Photopolymerized Injectable RGD-Modified Fumarated Poly(ethylene glycol) Diglycidyl Ether Hydrogels for Cell Growth

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Photopolymerized Injectable RGD-ModifiedFumarated Poly(ethylene glycol) DiglycidylEther Hydrogels for Cell Growth

Z. Seden Akdemir, Handan Akcakaya, M. Vezir Kahraman, Taskın Ceyhan,Nilhan Kayaman-Apohan, Atilla Gungor*

In this study, photopolymerized hydrogels of fumarated poly(ethylene glycol) diglycidyl-co-poly(ethylene glycol) diacrylate have been synthesized and modified with cell adhesionpeptide, Arg-Gly-Asp (RGD). The structural and mechanical properties of the hydrogels arefound to be poly(ethylene glycol) diacrylate (PEGDA) dependent. The percentage of gelation isincreased from 72 to 89 wt.-% when the amount of the crosslinker co-monomer (PEGDA) in thehydrogel formulation is increased from 20 to 40 wt.-%. In the present case, the equilibriummass swelling is decreased from 216 to 93%. The viscosities of the uncured formulations havealso been measured and likewise, the resultswere influenced by the increasing amount ofPEGDA that reduced the value from 83 to36 cP. The compressive modulus of the preparedhydrogels was improved with the addition of thePEGDA. Cell growth experiments have been per-formed by comparing the properties of the hydro-gels with and without RGD units. The resultsshow that RGD units enhance the adhesion ofcells to the surface of the hydrogels. SEM-EDSstudies reveal that nitrogen and calcium areproduced on the osteoblast-seeded surface ofthe scaffold within the culture time period.

Introduction

Tissue engineering is an interdisciplinary field that applies

the principles of engineering and biomedical sciences

Z. S. Akdemir, M. V. Kahraman, N. Kayaman-Apohan, A. GungorDepartment of Chemistry, Marmara University, 34722 Goztepe-Istanbul, TurkeyFax: þ90 216 3478783; E-mail: [email protected]. AkcakayaBiophysics Department, Istanbul University Capa-Istanbul, TurkeyT. CeyhanCevre Hospital Mecidiyekoy-Istanbul, Turkey

Macromol. Biosci. 2008, 8, 852–862

� 2008 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

toward the development of biological substitutes that

restore, maintain, or improve tissue or organ function.[1,2]

A tissue engineered implant is a biologic–biomaterial

combination in which cells are transplanted to penetrate

and proliferate in all directions to populate all regions of

the implant. Cell proliferation can be facilitated and enhanced

with the introduction of a cell adhesive sequence. The

properties of the implant such as biocompatibility,

porosity, biodegradability, and interconnectivity have an

impressive role in the formation of the new tissue.[3]

Recent developments have shown that hydrogels are

good candidates for tissue-engineered implants because of

their hydrophilic structure, which gives them physical

DOI: 10.1002/mabi.200700319

Photopolymerized Injectable RGD-Modified Fumarated Poly(ethylene glycol) . . .

characteristics similar to tissue.[4,5] A cell-transplanted

hydrogel acts as a temporary support system for the

growing new cells, which can reorganize into functional

tissue as the hydrogel degrades. Many synthetic hydrogels

have been investigated for tissue engineering applications

including poly(vinyl alcohol) (PVA),[6,7] poly(L-lactic acid)

(PLA), poly(glycolic acid) (PGA) and their copolymers,[8–10]

and poly(ethylene glycol) (PEG) and poly(ethylene glycol)-

based copolymers.[11–15]

PEG, also known as poly(ethylene oxide), has a very

hydrophilic nature that restricts the adhesion of proteins

and cells, but by using PEG in a co-polymer, it enhances the

biocompatibility of the composition and also enables

researchers to determine the cell attachment character-

istics of the prepared hydrogel.[2,16] Nevertheless, the

PEGylation method allows a protein, peptide, or non-

peptide molecule to be linked to PEG chains, to supply a

better hydrogel–cell interconnectivity.[17]

Injectable biomaterials are preferred in bone treatment

since they present a potential to minimize the invasive-

ness of some surgical techniques. Moreover, they can

easily be combined with the desired cells or growth

factors in a solution state prior to injection and have the

ability to fill the damaged area completely.[18–20] Photo-

polymerized hydrogel systems demand the above require-

ments. They elicit better temporal and spatial control over

the gelation process, minimal heat production, ability to

uniformly encapsulate cells, are operational at low

temperature, minimize the damage to the entrapped

bioactive agents or cells during hydrogel formation, and

are injectable in nature before in-situ polymeriza-

tion.[21–24] Practically, the low viscosity liquid hydrogel

formulation can be easily placed into complex shaped

areas through a small crease by injection and subse-

quently in-situ polymerization can be achieved by using a

fiber optic cable connected to a UV processor.[25] From this

aspect, hydrogels can be used to fill irregularly damaged

sections, to allow minimally invasive surgical procedures,

and to act as a facilitator to incorporate with cells.[2]

Poly(ethylene glycol) diacrylate (PEGDA) is a promising

tissue engineering material, because its fairly low-

viscosity prepolymer solution can be applied to tissue

and can be photopolymerized in situ to form a hydrogel

film on the tissue surface.

One of the prime goals of this study was to synthesize a

new PEG-based resin that has two repeating units,

poly(ethylene glycol) diglycidyl ether (PEGDGE) and

fumaric acid (FA), which are linked to each other by ester

bonds. Insertion of double bonds into a PEG chain enables

the formation of a crosslinked network by photopolymer-

ization in the presence of a radical initiator. Photopoly-

merizable hydrogel formulations that contain the new

synthesized resin are reinforced with PEGDA. Moreover,

PEGDA would adjust the viscosity of the formulations and



provides extra double bonds to the compositions, thereby

increasing the crosslink density.

Another task was to support the cell adhesion in the

prepared hydrogel matrix. Hence, we used Arg-Gly-Asp

(RGD) as a cell-adhesive peptide. RGD is a minimal peptide

sequence, found throughout the body including many

bone extracellular matrix proteins. It is stable because of

its short peptide sequence and can be coupled to the amine

of reactive-N-terminated PEG chains.[26,27] Poly(ethylene

glycol) monoacrylate can be modified with 1,10-carbonyl-

diimidazole (CDI) to yield amine-reactive acylimidazole-

terminated PEG chains, which can subsequently be

covalently bonded to RGD.

This paper represents the synthesis and the character-

ization studies of the poly(ethylene glycol) diglycidyl

fumarate (fumarated PEGDGE) and RGD-attached poly

(ethylene glycol) monoacrylate (RGD-PEGMA) resins. The

formulations composed of fumarated PEGDGE, PEGDA,

and RGD-PEGMA in water were crosslinked by photopoly-

merization techniques in the presence of a photoinitiator

to form a novel hydrogel. Before the photopolymerization

step, the viscosity of each formulation was determined.

Mechanical properties of the hydrogels were analyzed by

compressive testing. Furthermore, it was investigated

whether the hydrogels could serve as a tissue-engineering

scaffold.

Experimental Part

Materials

PEGDGE (Mn ¼ 526), poly(ethylene glycol) monoacrylate (PEGMA,

Mn ¼375), PEGDA (Mn ¼ 258), and FA, were purchased from

Aldrich Chemical Co. RGD was obtained from Sigma. CDI and

triphenyl phosphine (TPP) were provided by Fluka AG. The

photoinitiator 1-hydroxy cyclohexyl phenyl ketone (Irgacure-184)

was purchased fromCiba Speciality Chemicals. All other chemicals

were of analytical grade and were purchased from Merck AG.

Freshly double-distilled water was used throughout.

Synthesis of Fumarated PEGDGE

FA (6.2 g, 0.0532 mol, equivalent to 40% of the total epoxy content

in PEGDGE), was added to a three-neck round-bottomed flask

charged with 70 g (0.266 mol) of PEGDGE and 0.7 g of TPP (as a

catalyst, 1% w/w). The reaction was stirred mechanically under

nitrogen atmosphere at 80 8C until the acid value (mg KOH � g�1) of

the resinwas less than one. The acid value of the fumarated PEGDE

resin showed that all the acid groups had reacted with epoxy

groups. At the end of the reaction the resulting resin (PEGMA)was

analyzed by FT-IR spectroscopy. The epoxy content was deter-

mined by a titration method (ASTM D1652). Fumaric acid-

modified PEGDGE was prepared as shown in Scheme 1.

IR (NaCl): 3 400 (–OH), 1 720 (C––O), 1 650 (C––C), 1 250 cm�1

(C–O–C).

www.mbs-journal.de 853

Z. S. Akdemir, H. Akcakaya, M. V. Kahraman, T. Ceyhan, N. Kayaman-Apohan, A. Gungor

Scheme 1. Modification of PEGDGE with FA.

854

Activation of PEGMA with CDI

PEGMA (26.6 mmol) was transferred into a three-neck round-

bottomed flask filled with tetrahydrofuran (THF) and purged with

nitrogen. A solution was obtained by stirring magnetically. CDI

(29.3 mmol) was added to the reaction flask and the temperature

was raised to 40 8C and the mixture was stirred

overnight. THF was then distilled off by rotary

evaporation and the by-productswere removed by

dialysis against water. The hydroxy group of the

PEGMA was reacted with CDI to yield the

amine-reactive-N-acylimidazole.

IR (NaCl): 3 400 (imide), 1 755 (C––O str. for

imidazole), 1 720 cm�1 (C––O str. for acrylate).

Attachment of RGD into

N-Acylimidazole Tethered PEGMA

RGD (10mg) was dissolved in 10mL of 50� 10�3M

NaHCO3 solution (pH 8.2). N-Carbonylimidazole-

tethered PEGMA (0.8 g) was added dropwise to the

peptide solution and the mixture was gently

shaken at room temperature for 24 h. RGD-attached

PEGMA was frozen at �80 8C and lyophilized.

N-Acylimidazole-tethered PEGMA was coupled to

RGD as shown in Scheme 2.1H NMR (D2O/CD3OD, 1: 2): d¼2.4 (CH2–CH2–

CH2–NH–C(––NH)–NH2, arginine), 2.8 (–CH2–COOH,

aspartic acid), 3.0 (CH2–CH2–CH2–NH–C(––NH)–

NH2, arginine), 4.1–4.2 (–O–CH2–CH2–OH), 6.8–7.2

(acrylate).

Scheme 2. Attachment of RGD into N-acylimidazole tethered PEGMA.

Characterization of the Resins

The structure of fumarated PEGDGE and

N-acylimidazole-tethered PEGMA were analyzed

by FT-IR spectroscopy. RGD-PEGMA was charac-

terized using 1H NMR spectroscopy. 1H NMR

spectra were obtained using a Varian model

T-60 NMR spectrometer operated at 200 MHz.

FT-IR spectra were obtained on a Shimadzu 8300

FT-IR spectrophotometer.



Synthesis and Characterization of Hydrogels

Hydrogel formulations were prepared by mixing component A,

which consisted of an aqueous solution of the RGD-PEGMA

(10% w/v) with component B, which consisted of a mixture of

fumarated PEGDGE and PEGDA, at a total weight concentration of

18/82%. Component B was varied by changing the amount of the

PEGDA in the total mixture. Table 1 represents the composition of

the formulations. All of the formulations, which contain

Irgacure-184 (photoinitiator), were placed into a Teflon mold in

which the wells dimensions were 10 mm in diameter and 2 mm

in depth, to perform the UV-initiated crosslinking reaction. The

samples were irradiated for 10 min under a high-pressure UV

lamp (OSRAM 300 W, lmax¼ 365 nm) and the obtained hydrogels

were then removed from the wells.

Hydrogel samples were immersed in a large excess of distilled

water for a day to remove any unreacted monomers and residual

initiator. All of the hydrogels were dried in a vacuum oven at 30 8Cfor several days until reaching a constant weight. The dry gels

were weighed (Wi) and soaked in distilled water at room

temperature (20.0� 0.1 8C). The swollen gels were removed from

DOI: 10.1002/mabi.200700319


Table 1. The composition of the hydrogel formulations.

Hydrogel Component A Component B Photoinitiator

Water RGD-modified PEGMA FA-modified PEGDGE PEGDA g

g g g g

FPEG20 2.0 – 8.0 2.0 0.2

FPEG20-2 2.0 0.2 8.0 2.0 0.2

FPEG30-2 2.0 0.2 7.0 3.0 0.2

FPEG40-2 2.0 0.2 6.0 4.0 0.2

FPEG20-6 2.0 0.6 8.0 2.0 0.2

the distilled water, gently wiped with filter paper to remove any

water at the surface and weighed (Ws). Finally, all of the swollen

gels were kept at�80 8C overnight and then lyophilized. The dried

hydrogels were weighed (Wd).

The degree of gelation was calculated by using the formula:

Macrom

� 2008

Gelation ð%Þ ¼ ðWd=WiÞ � 100 (1)

The degree of swelling was calculated by using the formula:

Swelling ð%Þ ¼ ½ðWs �WdÞ=Wd� � 100 (2)

Degradation of Hydrogels

Degradation behavior of the hydrogels was investigated by the

following method. The vacuum dried hydrogels were weighed

(W0) and transferred into phosphate-buffered saline (PBS) (pH 7.4)

at 37 8C. Each pre-weighed hydrogel was kept in the buffer

solution until a specified time was reached and then removed.

Samples were dried under vacuum and thenmeasured to obtain a

final weight (Wdt). The weight loss of the hydrogels was calculated

by using Equation (3).

Total weight loss ð%Þ ¼ ðW0 �Wdt=W0Þ � 100 (3)

Rheometry

The viscosity of the prepared formulations was measured before

the photocrosslinking process using a Brookfield RV DV-IIþ Pro

Viscometer. The formulation was placed into a Teflon mold

(20mm in diameter and 70 mm in depth) and positioned on the

temperature-controlled plate. The temperature was set to 37 8C.The viscosity was measured with a spindle speed of 200 rpm.

Compressive Mechanical Testing

Compressive testing of hydrogels was conducted using a

mechanical testing system Zwick Roell BDO-FBO.5TH. For

compression measurement, hydrogels were prepared in a

ol. Biosci. 2008, 8, 852–862

WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

cylindrical shape by using a glass tube with a depth of 70mm and

14 mm in diameter. Each of the formulation solutions was poured

into these tubes and photopolymerized under a UV lamp for

10min. After polymerization, all of the cylindrical shaped samples

were removed from the glass tube and cut into 10 mm pieces.

Hydrogel cylinders were compressed in a wet state at a crosshead

speed of 5 mm �min�1 while stress and strain were monitored

throughout. The fracture strength and compressive strain was

evaluated by using software called Test Xper in Zwick Roell

BDO-FBO.5TH. The compressive modulus was calculated from the

linear region of the stress–strain curve of the compression test.

Differential Photocalorimetry

Photopolymerization of the hydrogel formulation was carried out

by Pyris Diamond DSC equipped with an EXFO Omni-Cure 2000

photo-DSC accessory. Filtered light (250–450 nm) with an

intensity of 20 mW � cm�2 at the tip of the light guide was used.

Approximately 150 mg of each sample was placed in the

aluminum differential scanning calorimetry (DSC) pans.

Heat flow vs. time curves were recorded in an isothermal mode

under a nitrogen flowof 20mL �min�1 at 37 8C. The rate of reactionin these experimentswas calculated by the following equation:[28]

Rp ¼ ðH=WÞ=DH (4)

where Rp is the rate of polymerization in s�1, H is the heat flow in

mW, W is the weight of monomer solution in mg and DH is the

enthalpy of the material in J � g�1.

The heat liberated during the polymerization reaction was

directly proportional to the number of vinyl groups reacted in the

system. By integrating the area under the exothermic peak, the

conversion of the vinyl groups (C) could be calculated by:[29]

C ¼ DHt=DHtheor0 (5)

where DHt is the heat evolved at time t, and DHtheor0 is the

theoretical heat for complete conversion.[30]

In vitro Studies of the Cell-Seeded Hydrogels

3T3 mouse fibroblasts were grown in Dulbecco’s Modified Eagle

Medium (DMEM)/F12 medium containing 10% fetal calf serum,



Figure 1. FT-IR spectra of a) Fumarated PEGDGE and b)N-carbonylimidazole tethered PEGMA

856

100 IU �mL�1 penicillin, and 0.1 mg �mL�1 streptomycin and were

kept in a controlled-atmosphere (5% CO2) incubator at 37 8C.Osteoblasts were obtained from the fourth and fifth passages of a

differentiation process of human bone marrow mesenchymal

stromal cells (MSCs). At subconfluence, MSCs were detached with

trypsin-EDTA and were grown in DMEM-low glucose (DMEM-LG)

with the addition of the osteogenic differentiation medium.[31]

For cell seeding, hydrogels with and without RGD peptide units

were prepared as described above. The hydrogels were sterilized

using 100% ethanol and then placed in 6-well tissue culture plates.

Cells were seeded on the surface of the hydrogels at an

approximate density of 105 cells �mL�1. After 24 h, the hydrogels

were taken out from the medium, rinsed twice with sterilized PBS

solution and the cells were then fixed with 2.5% glutaraldehyde

for 20 min. After fixation, the hydrogels were washed with PBS

solution once again. The lyophilization was performed at �100 8Cfor 20 min. Invert microscopy (Olympus UK) was used to

investigate cell–hydrogel interactions. The morphological topol-

ogy of the fibroblast-seeded surface images were taken on a

Philips XL30 ESEM FEG environmental scanning electron micro-

scope which made it possible to examine the materials in their

natural, uncoated state under low vacuum around 0.6–0.8 Torr

and at 10 kV. The samplemorphologies were investigated at room

temperature to reveal their natural surface characteristics by

using a special technique (gaseous low vacuum mode) that

allowed charge-free, gaseous secondary electrons. In addition,

scanning electron microscopy (SEM) was performed using a SEM

JOEL JSM-5910 LV. The specimens were prepared for SEM by

lyophilization at �80 8C and the dehydrated samples were placed

in liquid nitrogen until the SEM examination, which was

performed on the same day. An approximately 300 A gold coating

was applied using an Edwards S 150 B sputter coater.

Results and Discussion

Synthesis of Fumarated PEGDGE

PEGDGEwasmodified with fumaric acid in the presence of

triphenylphosphine, which is used as a catalyst. When the

amount of the epoxy groups reacted was calculated to be

40% of the total epoxy content, the reaction was

terminated. An FT-IR spectrum of the fumarated resin is

shown in Figure 1a. As seen in Figure 1a, a broad band

appears at 3 400 cm�1, which indicates the secondary

hydroxy (–OH) groups of the newproduct. A strong band at

1 720 cm�1 also illustrates ester bond formation because of

the reaction between fumaric acid and epoxy groups. The

band at 1 650 cm�1 belongs to the double bonds of the

fumaric acid, which is incorporated into the polymeric

chain during the reaction. One can also clearly see the

etheric C–O–C bond stretch at 1 250 cm�1.

Activation of PEGMA with CDI and Attachmentof RGD

CDI is often used for the coupling of amino acids or

peptides. It may easily react with the hydroxy group of the



PEGMA to produce a derivative that is sensitive to

nucleophiles. Therefore, at the end of the activation

reaction, it is sensible to expect the replacement of the

O–H stretching band with an imide band in the FT-IR

spectrum of theN-acylimide-terminated PEGMA. Figure 1b

is the spectrum of the N-acylimide-terminated PEGMA. At

3 400 cm�1 a newly formed imide band has appeared

while the –OH streching band has disappeared. Two

different carbonyl groups, one from the acrylate and the

other from the imidazole groups cause a split in the band

at 1 750 cm�1.

After RGD was attached to N-acylimide-terminated

PEGMA, 1H NMR analysis of the new product was

performed in a mixture of D2O and CD3OD (1: 2). As seen

in Figure 2, the 1H NMR spectrum of RGD-attached PEGMA

shows broad multiples that range from 2.4 to 3.0 ppm.

These multiples are attributed to the CH2–CH2–CH2–NH–

C(––NH)–NH2 group of arginine, which appears around

2.4 ppm, a multiplet of the –CH2–COOH of aspartic acid at

2.8 ppm, and a triplet of CH2–CH2–CH2–NH–C(––NH)–NH2

of arginine at 3.0 ppm. Ethylene glycol groups, –O–CH2–

DOI: 10.1002/mabi.200700319


Figure 2. Proton NMR spectra of RGD-attached PEGMA.

CH2–OH, show triplet peaks at 4.1 and 4.2 ppm. The proton

peaks of the acrylate group are at 6.8 and 7.2 ppm.

Gel Percentage and Equilibrium Mass Swelling of theHydrogels

Table 2 shows the gelation percentages and the equili-

brium mass swelling results of the hydrogels. It can be

seen in Table 2 that the hydrogel with 20 wt.-% of PEGDA

content has a 52 wt.-% gel fraction. It is thought that

various factors such as the formulation viscosity, the film

thickness, and initiator content may influence the gel

content. Because of the mobility restrictions that appear

upon gelation and ultimately vitrification of the

Table 2. Characterization of the hydrogels.

Sample

No.

Gel

percentage

Equilibrium %

mass swelling

Vis

wt.-%

FPEG20-2 52W 2.4 216W 4.1

FPEG30-2 62W 2.1 151W 4.0

FPEG40-2 69W 1.9 93W 3.7



UV-exposed sample, a decrease in conversion takes place.

In addition, because of the well-known screening effect,

UV light cannot penetrate sufficiently deeply into the

sample and high conversion could not be achieved. As seen

in Table 2, the gelation percentages increase from 52 to 69

wt.-% with an increase in PEGDA content in the hydrogel

formulation. This is because the double bonds of PEGDA

provide an additional contribution to the crosslinking

process. It is known that the degree of crosslinking and the

swelling behavior of the hydrogel are diversely related to

each other. Therefore, the decrease in swelling is expected

upon increasing the PEGDA content of the hydrogel

formulation. In the present case, the equilibrium mass

swelling is decreased from 216 to 93%.

cosity Fracture

strength

Compressive

modulus

Heat

release

cP N �mmS2 N �mmS2 J � gS1

85 0.11W 0.02 0.74W 0.05 20.74

49 0.35W 0.07 1.56W 0.16 22.86

36 1.09W 0.12 3.60W 0.23 25.93



858

Degradation

Degradable biomaterials are essential in orthopaedic

applications because second surgical intervention is not

necessary to remove the implant after a certain time.[8] In

several studies, it has been reported that the degradation

of the hydrogels could be measured by weighing the

swollen hydrogels in PBS at a certain time.[5,7,8] Anseth

et al.[7] synthesized a degradable and photocrosslinkable

hydrogel scaffold based on PVA and emphasized that

during degradation, hydrogels slowly hydrolyzed and the

structure of the network dramatically changed. This

behavior was reported because the network became

swollen and imbibed more water as the crosslinks in

the network were cleaved. However, in this study a

significant increase in water absorption of the hydrogel

samples was not observed during the first 20 weeks in PBS

media and they also exhibited no appreciable change in

their shapes. In Figure 3, the mass loss results calculated

according to Equation (3) by measuring the weight of the

dried hydrogels are given. As it can be seen from the graph,

the total mass loss of the FPEG20-2 hydrogel at the end of

the 52nd week was 73% while the other hydrogels

FPEG30-2 and FPEG40-2 were 54 and 42%, respectively.

The decrease in the degradation was probably because of

the existence of PEGDAunits.[32] In contrast to the constant

swelling ratio during degradation, after drying, the

hydrogels became very weak and crumbled upon hand-

ling, which is typical of degrading networks.

Solution Viscosity

The principle behind the design of an injectable biomater-

ial lies in the ability to syringe its uncured solution into a

damaged area in vivo easily and to fill the area completely.

Hence, the viscosity of the solution becomes important.

The viscosity of the uncured compositions was investi-

gated and the results are given in Table 2. The viscosities

were found to fall from 85 to 36 cP as the PEGDA content is

Figure 3. Degradation results of total mass loss of the hydrogels.



increased from 20 to 40%. The viscosity results demon-

strate that all formulations are suitable to use as an

injectable biomaterial.

Compressive Mechanical Testing

Hydrogels that are used in biomedical applications should

necessarily meet strict performance requirements, such as

significant mechanical strength when in contact with

body fluid. The mechanical properties of the hydrogels

were evaluated by compressive mechanical testing.

Compressive testing is preferred to determine themaximal

compressive stress that a sample can withstand. It

measures the stiffness of a given material, which reflects

the resistance of an elastic body against the deflection of

an applied force.[28] Stress–strain plots are given in

Figure 4. In Table 2, the compressive modulus and the

fracture strengths of the samples are given in the wet

state. The elastic modulus varied between 0.74� 0.05 and

3.60� 0.23 N �mm�2, while the fracture strength varied

between 0.11� 0.02 and 1.09� 0.12 N �mm�2. It is clearly

seen that the values increase in order. As it is well known,

the crosslinking density and heterogeneities such as

cyclization, unreacted oligomers, and trapped radicals

significantly affect the mechanical properties of the

hydrogels.[8] The hydrogel with 40 wt.-% PEGDA

(FPEG40-2) has an elastic modulus almost five times

greater than that with 20 wt.-% PEGDA (FPEG20-2). It is

thought that the greater reactive acrylate group fraction of

the former hydrogel makes it more crosslinked. In

addition, the molecular weight of PEGDA is lower than

themolecular weight of fumarated PEGDGE. Therefore, the

molecular weight between two crosslink points decreases

as the PEGDA content of the network increases. The tensile

strength of sample FPEG40-2 is ten times higher than that

of FPEG20-2. These results indicate that the mechanical

properties of the hydrogels depend on the crosslink density

of the network and the molecular weight between

crosslink points.

Figure 4. Stress–strain curves of cylindrical specimens loaded incompression test.

DOI: 10.1002/mabi.200700319


Figure 5. a) Conversion of double bonds and b) rate of polymeri-zation.

Differential Photocalorimetry

The rate of polymerization of hydrogel formulations was

studied with photo-DSC. The raw data obtained from

photo-DSC were the heat flow difference between a

reference cell and the cell that contained the hydrogel

formulation. A baseline correction for the heat flow was

performed in order to eliminate heat effects caused by

irradiation.[29] Figure 5a,b show the rate of polymerization

and double bond conversion versus time, respectively. As

seen in Figure 5a, FPEG20-2 showed a slightly lower rate of

polymerization compared to FPEG40-2. All formulations

display three distinct rate of polymerization stages that

are initially fast, then level off for a prolonged duration,

and then finally decrease due to the controlled propagation

and termination reactions. It has previously been reported

that the multifunctional monomers exhibit auto-

acceleration-induced behavior that leads to the very rapid

crosslinking process.[33] As seen in Figure 4b the double

bond conversion was obtained on the order of 0.7–0.9. At a



higher PEGDA content, the conversion reached 0.92. This

result demonstrates the lower reactivity and mobility of

the fumarated PEGDGE as a result of its higher molecular

weight and viscosity. Also the steric hinderance slows

down the addition of fumaric acid double bonds of the

fumarated PEGDGE resin to the polymer chain. Table 2 also

shows the reaction heat liberated during polymerization

and it varied between 20.74 and 25.93 J � g�1 with the

greatest value obtained for FPEG40-2. It can be seen in

Figure 5a that the crosslinking of the hydrogel formulation

is a very rapid process, in which gelation often occurs in a

short time. The maximum heat release occurs between 2.5

and 15.5 s after UV radiation has commenced. The short

times at which the maximum heat release and the low

heat release values obtained are encouraging for in-situ

curing applications to minimize adverse bone tissue

responses.[24] The heat of polymerization for poly(methyl

methacrylate) was reported as 54.75 kJ �mol�1 (547.5 J �g�1) and the curing time was given as 5–10 min.[34,35] The

intensity of the UV-light also had an effect on the in-situ

polymerization.[22,25] Amsden et al.[22] fabricated their

elastomer device by using 100mW � cm�2 long-wave

ultraviolet light (320–480 nm) at room temperature for

5min and reported that after photopolymerization more

than 80% of the bioactivity of the therapeutic proteins was

retained. A photo-DSC device that had filtered light

(250–450 nm) with an intensity of 20mW � cm�2 was

used for photopolymerization. It is clearly seen that the

intensity of the UV light is applicable for clinical use and

not harmful for cells or tissues.

In vitro Studies of the Cell Seeded Hydrogels

The main objective of this study was to design a new

injectable polymer that could be used as a scaffold for cell

growth. For this reason, first the interaction between the

hydrogels and cells was investigated by using invert

microscopy. Images of the 3T3 mouse fibroblasts, which

were taken after 48 h on the Petri dishes where they

remained during culture, can be seen in Figure 6.

The growth of 3T3 mouse fibroblasts on the hydrogel

surface was monitored by environmental scanning

electron microscopy (ESEM). The ESEM technique makes

it possible to examine materials in their natural, uncoated

state under low vacuum. During in-situ testing of

uncoated samples, the absence of charge formation in

this system provides an obvious advantage. The sample

morphologies were investigated at room temperature to

reveal their natural surface characteristics by using a

special technique (gaseous low vacuummode) that allows

charge-free, gaseous secondary electrons.[36] Attachment

of fibroblasts on the hydrogels having no RGD peptide unit

was very poor; however, cell adhesion was significantly



Figure 6. Light microscope images of the fibroblasts after 48 hinteraction with hydrogel sample (FPEG20) on the Petri disheswhere they remained during culture.

Figure 8. SEM micrographs of 3T3 fibroblasts cultured on thesurface of sample FPEG20-2, 24 h following seeding a) �500and b) �9 000 magnification.

860

enhanced for the hydrogel that had RGD peptide units

(FPEG20-2) (Figure 7). It was observed that the cells did not

vanish or lose their original structural shapes, and they

also adhered to the surface of the hydrogels. This is strong

proof that the hydrogels that have RGD units are

biocompatible materials and do not exhibit a toxic

character. Also, in Figure 7 it can clearly be seen that

some of the cells have started to divide.

The interaction of FPEG20-6 with 3T3 cells was also

investigated by SEM to improve resolution. SEM images

show that the surface of the hydrogel with RGD peptide

was cell adhesive and a large number of fibroblasts on its

surface are observed (Figure 8a,b). As seen in Figure 8b, a

spheroid cell with one to two filapodial extensions at high

magnification demonstrate cell spreading. It is known that

cell adhesion involves a sequence of four steps: cell

Figure 7. ESEM images taken 24 h after cell seeding and in-vitrogrowth of 3T3 mouse fibroblasts. Hydrogel with RGD peptideunits (FPEG20-2)



attachment, cell spreading, organization of an actin

cytoskeleton, and formation of focal adhesions.[37]

To further establish the effect of hydrogel RGD peptide

content on the cell adhesion, the hydrogels with 0–6 wt.-%

RGD-PEGMAwere prepared. The compositions are given in

Table 1. Osteoblasts were seeded onto the sterilized

hydrogel surface. As shown in Figure 9a, after 24 h only

rounded cells were observed on the hydrogel having no

RGD. However, some cell spreading is seen on the hydrogel

with 2 wt.-% RGD. In the case of the 6 wt.-% RGD-

PEGMA-containing hydrogel, the surface image has

changed significantly and a lace-like cell layer is observed

throughout the surface. A similar observation was

previously reported by Burdick and Anseth.[38] They

observed a significant increase in cell number and cell

spreading upon increase in the RGD concentration. It is

also observed that, while the hydrogel with the highest

RGD content (FPEG20-6) exhibits a very porous structure

(not shown here), all the pores are closed after the cell

DOI: 10.1002/mabi.200700319


Figure 9. SEM micrographs of osteoblasts cultured on the surfaceof a) FPEG20, b) FPEG20-2, and c) FPEG20-6 samples, 24 hfollowing seeding.

Table 3. SEM-EDS analysis results of the FPEG20-6 hydrogelsurface before cell seeding and after cell seeding.

Surface Spectrum Total

C N O Na Cl Ca

Empty

surface

55.45 – 39.18 0.57 4.80 – 100.00

Cell-seeded

surface

58.32 10.12 27.78 0.56 2.94 0.29 100.00

seeding experiment. This is attributed to the good

adhesion and spreading of cells inside the pores. Pre-

viously, He et al. reported that the collagen-coated electro-

spun polylactide showed enhanced endotheliazation and



increased cell viability and attachment because of its

fibrillar morphology.[39] Moreover, Woo et al. reported that

cell adhesion to fibrous scaffolds was different from that to

solid-walled scaffolds. They showed that abundant long

slender shaped osteoblasts were formed and they were

very difficult to distinguish from the polymeric scaffold.[40]

Scanning electron microscopy and energy dispersive

spectrometry (SEM-EDS) was used to identify the chemical

composition or the distribution of elements on the hydro-

gel surface before and after cell seeding for sample

FPEG20-6. As shown in Table 3, it was determined that

the RGD-modified hydrogel surface includes quite high

amounts of nitrogen and calcium atoms. This result

reveals that nitrogen and calcium are produced on the

osteoblast-seeded surface of the scaffoldwithin the culture

time period. However, nitrogen and calcium deposition

was not detected on the RGD-containing hydrogel scaffold

without cells. It is well known that Type I collagen is the

predominant component of a complex secreted extra-

cellular matrix that supports the process of tissue

mineralization to form calcified bone.[41] Osteoblasts can

expand on the hydrogel surface and, due to the osteoid

activity, collagen secretion takes place which increases the

nitrogen percentage on the hydrogel surface and miner-

alization, which can be determined by SEM-EDS studies.

Although this paper presents in-vitro studies, future

research will focus on the in-vivo study of the hydrogels,

which will be placed into the tibea of rats.

Conclusion

New injectable fumarated PEGDA-based polymer net-

works are prepared by a photopolymerization technique.

The swelling ratio of the hydrogels is decreased with the

increase in the PEGDA content, which, on the other hand,

enhances the crosslinking degree. The viscosities of the

formulations were dependent on the PEGDA content,

which reduces the viscosity significantly because of its low

molecular weight. The fracture strength and compressive



862

modulus of the hydrogels tested in this study are also

affected by the PEGDA content. Cell growth experiments

showed that the RGD peptide unit facilitated the adhesion

of the cells to the hydrogel surface.

Acknowledgements: This work was supported by TUBITAK TBAGProject no: 105T254. The authors thank Assoc. Prof. Dr. YusufMenceloglu for 1H NMR measurements.

Received: December 5, 2007; Revised: March 19, 2008; Accepted:March 25, 2008; DOI: 10.1002/mabi.200700319

Keywords: cell growth; hydrogels; photopolymerization; RGD

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DOI: 10.1002/mabi.200700319

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