from 4000 to

400 cm21 in the attenuated total reflection (ATR) mode

using an IR spectrophotometer (Thermo Fisher Co.).

X-ray diffraction (XRD) scans were performed by

mounting specimens (1 cm 3 1 cm) onto the X-ray diffrac-

tometer (Philips Xpert MPD) using Cu-Ka radiation (k 50.1541nm). The XRD curves of all specimens were gener-

ated using a current of 40 mA, an accelerating voltage of

45kV, and a scanning speed of 0.018/s from 2y 5 108–508.The curves were analyzed by Philips software (ProFit

V1.0).

Differential scanning calorimetric (DSC) experiments

were carried out to determine the change in crystallinity of

polymer because of degradation using a DSC instrument

(DSC Q 100, TA Instruments) in nitrogen atmosphere. The

aged scaffold samples (�10 mg) were ramped at a rate of

108C/min in the temperature range from 2258C to 2308C.

Mechanical Properties Evaluation

The tubular specimens were cut open and trimmed into rec-

tangular strips (20 mm 3 5 mm). Uniaxial tensile test was

performed using a universal tensile testing machine (INS-

TRON 5565, MA). A 100 N load cell was employed with

a loading rate of 0.1 N/min. The data were recorded using

a companion software (Bluehill 2) and the stress–strain

curves were generated for each specimen to determine the

modulus of elasticity from the slope of the initial linear

segment within 10% strain.

Statistical Analysis

The statistical analysis was performed based on ANOVA

and Tukey-Kramer tests (a 5 0.05) to determine difference

between any two groups in each pool of PBS uptake (%),

weight loss (%), and mechanical properties (JMP17, 2007

SAS Institute).

RESULTS

Two types of trilayered tubular protein/Maxon scaffolds

were prepared by multilayering sequential electrospinning

process. The scaffold containing the ‘‘buffer insoluble

elastin’’ (E) was coded as GE-GEM-GM and the other

containing the ‘‘buffer soluble elastin’’ (e) was coded as

Ge-GeM-GM. According to the manufacturer, buffer insol-

uble elastin was prepared by a neutral extraction method

developed by Partridge et al.32 and the buffer soluble elas-

tin was prepared by treating bovine neck ligament elastin

in hot oxalic acid solution followed by dialysis and lyophi-

lization.32 The processing methods can affect the physico-

chemical properties. From the thermogravimetric analysis,

the degradation onset temperature was found to be �2808Cfor insoluble elastin and �2008C for soluble elastin (data

not shown). Because there is dependence between molecu-

lar weight and degradation temperature as reported by

Heitz et al.,33 soluble elastin has relatively a low molecular

weight compared to insoluble elastin. Hence, the low mo-

lecular weight elastin is easily soluble in aqueous buffer

and the high molecular weight elastin is not. However, it

dissolves in polar organic solvent such as HFP and forms a

suspension-like solution at smaller concentrations and

therefore, it is spinnable with other polymers. At a low col-

lector rotation speed (400 rpm), the co-electrospun protein/

Maxon meshes have a randomly oriented fibrous struc-

ture13,25,34 with fibers in nano/micro diameters. The present

tubular scaffolds have a spatially designed trilayer structure

to mimic the trilayer of native blood vessels.25 The inner

lumen layer is designed purely of fibers of hydrogel pro-

teins (gelatin and elastin); the middle layer is made up of

blended fibers of gelatin, elastin, and polyglyconate

(Maxon); and the outer layer is made up of gelatin and

Maxon, that is, from inner to outer layers, the protein con-

tent in the chemical composition of fibers decreases. The

protein rich inner layer is designed to support the luminal

endothelial cell growth, whereas the polyglyconate rich

outer layers are to provide the mechanical and structural in-

tegrity to the grafts. The scaffolds have high porosity with

interconnected pores for the nutrient transport and high sur-

face area for the increased cell adhesion. Porosity calcula-

tions25,35 have shown that both scaffolds have relatively

high inherent porosities (82–84%).

Figure 1 demonstrates the PBS uptake by the scaffolds,

GE-GEM-GM and Ge-GeM-GM, during aging up to 30

days. The high porosity with interconnected pores of the

scaffolds and the increased hydrophilicity due to the pres-

137BIODEGRADATION OF PROTEIN/POLYGLYCONATE BLEND FIBERS


ence of proteins explain the higher values of PBS uptake

(%) by protein/Maxon blends compared to pure Maxon.25

PBS uptake (%) reached a maximum of 345% for GE-

GEM-GM graft and 260% for Ge-GeM-GM after 5 days

soaking. Afterwards, a decrease in PBS uptake percentage

was observed in the case of both grafts due to mass loss.

This is because of the fact that at some points the mass

loss due to the removal of fragmented products exceeds the

mass gain due to PBS uptake. The PBS uptake by GE-

GEM-GM graft remained more or less at the same level

(260%) up to 30 days whereas Ge-GeM-GM exhibited a

reduction to 208% after 20 days. This can be explained by

the increased mass loss from the sample due to degradation

after 20 days as seen in Figure 2.

Figure 1. PBS uptake (%) of the protein/polyglyconate tubular scaf-folds as a function of aging time. Figure 2. Mass loss (%) of the protein/polyglyconate tubular scaf-

folds as a function of aging time.

Figure 3. SEM images showing the effect of degradation on the fiber morphology of outer layer ofGE-GEM-GM graft at various aging period; (a) 5 days, (b) 10 days, (c) 20 days, and (d) 30 days.

Arrow points out cracks and fiber breaking.



The biodegradation is manifested by the materials loss

from both grafts because of aging in PBS medium (Figure

2). The GE-GEM-GM graft exhibited an increase in mass

loss (%) with time up to 15 days and then, attained a pla-

teau after 15 days. The maximum mass loss (%) was found

to be a mean value of 36.25%. However, Ge-GeM-GM

exhibited a gradual increase in mass loss up to 34.59%

with 20 days and finally escalated to 52% by 30 days. The

increased mass loss has a negative effect on PBS uptake

percentage as mentioned earlier. The significant decrease in

PBS uptake % at 10 days in the case of GE-GEM-GM

scaffold is due to the increased removal of degraded frag-

ments from scaffold to the medium. In the case of Ge-

GeM-GM scaffolds, this effect was found to be pronounced

only after 20 days.

Figures 3–5 illustrate the effect of aging in hydrolytic

medium on the fibrous mophology of the scaffolds. SEM

images in Figure 3(a–d) show the degradation of GM layer,

the outer layer for both GE-GEM-GM and Ge-GeM-GM

scaffolds. The individual fibers started to swell and fuse

together after 5 days of aging and to breakdown after 10

days. The biodegradation caused a significant number of

fiber-breaking by 20 days [Figure 3(c)] and the fibrous

morphology was destroyed significantly by 30 days of

aging [Figure 3(d)]. SEM images in Figure 4(a–d) were

captured on the lumen surface of GE-GEM-GM scaffold. It

clearly reveals that the inner GE layer of the scaffold has

lost its fibrous morphology by 5 days [Figure 4(a)] and was

fully dissolved or delaminated by 10 days of aging. How-

ever, the middle layer (GEM) has retained the fibrous mor-

phology throughout the aging period but, exhibited

macroscopic distortions as seen in Figure 4(d). When com-

paring image (c) against (a) and (b) in Figure 5, it is

obvious that the fiber morphology in (c) is almost identical

to that in (b). This means that the inner layer (Ge) of the

Ge-GeM-GM scaffold might have completely dissolved or

removed during the initial 5 days and exposing the middle

GeM layer. Although, the middle layer (GeM) of Ge-GeM-

GM graft has maintained its fibrous morphology without

much surface distortions, individual breaking and fiber

fusion can be seen in Figure 5(d,e) due to aging in PBS.

The effect of PBS aging of the scaffolds on the structural

and chemical characteristics was analyzed by IR, XRD, and

DSC. ATR-FTIR spectra of electrospun Maxon and Maxon-

protein scaffolds before and after aging are given in Figures

6–8. A typical spectrum of virgin electrospun Maxon shows

Figure 4. SEM images of lumen surface of GE-GEM-GM graft after aging for (a) 5 days, (b) 10

days, (c) 20 days, and (d) 30 days. Arrow points out cracks or individual fiber breaking.



characteristic peaks for both crystalline and amorphous

regions of PGA blocks [Figure 6(a)]. The bands at 846 and

716 cm21 can be attributed to the amorphous phase and the

bands at 973, 902, 787, 627, and 592 cm21 to the crystalline

phase as reported elsewhere.36 Also peaks at 1143 and 1084

cm21 can be assigned respectively to the C��O of ester and

oxymethylene groups of amorphous domains. Again, the

peak at 1084 cm21 has resulted from the overlapping of the

absorption bands of C��O of TMC segment at 1093 cm21 37

and C��O of PGA at 1077 cm21 to show an absorption at

1084 cm21, in between 1093 and 1077 cm21. Similarly, the

peak at 1185 cm21 in Maxon is due to the C��C��O asym-

metric stretching of ester in PGA units [Figure 6(a)]. The

protein blended scaffolds exhibited prominent peaks at 1643

cm21 for amide I, at 1536 cm21 for amide II, and around

1200 cm21 for amide III group, in addition to peaks for

Figure 5. Representative SEM images showing the effect of degradation on the fiber morphology ofinner layer of Ge-GeM-GM grafts at various aging period; (a) lumen layer before aging, (b) GeM middle

layer before aging, (c) 5 days, (d) 15 days, (e) 30 days. Arrows point out cracks and fiber breaking.



Maxon as shown in Figure 6(b,c). The amide I band for

C¼¼O stretching with characteristic frequencies in the range

from 1600 to 1700 cm21 is mainly associated with the

stretching vibrations on carbonyl groups along the polypep-

tide backbone and is a sensitive marker of polypeptide sec-

ondary structure.38,35 There are C��N stretching vibrations

and N��H deformations as minor vibration modes of amide

II and amide III bands as seen in Figure 6(c).

The aliphatic ester carbonyl peak (usually at 1743

cm21) and carbonate carbonyl peak (usually at 1725 cm21)

were appeared as a single band at 1738 cm21 in Maxon

and its protein blended scaffolds [Figure 6(a)]. The charac-

teristic bands for amorphous regions of PGA segments in

Maxon at 1143, 846, and 716 cm21 were disappeared com-

pletely in scaffolds hydrolyzed for 30 days [Figure 6(d)].

However, the peaks for the crystalline regions found to be

intact up to 20 days and decreased the intensity by 30

days. The ester carbonyl peak at 1738 cm21 exhibited a

significant reduction in intensity and appeared as a multip-

let by 30 days of aging [Figures 6(d) and 7(d)]. It should

be noted that the protein peaks for amide I (1643 cm21)

and amide II (1536 cm21) were also found to disappear or

exhibit reduction intensities drastically in samples aged for

30 days. After 20 days of aging, the peak at 1084 cm21

was shifted slightly to 1093 cm21 (C��O of TMC segment

appears at 1093 cm21),37 indicating the degradation of

Figure 6. ATR-FTIR spectra of (a) electrospun polyglyconate, (b) GE-GEM-GM outer layer, (c) GE-GEM-GM lumen layer, (d) GE-GEM-GM outer layer after 30 days aging. Peaks at 1143, 846, and

716 cm21 are represented by asterisks.

Figure 7. ATR-FTIR spectra of aged GE-GEM-GM scaffold; (a) GE-GEM-GM lumen layer 5 days,(b) 10 days, (c) 20 days, and (d) 30 days.



PGA segments in Maxon [Figures 6(a,d) and 8(d)]. The

spectra for the inner layer of scaffolds showed a drastic

decrease in intensity of ester carbonyl peak and peaks for

��COO and ��CH groups at 1416 cm21 in addition to the

disappearance of amorphous PGA peaks (1143, 846, and

716 cm21) (Figure 7). The peak for crystalline regions

appeared with a decreased intensity by 30 days of aging

suggesting the degradation of defective crystalline PGA

blocks as stated earlier. The degradation of ester groups

produce ��COOH ends groups and accelerates the ester hy-

drolysis. It is noteworthy that the amide I and II peaks

decreased intensity significantly by 10 days. This is either

due to the dissolution/disintegration of uncrosslinked pro-

teins or due to the peel-off of the protein based lumen

layer. This inference is supported by the observation of

maximum mass loss by 10 days. The IR spectra for the

outer layer GM layer (common for both scaffolds) of aged

scaffolds are given in Figure 8. The ester carbonyl exhib-

ited no significant reduction in intensity for carbonyl peak

up to 20 days; however, the peak intensity for ester car-

bonyl and crystalline PGA segments experienced a drastic

decrease by 30 days of aging. The amide I and II peaks

also showed a drastic decrease in intensity. The characteris-

tic peaks for amorphous PGA segments at 1143, 846, and

716 cm21 were disappeared by 20 days of aging (Figure

8). The gradual shifting of the C��O peak at 1084 cm21

towards higher wave number region (1093 cm21 for C��O

of TMC) can be clearly seen in Figure 8.

The XRD spectra of GE-GEM-GM and Ge-GeM-GM

scaffolds are given in Figures 9 and 10 respectively. The

Figure 8. ATR-FTIR spectra of aged Ge-GeM-GM scaffold; (a) Ge-GeM-GM outer layer 5 days,

(b) 10 days, (c) 15 days, (d) 20 days, and (e) 30 days aging.

Figure 9. XRD patterns of aged GE-GEM-GM scaffold; (a) virginscaffold; (b) in 5 days; (c) in 15 days; (d) in 30 days.

Figure 10. XRD patterns of aged Ge-GeM-GM scaffold; (a) virginscaffold; (b) in 10 days; (c) in 15 days; (d) in 30 days.



(110) and (020) planes of crystalline PGA blocks in the

PGA-TMC copolymer molecular chain are confirmed by

the reflection peaks respectively at 228 and 28.58.25 Aged

specimens showed relatively narrow peaks compared to vir-

gin samples (before aging). By examining the peak at 228,a trend of initial decrease in intensity by 5–10 days and

then a trend of increase in intensity can be observed due to

the removal of proteins and amorphous phase of Maxon.

The molecular chain reorganization and plasticizing effect

of medium may be responsible for the initial decrease in

crystallinity. The turning point of the trend for the GE-

GEM-GM scaffold was 10 days whereas that for Ge-GeM-

GM scaffolds was 15 days. The peak at 228 showed maxi-

mum intensity in the spectrum of specimens aged for 15

days for both grafts.

DSC scans (first heating) of the aged electrospun scaf-

folds are shown in Figures 11 and 12. It was observed that

the melting point of electrospun Maxon in Ge-GeM-GM

virgin scaffold decreased from 209 to 2018C because of 30

days of aging. DSC scan (second heating) showed an amor-

phous a soft segment Tg around 58C (Figure 12) and an

exothermic crystallization peak around 858C. The crystal-

linity (%) of aged samples can be determined from the ra-

tio of enthalpy of fusion to standard enthalpy of fusion

(DHof ) by assuming the heat of fusion of an infinite PGA

crystal as 139.1 J/g.39 From the DSC data (first heating)

GE-GEM-GM scaffolds, the apparent crystallinity (%) was

obtained as 33.11% for 5 days, 52.83% for 10 days, 37.05

% for 15 days, 31. 56 for 20 days, and 25. 77 % for 30

days (Figure 12). It is clear that crystallinity of the fibers

increased initially due to degradation of amorphous phase

and then decreased finally due to degradation of crystalline

phase. A similar trend was observed in the DSC data of

aged Ge-GeM-GM scaffolds also (Figure 11). The apparent

percentage of crystallinity was found to be 40.03% for 5

days, 45.13 % for 10 days, 43.25% for 15 days, 53.20%

for 20 days, and 35.58% for 30 days. The crystallinity of

virgin Ge-GeM-GM sample calculated from DSC experi-

ment was 22%. This shows a significant increase in crystal-

linity with aging and dissolution of the proteins as well as

the amorphous phase of Maxon. However, by 30 days, the

crystalline PGA segments also started to degrade and this

resulted in the decrease in percentage crystallinity. The en-

thalpy of fusion decreased drastically for samples aged for

30 days indicating the degradation of defective crystalline

regions in Maxon.

The biodegradation has a deteriorative effect on me-

chanical properties of the scaffolds. Tables I and II give

the tensile test results of GE-GEM-GM scaffold and Ge-

GeM-GM scaffold, respectively. Though the weight-ratio of

Maxon to protein is identical in both scaffolds, Ge-GeM-

GM exhibited higher tensile properties (tensile strength of

2.11 6 0.53 MPa and tensile modulus of 27.38 6 7.03)

Figure 11. DSC scans of aged Ge-GeM-GM scaffold. The melting

point decreases with increasing aging period. [Color figure can be

viewed in the online issue, which is available at www.interscience.wiley.com].

Figure 12. DSC scans of aged GE-GEM-GM scaffold. Prolonged

aging up to 45 days decreased the crystallinity and melting point.

[Color figure can be viewed in the online issue, which is available atwww.interscience.wiley.com].

TABLE I. Mechanical Properties of Aged GE-GEM-GM Scaffold

Aging

Period

(day)

Tensile

Strength (MPa)

Failure

Strain (%)

Tensile

Modulus

(MPa)

0 0.76 6 0.07 45 6 4.0 3.72 6 0.27

5 0.26 6 0.11 28 6 7.0 1.32 6 0.92

10 0.27 6 0.07 16 6 0.5 2.12 6 0.36

20 0.02 6 0.01 1.0 6 0.2 –

30 Physically cracked – –

TABLE II. Mechanical Properties of Aged Ge-GeM-GM Scaffold

Aging

Period

(day)

Tensile

Strength (MPa)

Failure

Strain (%)

Tensile

Modulus

(MPa)

0 2.11 6 0.53 71 6 20 27.38 6 7.03

5 1.68 6 0.31 83 6 22 14.33 6 3.34

10 1.02 6 1.00 24 6 7.6 15.01 6 2.61

20 1.06 6 0.28 1.2 6 0.18 –

30 Physically cracked – –



than GE-GEM-GM (tensile strength of 0.76 6 0.07 MPa

and tensile modulus of 3.72 6 0.27). This could be attrib-

uted to the presence of soluble elastin in the latter. The

soluble elastin with relatively lower molecular weight

chains may form a phase-mixed blend system in protein/

Maxon fibers in Ge-GeM-GM scaffolds. The high molecu-

lar weight insoluble elastin can cause some immiscibility

of the proteins with synthetic polymers in the blended-

fiber-structure and the resulted phase-separated blend struc-

ture may be responsible for the inferior mechanical proper-

ties. In vitro biodegradation impacted negatively to the

mechanical properties of both scaffolds. The tensile

strength and failure strain (%) of the aged tubular scaffolds

decreased with time. The decrease in mechanical properties

was more pronounced in the case of GE-GEM-GM scaffold

than Ge-GeM-GM scaffold. The tensile strength decreased

by more than 90% by 20 days for GE-GEM-GM scaffold

whereas that decreased only by 50% for Ge-GeM-GM scaf-

fold for the same period, that is, Ge-GeM-GM scaffold

retained 50% or more of its tensile strength up to 20 days

of PBS aging. The decrease in the failure strain (%) in the

case of both scaffolds is due to the disintegration and re-

moval of elastic proteins from the blend fibers.

DISCUSSION

For a tissue engineering scaffold, it is supposed that the

rate of degradation of the scaffold should balance the rate

of formation of new tissues. However, for engineering of

functionally dynamic tissues such as vascular and cardiac

tissues, the scaffolds have to undergo the intended cyclic

loading. Because biodegradation has a negative effect on

the mechanical and structural integrity of the scaffolds, it is

essential to study a correlation between materials degrada-

tion in vitro and mechano-morphological characteristics

before implantation in animals. It is known that most of the

biodegradable polyesters undergo degradation by hydrolysis

mechanism initialed by different hydrolyzing agent such as

water, ions, and enzymes. At first, water penetrates into the

materials and then the hydrolytic agents and pH of the me-

dium accelerate the bulk degradation of swollen materials.

Electrospun fibers have high surface area to volume ratio

and electrospun scaffolds have high porosity with well-

interconnected pores for water absorption. The protein-

blending and multilayering have further decreased the

nanofiber diameters and increased the surface area and po-

rosity in comparison with electrospun pure Maxon.25 This

explains the increased percentage of PBS medium uptake

by the trilayer scaffolds. Moreover, the hydrophilicity of

the proteins and hygroscopic nature of glycolide segments

in Maxon also contribute towards the increased PBS uptake

by the scaffolds.

It is known that the hydrolytic degradation of polyglyco-

nate occurs through a bulk degradation process, if the gly-

colide content is higher than 20 mol %.29 Again, the

kinetics of ester hydrolysis follows a first order dependence

on carboxylic end groups,28 whose concentration increases

in aging medium and thereby induces an auto catalytic

effect. Because of the large surface area and high hydrophi-

licity of protein blended Maxon, the rate of degradation is

expected to be higher compared to that for the degradation

of suture macrofilaments. Polyglyconate (Mw 5 86,000) in

the form of sutures of 0.553 mm diameter and 90 cm

length have exhibited a weight loss of 20% after 18 days

of incubation in PBS (pH 7.4) medium.28 The higher per-

centage of mass loss in the present scaffolds can be attrib-

uted to the high surface area and porosity of electrospun

scaffolds The interconnectivity of pores helps to diffuse the

fiber-fragments out from the rest of the scaffolds. Because

of the PBS uptake the individual fibers swell and fuse to-

gether in the initial stages and finally lead to fiber-breaking.

The analysis of SEM images leads to conclusion that by

10–15 days of aging fibrous morphology of the scaffolds

changed drastically and the inner protein layer dissolved or

delaminated from the tube. This can be correlated to the

initial high increase in the rate of mass loss up to 10 days

in the case of both scaffolds.

Maxon is an ABA type block copolymer with a seg-

mented nature having amorphous and crystalline blocks of

different chemical constituents. The middle soft segment is

constituted of a randomly distributed glycolide and tri-

methylene carbonate units and the hard segment is basi-

cally composed of glycolide units as the two end blocks of

the copolymer. Usually, amorphous regions are more sus-

ceptible to hydrolysis as water can penetrate easily there

than in highly packed and densely ordered crystalline

regions. In the case of Maxon, a competitive effect may

exist as a result of its chemical nature having highly hygro-

scopic polyglycolide crystalline segments.28 Therefore, the

initial increase in the crystallinity observed in the DSC

scans can be attributed due to the degradation of amor-

phous phase and further decrease in the crystallinity due to

the degradation of crystalline phase with prolonged time

(�1 month). The samples aged for 45 days further contin-

ued the decrease in crystallinity to 22.46%. In a recent

study of in vitro hydrolysis of Maxon monofilaments, Zur-

ita et al.28 have shown that the weight-average molecular

weight of Maxon clearly decreased to a value lower than

20,000 within 28 days of aging in PBS. This type of chain-

scission and fragmentation can be envisaged in the present

elctrospun Maxon scaffolds through decrease in weight and

melting point (from 208 to 2018C).In vitro biodegradation can be substantiated by IR func-

tional group analysis of aged scaffolds. The disappearance

or the loss of intensity of the peaks at 1143 and 1084 cm21

and the bands for amorphous regions (846 and 716 cm21)

clearly explains the early stages of hydrolysis of PGA and

its copolymers (Figures 6 and 7). Similar results have been

reported for the degradation of injection molded disks of

PGA, the major constituent of Maxon (67.5%). Chu et al.36

have shown that bands for amorphous region around 850,

753, 713, and 560 cm21 exhibited a drastic decrease in the



intensity than the bands for crystalline region around 972,

901, 806, 627, and 590 cm21 during the degradation of

PGA. The absorption band at 1244 cm-1 can be assigned to

the O��C��O asymmetric stretching of carbonate in TMC

units.40 And the intensity of which is not decreased signifi-

cantly in the present scaffolds, suggesting the relative sta-

bility of TMC units in the copolymer. The observed peak-

shift from 1084 to 1093 cm21 at 20 days of aging (Figure

6) could be due to the domination of TMC unit after the

disappearance of amorphous PGA segments.37 Previous

studies41–43 have shown the TMC based polymers degrade

slowly and the rate of hydrolysis of TMC based copoly-

mers is mainly depends on the hydrolysis of ester content

in them as the ester linkages are more susceptible to hydro-

lysis than carbonate linkages.

The vascular scaffolds for use in vivo need to provide a

high degree of graft-arterial compliance in addition to the

structural integrity. Therefore, biodegradation and its

effects on elastic and fatigue properties of prosthetic grafts

must be evaluated in vitro for a substantial period of time

to maximize the likelihood of clinical success.44 The struc-

tural and mechanical integrity of the graft should be main-

tained until mature vascular tissue is formed. As the

degradation of scaffolds proceeds, the extracellular matrices

secreted by the cells will be integrated into the cell/scaffold

construct. However, we have studied the degradation of

protein blended Maxon scaffolds in an acellular medium

(PBS) to evaluate the in vitro degradation behavior on scaf-

fold properties. The studies show that the tensile properties

of scaffolds decreased due to degradation. However, the

tensile strength of Ge-GeM-GM scaffold after 20 days

incubation in PBS are comparable to that for native femo-

ral artery (tensile strength of 1–2 MPa).13 The mechanical

disintegrations developed in the present scaffolds are in ac-

cordance to the gradual loss of fibrous morphology due to

fiber-breaking and weight loss due to the removal of frag-

mented products. For the GE-GEM-GM scaffold, SEM

images confirmed an escalated degradation after 10 days

and this resulted in decrease in mechanical properties (Ta-

ble I). The Ge-GeM-GM scaffolds have shown the fibrous

morphology up to 20 days in both the middle and outer

layers and sustained its stress, even though strain continu-

ously decreased because of fiber-breaking and dissolution

of uncrosslinked proteins (Table II). The increase in the

crystallinity during aging also contributed to the brittleness

of the scaffolds (tested under dry conditions). Under physi-

ological conditions, hydration effect may contribute posi-

tively to the elasticity. The overall results show that protein

blended polyglyconate tubular scaffolds show a deteriora-

tion in mechanical and structural properties after 15–20

days in vitro incubation in PBS. The dissolution of uncros-

slinked proteins (42 wt %) increased the rate of mechanical

and structural damage in the present scaffolds. Under invivo situation, deposition of native extracellular matrix pro-

teins by the cells and the de novo tissue formation may

compensate to retain the mechanical properties. However,

chemical crosslinking of proteins could be an alternative to

improve the mechanical and structural properties as well as

the degradation behavior of protein blended scaffolds.45,46

For vascular grafts, other mechanical properties such as

compliance matching, burst pressure, and suture retention

strength are equally important to tolerate suturing and

changing arterial pressures. Therefore, further studies on

the effect of in vitro aging on the biomechanical properties

will be undertaken in future using chemically crosslinked

protein blended scaffolds.

CONCLUSION

Studies on the in vitro aging behavior of two types of pro-

tein/polyglyconate (Maxon) fibrous tubular scaffolds having

a designed trilayer structure are reported for the first time.

Under hydrolytic aging conditions (PBS), both scaffolds

exhibited maximum PBS uptake within 5 days and mass

loss within 10–15 days. XRD spectra and DSC scans

revealed an increase in the crystallinity of Maxon after 10–

15 days of aging indicating the degradation of amorphous

phase in synthetic polymer and dissolution of uncrosslinked

proteins. The IR spectral analysis of aged scaffolds showed

that the characteristic peaks for amorphous PGA segments

at 1143, 846, and 716 cm21 decreased intensity after 15

days and disappeared completely by 20 days. The peak for

crystalline regions remained with a decreased intensity by

30 days of aging, suggesting the degradation of defective

crystalline PGA blocks. The degradation in hydrolytic

media affected the mechanical properties of scaffold con-

taining insoluble elastin than that containing soluble elastin.

Ge-GeM-GM scaffold retained 50% or more of its tensile

strength up to 20 days of PBS aging. The tensile strength

(1.2 MPa) exhibited by Ge-GeM-GM scaffold after 20 days

incubation in PBS is comparable to that for native femoral

artery. However, detailed evaluation on the effect of degra-

dation on compliance, burst pressure, and suture retention

is required for the potential use of these protein blended

Maxon grafts.

Thanks are due to Dr. D. Cakir of Department of Prosthodon-tics, Drs. D.R. Dean and R. Griffin of Department of MaterialsScience and Engineering and Dr. S.A. Catledge of Department ofPhysics for the technical support and instrumental training. Theauthors also acknowledge valuable discussions with Dr. A. Eber-hardt of Department of Biomedical Engineering.

REFERENCES

1. Niklason LE, Gao J, Abbott WM, Hirschi KK, Houser S,Marini R, Langer R. Functional arteries grown in vitro.Science 1999;284:489–493.

2. Hoenig MR, Campbell GR, Rolfe BE, Campbell JH. Tissueengineered blood vessels: Alternative to autologous grafts?Arterioscler Thromb Vasc Biol 2005;25:1128–1134.

3. Tiwari A, Salacinski HJ, Punshon G, Hamilton G, SeifalianAM. Development of a hybrid cardiovascular graft using atissue engineering approach. FASEB J 2002;16:791–796.



4. Jeong SI, Kim SY, Cho SK, Chong MS, Kim KS, Kim H,Lee SB, Lee YM. Tissue-engineered vascular grafts composedof marine collagen and PLGA fibers using pulsatile perfusionbioreactors. Biomaterials 2007;28:1115–1122.

5. Buttafoco L, Engbers-Buijtenhuijs P, Poot AA, Dijkstra PJ,Vermes L, Feijen J. Physical characterization of vasculargrafts cultured in a bioreactor. Biomaterials 2006;27:2380–2389.

6. Barnes CP, Sell SA, Boland ED, Simpson DG, Bowlin GL.Nanofiber technology: Designing next generation of tissue en-gineering scaffolds. Adv Drug Deliv Rev 2007;59:1413–1433.

7. Liao S, Li B, Ma Z, We H, Chan C, Ramakrishna S. Biomi-metic electrospun nanofibers for tissue regeneration. BiomedMater 2006;1:45–53.

8. Pham QP, Sharma U, Mikos AG. Electrospinning of poly-meric nanofibers for tissue engineering applications: Areview. Tissue Eng 2006;12:1197–1211.

9. Thomas V, Dean D, Vohra YK. Nanostructured biomaterialsfor regenerative medicine. Curr Nanosci 2006;2:155–177.

10. Li M, Mondrinos MJ, Gandhi MR, Ko FK, Weiss AS, LelkesPL. Electrospun protein fibers as matrices for tissue engineer-ing. Biomaterials 2005;26:5999–6008.

11. Teo WE, Kotaki M, Mo XM, Ramakrishna S. Porous tubularstructures with controlled fiber orientation using a modifiedelectrospinning method. Nanotechnology 2005;16:918–924.

12. Boland ED, Mathews JA, Pawlowski KJ, Simpson DG, WnekGE, Bowlin GL. Electrospinning collagen and elastin: Prelim-inary vascular tissue engineering. Front Biosci 2004;9:1422–1432.

13. Sell SA, McClure MJ, Barnes CP, Knapp DC, Walpoth BH,Simpson DG, Bowlin GL. Electrospun polydioxanone-elastinblends: Potential for bioresorbable vascular grafts. BiomedMater 2006;1:72–80.

14. Stitzel J, Liu J, Lee SJ, Komura M, Berry J, Soker S, Lim G,Dyke MK, Czerw R, Yoo JJ, Atala A. Controlled fabricationof a biological vascular substitute. Biomaterials 2006;27:1088–1094.

15. Gan ZH, Fung JT, Jing XB, Wu C, Kuliche WK. A novellaser light-scattering study of enzymatic biodegradation ofpoly(e-caprolactone) nanoparticles. Polymer 1999;40:1961–1967.

16. Wu C, Gan Z. A novel method of studying polymer biodegra-dation. Polymer 1998;39:4429–4431.

17. Chen DR, Bei JZ, Wang SG. Polycaprolactone microparticlesand their biodegradation. Polym Degrad Stab 2000;67:455–459.

18. Zong XZ, Ran SF, Kim KS, Fang DF, Hsiao BS, Chu B.Structure and morphology changes during in vitro degradationof electrospun poly(glycolide-co-lactide) nanofiber membrane.Biomacromolecules 2003;4:416–423.

19. You Y, Min BM, Lee SJ, Lee TS, Park WH. In vitro degrada-tion behavior of electrospun polyglycolide, polylactide, andpoly(lactide-co-glycolide). J Appl Polym Sci 2005;95:193–200.

20. You Y, Lee SW, Youk JH, Min BM, Lee SJ, Park WH. Invitro degradation behaviour of non-porous ultra-fine poly(gly-colic acid)/poly(L-lactic acid) fibres and porous ultra-fine poly(glycolic acid) fibres. Polym Degrad Stab 2005;90:441–448.

21. Buttafoco L, Kolkman NG, Poot AA, Dijkstra PJ, Vermes I,Feijen J. Electrospinning collagen and elastin for tissue engi-neering small diameter blood vessels. J Control Release2005;101:322–324.

22. Stankus JJ, Guan J, Wagner WR. Fabrication of biodegradableelastomeric scaffolds with sub-micron morphologies. J BiomedMater Res A 2004;70:603–614.

23. Matthews JA, Simpson GD, Wnek GE, Bowlin GL. Electro-spinning of collagen nanofibers. Biomacromolecules 2002;3:232–238.

24. Vaz CM, Tuijl SV, Bouten CVC, Baaijens FPT. Design ofscaffolds for blood vessel tissue engineering using a multi-layering electrospinning technique. Acta Biomater 2005;1:575–582.

25. Thomas V, Zhang X, Catledge SA, Vohra YK. Functionallygraded electrospun scaffolds with tunable mechanical proper-ties for vascular tissue regeneration. Biomed Mater 2007;2:224–232.

26. Veith FJ, Gupta SK, Ascer E, White-Flores S, Samson RH,Scher LA, Towne JB, Bernhand VM, Bonier P, Flinn WR,Astelford P, Yao JST, Bergan JJ. Six years prospective multi-center randomized comparison of autologous saphenous veinand expanded polytetrafluoroethylene grafts in infrainguinalarterial reconstructions. J Vasc Surg 1986;3:104–114.

27. Tomihata K, Suzuki M, Ikada Y. The pH dependence ofmonofilament sutures on hydrolytic degradation. J BiomedMater Res Appl Biomater 2001;58:511–518.

28. Zurita R, Franco L, Puiggali J, Rodriguez-Galan A. Thehydrolytic degradation of a segmented glycolide-trimethylenecarbonate copolymer (MaxonTM). Polym Degrad Stab 2007;92:975–985.

29. Jie C, Zhu KJ, Shilin Y. Preparation, characterization and bio-degradable characteristics of poly(1,3-trimethylene carbonate-co-glycolide). Polym Int 1996;41:369–375.

30. Metz SA, Chegini N, Masterson BJ. In vivo and in vitro deg-radation of monofilament absorbable sutures PDS and Maxon.Biomaterials 1990;11:41–45.

31. Kangas J, Paasimaa S, Makela P, Leppilahti J, Tormala P,Waris T, Ashammakhi N. Comparison of strength properties ofpoly-L/D-lactide (PLDLA) 96/4 and polyglyconate (Maxon1)sutures: In vitro, in the subcutis, and in the achilles tendon ofrabbits. J Biomed Mater Res 2001;28:121–126.

32. Partridge SM, Davis FF, Adair ST. The chemistry of connec-tive tissues 2: Soluble proteins derived from partial hydrolysisof elastin. Biochem J 1955;61:11–21.

33. Heitz M, Carrasco F, Overend RP, Chornet E. Correlationbetween polystyrene molecular weights and a characteristictemperature derived from the thermogravimetric weight losscurves. Thermochim Acta 1989;142:83–88.

34. Thomas V, Jose MV, Chowdhury S, Sullivan JF, Dean DR,Vohra YK. Mechano-morphological studies of aligned nanofi-brous scaffolds of polycaprolactone fabricated by electrospin-ning. J Biomater Sci Polym Ed 2006;17:969–984.

35. Thomas V, Dean D, Jose MV, Mathew B, Chowdhury S,Vohra YK. Nanostructured biocomposite scaffolds based oncollagen coelectrospun with nanohydroxyapatite. Biomacro-molecules 2007;8:631–637.

36. Chu CC, Zhang L, Coyne LD. Effect of gamma-irradiationand irradiation temperature on hydrolytic degradation of syn-thetic absorbable sutures. J Appl Polym Sci 1995;56:1275–1294.

37. Wang H, Dong JH, Qiu KY. Synthesis and characterization ofABA-type block copolymer of poly(trimethylene carbonate)with poly(ethylene glycol): Bioerodible copolymer. J PolymSci Part A: Polym Chem 1998;36:695–702.

38. Payne KJ, Veis A. Fourier transform IR spectroscopy of colla-gen and gelatin solutions: Deconvolution of the amide I bandfor conformational studies. Biopolymers 1988;27:1749–1760.

39. Cohn D, Younes H, Marom G. Amorphous and crystallinemorphologies in glycolic acid and lactic acid polymers. Poly-mer 1987;28:2018–2022.

40. Wang H, Dong JH, Qiu KY, Gu ZW. Synthesis of poly(1,4-dioxan-2-one-co-trimethylene carbonate) for application indrug delivery systems. J Polym Sci Part A: Polym Chem1998;36:1301–1307.

41. Zhu KJ, Hendren RW, Jensen K, Pitt CG. Synthesis, proper-ties, and biodegradation of poly(1,3-trimethylene carbonate).Macromolecules 1991;24:1736–1740.



42. Albertsson AC, Eklund M. Influence of molecular structureon the degradation mechanism of degradable polymers: Invitro degradation of poly(trimethylene carbonate), poly(tri-methylene carbonate-co-caprolactone), and poly(adipic anhy-dride). J Appl Polym Sci 1995;57:87–103.

43. Pego AP, Poot AA, Grijpma DW, Feijen J. In vitro degrada-tion of trimethylene carbonate based copolymers MacromolBiosci 2002;2:411–419.

44. L’Heureux N, Dusserre N, Marini A, Garrido S, de la FuenteL, McAllister T. Technology insight: The evolution of tissue-

engineered vascular grafts- From research to clinical practice.Nat Clin Pract Cardiovasc Med 2007;4:389–395.

45. Bigi A, Cojazzi G, Panzavolta S, Rubini K, Roveri N. Me-chanical and thermal properties of gelatin at different degreesof glutaraldehyde crosslinking. Biomaterials 2001;22:763–768.

46. Sung HW, Chang Y, Chiu CT, Chen CN, Liang HC. Cross-linking characteristics and mechanical properties of a bovinepericardium fixed with a naturally occurring crosslinkingagent. J Biomed Mater Res 1999;47:116–126.



In vitro biodegradation of designed tubular scaffolds of electrospun protein/polyglyconate blend...

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