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Biomaterials 43 (2015) 61e70

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Biomaterials

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Mitigation of hypertrophic scar contraction via an elastomericbiodegradable scaffold

Elizabeth R. Lorden a, Kyle J. Miller b, Latif Bashirov b, Mohamed M. Ibrahim b,Ellen Hammett a, Youngmee Jung c, Manuel A. Medina b, Ali Rastegarpour b,Maria A. Selim d, Kam W. Leong a, e, **, Howard Levinson b, d, *

a Department of Biomedical Engineering, Duke University, Box 90281, Durham, NC 27708, USAb Division of Plastic and Reconstructive Surgery, Department of Surgery, Duke University Medical Center, Box 3181, Durham, NC 27710, USAc Biomedical Research Institute, Center for Biomaterials, Korea Institute of Science and Technology, Seoul 136-791, Republic of Koread Department of Pathology, Duke University Medical Center, Box 3712, Durham, NC 27710, USAe Department of Biomedical Engineering, Columbia University, New York, NY 10027, USA

a r t i c l e i n f o

Article history:Received 18 October 2014Accepted 9 December 2014Available online 7 January 2015

Keywords:Animal modelBiodegradationBurnCollagenMechanical propertiesMicrostructure

* Corresponding author. Duke University Medical C27710, USA. Fax: þ1 919 681 7340.** Corresponding author. 351 Engineering TerraAmsterdam Avenue, Columbia University, New York,854 8725.

E-mail addresses: kwl2121@columbia.edu (K.W. Lduke.edu (H. Levinson).

http://dx.doi.org/10.1016/j.biomaterials.2014.12.0030142-9612/© 2014 Elsevier Ltd. All rights reserved.

a b s t r a c t

Hypertrophic scar (HSc) occurs in 40e70% of patients treated for third degree burn injuries. Current burntherapies rely upon the use of bioengineered skin equivalents (BSEs), which assist in wound healing butdo not prevent HSc contraction. HSc contraction leads to formation of a fixed, inelastic skin deformity.We propose that BSEs should maintain their architecture in the wound bed throughout the remodelingphase of repair to prevent HSc contraction. In this work we study a degradable, elastomeric, randomlyoriented, electrospun micro-fibrous scaffold fabricated from the copolymer poly(L-lactide-co-ε-capro-lactone) (PLCL). PLCL scaffolds displayed appropriate elastomeric and tensile characteristics for im-plantation beneath a human skin graft. In vitro analysis using human dermal fibroblasts demonstratedthat PLCL scaffolds decreased myofibroblast formation as compared to an in vitro HSc contraction model.Using a validated immune-competent murine HSc contraction model, we found that HSc contraction wassignificantly greater in animals treated with standard of care, Integra, as compared to those treated withcollagen coated-PLCL (ccPLCL) scaffolds. Finally, wounds treated with ccPLCL were significantly less stiffthan control wounds at d30 in vivo. Together, these data suggest that scaffolds which persist throughoutthe remodeling phase of repair may represent a clinically translatable method to prevent HSccontraction.

© 2014 Elsevier Ltd. All rights reserved.

1. Introduction

Dermal scarring affects more than 80 million people worldwideannually [1]. For example, over 4.4 million people are injured inmotor vehicle accidents, thousands of our nation's warriors arewounded in military excursions, and over 2.4 million patients areburned [2,3]. TheWorld Health Organization states that “there is nodoubt that the social and medical costs of burns are significant.

enter, Box 3181, Durham, NC

ce, Mail Code 8904, 1210NY 10027, USA. Fax: þ1 212

eong), howard.levinson@dm.

Economic impact of burns includes lost wages, and the costs relatedto deformities from burns, in terms of emotional trauma and lostskill.” In severe burns (~28,000 patients/year in the United States),the incidence of hypertrophic scar (HSc) is 40e70% [1,4]. HSc arefirm, raised, red, itchy scars that are disfiguring and can have a se-vere impact on quality of life. 70% of HSc occur across joints or otherareas of high tension in the body, resulting in a scar contracturesthat restrict range-of-motion [5]. Current preventative therapies forscar contracture are ineffective, and patients requiring interventionundergo at least four corrective surgeries on average [1,5].

HSc develops during the first 4e8 weeks following injury andcontinues to mature and contract throughout the remodeling phaseof repair for as long as six months post trauma; however,commercially available bioengineered skin equivalents (BSE)degrade and are remodeled in thewound bed within 1e4 weeks [6].There are currently three BSEs on themarket which are approved for

E.R. Lorden et al. / Biomaterials 43 (2015) 61e7062

third degree burn treatment by the Food and Drug Agency as well asthe Center for Medicare and Medicaid Services: Epicel, TransCyte,and Integra. Of these, Integra is the most widely used for treatmentdue to its cost and ease of use. Integra Dermal Regeneration Tem-plate (Integra LifeSciences, Plainsboro, NJ) is a bovine collagen-basedlyophilized matrix covered by a silicone dressing. The biologic layerremains in the wound bed, while the silicone layer is removed priorto application of skin graft. Previous work studying collagen-basedscaffolds for dermal regeneration has shown that prolonging thehalf-life of the material in the wound bed has a profound impact onminimizing wound contraction; however, the longest scaffold half-life tested in these studies was 2e4 weeks [7]. We hypothesizethat to improve mitigation of HSc, the scaffold should be present inthe wound bed throughout the remodeling phase of repair whenHSc occurs.

In comparison to biologic biomaterials, synthetic biomaterialshave advantages in that they possess tunable mechanical proper-ties and biodegradation rates. In searching for a suitable syntheticelastomer for this application, we looked for favorable character-istics of mechanical strength, elasticity, and biodegradability. Weselected the copolymer poly(L-lactide-co-ε-caprolactone)s (PLCL),synthesized from a 50:50 ratio of poly(lactic acid) (PLA) and poly(ε-caprolactone) (PCL) [8e12]. The FDA-approved PLA and PCL havebeen extensively studied for tissue engineering applications [13].Although neither PLA nor PCL is elastomeric, PLCL displays elasto-meric characteristics due to the phase separation of the crystallinePLA and the amorphous PCL segments, creating hard and soft do-mains somewhat akin to that observed in elastomeric poly-urethanes [8e10]. Synthetic materials have additional advantagesover biologics with respect to ease of handling, long shelf-life, lowcost, and well-defined physicochemical properties.

Application of mechanical load has been shown to initiate HScformation in mice [14]. Thus, it has been suggested that stress-shielding cells from transmission of mechanical load could assistin mitigating HSc [15]. Mechanical tension and inflammatory cy-tokines (primarily transforming growth factor-beta (TGFb))secreted following dermal injury cause resident fibroblasts todifferentiate into myofibroblasts [16]. Myofibroblasts are distin-guished from fibroblasts by (1) the presence of a contractileapparatus, similar to that of smooth muscle cells, and (2) the neo-expression of an actin isoform found in vascular smooth musclecells, a-smooth muscle actin (aSMA) [17]. When de novo aSMA isincorporated into stress fibers, myofibroblasts produce strongcontractile force and physically contract the wound or scar bed. Toachieve tissue level contraction, single myofibroblasts join stressfibers at sites of adherens junctions, resulting in cytoskeletalalignment and the formation of a coordinated cellular syncytium[18]. This conglomeration allows myofibroblasts to multiply theircontractile forces along the axis of cell alignment, and coincideswith the alignment of the extracellular matrix (ECM). This isobserved in scar tissue as linear arrays of ECM, as compared torandomly oriented ECM in uninjured skin.

To encourage random cell alignment, and hence disordered ECMdeposition, we electrospun PLCL into a randomly oriented micro-fibrous scaffold [19e21]. We first studied the behavior of humandermal fibroblasts on this scaffold using the fibroblast populatedcollagen lattice (FPCL) assay (22), which has been used to modelHSc contraction [16,22,23]. This was followed by in vivo evaluationusing a murine HSc contraction model recently established in ourlaboratory; our model is the only immune-competent murinemodel that mirrors the human condition in terms of causality [24].Together our in vitro and in vivo data suggest that slowly degradingelectrospun scaffolds prevent myofibroblast activation associatedwith HSc, and provide the necessary longevity to prevent HSccontraction in vivo.

2. Materials and methods

2.1. Synthesis of PLCL

PLCL (50% LA, 50% CL) was synthesized as described elsewhere [8]. Briefly, L-lactide (100 mmol; Purac; Lincolnshire, IL, USA) and ε-caprolactone (100 mmol;Sigma; St. Louis, MO, USA) were polymerized at 150 �C for 24 h in the presence ofstannous octoate (1 mmol, Sigma) as a catalyst. After being dissolved in chloroform,the polymer was precipitated in methanol, then dried under a vacuum for 72 h andstored in vacuum pack at �20 �C.

2.2. Fabrication & analysis of electrospun PLCL

PLCL scaffolds were fabricated using continuous single fiber electrospinning todeposit a 3D matrix of fibers on a rotating grounded mandrel using a customspinning apparatus. PLCL was dissolved 14% (w/w) overnight in dichloromethane.Random fibers were spun at a flow rate of 3 ml/h with a voltage of 8 kV at a distanceof 13 cm from the mandrel, which was rotating at ~70 revolutions per minute.Ambient temperature was 22 �C with 43% humidity. Following spinning, fibers wereremoved from the mandrel and residual solvent was removed by air drying for 72 h.Fiber characteristics and scaffold thickness were analyzed using scanning electronmicroscopy (FEI XL30 SEM-FEG, Hillsboro, OR, USA).

2.3. Oxygen plasma treatment & collagen coating methods

Samples were placed inside of a plasma asher (Emitech K-1050X, Montigny-le-Bretonneu, France) and treated with reactive oxygen plasma for 45 s at 100 W toimprove hydrophilicity prior to cell culture and prepare for covalent collagencoating. Following treatment, samples were immediately immersed in sterile waterand subsequently sterilized in 70% ethanol for 20 min. Samples were rinsed thor-oughly with water following sterilization. Covalent collagen coating was performedby EDC/NHS chemistry as previously described [25]. This well-characterizedmethodis biocompatible, non-cytotoxic, and does not include a linker-arm. Carbodiimide isnot incorporated into the covalent-linkage, allowing the collagen to directly coatscaffold. This method does not modify scaffold morphology and generates a uniformcollagen coating covering fibers throughout the depth of the scaffold. Scaffolds incollagen coated PLCL (ccPLCL) groups were covalently coated with bovine type-1collagen (Nutragen, Advanced Biomatrix, San Diego, CA, USA) prior to in vivo im-plantation. Contact angle analysis before oxygen plasma treatment, after oxygenplasma treatment, and after collagen coating treatment was carried out on PLCLfilms using a goniometer as previously described [26].

2.4. Permeability measurement

In order to examine the impact of collagen coating on scaffold permeability,effective hydraulic permeability of PLCL and ccPLCL scaffolds was measured ac-cording to ASTM protocol F2952 using a custom-built flowmeter as previouslydescribed [27]. In brief, a “scaffold sandwich”was constructed between two siliconegaskets using a 100 mm scaffold and a fine stainless steel mesh. The “sandwich”wasplaced between mount pieces and a watertight seal was formed by applying lightpressure using a threaded screw housing unit. A 50 mL pipette was suspendedhorizontally 32 cm above the flow chamber and connected to the specimen mountusing tubing. Phosphate buffered saline (PBS) fluid flowwas equilibrated for 15 minbefore timing, and was measured at 1 mL intervals for the first 5 mL, then 5 mLintervals for the next 10 mL in order to ensure consistent flow. The total duration offlow (15 mL) was measured and used to obtain the Darcy coefficient, t, and averagepore diameters as previously described [27].

2.5. Static tensile testing of electrospun scaffolds as compared to skin and scartissues

Static tensile testing was carried out as described in ASTM D3822-07. Humanand murine skin and scar tissues were tested as previously described [24]. In breif,scar and uninjured skin samples were gathered from human and murine donors.Uninjured murine tissue samples were collected from the dorsum of 10e12 week-old C57BL/6 mice. Murine scar tissue was taken from contracted d30 skin graftedmice. Tissue from five murine donors was used, with three biological replicates perdonor. Human skin samples were donated from discarded human tissue from DukeHospital operating rooms under exemption by Duke institutional review board.Uninjured human skin was gathered from breast resection, while scar tissue wastaken from excised keloid, radiated forearm, and rejected skin graft. Tissue fromthree human donors was used with 3e5 biological replicates per donor. All humanand murine tissues were kept moist on damp gauze between collection and me-chanical testing, and analyzedwithin 1e5 h of collection. Prior to testing, underlyingtissue was removed and samples were cut into uniform strips. PLCL, ccPLCL, Integra,and tissue samples were cut to 5 cm � 5 mm strips using a scalpel and surgicalscissors and loaded with a 5 mm gap between clamps. Samples were analyzed on amicrostrain analyzer (MSA) (RSA II, TA Instruments, New Castle, DE, USA) at a rate of0.1 mm/s at room temperature (23 �C) until failure. The initial elastic modulus(within the first 0e200% strain) was analyzed for each sample. The lower elasticmodulus was selected for analysis because this strain range best mimics strains that

E.R. Lorden et al. / Biomaterials 43 (2015) 61e70 63

are experienced in the body. Ultimate tensile strength and elongation at break weretaken from the data set of each sample.

2.6. Fatigue testing of PLCL and ccPLCL scaffolds

Fatigue properties of scaffolds were measured to determine whether the sam-ples would be able to withstand continuous cyclic loading, such as that encounteredacross joints in the body. PLCL and ccPLCL samples were loaded on the MSA andimmersed in 37 �C PBS using a heated sample cup. Samples were allowed toequilibrate for 1 h, then extended to 10% strain, and oscillated over 10% strain at 1 Hzfor 24 h according to ASTM protocol D3479/D3479M-12.

2.7. Cell culture

Adult normal human dermal fibroblasts (NHDFs) (Lonza, Basel, Switzerland)were cultured in high glucose Dulbecco's Modified Eagle's Medium (GIBCO 11960-044) (Invitrogen, Grand Island, NY, USA) supplemented with 10% Premium Selectfetal bovine serum (FBS) (Atlanta Biologicals, Lawrenceville, GA, USA), 25 mg/mLgentamicin (Invitrogen), and 1� GlutaMAX, non-essential amino acids, sodiumpyruvate, and b-mercaptoethanol (Invitrogen), at 37 �C and 5% CO2. NHDFs werepassaged a maximum of six times prior to experimentation.

2.8. Contraction studies

Media was prepared fresh on d1 (10% FBS þ 5 ng/mL TGFb), fresh media wasadded at d2.5, and the trial was stopped at d5 for analysis. Following sterilization,PLCL and ccPLCL scaffolds (8mm round, 60 mm thick) were washed with PBS and leftin 10% FBSmedia for 24 h prior to cell culture studies. Prior to seeding, scaffolds wererinsed thoroughly with PBS, placed in a 48 well suspension plate, and seeded with200,000 cells in 500 mL media. Cells were allowed to attach for 1 h before addition of500 mL media.

FPCLs were constructed as previously described [28,29]. In brief, FPCLs wereprepared in triplicate by combining the following materials in the correspondingorder: 50 mL of 5xPBS (pH 8.5, 23 �C) was combined with 200 mL bovine type Icollagen (6 mg/ml, pH 5, 23 �C) (Nutragen, Advanced Biomatrix, San Diego, CA, USA),and 750 mL cell suspension (590,000 cells/mL). FPCLs were cast in small, pre-warmed, triplicate 35 mm � 10 mm tissue culture polystyrene dishes (250mL/dish). Dishes were placed in a 37 �C cell culture incubator for 1 h to allow FPCLs tosolidify before slowly adding 2.5 mL media. FPCLs were allowed to remain attachedto the cell culture dish throughout the study [30].

2.9. Viability analysis

Cellular viability analysis was performed using LifeTechnologies Live/DeadViability/Cytotoxicity Kit for Mammalian Cells (Invitrogen). Live/dead staining on d5NHDF-seeded PLCL and ccPLCL scaffolds was conducted according to manufacturerinstructions. d5 scaffolds were mounted and imaged on an inverted fluorescentmicroscope (Eclipse TE2000-U, Nikon, Tokyo, Japan) within 3 h. Images wereanalyzed by counting labeled cells using Image J software (NIH).

2.10. Immunocytochemistry

Cells were fixed with 4% paraformaldehyde in PBS for 15 min at room temper-ature, permeabilized, and stained in a blocking solution containing the antibodies,0.03 g/ml bovine serum albumin (BSA, Sigma), 10% goat serum (Sigma), and 0.3%Triton X-100 (Sigma) in PBS. Samples were blocked for 3 h prior to staining. Primarystain for aSMA (1:100, Abcam, Cambridge, MA, USA) was conducted for 18 h at 4 �C,the samples were then washed before incubation with Alexa Fluor 594 anti-mousesecondary antibody (1:200, Invitrogen) for 4 h at room temperature. Cell nuclei werestained with 4,6-Diamidino-2-phenylindole (DAPI) (1:5000, Invitrogen) and theactin cytoskeletonwas stained with phalloidin 488 (1:200, Invitrogen). The sampleswere then mounted in Fluoro-Gel (Electron Microscopy Sciences, Hatfield, PA, USA)for fluorescent imaging.

Images for analysis of aSMA presence in immunocytochemistry were acquiredusing an inverted confocal microscope (Zeiss LSM 510, Oberkochen, Germany) with10� magnification. Three images from each of three replicates were collected foreach condition, resulting in a minimum of 2000 cells for each condition for analysisusing Image J software. Nuclei quantification and area of aSMA positive stress fiberswere analyzed using a size exclusion to disregard non-cell debris and aSMA local-ized to the cytosol or focal adhesions. Quantification of aSMA positive stress fiberswas achieved by dividing the area of aSMA by the number of nuclei present in theimage and normalizing to staining in FPCLs as 100.

3. Surgical methods

Female C57BL/6 mice, 10-12 weeks-old, weighing 18e23 g(Jackson Laboratories, Bar Harbor, ME), were used as wild typemicethroughout the study. Mice were housed individually for the firstthree weeks, and in groups of four for the final week of study. Fe-male mice were selected to reduce the risk of aggression-related

injury when transferred to group housing; only one sex was usedin order to reduce secondary variables related to hormonal differ-ences between male and female mice. Mice were housed underprotocols approved by the Institutional Animal Care and UseCommittee of Duke University. Surgical methods were performedas described in our previous work; in all instances, skin graft refersto the use of a split thickness skin graft fashioned from the skintissue of two donor mouse ears [24]. In brief, a third degree burnwas created on the dorsum of the donor mouse, the burn was leftfor 3d, and a 14 mm diameter circle over the burn site was excisedfor recipient skin grafting. Treatment groups included: (1) skingraft without placement of scaffold material, (2) skin graft overIntegra, (3) skin graft over PLCL, (4) skin graft over ccPLCL. Scaffolds110 mm in thickness ± collagen coating were cut into circles using a1.2 cm diameter arch steel hole punch (VWR International, Radnor,PA) prior to sterilization. Integra was cut using a 1.2 cm punchimmediately before surgery and removed from the silicone backing.Integra, sterile PLCL, or sterile ccPLCL scaffolds were laid into thewound bed and sutured just beneath the skin edges with fourmattress stitches of 6-0 silk suture. Donor ear skins were laid overthe excised burn wound after the implantation of the relevantscaffold to fashion a skin graft. The murine skin grafts were thensecured with a padded bolster. The bolster was removed on post-operative d3.

3.1. Analysis of electrospun scaffold mechanical properties anddegradation on d30 in vivo

The mice were euthanized and tissue was collected on post-operative d30 in all mice. The collected tissues from d30 micetreated with ccPLCL, Integra, or skin graft alone (scar tissue) werecut into three pieces (Fig. 5A). The two peripheral tissue specimenswere preserved in 10% formalin and embedded in paraffin wax forhistological analyses. The center piece was immediately taken toMSA for static tensile testing. Tissue explants were kept moist be-tween collection and mechanical testing, and analyzed within 2 hof collection.

PLCL was extracted from d30 excised tissue samples by incu-bation with chloroform overnight and tested using nuclear mag-netic resonance (NMR) and gas permeation chromatography aspreviously described [8]. The component composition of PLCL wasanalyzed using a 400 MHz 1H NMR (Varian, USA). Spectra wereobtained using 1% (w/v) solutions in CDCL3 and the compositionscalculated from these relative intensities. Molecular weights weredetermined using a gel permeation chromatography (GPC, ViscotekTDAmax, Malven Instruments Ltd, UK). Chloroformwas used as themobile phase at a flow rate of 1 mL/min. Calibrationwas performedusing polystyrene standards to determine number-average andweight-average molecular weights (Mn and Mw).

3.2. Immunohistochemistry

Routine staining was carried out as previously described [24].Primary antibodies used included: F4/80 (1:1500, eBioscience, SanDiego, CA, USA) and anti-CD31 (1:50, Abcam). Secondary antibodiesincluded: biotinylated rabbit anti-rat (1:200, Vector Laboratories,Burlingame, CA, USA), biotinylated goat anti-rabbit (1:50, VectorLaboratories), avidin-biotin complex reaction (Vector Laboratories),DAB substrate solution (BiocareMedical, Concord, CA, USA). Stainedslides were visualized by use of a Nikon eclipse E600 microscopeand images were captured with a Nikon DXM 1200 digital cameraunder the same settings. Quantification of CD31 images was per-formed using five HPF images per section across a minimum of fivemice. Only positively stained vessels inside the perimeter of thescaffolds were quantified. Fiber diameter of ccPLCL scaffolds at d30

E.R. Lorden et al. / Biomaterials 43 (2015) 61e7064

in vivo was calculated using H&E images, one HPF per mouse and aminimum of ten measurements were used.

4. Statistical analysis

Unless otherwise stated, the following applies to all experi-mental results. Gaussian data have been presented asmean ± standard error of the mean. Two-way ANOVA followed bystudents t-test was carried out to discern differences betweengroups, with significance at p < 0.05. In the case of non-Gaussianresults, data were presented in box and whisker format. Signifi-cance was analyzed via two-way ANOVA followed by Man-neWhitney test. Mechanical and permeability analysisexperiments were carried out with n� 5, in vitro experiments werecarried out in triplicate at least three separate times, and in vivoexperiments were conducted with n � 5 per treatment group.

5. Results

5.1. Scaffold characteristics

Scaffolds were spun to thickness of 60 ± 10 mm for in vitrostudies and 100 ± 11 mm for in vivo studies; fiber diameter was5.6 ± 0.70 mm for all samples (Fig. 1A,B). We chose to use a fiberdiameter in this range because it has been associated with anti-fibrotic events in coronary tissue [31]. Covalent collagen coatinguniformly covered fibers throughout the depth of the scaffolds(Fig. 1C) and did not significantly modify fiber diameter(6.5 ± 0.66 mm) or scaffold morphology (Fig. 1D). Contact angleanalysis before oxygen plasma treatment, after oxygen plasmatreatment, and after collagen coating treatment on PLCL filmsshowed progressive hydrophilicity of 72 ± 6�, 60 ± 1�, and 44 ± 8�,respectively (Fig. 1E).

Fig. 1. Characterization of PLCL Scaffolds. Scaffolds were imaged using SEM; fiber diametermeasured via cross-sectional imaging (B). Scaffolds were covalently coated with Bovine tycroscopy (C) to analyze coating efficiency. Collagen coating did not fill pores or alter scaffoldPLCL films, following collagen coating films were significantly more hydrophilic than untreatDarcy coefficient of permeability was not significantly different between untreated PLCL an

5.2. Permeability analysis

PLCL scaffolds exhibited a mean Darcy permeability constant, t,of 20.0 ± 1.2 mm2, compared to 19.7 ± 1.6 mm2 for ccPLCL scaffoldswith no statistically significant difference (Fig. 1F). The corre-sponding average pore sizes were 4.54 ± 0.13 mm and4.50 ± 0.18 mm, respectively. Similarly, measured pore size usingSEM images found an insignificant difference in average pore size ofPLCL scaffolds (40.8 ± 2.8 mm) and ccPLCL scaffolds (36.3 ± 2.3 mm),further confirming that the collagen coating did not affect scaffoldpore size. The discrepancy between derived average pore size andSEM measured pore size has been previously described whencalculating pore size of electrospun scaffolds using this technique[27]. Sell et al. postulated that these difference may stem from thepresence of “faux pores” and blind pouches within the scaffoldswhich are not visible via SEM measurement but affect the effectivepore size gathered from flowmeter measurement.

5.3. Mechanical properties

The mechanical properties of PLCL and ccPLCL scaffolds wereanalyzed using MSA and compared to those of human skin, humanscar, and Integra (Fig. 2A). The elastic modulus of uncoated PLCL(3.6 ± 0.22 kPa) was significantly lower than that of Integra(6.5 ± 0.91 kPa), human skin (17 ± 1.6 kPa), and scar (55 ± 14 kPa).Elastic modulus of ccPLCL (8.3 ± 0.67 kPa) was not significantlydifferent from that of Integra, suggesting that the combined processof oxygen plasma treatment and collagen coating increased thePLCL scaffold stiffness (Fig. 2B). Oxygen plasma treatment is knownto generate a higher crosslinking density within the first fewthousand angstroms of polymer surface, which results in a localincrease in hardness and likely attributes to the modest increase inelastic modulus of ccPLCL [32].

values were measured from top-view images (A) while scaffold overall thickness waspe 1 collagen, collagen was stained using AlexaFluor488 and imaged via confocal mi-topography when imaged on confocal (C) or SEM (D). Contact angle was analyzed on

ed PLCL films. Permeability analysis was carried out on PLCL scaffolds 110 mm thick, thed ccPLCL. Scale bar is set to 100 mm for all images.

Fig. 2. Mechanical analysis of PLCL and ccPLCL as compared to Integra, human skin, and human scar. (A) Tensile stressestrain curves of human scar (dark green), human skin (lightgreen), Integra (purple), ccPLCL (dark blue), and untreated PLCL scaffold (light blue). (B) Elastic moduli of PLCL and ccPLCL are less than or equal to that of human skin and scar,suggesting that these materials will not impede patients' range of motion if placed across a joint. (C) Unlike Integra, PLCL and ccPLCL scaffolds have similar ultimate tensile stress tohuman skin and scar. (D) Unlike Integra, PLCL and ccPLCL have elongation at break significantly greater than that of human skin and scar, suggesting that these materials will not faildue to tensile forces placed on the skin. (AeD) (E) Fatigue properties of PLCL scaffolds. Storage modulus (black), loss modulus (light gray), and tan D (dark gray) are all preservedover 15,000 cycles. (F) Fatigue properties of ccPLCL scaffolds. Storage modulus (black), loss modulus (light gray), and tan D (dark gray) are all preserved over 15,000 cycles. Together,these data show that the clinical standard of care, Integra, is mechanically inappropriate for use beneath skin graft; however, PLCL and ccPLCL scaffolds display appropriatemechanical properties for this application. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

E.R. Lorden et al. / Biomaterials 43 (2015) 61e70 65

The ultimate tensile stress (UTS) for both PLCL (0.97 ± 0.10 GPa)and ccPLCL scaffolds (1.3 ± 0.17 GPa) was significantly greater thanthat of Integra (0.26 ± 0.020 GPa). While PLCL showed lower UTSthan that of human skin (2.6 ± 0.040 GPa), or human scar(2.7 ± 0.50 GPa), ccPLCL did not (Fig. 2C). In contrast, the elon-gation at break (EB) for PLCL (1100 ± 93 kPa) and ccPLCL(1300 ± 170 kPa) scaffolds was significantly higher than the valuesobtained for Integra (75 ± 4.4 kPa), human skin (200 ± 16 kPa),and human scar (140 ± 16 kPa) (Fig. 2D). The storage modulus, lossmodulus, and tanD of PLCL and ccPLCL scaffold subjected 10%strain at 1 Hz showed negligible deterioration over 24 h, or 15,000cycles (Fig. 2E). This fatigue testing condition was chosen toemulate joint range of motion at the pace of walking [33].

Taken together, the mechanical properties of PLCL and ccPLCLscaffolds relative to human skin and scar e lower modulus,comparable UTS, and higher elongation at break e suggest thatthey would be suitable materials for placement beneath a skingraft.

5.4. In vitro analysis of HSc-related outcomes

Immunostaining was used to study cell morphology and as asemi-quantitative method to analyze expression of aSMA in stressfibers. Cells seeded in FPCLs displayed distinct actin stress fiberswith heavy aSMA incorporation (Fig. 3AeC). Cells seeded in PLCLscaffolds displayed larger size, with more diffuse actin stress fibersand little aSMA incorporation (Fig. 3DeF). At d5 following seeding,significantly fewer cells had converted to myofibroblast phenotypein PLCL scaffolds (34 ± 6.2%), as compared to those seeded in FPCLs(normalized to 100%) (Fig. 3G). Cells remained viable in PLCL andccPLCL scaffolds (Fig. 3H) with 97 ± 1.3% viable cells at d5 in PLCLscaffolds.

5.5. In vivo application of PLCL & ccPLCL scaffolds

All data presented below are described in terms of percentage oforiginal wound size, where 100% is the wound size at d3 (afterremoval of the post-operative bolster), and a fully contractedwound would be described as 0% of its original size. All sampleswere applied beneath the skin graft. Murine wounds treated withskin grafts alone contracted to 47 ± 2.0% at d30 (Fig. 4A), whilewounds treated with Integra contracted to 28 ± 1.8% (Fig. 4B).Wounds treatedwith uncoated PLCL scaffolds remained at 68 ± 11%at day 21. However, partial scaffold extrusion and skin graft deathbegan at d21 and continued until the end of the study, resulting inwound contraction down to 22 ± 11% by d30 (Fig. 4C). In contrast,wounds treated with ccPLCL scaffolds showed significantlydecreased contraction, down to 95 ± 5.8% at d30 (Fig. 4D). Fig. 4Eshows the rate of wound contraction over the test period. Uponextraction of the tissue from themice on d30, ccPLCL scaffolds wereobserved to integrate with host tissue beneath the skin grafts,indicating that it out-performed all other groups in delaying HSccontraction (Fig. 4F).

5.6. Physicochemical properties of scaffolds at d30 afterimplantation

Tensile testing of explanted ccPLCL from d30 mouse studiesdisplayed a similar elastic modulus to scaffolds prior to implanta-tion (Fig. 5B). d30 ccPLCL explants exhibited an elastic modulus(5.7 ± 2. kPa) significantly lower than tissue alone in both mouseskin (36 ± 3.3 kPa) and scar (80 ± 17 kPa) samples, although notsignificantly different from that prior to implantation(9.7 ± 0.20 kPa).

The molecular weight of the scaffolds decreased by 49% fromMn ¼ 151 kDa to Mn ¼ 73.2 ± 6.5 kDa over the implanted period of

Fig. 4. Wound contraction and scaffold incorporation at day 30. (AeD) Gross appearance of wounds treated with (A) skin graft alone, (B) Integra beneath skin graft, (C) PLCL scaffold,and (D) ccPLCL scaffold, at days 7, 9, 14, 21, and 30 following surgery. (E) Wound contraction curves derived from measurements of wounds shown in AeD. (F) ccPLCL scaffoldsbeneath skin graft immediately following excision from wound bed on d30.

Fig. 3. Cellular interactions with scaffolds. (AeF) Overlay of DAPI (blue, left), F-actin (green, center), and aSMA (red, right) in (AeC) FPCLs and (DeF) PLCL scaffolds. (G) Significantlyless aSMA was present in immunostaining in PLCL scaffolds than in FPCLs. (H) Cells remained viable in PLCL scaffolds. (For interpretation of the references to color in this figurelegend, the reader is referred to the web version of this article.)

E.R. Lorden et al. / Biomaterials 43 (2015) 61e7066

Fig. 5. Mechanical properties of excised ccPLCL scaffolds on d30 in vivo. (A) Wounds were cut into three distinct sections for analysis. (B) Tensile elastic modulus of explantedccPLCL from day 30 wounds displayed similar elastic modulus to ccPLCL prior to implantation (p > 0.05). Tensile elastic modulus of explanted PLCL was significantly less than mouseskin or scar tissues (p < 0.05), suggesting that the explanted scaffold þ tissue maintains mechanical properties more similar to the scaffold than to the murine tissue.

E.R. Lorden et al. / Biomaterials 43 (2015) 61e70 67

30 days. NMR analysis shows that the lactide (LA) moiety wasnoticeably more rapidly decreased than the caprolactone (CL) units.The molar ratio was as expected since PLA is known to degrademore rapidly than PCL. The mole fraction of LA decreased from 50%to 44% in 30 days, while that of CL increased from 50% to 56%.Table 1 displays results of NMR and GPC analysis.

5.7. ccPLCL scaffolds incorporate into host tissue and maintainarchitecture in vivo

Fig. 6A-D shows representative images of histological staining inIntegra and ccPLCL scaffold treated wounds on d30. Fiber diameterof ccPLCL scaffolds at d30 in vivo (5.5 ± 0.33 mm) was not signifi-cantly different from initial fiber diameter prior to implantation.Throughout all samples, H&E stained Integra and ccPLCL scaffoldson d30 exhibited acute inflammation (Fig. 6A,B). ECM alignmentappeared more prevalent in Integra treated samples than in ccPLCLtreated samples (Fig. 6A,B, data not quantified). Negligible fibrouscapsule formation was present in ccPLCL treated mice; no fibrouscapsule was visible in Integra treated mice as the implant was nolonger discernable within the wound bed. CD31 staining demon-strated vessel ingrowth into ccPLCL scaffolds and Integra. Multi-nuclear giant cells and neutrophils were visible at the boundaryof the ccPLCL scaffolds with the tissue (Fig. 6B). F4/80 stainingconfirmed the presence of macrophage within the giant cells inwounds treated with Integra and ccPLCL (Fig. 6C). Cells penetratinginto the ccPLCL scaffold and Integra granulation tissue includedhistocytes and foreign body giant cells. Quantification of CD31outlined vessels showed that angiogenesis into ccPLCL (6.1 ± 0.54vessels/HPF) and Integra (6.5 ± 0.90 vessels/HPF) was not signifi-cantly different; vessels were observed spanning both scaffoldmaterials as well as passing through the material (Fig. 6D).

Table 1Analysis of explanted scaffold material. Mn, Mw, and PDI were gathered using GPC,LA:CL ratio was found via NMR.

Mn (kDa) Mn (%) Mw (kDa) Mw (%) PDI LA:CL

Initial 151 100.0 266 100.0 1.8 50:50d30 Explant 73 48.6 135 50.7 2.0 44:56

6. Discussion

This study highlights the role of scaffold longevity in the list ofrelevant design parameters for BSEs aimed at mitigating HSccontraction. In this work, we employed a synthetic, biodegradable,elastomeric, electrospun scaffold to study the effect of scaffoldlongevity on the development of HSc contraction in vivo. Indeveloping a scaffold for use in HSc prevention, we focused on theproperties of healthy skin as design parameters. First, we consid-ered the random nature of the ECM which inspired us to chooseelectrospinning as our fabrication method. Next, we considered themechanical properties of human skin, including its tensile andviscoelastic characteristics, guiding us to select a viscoelastic syn-thetic polymer. Finally, we hypothesized that the optimal scaffoldshould last through the remodeling phase of repair when HSc oc-curs, which prompted us to study a slow-degrading PLA-PCL blend.To improve cellescaffold interactions, we covalently attachedbovine collagen to the scaffold prior to in vivo implantation. Toensure adequate time for HSc stabilization in skin-grafted murinewounds, we followed all treatment groups out for 30 days followingimplantation [24]. While wound healing studies are often con-ducted over 14 days, this extended timeline allowed us to gain anearly determination of the effects of scaffold degradation on scarcontraction. In comparison to previous work using biodegradablepolymers for BSE fabrication, this biodegradable elastomer dem-onstrates appropriate mechanical properties for implantationbeneath skin grafts and is capable of repetitively undergoingphysiologically relevant strain and relaxation without enteringplastic deformation. Most importantly, this synthetic elastomerpossesses a degradation rate on the scale of six months in vivo,allowing it to maintain its architecture throughout the remodelingphase of repair [6,34].

While traditional BSEs are attractive, their degradation rates andmechanical properties are difficult to design in practice. Wepostulate that the mechanical properties of human skin are criticalto consider during the design of a scaffold for HSc prevention.Scaffolds with tensile elastic moduli greater than human skin mayinhibit joint motion, similar to how stiffened scar inhibits motion[24,35]. Therefore, BSEs should possess an elastic modulus less thanor equal to that of human skin. Further, possessing elongation at

Fig. 6. Histological analysis of implanted materials at d30. (A, left) Low power view ofIntegra treated wounds shows the presence of skin grafts (epidermis, e) and underlyingdermis (d) and dermal scar (s); no evidence of Integra is identified. (A, right) ccPLCLtreated wounds with overlying skin graft (e), dermis (d) and ccPLCL scaffold (p, out-lined); architecture of PLCL scaffold remains intact. (B, left) High power view ofepidermis and underlying scar in Integra treated wound no evidence of Integra archi-tecture. (B, right) Cellular infiltration into PLCL scaffold at d30, ccPLCL (p) architecturecan be noted by the existence of white space where the scaffold is present. (C) Presenceof foreign body giant cells in both Integra treated wound (left) and ccPLCL treatedwounds (right) is confirmed by F4/80 pan-macrophage membrane stain (brown).Foreign body giant cells are noted by the presence of black arrows. (D) Angiogenesisinto wounds is analyzed by staining the endothelial lining of neo-vessels in the woundbed using CD31 stain (brown). Black arrows indicate examples of representativeCD31 þ vessels. (E) Quantitative analysis of CD31 stained sections is graphically rep-resented. Number of CD-31 positive vessels were counted in five 40� high power fields(HPF). There is no significant difference in vascularity in ccPLCL Scaffold compared toIntegra-treated wounds at d30 in vivo. (For interpretation of the references to color inthis figure legend, the reader is referred to the web version of this article.)

E.R. Lorden et al. / Biomaterials 43 (2015) 61e7068

break and ultimate tensile stress less than that of human skin maycause a scaffold to rupture beneath the skin prior to completion ofhealing. Along with its tensile characteristics, skin is viscoelastic innature, allowing it to stretch and relax across joints repetitively[35]. Burns that occur across joints, especially those of the upperbody, are the most common location for HSc to occur [1]. In fact,joint motion during healing is likely a driving factor of HSc [36].Thus, it is important that BSEs placed in the wound bed aredesigned with the appropriate elastomeric properties to withstandrepeated expansion and relaxation cycles across joints.

In designing a synthetic scaffold for permanent placement in thewound bed, it is critical that cells remain viable in the scaffold andthat the scaffold does not exacerbate scarring. Cells seeded inccPLCL scaffolds remain viable in vitro confirming that electrospunscaffolds provide a suitable cellular micro-environment. To inves-tigate the potential role of the scaffold in scarring, we studiedfibroblast to myofibroblast transition in PLCL scaffolds as comparedto the in vitro wound contraction model, the FPCL. Once seeded inthe FPCL, fibroblasts are able to contract the collagen fibrils in thegel similarly to how they would contract scar granulation tissue.The fraction of cells which converted into myofibroblasts in thismodel was analyzed by immunocytochemical staining for aSMA,which is a marker for myofibroblast formation. Using this semi-quantitative method, we found significantly more aSMA in theFPCL than was present in PLCL scaffolds. These data suggest thatelectrospun PLCL scaffolds mitigate cellular processes associatedwith HSc.

Human HSc has typically been studied in immune-compromised mice, in which multiple models have been devel-oped; however, due to the importance of the immune system inwound healing, we have chosen not to use these models [37e42].Two porcine models are available to study HSc: the female redDuroc pig, and the Yorkshire pig [43]. Swine arguably represent thebest model for human wound healing, with 78% agreement be-tween human and porcine wound healing, however pre-clinicalstudies in swine are costly [44]. We chose to test PLCL and ccPLCLscaffolds in our recently developed murine model of HSc contrac-tion [24]. While the murine model employed in this work directlyfollows the human condition in terms of treatment and contractingscar, it should be noted that the model does not reflect the prolif-erative component or timeline of human HSc. The term HSc, as itapplies to humans, describes a constellation of scar symptoms (red,raised, itchy, contracted scars), whereas the murine model onlyexhibits the contraction phenotype. Over time the murine skingrafts become flat and pale, as is observed in some but not all hu-man HSc [45,46]. The difference between these two scarring statesmay be due to a diminished level of mechanical tension; humanskin exists in a resting state of tension, unlike murine skin, which isloose. Further, the model employed in these studies is an expeditedHSc contraction model where contraction ensures for 14d and thenstabilizes, whereas in humans contraction can last for 6e12months. The PLCL copolymer used in these studies has been shownto persist in canine patients for up to 12 months, suggesting that itwill last through the contraction phase of scarring if translated touse in humans [34].

While uncoated PLCL scaffolds were rejected between d21-d28,ccPLCL scaffolds integrated into the wound tissue and preventedHSc contraction. This is likely because synthetic polymer implantsdo not offer integrin binding sites and often require surfacetreatment prior to implantation. PLCL scaffolds are hydrophobic innature, encouraging the likelihood for random protein adsorptionand eventual extrusion. The extrusion process is caused by spon-taneous and uncontrollable adsorption of proteins from the blood,lymph, and wound exudate to hydrophobic implant surfaces [47].Implantation of synthetic polymer structures into open wounds

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without integrin binding sites can lead to foreign body encapsu-lation and extrusion. To decrease the likelihood of random proteinadsorption and encourage cellescaffold interactions throughintegrin binding sites, collagen was covalently attached to thesurface of PLCL scaffolds. There is concern that protein coating ofelectrospun polymers may restrict effective aqueous fluid flow,thus disrupting transport of fluids and nutrients through scaffoldsin vivo [48]. However, despite the presence of a protein coating, anegligible change in effective permeability was found betweenccPLCL and PLCL scaffolds. Following in vivo implantation, ccPLCLscaffolds integrated into the tissue beneath the skin graft. Thebeneficial impact of the collagen coating could be explainedby favorable changes in scaffold hydrophilicity or the introductionof cell-binding motifs. These data are in agreement with theliterature on the importance of introducing integrin binding sitesto the surfaces of intra-dermal, hydrophobic, synthetic polymerimplants.

Upon removal of wound tissue from the mice, ccPLCL scaffoldscould be discerned beneath skin grafts. Tensile testing of explantedccPLCL from d30 mouse studies displayed a similar elastic modulusto scaffolds prior to implantation (Fig. 5B). d30 ccPLCL explants alsoexhibited an elastic modulus significantly lower than tissue alonein both mouse skin and scar samples. These data suggest that thescaffold is preventing stiffening associated with HSc scar formation.Scar formation is also associated with alignment of the ECM, as isseen in d30 Integra treated samples via H&E staining. Conversely,cells in ccPLCL scaffolds display random nuclear orientation andECM alignment at d30 via H&E staining (Fig. 6B). The presence of aforeign body reaction, characterized by neutrophil andmacrophageinfiltration, can be seen in both treatment groups. The acuteinflammation present in these sections is confirmed by F4/80staining and is commonly observed following intra-dermal im-plantation of a foreign body. Vessel infiltration is not significantlydifferent between Integra and ccPLCL scaffolds, suggesting thatboth materials are integrating well with the host tissue andallowing lymphogenesis and/or angiogenesis to occur. Overall, thehistological data suggest the presence of an acute inflammatoryreaction in ccPLCL treated mice along with vessel infiltration intothe scaffold and maintenance of randomly oriented ECM. Theseoutcomes are all necessary to maintain skin graft health and pre-vent the formation and subsequent contraction of HSc.

ccPLCL scaffolds significantly prevented HSc contraction in vivoas compared to Integra or placement of a skin graft alone. Integratreated mice maintained 75% of the original wound area until dayeight, after which scars rapidly contracted (Fig. 4E). The contractionseen after day eight could be associated with partial necrosis of skingrafts on Integra-treated wounds due to the immediate placementof skin grafts over Integra [49]. While collagen-based scaffolds canbe successful following immediate skin graft placement, Integrarequires a two-step implantation procedure [50]. During clinicaluse, Integra is placed and allowed to integrate into the wound bedfor approximately two weeks prior to placement of the skin graft.This delay allows a period for vascular infiltration of the Integramatrix; however, this method increases risk of infection and re-quires a second operation [51]. In general, collagen scaffolds, suchas Integra, are associatedwithwound contraction and scarring [52].Collagen is degraded and remodeled by collagenases and otherproteases within the wound bed with a half-life on the order ofdays to weeks depending on the crosslinking method. It isconceivable that after day eight of implantation, Integra begins tolose its architecture and mechanical properties, allowing rapid scarcontraction [53]. A similar rapid contraction after day eight hasbeen shown in the literature following treatment of full thicknessdermal wounds with electrospun collagen scaffolds in a guinea pigmodel [54]. Together, these data suggest that rapid scar contraction

after day eight is related to the longevity of collagen scaffolds in thewound bed, rather than the fabrication technique. Wounds alsorapidly contracted following extrusion of uncoated PLCL scaffolds,which began after d21 in vivo (Fig. 4E). Rapid wound contraction iscommonly seen following removal of splinting materials frommurine wounds, suggesting that Integra and non-incorporatedscaffolds may act as temporary splinting materials to keep thewound open until they either extrude or lose their architecture viadegradation [55].

HSc contraction develops progressively over the course ofseveral weeks inmice andmultiple months in humans; therefore, itis feasible that a loss of scaffold architecture prior to the completionof the remodeling phase of repair could allow for delayed cellularalignment and the formation of a mechanically-coordinatedcellular syncytium, thus increasing the likelihood of HSc contrac-tion. The structural support provided by collagen-based scaffolds,such as Integra, has been suggested to rely upon the pore archi-tecture of the scaffold, rather than the elastic modulus of the ma-terial [53]. Our data support the hypothesis that elastic modulus ofthe implant does not dictate its ability to resist wound contraction;the elastic modulus of ccPLCL electrospun scaffolds is within thesame order of magnitude as Integra when tested in the early strainregion. However, the effects onwound contraction between Integraand ccPLCL scaffolds are drastically different at d30 in vivo. At d30following implantation, the architecture of ccPLCL scaffolds re-mains clearly defined in the wound bed histology, whereas the d30the architecture of Integra is not visible via histological staining.Lyophilized collagen-GAG based scaffolds, similar to Integra,degrade over the order of several days to several weeks [56]. Thisshort degradation time can be exacerbated by patient-to-patientvariability in the wound healing process, leading to the possibilityof premature degradation prior to proper healing. Indeed, rapiddegradation and resorption of a BSE prior to the completion of theremodeling phase of repair leads to increased likelihood of scarformation [57]. We hypothesize that the critical difference leadingto the success of ccPLCL scaffolds in preventing HSc contraction isthe residence time of the scaffold architecture in the wound bed.The PLCL material employed in this study will retain its structure inthe wound bed throughout the remodeling phase of repair, havingonly lost 51% of its number-average molecular weight at 30 daysin vivo. Future studies are required to confirm this hypothesis andwill explore longer time points to evaluate the response of thetissue to degradation of the polymer.

7. Conclusions

Loss of scaffold architecture prior to the completion of theremodeling phase of repair is likely a major cause for HSccontraction. This study suggests that the relative maintenance ofintegrity, rather than the absolute mechanical properties of theccPLCL scaffold, is responsible for improved mitigation of HSccontraction.

Acknowledgments

The authors would like to acknowledge support from the PlasticSurgery Foundation Endowment Grant, the Duke UniversityDepartment of Surgery, and the Duke University Department ofBiomedical Engineering.

The authors would like to thank Bruce Klitzman, PhD, for hiscollaboration and technical assistance. The authors would also liketo thank Eugenia H. Cho and Carlos Quiles-Torres for their technicalassistance, and Gloria Adcock for her assistance with tissueprocessing.

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