Jamie L Ifkovits, Robert F Padera and Jason A Burdick- Biodegradable and radically polymerized...

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IOP PUBLISHING BIOMEDICAL MATERIALS

Biomed. Mater. 3 (2008) 034104 (8pp) doi:10.1088/1748-6041/3/3/034104

Biodegradable and radically polymerized

elastomers with enhanced processingcapabilities

Jamie L Ifkovits1, Robert F Padera2 and Jason A Burdick1

1 Department of Bioengineering, University of Pennsylvania, Philadelphia, PA 19104, USA2 Department of Pathology, Brigham and Women’s Hospital, Harvard Medical School, Boston,

MA 02115, USA

E-mail: [email protected]

Received 26 November 2007

Accepted for publication 13 February 2008

Published 8 August 2008

Online at stacks.iop.org/BMM/3/034104

Abstract

The development of biodegradable materials with elastomeric properties is beneficial for a

variety of applications, including for use in the engineering of soft tissues. Although others

have developed biodegradable elastomers, they are restricted by their processing at high

temperatures and under vacuum, which limits their fabrication into complex scaffolds. To

overcome this, we have modified precursors to a tough biodegradable elastomer, poly(glycerol

sebacate) (PGS) with acrylates to impart control over the crosslinking process and allow for

more processing options. The acrylated-PGS (Acr-PGS) macromers are capable of

crosslinking through free radical initiation mechanisms (e.g., redox and photo-initiatedpolymerizations). Alterations in the molecular weight and % acrylation of the Acr-PGS led to

changes in formed network mechanical properties. In general, Young’s modulus increased

with % acrylation and the % strain at break increased with molecular weight when the %

acrylation was held constant. Based on the mechanical properties, one macromer was further

investigated for in vitro and in vivo degradation and biocompatibility. A mild to moderate

inflammatory response typical of implantable biodegradable polymers was observed, even

when formed as an injectable system with redox initiation. Moreover, fibrous scaffolds of

Acr-PGS and a carrier polymer, poly(ethylene oxide), were prepared via an electrospinning

and photopolymerization technique and the fiber morphology was dependent on the ratio of

these components. This system provides biodegradable polymers with tunable properties and

enhanced processing capabilities towards the advancement of approaches in engineering soft

tissues.

(Some figures in this article are in colour only in the electronic version)

1. Introduction

The well-known tissue engineering paradigm accounts for

the importance of scaffolds, cells, and growth factors and

combinations of these components for the successful design

and integration of constructs into living systems to enhance

tissue regeneration [1]. It is generally believed that cells either

delivered or from surrounding tissues receive necessary cues

from their microenvironment, which consists of both matrix(e.g., mechanics, chemistry) and soluble factors [2, 3]. With

this in mind, the chemical and physical properties of scaffolds

are of vital importance in controlling cellular behaviors (e.g.,

differentiation, matrix production) and in the overall success

of the construct [2, 4–9].

Scaffolds may comprise natural enzymatically degradablebiopolymers (e.g., hyaluronic acid) or synthetic polymers

(e.g., polyurethanes), which are typically biodegradable,

depending on the desired application and in vivo environment

[10, 11]. One advantage to using synthetic polymersis the ability to tailor scaffold mechanical properties

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Biomed. Mater. 3 (2008) 034104 J L Ifkovits et al

and degradation kinetics through chemistry and processing

[2, 8, 12]. Synthetic materials frequently have moduli of the

order of MPa and GPa [8, 9, 11, 13], whereas the elastic

modulus of many tissues is of the order of Pa and kPa [9, 13].

While hydrogels may have moduli of the order of many native

tissues [14], they are by definition hydrophilic materials,

which generally implies minimal protein adsorption, andconsequently, minimal cell attachment compared to their

hydrophobic polymeric counterparts [15]. Furthermore, many

tissues exhibit elasticity such that they canfunctionand recover

in the mechanically dynamic environment that exists in the

body [6, 10, 13, 16]. Therefore, investigators have been

motivated to synthesize and develop novel materials that better

mimic the stiffness and elasticity of native tissues [2, 6, 10,

13, 16–18].

Wang and colleagues [16] synthesized a tough

biodegradable elastomer, poly(glycerol sebacate) (PGS), that

has potential for the engineering of soft tissues due to its

mechanical properties and biocompatibility [13, 16, 19].

However, the curing of PGS requires high temperature and

vacuum conditions [16, 19], which makes processing into

complex scaffolds difficult and in vivo crosslinking impossible.

Recently, Nijst et al [18] reported the modification of

free hydroxy groups on the PGS prepolymer with acrylate

functional groups to form an acrylated PGS (Acr-PGS)

macromer. The introduction of the acrylate functionality

introduces control over the crosslinking and thus expands

upon the current processing options of this elastomer. We

also report here the modification of PGS precursors with the

acrylate functionality, but investigatethe role of bothmolecular

weight and degree of acrylate substitution on the formed

networks properties. Since crosslinking of the vinylic bondson the Acr-PGS can occur via both redox and photo-initiated

free radical polymerizations [20], we also explore these

different mechanisms for enhanced processing of the networks

(e.g., injectability and electrospinning). Additionally, these

networks could be crosslinked using Michael-type addition

reactions with the addition of a multifunctional nucleophile

(e.g., di-thiol), but that approach is not shown here.

Redox initiation is commonly used clinically to

polymerize poly(methyl methacrylate) (PMMA)-based bone

cements in vivo [21–23]. Generally, a bi-component system,

such as that of benzoyl peroxide and N,N-dimethyl-p-

toluidine, is used to promote initiation upon the mixingof the two components, allowing for injectable applications

[21–23]. With this process, it is possible for the minimally

invasive delivery of injectable polymer formulations to remote

areas of the body where light penetration may not be possible

[24]. Furthermore, injectable polymers can be formed directly

in a defect and good contact between the polymer and

surrounding tissue is possible [11, 25]. Therefore, injectable

polymers are useful for a wide range of applications including

drug delivery vehicles, tissue adhesives, and as tissue barriers

and scaffolds [25–28].

Photoinitiated polymerizations are also useful for a

wide range of applications, particularly due to the spatial

and temporal control afforded during processing [11, 29].Additionally, polymerizationexothermsand gelation timescan

be controlled by simply varying the initiator concentration and

light intensity [11]. These advantages of photopolymerization

make it possible for the formation of complex scaffolds in both

2D and 3D [30–32]. We recently used photopolymerization

and electrospinning to generate fibrous scaffolds with both

isotropic and anisotropic structures from biodegradable

polymers [33]. Electrospinning is a technique in which amat of continuous fibers is created by applying a voltage

to a polymer solution [34, 35]. Fibrous scaffolds may be

advantageous for tissue engineering due to their high surface-

to-mass ratio and their ability to mimic the native extracellular

matrix size and architecture [2, 34]. The ability to crosslink

these elastomers using free radical polymerizations now opens

up the possibility of processing into fibrous scaffolds.

The overall objective of this work was to develop

radically polymerizable macromers that form biodegradable

elastomeric networks upon crosslinking. To accomplish this,

a range of Acr-PGS macromers were synthesized and their

network properties (e.g., degradation and mechanics) were

characterized upon crosslinking. The reaction behaviors

for both redox and photo-initiated polymerizations and the

tissue response to these biodegradable elastomers, including

injectable formulations, were also investigated. Finally, the

processing of the Acr-PGS macromers into fibrous scaffolds

via electrospinning was explored.

2. Experimental details

2.1. Macromer synthesis and characterization

All reagents were used as received from Sigma-Aldrich

(St Louis, MO) unless otherwise noted. The prepolymer wasformed via the condensation reaction of equimolar amounts of

glycerol (ThermoFisher Scientific, Waltham, MA) and sebacic

acid. Thereagentswere combined andstirred at 120◦Cundera

nitrogen atmosphere for approximately 2 h and then a vacuum

of 19 mbar was applied for various amounts of time (26–

63 h) to obtain prepolymers of varying molecular weights.

For acrylation, the prepolymer was dissolved in methylene

chloride (ThermoFisher Scientific) containing triethylamine

(TEA, equimolar to acryloyl chloride) and 500 ppm

4-methoxyphenol (inhibitor, 10 wt% in methylene chloride).

Various molar ratios of acryloyl chloride (1:10 v/v

in methylene chloride) were dripped into the solution. Thesevalues were calculated using the estimation that two of the

three hydroxy groups present in glycerol reacted with the

sebacic acid and provide a range of overall acrylations.

An additional 500 ppm 4-methoxyphenol was added to the

reaction chamber and a rotary evaporator (40 ◦C, 450 mbar)

was used to remove the methylene chloride. Ethyl acetate

was added to the reaction flask and the solution was vacuum

filtered to remove the TEA salts and washed three times

with 10 mM hydrochloric acid (ThermoFisher Scientific).

Ethyl acetate was removed via rotovapping (40 ◦C, 99 mbar)

to leave a viscous liquid, which was redissolved in methylene

chloride and stored at 4 ◦C. The prepolymer and macromer

molecular weights and chemical structures were verifiedusing GPC (Waters GPC System, Milford, MA) and 1H

2




NMR spectroscopy (Bruker Advance 360 MHz, Bruker,

Billerica, MA) and the results are outlined in figure 1.

2.2. Reaction characterization

Macromers were mixed with 0.5 wt% of the photoinitiator

2,2-dimethoxy-2-phenylacetophenone (DMPA, 10 wt%

methylene chloride) and excess methylene chloride was

removed by rotovapping. The reaction behavior was

monitored using real time attenuated total internal reflection

Fourier transform infrared spectroscopy with a zinc selenium

crystal (ATR-FTIR, Nicolet 6700, Thermo Electron, Waltham,

MA). A sample of the macromer with photoinitiator was

placed directly on the surface of the crystal, covered with

a glass coverslip and monitored in real time with exposure

to ultraviolet light (1.5 mW cm−2, 365 nm Omnicure S100,

EXFO, Quebec). In the case of redox initiation, samples

were prepared by the addition of 1.0 wt% benzoyl peroxide

(BPO, 10 wt% in methylene chloride) or N,N-dimethyl-

p-toluidine (DMPT) to the macromer/methylene chloridesolution. Methylene chloride was removed by rotovapping

and the samples were dissolved in 200 proof ethanol and

transferred to a dual barrel syringe (PlasPak Industries, Inc,

Norwich, CT), which was subsequently placed into a 60 ◦C

oven overnight to remove any excess ethanol. Samples

were ejected from the dual barrel syringe directly onto the

surface of the crystal. Reaction conversion was determined

by monitoring the change in the vinylic double bond peak

(∼1635 cm−1); however, direct quantification was not possible

due to overlapping adjacent peaks and an unstable baseline.

The gelation times and reaction exotherms were quantified

using a slowly stirring stirbar (60 rpm) in a vial post-mixingand a thermocouple temperature probe.

2.3. Degradation and material property characterization

For sample fabrication, the macromer/initiator solutions were

poured into a 50 × 15 × 1 mm teflon mold and placed

in an oven at 60 ◦C overnight. The construct was then

covered with a glass slide and polymerized with exposure to

∼10 mW cm−2 365 nm ultraviolet light (Blak-Ray, Ultraviolet

Products, Upland, CA) for 10 min. Polymer disks (1 mm thick,

5 mm diameter) were punched from the resulting polymer

slabs. To monitor in vitro degradation, samples were weighed,

submerged in 150 mM NaCl PBS, and placed on an orbitalshaker at 37 ◦C. At each time point (2, 4 and 8 weeks) samples

(n = 3) were removed, lyophilized (Freezone 4.5, Labconco,

Kansas City, MO) for 24 h, and weighed to determine mass

loss. For mechanical testing, strips (15 × 5× 1 mm) were cut

from the slabs and tensile testing was conducted on an Instron

5848 mechanical tester (Norwood, MA) with a 500 N load cell

at a strain rate of 0.1% s−1.

2.4. In vivo tissue response

Animals were cared for according to a protocol approved

by the University of Pennsylvania Institute for Animal and

Use Committee. Photopolymerized polymer slabs wereprepared as described above. Polymer discs (1 mm thick,

5 mm diameter) were punched and submerged in ethanol.

Ethanol was evaporated off and the discs were placed under

a germicidal ultraviolet lamp for 30 min. Redox initiated

macromer solutions were loaded into a sterile dual barrel

syringe followed by exposure to the germicidal lamp for

30 min. Precrosslinked and preweighed discs (n = 4 per

time point, 4 discs per animal) were implanted subcutaneouslyinto the dorsal pocket of male Sprague–Dawley rats. Redox

initiator loaded macromer solutions were also injected (n= 4)

into the dorsal pocket of a male Sprague–Dawley rat. The

animals were sacrificed at various time points (2, 4 and

8 weeks) and the polymer samples and surrounding tissue

was collected and fixed with 10% formalin for 24 h. Standard

hemotoxylin and eosin (H&E) staining of paraffin embedded

sections was usedto investigate the tissue response. Additional

samples were removed to monitor in vivo degradation

behavior. All tissue was excised from the sample prior to

lyophilization to obtain the sample dry weight.

2.5. Electrospinning

An electrospun Acr-PGS scaffold was prepared by first

dissolving the photoinitiator/macromer in 90% ethanol

(50 wt%). To improve electrospinning potential, the

macromer/ethanol solution was combined with varying

percentages of 10% poly(ethylene oxide) (PEO) (200 kDa,

Polysciences, Warrington, PA) in 90% ethanol. Various

solutions containing different ratios of the Acr-PGS/ethanol

and PEO/ethanol were electrospun in a horizontal setup using

a flow rate of 2.5 mL h−1, distance to collection plate of

15 cm and a +15 kV applied voltage (ES30, Gamma HighVoltage, Ormond Beach, FL). The scaffolds were crosslinked

post-electrospinning using an ultraviolet lamp (Blak-Ray,

Ultraviolet Products, Upland, CA) in a nitrogen atmosphere.

Scaffolds were gold sputter coated and viewed using scanning

electron microscopy (Penn Regional Nanotech Facility, JEOL

6400 SEM, Tokyo, Japan).

3. Results and discussion

When designing a scaffold for tissue engineering, there

are several design criteria (e.g., mechanics, degradation,

biocompatibility) to keep in mind. Many believe that it

is important to closely match the biomaterial mechanical

properties with those of the surrounding native tissue to assist

in the gradual transfer of stresses from implant to the newly

formed tissue [7, 16]. Specifically, the elasticity of tissues

is often overlooked in material design, yet biodegradable

elastomers may fill that need. Elastomers are generally defined

as lightly crosslinked polymers that easily and quickly undergo

large, reversible deformations with complete recovery [20].

These important features of elastomers (e.g., PGS) make

them attractive materials to alleviate the compliance mismatch

problem that often exists with synthetic polymeric implants,

particularly in the dynamic environment of the human body

[16].

3




( A)

( B)

Figure 1. Synthetic scheme ( A) and representative 1H NMR spectra of Acr-PGS macromers ( B). Peak letters correspond to those in themacromer structure above. Unlabeled peaks correspond to protons from the initiator, inhibitor or unreacted glycerol (peak d).

Table 1. Summary of Acr-PGS macromers synthesized andinvestigated.

Macromer M n (kDa) M w (kDa) % acrylation

1 2.01 4.06 88.02 2.01 4.06 21.73 2.01 4.06 9.04 3.40 7.02 21.7

5 4.91 22.6 17.96 5.33 23.5 9.6

3.1. Macromer synthesis and network characterization

Our approach is to fabricate tissue engineering scaffolds

using a modified PGS prepolymer (i.e., Acr-PGS) that can be

crosslinked under mild and physiologic conditions. The first

step in the synthesis of the Acr-PGS macromer (figure 1( A))

is the combination of trifunctional glycerol and difunctional

sebacic acid via a polycondensation reaction in a 1:1 molar

ratio for varying amounts of time (26–63 h). The starting

reagents were chosen because they are naturally present in

the body and have been previously approved by the US Food

and Drug Administration for medical applications [16]. The

M w of the prepolymer as defined by GPC increased with

reaction time and ranged from ∼4.06 kDa to 23.46 kDa,

illustrating the tunability of molecular weight (table 1).

As with most condensation reactions, prepolymers were

polydisperse (2.01–4.60) and generallyincreasedwith reaction

time. Multifunctional Acr-PGS macromers were prepared by

reaction of the prepolymer with varying amounts of acryloyl

chloride. These amounts were defined assuming that two

of the three hydroxy groups on the glycerol reacted with thesebacic acid during the condensation reaction and were chosen

to provide a range of acrylations. The Acr-PGS% acrylations

were determined using 1H NMR (table 1).

As seen in figure 1, peaks at ∼1.3, 1.6 and 2.3 ppm

correspond to the protons in the olefin chain from sebacic acid

and multiplets at∼4.2and 5.2ppm correspond to theprotons in

the glycerol. The peaks at ∼5.9, 6.1 and 6.3 ppm correspond

to those of the functional acrylate group. The % acrylation

was determined by comparing the actual number of acrylate

groups with the theoretical values for 100% incorporation of

the acrylate group into the prepolymer. The % acrylation

values range from ∼9.6 to ∼88.0% (table 1). These six

Acr-PGS macromers represent a range of molecular weights

and acrylations, and thus, provide insight into the relationships

between macromer structure and network properties.

The introduction of the acrylate functional groups was

also confirmed using ATR-FTIR by visualization of the

characteristic absorption of the acrylate group at ∼1635 cm−1

(figure 2). In general, the intensity of this absorption increased

as the % acrylation increased. For example, figure 2( A)

displays the characteristic acrylate absorption for macromer 1and is representative of a high % acrylation (88.0%), whereas

figure 2( B) displays the characteristic acrylate absorption for

macromer 6 and is representative of a low % acrylation

(9.6%). A large difference in the acrylate absorption intensity

is observed when comparing the two spectra.

3.2. Network formation

The photopolymerization reaction was investigated by

introducing 0.5 wt% photoinitiator (DMPA) into the Acr-PGS

macromer and monitoring the consumption of the acrylate

group peak in real time with exposure to 365 nm ultravioletlight (figure 2( A)). The maximum conversion occurred after

4




( A)

( B)

Figure 2. Consumption of the acrylate peak with time duringphotopolymerization of macromer 1 (0.5 wt% DMPA, light intensity∼1.5 mW cm−2) ( A). Consumption of the acrylate peak with timeduring the redox initiated polymerization of macromer 6 (1.0 wt%

BPO and DMPT) ( B).

∼8 min. The large difference in acrylate absorption between

initial and final time points and the near baseline level at the

final time point indicate a high level of conversion of the

acrylate group to crosslinks. The redox-initiated crosslinking

was also monitored by introduction of 1.0 wt% of the

bi-component BPO and DMPT initiation system, which is

commonly used in bone cements [21–23] (figure 2( B)). The

maximum reaction conversion occurred after∼20 min. Again,

the difference in intensity between the initial and final time

points indicates high conversion values.

The gelation time was defined as the point when a slowly

stirring stirbar was stopped after injection of the macromer

into a vial. For this system, gelation occurred at ∼5 min

and a maximum temperature of ∼30 ◦C (starting from room

temperature) was observed. The maximum conversion and

gelation time can be tailored by altering the amount of initiator

incorporated into the macromer system, as is the case with

similar bi-component initiator systems [24, 36], depending

on the application. For example, delivery of this material

to the heart through a catheter might require slower gelation

than direct injection into a defect. The minimal increase in

temperature is also important if this polymer is to be used

as an injectable formulation to prevent temperature-induced

tissue necrosis.

( A)

( B)

Figure 3. Representative tensile stress vs. elongation plots fornetworks formed from the Acr-PGS macromers with various MWsand acrylations ( A). Young’s modulus (black) and % strain at break (white) for networks formed from the various synthesizedmacromers ( B). Further information on the macromers can be found

in table 1.

3.3. Network mechanical properties

Polymer slabs for mechanical and degradation analysis were

prepared using photopolymerization. Typical tensile stress

versus % elongation relationships for networks formed from

Acr-PGS macromers are shown in figure 3( A). It is important

to note that many of the samples broke at the clamp, and

thus, could lead to lower than actual values for the %

strain at break. Young’s modulus was determined from the

slope of the linear portion of the plot (<20% strain) and

varied (∼0.15–30 MPa) depending on the Acr-PGS macromer(figure 3( B)). The % strain at break also varied (∼5–200%)

depending on the Acr-PGS macromer used for network

formation. In general, Young’s modulus increased as the

degree of acrylation increased for a given molecular weight.

As expected, the % strain at break increased as Young’s

modulus and % acrylation decreased for a given molecular

weight. Furthermore, Young’s modulus and % strain at

break increased with increasing molecular weight for similar

degrees of acrylation. As seen in figure 3( B), macromers 3

and 6 have similar Young’s moduli (∼150 kPa) but their %

elongation at break varied by almost an order of magnitude.

Based on these results, it important to note that not all of

the macromers formed elastomeric networks. Therefore,

5




Figure 4. In vivo (black) and in vitro (white) degradation results at2, 4 and 8 weeks for networks formed from macromer 6.

macromer 6 was selected for the remaining studies since its

mechanical properties are the most elastomeric.

Relationships between macromer structure and network properties can easily be drawn from these data. For instance,

increased % acrylation leads to an increase in the number

of crosslinks formed, which is associated with an increase

in the modulus of a resulting material and a decrease in

the ability to elongate before failure. Additionally, more

elastomeric-like features are obtained as the molecular weight

of the prepolymer is increased. This is a clear demonstration

that small modifications during synthesis can lead to drastic

differences in the properties of the resulting material. Thus,this same backbone chemistry can be used to develop materials

suitable for a wide range of applications. For example, a more

elastic and softer material (e.g., Young’s modulus∼

150 kPa)may be more ideal for cardiac tissue engineering, whereas

a less elastic and a stiffer material (e.g., Young’s modulus

∼30 MPa) may be more ideal for bone tissue engineering.

Although not investigated here, dynamic fatigue testing would

be necessary to illustrate the potential of these materials to

withstand the dynamic in vivo environment (e.g., beating of

the heart).

3.4. Network degradation and in vivo tissue response

The in vivo and in vitro mass loss of photopolymerized samples

reached a maximum of ∼37% and ∼33%, respectively, at

8 weeks (figure 4). The in vivo mass loss was potentiallygreater because of its location in a more dynamic environment

where there is more fluid exchange to remove any degradation

products from the implant region and due to the presence

of enzymes compared to the in vitro environment. Also,

the sample preparation was slightly different due to the

sterilization of the in vivo samples. Based on these mass

loss data, it is suspected that this material would be suitablefor a variety of tissue engineering applications. Since redox-

intiated samples could not be massed prior to in vivo injection,

the degradation profile could not be monitored. However, it

is anticipated that it would be comparable to its photoinitiated

counterpart if similar conversions are reached.

After two weeks of implantation, the host reaction tothe polymer discs comprised granulation tissue, with new

( A)

(C )

( B)

( D)

Figure 5. H&E staining for in vivo tissue response to networksformed from macromer 6 at 2 weeks ( A), 4 weeks ( B), 4 weeks viainjectable redox initiation (C ) and 8 weeks ( D). (T: tissue; P:polymer; scale bar = 100 µm).

blood vessels, loose connective tissue formation and mild

chronic inflammation. Macrophages and foreign body giant

cells were present at the polymer-tissue interface (figure 5).

At the 4 and 8 week time points, a thin fibrous capsule is

present around the implant with minimal associated chronic

inflammation (figures 5( B) and ( D), respectively). There isno evidence of inflammation or necrosis within the adjacent

subcutaneous fibroadipose tissue, skin adnexal structures or

deep skeletal muscle. This represents a typical host response

to a biocompatible material. The reaction to the injected

polymer (figure 5(C )) contained slightly more perivascular

chronic inflammation in the surrounding host tissue, but

without evidence of necrosis or tissue damage. This slight

difference in response to the injected polymer may be due to

a mild toxicity associated with the initiators or differences in

the polymer configuration and surface area as compared to the

photoinitiated samples.

3.5. Electrospinning into fibrous scaffolds

Electrospinning has gained much attention in recent years

as a method for generating fibrous scaffolds [34, 35].

Fibrous scaffolds are thought to be advantageous since they

closely mimic the architecture and size-scale of the native

extracellular matrix [2, 34]. Many synthetic and natural

polymers have been successfully electrospun to date [34, 37];

however, the need to electrospin more polymers with a range

of material properties (e.g., mechanics and degradation),

especially those of elastomers, still exists. Due to the low

molecular weight and high polydispersity of Acr-PGS, it was

necessary to modify techniques to electrospin the macromers.Tan and colleagues recently used photopolymerization and

6




( A)

( B)

(C )

Figure 6. SEM images of electrospun Acr-PGS:PEO scaffolds atratios of 30:70 ( A), 40:60 ( B) and 50:50 (C ). Scale bar = 50 µm.

electrospinning to create isotropic and anisotropic scaffolds

from low molecular weight biodegradable macromers by

using 200 kDa PEO as a carrier polymer [33]. The

same processing protocol was followed to prepare mats of

electrospun and photocrosslinkable Acr-PGS/PEO containing

a photoinitiator. Mats were electrospun at various ratios of

Acr-PGS to PEO solutions and crosslinked with ultraviolet

light prior to visualization under SEM (figure 6). A ratio of

30:70 PGS:PEO (figure 6( A)) produced a mat with the best

mechanical integrity and most distinct fibers, which is lost

when the PEO concentration is too low (figures 6( B) and (C )).

At this point, a thorough characterization of mechanics and

cellular interactions has not been performed on these scaffolds,

but our ability to obtain the proper structures motivates further

exploration of these materials.

4. Conclusions

In this study, radically polymerized networks with tunable

mechanical properties were successfully synthesized and

characterized. Notably, an increase in Young’s modulus with

increasing acrylation, as well as an increase in the % strain

at break with increasing molecular weight were observed,

indicating that these properties canbe tuned through the design

of the macromer and that networks with elastomeric properties

can be obtained. The reaction behavior was rapid and

reached high conversions with both redox and photoinitiated

polymerizations. The networks also degraded more rapidly

in vivo and only mild inflammation was seen, even with

injectable formulations. Moreover, this polymer system could

be processed into fibrous scaffolds using electrospinning and

PEO as a carrier polymer. This biodegradable and elastomericpolymer system can be further explored for the engineering

of numerous tissues where elasticity is an importantparameter.

Acknowledgments

The authors would like to acknowledge Joshua S Katz for

helpful synthetic discussions and Cindy Chung for assistancewith animal surgeries. This work was funded by the AmericanChemical Society Petroleum Research Fund, a pilot grant fromthe National Science Foundation Materials Research Scienceand Engineering Center at the University of Pennsylvania, andan Ashton Fellowship to JLI.

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Jamie L Ifkovits, Robert F Padera and Jason A Burdick- Biodegradable and radically polymerized...

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